Proceedings - Umeå universitet

Proceedings of the International Federation for
Medical & Biomedical Engineering
13 th Nordic Baltic Conference on Biomedical
Engineering and Medical Physics
13NBC 2005
June 13 th – 17 th , 2005, Umeå, Sweden
Regional Meeting of IFMBE
Editors
Ronnie Lundström
Britt Andersson
Helena Grip

IFMBE PROCEEDINGS
ISSN 1680-0737
Volume 9, 2005
13 th Nordic Baltic Conference on Biomedical Engineering and Medical Physics
The International Federation for Medical and Biomedical Engineering, IFMBE, is a federation of national and
international organizations which represent national interests in medical and biological engineering.
The objectives of the IFMBE are scientific, technological, literary, and educational.
Within the field of medical, clinical, and biological engineering, IFMBE’s aims are to encourage research and
the application of knowledge, and to disseminate information and promote collaboration.
IFMBE Officers
President: Joachim H. Nagel, Vice-President: Makoto Kikuchi, Past-President: Dov Jaron,
Treasurer: Shankar Muthu Krishnan, Secretary-General: Ratko Magjarevic
http://www.ifmbe.ore
Published by: Swedish Society for Medical Engineering and Medical Physics
Print: Informationsenheten, Västerbottens Läns Landsting, Umeå, Sweden
Issue: 250
Copyright © 2005 IFMBE, SSMEMP and SFfR
All rights reserved. This book or any part thereof may not be reproduced in any form or by any means without
written permission from the publishers.
Enquiries: Swedish Society of Biomedical Engineering and Medical Physics
Previous Editions:
6th Asian-Pacific Conference on Medical and Biological Engineering “APCMBE 2005”, Vol. 8, 2005, Tsukuba,
Japan, CD.
Kuala Lumpur International Conference on Biomedical Engineering “BIOMED 2004”, Vol. 7, 2004, Kuala
Lumpur, Malaysia.
X Mediterranean Conference on Medical and Biological Engineering “MEDICON and HEALTH
TELEMATICS 2004”, Vol. 6, 2004, Ischia, Italy, CD.
3rd Latin - American Congress on Biomedical Engineering “III CLAEB 2004”, Vol. 5, 2004, Joao Pessoa,
Brazil, CD.
World Congress on Medical Physics and Biomedical Engineering “WC2003”,Vol. 4, 2003, Sydney, Australia,
CD.
2nd European Medical and Biological Engineering Conference, EMBEC'02", Vol. 3, Parts 1 & 2, 2002, Vienna,
Austria.
12NBC "12th Nordic Baltic Conference on Biomedical Engineering and Medical Physics", Vol. 2, 2002,
Reykjavik, Iceland.
MEDICON 2001 - "IX Mediterranean Conference on Medical Engineering and Computing", Vol. 1, Parts 1 & 2,
2001, Pula, Croatia.
For ordering Proceedings from the IFMBE Proceedings Series, please contact the IFMBE Secretariat at:
info@ifmbe.org
II

Welcome to 13NBC, June the 13 th to 17 th 2005 in Umeå
This series of NBC conferences has proven to be a forum of great value for biomedical engineers and physicists
in this rapidly developing field of science and technology. The exchange of ideas and discussions around
technical advances supports the ongoing development of the healthcare and makes the NBC conferences an
event worthwhile to attend.
The IFMBE Proceedings volume from this Conference will hopefully reflect the exchange of knowledge,
experience and new ideas which has been taking place here in Umeå. There are people coming from more than
30 different countries to present and discuss results from research and development activities. Without doubt, the
content of the present proceedings volume clearly manifest the rapid evolution in the field of Biomedical
Engineering and Medical Physics.
This NBC conference will be held in Umeå, a very nice city located in the middle of northern Sweden, known as
the city of birches. The great river of Umeå that winds its way through city, the bright summer nights, the green
parks and the many birch trees along the streets give this city its own character. The undisturbed environment
offers you to combine the cultural life in the city with nature experiences such as hiking, cycling, water rafting
and canoeing. Umeå is also the largest centre for higher education and administration in northern Sweden, with
many research activities carried out at Umeå University and Umeå University Hospital. For example, the Centre
for Biomedical Engineering and Physics (CMTF) was inaugurated in 2001 at Umeå University. The center
promotes research and development in that area. The industrial sector in Umeå is dynamic and steadily growing,
consisting of major industries and smaller enterprises representing biomedical engineering, biotechnology, IT,
and other industrial products. This makes Umeå a model city for modern advanced education and research, and
thus a very suitable place for NBC'05
On behalf of the organizing committee for 13 th Nordic Baltic Conference on Biomedical Engineering and
Medical Physics I have the honour and pleasure to welcome you to Umeå. I express my warmest and deepest
thanks for your participation and contribution. In addition, I would also like to take the opportunity to address
my gratitude to all individuals who have been involved in organizing and preparing for this event. You have
made this conference possible through your dedicated and hard work.
At last, I wish you all a pleasant, fruitful and enjoyable stay in Sweden.
Ronnie Lundström
Editor and chair of the organizing committee of 13NBC 2005
PREFACE
The Proceedings include papers of the 13 th Nordic Baltic Conference of Biomedical Engineering and Medical
Physics, 13NBC 2005, held in Umeå in June 2005. This proceeding is the 9 th volume in the series of IFMBE
proceedings from conferences that are co-organized or sponsored by the federation.
In this proceeding are published papers from 9 well-known keynote speakers followed by about 170 papers by
scientist from about 32 countries. Thanks to all of you. I would like to thank everyone that has contributed to
this conference proceeding by submitting abstracts of their on-going research. I would also like to thank all the
members of the local and international scientific committees that have worked with the scientific program and
reviewed the papers.
On behalf of the scientific committee
Britt Andersson
Editor and chair of the scientific committee of 13NBC 2005
III

ORGANIZED BY
SSBEMP - Swedish Society for Biomedical Engineering and Medical Physics
SFfR - Swedish Society of Radiation Physics
HOSTED BY
CMTF - Centre for Biomedical Engineering and Physics at Umeå University
REGIONAL MEETING OF
IFMBE – International Federation for Medical and Biological Engineering.
IN CO-OPERATION WITH
IFMBE – International Federation for Medical and Biological Engineering.
Societies for Biomedical Engineering and Medical Physics in the Nordic and Baltic countries.
Organizing Committee
Ronnie Lundström, Sweden (Chair)
Per-Åke Ericson, Sweden
Olof Lindahl, Sweden
Helena Grip, Sweden
Birgitta Lanhede, Sweden
Per Ackeberg, Sweden
Maria Lindén, Sweden
Heikki Teriö, Sweden
Per Ask, Sweden
Stéfan B Sigurdsson, Iceland
Tarmo Lipping, Estonia
Jaanus Lass, Estonia
Timo Jämsä, Finland
Kalle Kepler, Estonia
Sergey Popov, Latvia
Local Scientific Committee
Britt Andersson, Sweden (Chair)
Heikki Tölli, Sweden
Haibo Li, Sweden
Olof Lindahl, Sweden
Adi Anani, Sweden
Fredrik Georgsson, Sweden
Britta Sethson, Sweden
Mikael Karlsson, Sweden
Anders Eklund, Sweden
Urban Wiklund, Sweden
Stefan Karlsson, Sweden
Heikki Teriö, Sweden
Karin Wårdell, Sweden
Per Ask, Sweden
Ronnie Wirenstam, Sweden
International Scientific Committee
Sauli Savolainen, Finland
Leif Sörnmo, Sweden
Magnus Borga, Sweden
Ewert Bengtsson, Sweden
Jerker Delsing, Sweden
Metin Akay, USA
Janis Spigolis, Latvia
Kalju Meigas, Estonia
Arunas Lukosevicius, Lithuania
Björn-Erik Erlandson, Sweden
Noam Alperin, USA
Karin Roeleveld, Norway
Hans Åhlfelt, Sweden
Ronnie Lundström, Sweden
Jan Persson, Sweden
Maria Lindén, Sweden
Lennart Johansson, Sweden
Åke Öberg, Sweden
Kjell Hansson Mild, Sweden
Tomas Jansson, Sweden
Solange Akselrod, Israel
Richard Kitney, UK
Toril A. Nagelhus Hernes, Norway
John G Webster, USA
George Tesar, Sweden
Alan Murray, UK
Joachim H. Nagel, Germany
Inger-Lena Lamm, Sweden
Arunas Lukosevicius, Lithuania
IV

Table of contents
KEYNOTE LECTURES
THE COMPUTER ASSISTED FUTURE; NEW POSSIBILITIES IN EDUCATION, SIMULATION 1
AND MINIMALLY INVASIVE THERAPY
T Hernes, R Mårvik, T Langø, J Kaspersen, T Selbekk, G Tangen, G Unsgård, H Myhre
WEB-BASED EHEALTH AND THE FUTURE ROLE OF MOLECULAR BIOLOGY 4
R Kitney
SCIENTIFIC ENTREPRENEURSHIP AND MARKET REALITIES 5
G Tesar
FROM A SCIENTIFIC STUDY, TO A MANUSCRIPT, TO ITS PUBLICATION 6
A Murray
COMPLEXITY OF RESPIRATORY NEURAL NETWORK DURING MATURATION 8
M Akay
TISSUE ABLATION: DEVICES AND PROCEDURES 9
J G Webster
HEART RATE VARIABILITY: MODELS, METHODS, AND APPLICATIONS IN STRESS TESTING 11
P Laguna
THE EUROPEAN HIGHER EDUCATION AREA IN BIOMEDICAL ENGINEERING 12
– ACHIEVEMENTS, TRENDS AND DEVELOPMENTS
J H Nagel
EDUCATION AND TRAINING OF THE EUROPEAN MEDICAL PHYSICIST - ROLES AND 13
RESPONSIBILITIES IN RADIATION PROTECTION AND SAFETY.
I-L Lamm
HEALTHCARE ASSESSMENT AND CLINICAL ENGINEERING
TRENDS IN HEALTH CARE ASSESSMENT 14
J Persson
BUSINESS DEVELOPMENT OF BIOMEDICAL ENGINEERING INVENTIONS 15
O Lindahl
TUBERCULOSIS-THE SILENT KILLER OF DEVELOPING WORLD AND WHERE WE ARE? 17
M Rahman
ADOPTION OF MEDICAL DEVICES: THE NEONATAL INTENSIVE CARE UNIT 18
AS A CASE STUDY
K Roback, P Gäddlin, N Nelson, Jan Persson
THE NEED OF PROCEDURES COMPILED WITH MDD TO INVESTIGATE MEDICAL 20
DEVICE FAILURES INVOLVING PATIENT INJURY
H Gilly
IMPROVEMENT OF PATIENT SAFETY IN PRACTICE - EXAMPLES OF PREVENTION 21
OF ACCIDENTS
H Teriö
EDUCATIONAL DEVELOPMENT AND CURRICULUM PLANNING WHEN USING 23
SIMULATORS.-USEFUL TOOLS FOR TRAINING DOCTORS AND ENGINEERS.
K Mäkinen, L Felländer-Tsai, P Ström, A Kjellin, T Wredmark, L Hedman
EARLY EXPOSURE TO HAPTIC FEEDBACK ENHANCES PERFORMANCE IN IMAGE 25
GUIDED SURGICAL SIMULATOR TRAINING - A PROSPECTIVE RANDOMIZED CROSS
OVER STUDY IN SURGICAL RESIDENTS
L Felländer-Tsai, K Mäkinen, P Ström, A Kjellin, T Wredmark, L Hedman
II

Table of contents
EXPOSURE TO MAGNETIC FIELDS OF THERAPEUTIC STAFF DURING TMS/RTMS 26
TREATMENTS
R Lundström, E F Karlström, O Stensson, K Hansson Mild
PRELIMINARY REFERENCE LEVELS FOR DIAGNOSTIC RADIOLOGY IN ESTONIA 29
K Kepler, A Servomaa, I Filippova
EXPERIENCE FROM TEACHING BIOMEDICAL ENGINEERING TO HEALTH CARE PERSONAL 31
M Folke, A Jonsson
MEDICAL INFORMATICS
IMPLEMENTATION OF GLUCOSE-INSULIN CONTROL IN H2/HINF SPACE USING 33
MATHEMATICA
L Kovacs, B Palancz, Z Almassy, Z Benyo
A PILOT STUDY OF DEVELOPING THE LABORATORY INFORMATION SYSTEM OF A 35
COMMERCIAL LABORATORY
S Tang, J Lin, P Li, Y Huang, S Young
TELEMEDICINE AND PATIENT DATA MANAGEMENT
BIOMEDICAL SENSOR NETWORK ARCHITECTURE BASED ON TCP/IP 37
C López, J C Tejero-Calado, A Bernal, M A López, G Quesada, J Lorca
GAP ANALYSIS BUSINESS IMPACT OF MODEL DRIVEN ARCHITECTURE (MDA) ON 39
TELEMEDICINE HEALTHCARE SOLUTION
R Nabiev, N Tariq, S Jonsson, H Teriö, D Andersson
WEB-BASED SUPPORT TO ENHANCE SELF-MANAGED DIABETES CARE 41
M Psaros, A Ekbom-Schnell, A Vidmark, S Koch
COMPARISON OF MODEL DRIVEN ARCHITECTURE (MDA) BASED TOOLS USING 43
TELEMEDICINE HEALTHCARE SYSTEM
N Tariq, S Jonsson, R Nabiev, H Teriö, D Andersson
A LOW COST ECG MONITORING SYSTEM EMPLOYING TELEMEDICINE 45
Mamun Reaz
A TECHNICAL PLATFORM FOR REMOTE AUSCULTATION AND REAL-TIME MONITORING 47
OF PHYSIOLOGICAL PARAMETERS
J Skönevik, P Hallberg, A Müller, R Lundström, U Wiklund, S Karlsson, U Wiklund
WEARABLE MOBILE TECHNOLOGY FOR HEALTHCARE 49
O Atzmon, D Shklarski, S Jonsson, H Teriö
IMPLEMENTATION OF A TECHNICAL SYSTEM IN DISTRIBUTED CARE- ATTITUDES 50
AND POSSIBILITIES
L Rattfält, C Hagström, M Lindén, P Hult
SPEX-SPREADING EXCELLENCE IN HEALTHCARE-A TELEMEDICINE PROJECT 52
E Fridén, B Eklund, B Gerdin, P Gustafsson, K Eriksson
TECHNICAL CONSIDERATIONS AND WORKAROUNDS INTEGRATING FUSION 54
IMAGE DATA INTO A TRADITIONAL PACS ENVIRONMENT
J Leal
III

Table of contents
NEURAL ENGINEERING
BOLD SIGNAL ANALYSIS DURING ACOUSTIC STIMULATION OF THE BRAIN AT 3 TESLA 56
S Casciaro, D Zacà, R Bianco, S Neglia, F Esposito, F Di Salle, A Distante
THE MAGNITUDE-SQUARED COHERENCE IN THE DETECTION OF STIMULATION/NO- 58
STIMULATION TRANSITIONS
D B Melges, A F Infantosi, M Cagy, A M Miranda de Sá
MULTIVARIATE SPECTRAL ANALYSIS APPLIED TO THE EEG DURING RHYTHMIC 60
STIMULATION – A COHERENCE-BASED APPROACH
A M Miranda de Sá, M Cagy, D Melges, A F Infantosi
SIMULATIONS OF RADIO-FREQUENCY LESIONS WITH VARYING BRAIN ELECTRODE 62
DIMENSIONS
J D Johansson, J Wren, O Eriksson, D Loyd, K Wårdell
IMPROVING COHEN’S MODEL FOR BOLD FUNCTIONAL RESPONSE 64
S Casciaro, S Neglia, G Palma, D Zacà, R Bianco, E Casciaro, A Distante
PSYCHO-ACOUSTIC EXPERIMENTS OF THE PERCEPTION OF THE MISSING FUNDAMENTAL 66
T Matsuoka, D Konno
BIOMEDIA
BIOMEDIA 68
A Anani
FIGURE CHECKER 69
N Anani, H Li
HAND BASED PERSONAL VERIFICATION SYSTEM 71
S Anani, H Li
A PHYSIOLOGICAL SIGNAL BASED PERSONAL VERIFICATION SYSTEM, A PILOT STUDY 73
H Begic, A Anani
IMPROVEMENT IN BRAILLE READING USING A FINGER COVER 75
K Doi, S Shinohara, H Fujimoto
HAPTIC VIBROTACTILE: DRIVER ASSISTANT SYSTEM 77
L Liu
TACTILE VIDEO 78
L Liu
CARDIOVASCULAR ENGINEERING
THE INTELLIGENT STETHOCOPE AS A TOOL IN MODERN HEALTH CARE 79
P Hult, C Ahlstrom, L Rattfält, C Hagström, N-E Pettersson, P Ask
A DSP SYSTEM FOR REAL-TIME ANALYSIS OF PERIPHERAL VESSELS FROM 81
SEQUENCES OF ECHOGRAPHIC IMAGES
F Faita, V Gemignani, M Demi
MYOCARDIAL PERFUSION ASSESSMENT USING AN ECG TRIGGERED LASER DOPPLER 83
TECHNIQUE
C Fors, M Karlsson, H Casimir-Ahn, K Wårdell
WALL BACK FLOW IN HUMAN AORTA: INFLUENCE OF GEOMETRY 85
J Svensson, R Gårdhagen, D Loyd, M Karlsson
IV

Table of contents
MODELING OF ACTION POTENTIAL PROPAGATION IN CARDIAC CELLS DISPERSED IN 87
HETEROGENEOUS MEDIA
I Haq, L Al-Kury, Z Hani, T Musallam
A MICROSYSTEM FOR MONITORING HEART MOTION 89
L Hoff, O J Elle, M J Grimnes, S Halvorsen, H J Alker, E Fosse
MITRAL VALVE OPENING IN THE FAILING HEART 91
K Kindberg, M Karlsson
INSTRUMENTATION FOR REPRODUCING THE POSITIONING OF A PERSON IN 93
BALLISTOCARDIOGRAPHIC MEASUREMENT
L Leppäkorpi, R Sepponen
BLOOD PRESSURE MEASUREMENT 95
F E Smith, A Haigh, J Wild, E J Bowers, S T King, J Allen, D Zheng, P Langley, A Murray
PRELIMINARY CLINICAL RESULTS FROM A COMPUTER-ASSISTED CORONARY 97
ANGIOPLASTY BALLOON INFLATION DEVICE
T Olbrich, A Murray
DEVELOPMENT OF IMPLANTABLE CARDIAC MEASUREMENT DEVICE - MODELING 99
APPROACH
J Väisänen, J Hyttinen, J Malmivuo
A NOVEL ISOLATED MEASUREMENT SYSTEM FOR BIO-POTENTIALS 101
K Piipponen, R Sepponen
PARAMETRIC ASSESSMENT OF HRV IN CONGENITAL CENTRAL HYPOVENTILATION 103
SYNDROME: A CASE REPORT
T Princi, A Accardo, D Peterec
DIFFERENT APPROACHES OF MEASUREMENT OF HEMODYNAMIC BY ELECTRICAL 105
IMPEDANCE PLETHYSMOGRAPHY METHOD
A Stankus
MEDICAL ULTRASOUND
AN APPROACH OF INDIVIDUAL DESIGN OF EFFECTIVE MODAL COUPLING IN 107
PIEZOELECTRIC ELEMENTS
D Kybartas, A Lukoševicius
ULTRASOUND SIGNAL ENHANCEMENT VARYING MICROBUBBLE CONCENTRATION 109
AT VERY LOW MECHANICAL INDICES
S Casciaro, R Palmizio Errico, F Conversano, A Maffezzoli, A Sannino, A Distante
FREQUENCY EFFECTS ON ECOCONTRAST AGENTS SIGNAL BEHAVIOUR AT LOW 111
MECHANICAL INDICES
R Palmizio Errico, S Casciaro, F Conversano, E Casciaro, G Palma, A Distante
ULTRASOUND CONTRAST FOR PERFUSION STUDIED 113
M Ressner, A Kvikliene, L Hoff, R Jurkonis, T Jansson, B Janerot-Sjöberg, A Lukosevicius, P Ask
AN ULTRASONIC METHOD FOR DETECTION OF FLUID PROPERTIES IN THE 115
PARASANAL SINUSES
T Jansson, P Walfridsson, P Sahlstrand-Johnson, H Persson, N Holmer, M Jannert
NEW IMPROVED METHOD FOR 2-D ARTERIAL WALL MOVEMENT MEASUREMENTS 117
M Cinthio, Å Rydén Ahlgren, T Jansson, H Persson, K Lindström
ULTRASONIC ATTENUATION IMAGING USING COHERENT PROCESSING IN ULTRASONIC 119
COMPUTER TOMOGRAPHY
A Filipik, R Jirik, J Jiri
V

Table of contents
LUNG FRACTIONAL MOVING BLOOD VOLUME EVALUATED WITH POWER DOPPLER 121
ULTRASOUND IN NORMALLY GROWN AND GROWTH RESTRICTED FETUSES
E Hernandez-Andrade, A Thuring-Jönsson, T Jansson, G Lingman, K Marsál
VERSATILE MICROCHIP UTILISING ULTRASONIC MANIPULATION OF MICROPARTICLES 123
M Nilsson, T Lilliehorn, L Johansson, M Almqvist, U Simu, S Johansson, T Laurell, J Nilsson
INVESTIGATION OF THE FETAL HEART CIRCULATION IN AN ANIMAL MODEL USING 125
CONTRAST ENHANCED ULTRASOUND
T Jansson, E Hernandez-Andrade, G Lingman, P Malcus, D Ley, K Marsál
BIOMEDICAL INSTRUMENTATION
RESONANCE SENSORS FOR BETTER HEALTH CARE 127
O Lindahl
EYE-PRESSURE MEASUREMENT WITH APPLANATION RESONANCE TONOMETER 128
A Eklund, P Hallberg, K Santala, T Bäcklund, C Lindén
DETECTION OF PROSTATE CANCER WITH A RESONANCE SENSOR 130
V Jalkanen, B Andersson, A Bergh, B Ljungberg, O Lindahl
AN IMPROVED RESONANCE SENSOR SYSTEM FOR DETECTING CANCEROUS TISSUE 132
IN THE PROSTATE
P Lindberg, B Andersson, A Bergh, B Ljungberg, O Lindahl
REALTIME WIRELESS MEASUREMENT OF MECHANICAL DATA 134
J Delsing, J Lindblom, D Sjölund, P Lindgren
TEMPERATURE INDEPENDENCE OF AN ELECTRO ACOUSTIC CAPNOGRAPH 136
M Folke, B Hök
EVALUATION OF A SURGEON-CENTERED LAPAROSCOPIC SURGICAL TOOL DESIGN 138
M Hallbeck, D Oleynikov
SKIN TEMPERATURE EFFECTS ON SKIN BLOOD FLOW AT AREAS PRONE TO 140
PRESSURE SORE DEVELOPMENT
A Jonsson, M Lindgren, M Lindén
BLUETOOTH ECG MONITORING SYSTEM 142
J C Tejero-Calado, C López, A Bernal, M López, G Quesada, J Lorca
WIRELESS WEARABLE EMG AND EOG MEASUREMENT SYSTEM FOR 144
PSYCHOPHYSIOLOGICAL APPLICATIONS
N Nöjd, M Puurtinen, P Niemenlehto, A Vehkaoja, J Verho, T Vanhala, J Hyttinen, M Juhola,
J Lekkala, V Surakka
AN INFLATABLE HIP PROTECTOR FOR PREVENTING INJURIES 146
T Tamura, T Yoshimira, M Sekine
ACOUSTIC MONITORING OF LUNG SOUNDS FOR THE DETECTION OF ONE 148
LUNG INTUBATION
S Tejman-Yarden, L Weizman, A Zlotnik, J Tabrikian, A Cohen, G Gurman
ELECTROMUSCULAR INCAPACITATING DEVICES 150
J G Webster
A WIRELESS USB BASED VIEWING BOX FOR CEPHALOGRAM ANALYSIS 152
Kuo-Sheng Cheng, Ching-Lin Li, Cheng-Yu Cheng, Wen-Hung Ting, Yen-Ting Cheng
TISSUE PERMEABILIZATION STUDIED ON A MATHEMATICAL MODEL OF A 154
SUBCUTANEOUS TUMOR IN SMALL ANIMALS
N Pavselj, Z Bregar, D Cukjati, D Batiuskaite, L M Mir, D Miklavcic
VI

Table of contents
PRODUCTION OF NANO/MICRO BECLOMETHASONE DIPROPIONATE PARTICLES USING 156
SUPERCRITICAL CARBON DIOXIDE
M R Golriz, E Matida, B M Andersson
AEROSOL DEPOSITION SIMULATION IN AN IDEALIZED MOUTH 158
E A Matida, W H Finlay, M R Golriz
CHANGE OF ARTERIAL PULSE WAVE IN PATIENTS WITH HYPERLIPIDAEMIA 160
I Hlimonenko, K Meigas, R Vahisalu
DEVELOPMENT AND OPTIMISATION OF A NOVEL SKIN IMPEDANCE INSTRUMENT 162
I Bodén, L Noren, Å Wisten, P Geladi, J Nyström, B Lindholm-Sethson
DEVICE FOR CONTINUOUS BLOOD PRESSURE MEASUREMENTS 164
K Meigas, J Lass, D Karai, R Kattai, J Kaik
LATENCY OF RANDOM SEARCH SACCADES 166
N Ramanauskas, G Daunys, V Laurutis, V Vysniauskas
WIRELESS SYSTEM FOR REAL-TIME RECORDING OF HEART RATE VARIABILITY 168
M Karlsson, F Ragnarsson, N Östlund, U Edström, T Bäcklund, J S Karlsson, U Wiklund
WAYS TO DECREASE THE ADHESION OF PSEUDOMONAS AERUGINOSA BACTERIA 170
TO SURFACES OF ENDOTRACHEAL TUBES
M Ramstedt, H J Mathieu
BIOMEDICAL OPTICS
IN VITRO IMAGING OF HUMAN CARTILAGE - CONTRAST IMPROVEMENT BY OPTICAL 172
WAVELENGTH SELECTION
A Johansson, T Sundqvist, J-H Kuiper, P Å Öberg
CLEARANCE VARIATIONS MONITORED BY ON-LINE UV-ABSORBANCE DURING 174
HAEMODIALYSIS
F Uhlin, I Fridolin, M Magnusson, L-G Lindberg
MEASUREMENTS OF ALA METHYLESTER DIFFUSIVITY IN NORMAL SKIN IN VIVO: 176
A PILOT STUDY
N Yavari, H.R Mobini Far, N Yavari, N Gustavsson, F Torabi, C Anderson, B Danielsson, P O Larsson, K
Svanberg, S Andersson-Engels
ADVANCED FIBRE-OPTIC SPECTROMETRY TECHNIQUE FOR SKIN REFLECTANCE 178
AND FLUORESCENCE DIAGNOSTICS
J Spigulis, L Gailite, I Kuzmina, A Lihachev
DEVELOPMENT OF OPTICALLY FUNCTIONAL FIBER-REINFORCED COMPOSITE 180
FOR MEDICINE AND DENTISTRY
J Lehtinen, T Laurila, T Reivonen, E Levänen, T Mäntylä, L Lassila, S Tuusa, P Kienanen, P Vallittu, R
Hernberg
LIPID DIFFUSION IN COCHLEAR MEMBRANES BY FRAP 182
J Boutet de Monvel, W E Brownell, M Ulfendahl
DIFFUSE REFLECTANCE SPECTROSCOPY OF SKIN PATHOLOGIES 184
I Kuzmina, L Gailite, A Lihachev, R Karls, J Spigulis
STUDY ON OPTICAL QUALITY OF INTRAOCULAR LENSES 186
A F Shkapa, S M Kulikov, L V L'vov, A N Manachinski, S A Sukharev, L I Zykov
EXPERIMENTAL EYE MODEL FOR INTRAOCULAR LENS TESTS AND DEMONSTRATIONS 188
L I Zykov, S M Kulikov, L V L'vov, A N Manachinski, S A Sukharev, A F Shkapa, S N Bagrov
OPTICAL COHERENCE TOMOGRAPHY IN HIGH RESOLUTION IMAGING 190
B Kudimov, G Dobre, A Podoleanu
VII

Table of contents
BIOMEDICAL IMAGING TECHNIQUES
CENTER FOR MEDICAL IMAGE SCIENCE AND VISUALIZATION (CMIV) - A UNIQUE 192
CROSS-DISCIPLINARY ENVIRONMENT FOR MEDICAL IMAGE PROCESSING RESEARCH
M Borga
CORRELATION CONTROLLED ADAPTIVE FILTERING FOR FMRI DATA ANALYSIS 193
J Rydell, H Knutsson, M Borga
SEGMENTATION OF VELOCITY ENCODED CARDIAC MAGNETIC RESONANCE IMAGES 195
E Bergvall, H Arheden, G Sparr
MORPHONS AND BRAINS - NON-RIGID REGISTRATION FOR ATLAS-BASED 197
A Wrangsjö, H Knutsson
GENERATION OF PATIENT SPECIFIC BONE MODELS FROM VOLUME DATA 199
USING MORPHONS
J Pettersson, H Knutsson, M Borga
MEASURING THE MOTION PATTERNS OF THE HEARING ORGAN USING A THREE 201
DIMENSIONAL OPTICAL FLOW METHOD
M von Tiedemann, A Fridberger, M Ulfendahl, J Boutet de Monvel
MICROWAVE PROBING OF COMPLEX DIELECTRIC BODIES 203
P Norin, T Gunnarsson, D Åberg, P-O Risman
BRAIN VENOGRAPHY WITH NMR BOLD CONTRAST AT 3T 205
D Zacà, S Casciaro, R Bianco, S Neglia, E Casciaro, T Scarabino, A Distante
THE INFLUENCE OF CORNEA TO FORMATION OF THE 2D IRIS IMAGE 207
E Paliulis, G Daunys
HIGH RESOLUTION VENOGRAPHY AS A VASCULAR MASK FOR ACTIVATION 209
SITES OF AUDITORY CORTEX AT 3T
S Casciaro, D Zacà, G Palma, R Bianco, S Neglia, E Casciaro, A Distante
BPA QUANTIFICATION AND DETECTION IN PHANTOMS USING THREE DIMENSIONAL 1H 211
MAGNETIC RESONANCE SPECTROSCOPIC IMAGING
M Timonen, S Savolainen, S Heikkinen
BIOSENSORS
AN INTRODUCTION TO BIOSENSORS 213
B Lindholm-Sethson
CUSTOMIZING THE COMPUTER SCREEN PHOTO-ASSISTED TECHNIQUE FOR 215
EVALUATING QUICK DIAGNOSTIC TESTS
D Filippini, I Lundström
EXTRACTION AND ACTIVITY MEASUREMENT OF HUMAN 11ß-HSD2 FOR USE IN A 217
SALIVA CORTISOL BIOSENSOR
M Sandström, T Shen
BIOSENSORS BASED ON MEMBRANE ORGANISATION 219
P Geladi, A Nelson, K Bradley, B Lindholm-Sethson
ORIENTED IMMOBILISATION OF ANTIBODIES FOR IMMUNOSENSING 221
I Vikholm-Lundin, W M Albers
BIOCOMPATIBLE PACKAGING OF A THREE-AXIS MICRO ACCELEROMETER 223
K Imenes, K Aasmundtveit, L Hoff, O J Elle
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SPORTS AND REHABILITATION BIOMECHANICS
POSSIBILITIES AND CHALLENGES OF MULTI-CHANNEL SURFACE EMG IN 225
SPORTS AND REHABILITATION
K Roeleveld, J S Karlsson, C Grönlund, A Holtermann, N Östlund
MUSCLE ARCHITECTURE AND FIBRE TYPES USING SPATIOTEMPORAL 227
INFORMATION OF PROPAGATING MOTOR UNIT ACTION POTENTIALS
RECORDED BY 2-D MULTICHANNEL SURFACE EMG
C Grönlund, K Roeleveld, S Karlsson
GLOBAL MUSCLE ACTIVATION IN SUSTAINED CONTRACTIONS 229
A Holtermann, K Roeleveld
MYOFEEDBACK ISSUES IN REHABILITATION OF WORK-RELATED MUSCULAR 231
PAIN A REVIEW
L Sandsjö
EFFECT OF INTERELECTRODE DISTANCE ON MEASURING SENSITIVITY FOR FACIAL EMG 233
M Puurtinen, J Hyttinen, J Malmivuo
NEAR INFRARED SPECTROSCOPY FOR MEASURING MUSCLE OXYGENATION 235
A Crenshaw, B R Jensen
APPLICATION OF RECIPROCAL ORTHOTIC SYSTEMS WITH ELECTRICAL STIMULATION 237
OF MUSCLES IN CHILDREN WITH SPINAL DISEASES
E Dukendjiev, V Mihnovich
COMPARING DYNAMICS OF SKI JUMPING AND DRY IMITATION JUMPS BY COMPUTER 239
SIMULATION
G J Ettema, S Bråten
DIMENSIONAL MODELING OF HUMAN BODY UNDER VERTICAL AND HORIZONTAL 241
VIBRATION
M Behzad, P Hejazi Dinan, M Rasolzadeh, F Farahmand
STUDY ON THE NET JOINT MOMENTS OF THE LOWER LIMB IN NORMAL AND ABOVE 243
KNEE AMPUTED SUBJECTS
P Hejazi Dinan, T Rezaeian, M Behzad
PLANAR GEOMETRY OF FEET IMPRINTSAS THE REFLECTION OF THE STRUCTURE 245
AND CONDITION OF LOCOMOTIVE SYSTEM
E Dukendjiev, T Ogurtsova
PROPRIOCEPTIVE ABILITY WITHIN PATIENTS WITH WHIPLASH ASSOCIATED DISORDERS 247
H Grip, G Sundelin, S Karlsson
A PILOT STUDY TO ESTIMATE THE LACTATE THRESHOLD USING AN ELECTRO 249
ACOUSTIC SENSOR
M Folke, L Gullstrand, B Hök
MONITORING HEALTH AND ACTIVITY BY SMARTWEAR 251
L Berglin, M Ekström, M Lindén
THE ASSESSMENT OF THREE DIMENSIONAL ANKLE JOINT FORCES DURING THE 253
POSTURAL BALANCE CONTROL MOVEMENT
M J Seo, H Choi
THE EFFECTS OF EXTERNAL LOAD ,TRUNK AND KNEE POSITION ON THE TRUNK 255
MUSCLES ACTIVITY
S Kahrizi, M Parnianpour, S M Firoozabadi, A Kazemnejad
AN INVERSE KINEMATIC MODEL OF THE OSTEOPOROTIC SPINE 257
Z Yang, J Griffith, P Leung, M Pope, R Lee
DYNAMIC CONTROL OF THE TRUNK IN OBESE SUBJECTS 259
C Loong, S Yeung, S Kwong, N O'Dwyer, R Lee
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A TECHNIQUE FOR EVALUATING INFANTS MUSCLE ACTIVITY USING SURFACE EMG 261
N Östlund, S Håkansson, U Edström, S Karlsson
BIOMECHANICS
APPLICATION OF RESISTANCE TESTING FOR IDENTIFYING SHUNT RESPONDERS 263
IN IDIOPATHIC NORMAL PRESSURE HYDROCEPHALUS
A Marmarou
USING THE PULSATILITY OF BLOOD AND CSF FLOWS TO PROBE THE BIOMECHANICAL 264
STATE OF THE CRANIOSPINAL SYSTEM
Alperin Noam
ESTIMATION OF CSF OUTFLOW CONDUCTANCE REPEATABILITY AND PRECISION 266
N Andersson, J Malm, T Bäcklund, A Eklund
COCHLEAR AQUEDUCT PATENCY IN TYMPANIC MEMBRANE DISPLACEMENT 268
MEASUREMENT FOR INTRACRANIAL PRESSURE ASSESSMENT
S Shimbles, K Banister, C Dodd, D Mendelow, I Chambers
COMPUTERISED ASSESSMENT OF CSF DYNAMICS IN HYDROCEPHALIC PATIENTS 270
P Smielewski, M Czosnyka, Z Czosnyka, J Pickard
RESTING PRESSURE AND ACCURACY OF CEREBROSPINAL FLUID INFUSION TEST 272
A Eklund, N Andersson, J Malm
INTRACRANIAL PRESSURE MEASUREMENT VIA LUMBAR SPACE 274
N Lenfeldt, L–O Koskinen, A Ågren-Wilsson, T Bergenheim, J Malm, A Eklund
NON-NEWTONIAN EFFECTS ON WALL SHEAR STRESS IN A HUMAN AORTA WITH 275
COARCTATION AND DILATATION
R Gårdhagen, J Svensson, D Loyd, M Karlsson
ERRORS IN APPLANATION TONOMETRY RELATED TO REFRACTIVE SURGERY 277
P Hallberg, K Santala, T Koskela, O Lindahl, C Lindén, A Eklund
TRANSMURAL MYOCARDIAL STRAIN DISTRIBUTION THEORETICAL RESULTS AND 279
IN VIVO DATA
K Kindberg, M Karlsson
SPRING-DAMPER MODEL FOR PROSTATE TISSUE 281
N Norén, B Andersson, A Bergh, B Ljungberg, O Lindahl
MODELLING OF PERIPHERAL BLOOD FLOW DYNAMICS: COMPARISON WITH FINGER 283
PHOTOPLETHYSMOGRAPHY SIGNALS
U Rubins, J Spigulis
THE EFFECT OF MEMBRANE MOVING PATTERN ON THE BLOOD FLOW IN A SAC-TYPE 285
VENTRICULAR ASSIST DEVICE
F Firouzi, N Fatouraee, S Najarian
BONE MINERAL DENSITY IN RELATION TO FRACTURAL ANALYSIS BY 287
BIOMECHANICAL PROPERTIES
M Mokhtari-Dizaji, MR Dadras, B Larijani, G Torkaman, A Kazem-Nejad
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BIOMEDICAL SIGNAL PROCESSING
HEART RATE VARIABILITY IN FAMILIAL AMYLOIDOTIC POLYNEUROPATHY 289
U Wiklund
A WAVELET-BASED TECHNIQUE FOR BASELINE WANDER CORRECTION IN ECG AND 291
MULTI-CHANNEL ECG
A Khawaja, S Sanyal, O Dössel
EVALUATION OF AN EFFICIENT METHOD FOR HANDLING ECTOPIC BEATS IN HRV 293
K Solem, P Laguna, L Sörnmo
ACCURACY LIMITATIONS OF THE DOPPLER ULTRASOUND SIGNAL IN RELATION TO 295
FHR VARIABILITY ANALYSIS
J Wrobel, J Jezewski, K Horoba, A Gacek
OPEN-LOOP MODEL OF EQUINE HEART CONTROL 297
J Holcik, P Kozelek, M Jirina, J Hanak, M Sedlinska
ADAPTIVE MULTICHANNEL FILTER FOR HEART BEAT DETECTION 299
F Ragnarsson, N Östlund, U Wiklund
OTOACOUSTIC EMISSIONS TIME FREQUENCY MAPPING BASED ON THE ENSEMBLE 301
CORRELATION AND HILBERT-HUANG TRANSFORM
A Janusauskas, A Lukosevicius, V Marozas, L Sörnmo
WAVELET-BASED FEATURE EXTRACTION 303
I Rodríguez Carreño, M Vuskovic
WHEEZE ANALYSIS AND DETECTION WITH NON-LINEAR PHASE-SPACE EMBEDDING 305
C Ahlstrom, P Ask, P Hult
POSITION APPROXIMATION FROM ACCELERATION DATA OF AN ISCHEMIC PIG HEART, 307
SYNCHRONIZED WITH THE ELECTROCARDIOGRAPH SIGNAL
L Fleischer, L Hoff, O Elle, S Halvorsen, E Fosse
BIOIMPEDANCE ANALYSIS OF CARDIAC FUNCTION USING IN-VIVO EXPERIMENTS, 309
MEASUREMENT AND SIMULATION
R Gordon, A Haapalainen, P Kauppinen, R Land
CUT-OFF MECHANICAL INDEX FOR EFFECTIVE CONTRAST IMAGING 311
F Conversano, S Casciaro, R Palmizio Errico, E Casciaro, C Demitri, A Distante
COMPUTER AIDED FETAL MONITORING SYSTEM USING NONINVASIVE 313
ELECTROCARDIOGRAM
A Matonia, T Kupka, K Horoba, J Jezewski, P Labaj
CHARACTERIZATION METHODOLOGY FOR SOUND ABSORPTION OF SYNTHETIC 315
MATERIALS IN ULTRASOUND IN VITRO STUDIES
S Casciaro, R Palmizio Errico, F Conversano, E Casciaro, L Ostuni, A Distante
HEART SOUND CONFIDENCE INTERVALS ESTIMATION BASED ON THE 317
IEFE COEFFICIENTS
R Osman, S Hussien, A Babiker
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MEDICAL PHYSICS
CLINICAL MR SPECTROSCOPY, PAST, PRESENT AND FUTURE A REVIEW 319
J Hauksson
IMPLEMENTING PATIENT SELF-EVALUATED RECTAL PROBLEMS IN NTCP MODEL 320
FOR LOCALIZED PROSTATE CANCER PATIENTS TREATED WITH EXTERNAL BEAM
RADIOTHERAPY
S Åström
QUALITY FACTOR VS. FIVE-POINT SCALE IN EVALUATION OF CLINICALY 321
ACCEPTABLE IMAGE QUALITY IN CHEST CT
O Sveljo, B Reljin, Z Markovic, M Lucic, O Adjic, M Prvulovic
CALCULATING DOSE OUTPUT AT OFF-AXIS POSITIONS IN PHOTON BEAMS USING A 323
LATERALLY BEAM QUALITY DEPENDENT PENCIL KERNEL MODEL
J Olofsson, T Nyholm, A Ahnesjö, M Karlsson
THE USE OF CARCINOGENESIS RISK ESTIMATION MODELS FOR RADIOTHERAPY 324
A Daşu, I Toma-Daşu, J Olofssonp, M Karlsson
CORRECTION FOR SCATTER AND SEPTAL PENETRATION IN 123I BRAIN SPECT 325
IMAGING A MONTE CARLO STUDY
A Larsson, M Ljungberg, S Jacobsson Mo, K Åhlström Riklund, L Johansson
ERROR ESTIMATION IN DOSE CALCULATION FOR VERIFICATION PURPOSES 326
T Nyholm
INTRA-OPERATIVE MONITORING WITH NEEDLE ELECTRODES IN CONJUNCTION 327
WITH SURGERY,INCLUDING ELECTROSURGICAL UNIT (ESU).A HAZARD ? !
P Fällmar, T Winkler, R Flink
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Keynote
THE COMPUTER ASSISTED FUTURE; NEW POSSIBILITIES IN
EDUCATION, SIMULATION AND MINIMALLY INVASIVE THERAPY
T. Hernes 1, 2 , R. Mårvik 3 , T. Langø 2 , J. Kaspersen 2 , T. Selbekk 2 , G. Tangen 2 , G. Unsgård 4 , H.
Myhre 5
1 Medical Faculty, NTNU, Trondheim, Norway
2 Health research, SINTEF, Trondheim, Norway
3 NSALK, St Olavs Hospital, Trondheim, Norway
4 Dept of neurosurgery, St Olavs Hospital, Trondheim, Norway
5 Dept of Surgery, St Olavs Hospital, Trondheim, Norway
Toril.N.Hernes@sintef.no
Abstract
More minimally invasive surgery procedures are now
being introduced into clinical practice. These
procedures have many advantages due to improved
treatment, faster recovery and more efficient use of
resources. Since most of these operations are
performed through small incisions, the surgeon is not
able to palpate the organs and special instruments
and methods are required. Robotics and sophisticated
navigated image guidance technology are therefore
emerging in various clinical applications. The demand
for special skills, improved competence, training and
experience of the operator is increasing with the
introduction of new and advanced technology. In
order to be able to perform the procedure successfully
with improved outcome, simulators and training
facilities are introduced. In the interdisciplinary
research collaboration at St Olavs University Hospital
in Trondheim we have developed technology for
image guided education, training and surgical
treatment were navigation and multimodal and
intraoperative imaging are integrated. We have
established facilities for clinical research in minimally
invasive neurosurgery, endovascular therapy and
laparoscopy. The present paper shows examples from
clinical studies using future, commercial and still noncommercial,
solutions and methods of navigated
advanced visualization techniques in minimally
invasive image guided surgery and training. “The
Operating Room of the Future” will be presented.
Introduction
The amount of elderly people will increase considerably
during the next decades [1]. This again increases the
demand for new treatment technology and methods that
fasten the recovery and makes it easier for the patients to
take care of themselves and take part in the social and
working life after treatment. Clinical treatment is now
going from open surgery to minimally or non invasive
procedures, because the advantages of these techniques
are many both due to reduced patient recovery time and
reduced overall costs. New technologies that integrate
navigation technology and advanced visualisation
technology are now introduced [2]. Intraoperative
imaging as MR and Ultrasound, intraoperative CT and
angiolaboratories are emerging in order to cope with the
changes that occur during interventions. This makes the
surgeon still able to guide the procedure minimally based
on updated images and not by direct insight as in open
surgery [3]. The benefits for the patients of the
introduction of new methods and technology are also
starting to be scientifically proved [4]. But new advanced
technology also demands more training and simulators
for improving skills and performance of specific surgical
procedures. Also, special clinical facilities for training,
education and innovation where new methods and
clinical research can be developed based on the demands
in the clinic are needed.
Methods
We have during more than 10 years achieved experience
with an established interdiciplinary research collaboration
were research scientists with technological background
are joining surgeons and radiologists in the
operating/interventional room. We have developed a
navigation system (CUSTUS X, SINTEF, Trondheim,
Norway) that integrates preoperative image data and
intraoperative updates based on ultrasound, MR and CT
data. Various 2D, 3D and multimodal and multifunctional
displays can be monitored and selected using surgical
tools, pointers and endoscopes equipped with optical
(Polaris, NDI, Waterloo, Canada) and magnetic (Aurora,
NDI, Waterloo, Canada) position devices. Focus has been
put on minimally invasive treatment in neurosurgery,
endovascular and laparoscopic surgery, where the system
has been used in various clinical studies for planning,
guidance and postprocessing/evaluation of the treatment.
In feasibility studies during laparoscopic procedures,
tracking of the endoscope was performed. We have also
tested the system in various clinical studies in ”The
Operating Room of the Future” at St Olavs University
Hospital; a completely new innovative facility for
integrated clinical and technological research activity. In
neurosurgery we have used an additional intraoperative
IFMBE Proc. 2005;9: 1

Keynote
3D ultrasound navigation system , SonoWand (MISON
AS, Trondheim, Norway) in more than 200 tumor
resections and vascular lesions in the brain.We have also
tested the navigation system during training of medical
students in anatomy with 2D and 3D image displays in
combination with corresponding cadaver dissections.
Also in planning and follow up of patients treated for
abdominal aortic aneurysms, 3D virtual teleconsultations
were tested due to image quality, speed, functionality and
advantages for the patient as well as for hospital owners.
Results
Our image guided navigation system has been
successfully used in many of the clinical applications
tested. In laparoscopic surgery the system was important
because it helped the surgeon in guiding the surgical
instrument into correct position and to virtually ”see”
beyond the surface of the organs in cases where the
surgeon was not able to palpate the organs (Figure 1).
Figure 2: Navigation and 3D visualization technology
during anatomy courses for medical students. 3D virtual
models were used in combination with cadavers during
rehearsel of needle pinprick/incision.
Figure 1: Intraoperative image guided therapy using
navigation and advanced visualisation technology for
guiding surgical procedures in laparoscopic surgery.
This again improved the safety and precision of the
procedures. The need for intraoperative imaging was
unquestionable, and updates based on intraoperative 2D
and 3D ultrasound (neurosurgery) and CT (endovascular
treatment) demonstrated improved resection control and
guidance of the procedure. Teleconsultation made it
possible for the patient to be followed up by the local
hospital with an additional expert radiologist evaluation
at the university hospital. This was both convenient for
the patients as well as cost reducing for hospital owners.
In training and in rehearsal of practical clinical
procedures as for example during needle
pinprick/inscision the use of navigated 3D models in
combination with cadavers improved the enthusiasm of
the students and made the students feel more safe and
comfortable when performing the procedure (Figure 2)
Discussion
Also other research groups have demonstrated navigated
intraoperative imaging and guidance as a tool for
improved minimally invasive therapy [5]. Many of the
methods make it easier for the operator and safer to
perform procedures that have not earlier been performed.
Advanced real time 3D imaging are introduced for
improved diagnostic, functional imaging and image
guidance [6]. In addition also non-invasive treatment as
focused ultrasound and RF treatment are now emerging.
We believe that minimally invasive treatment is here to
come, and our technological solution is one of many
solutions that have been presented. New technology
demands training and enhancement of special skills, not
only generated and demonstrated in the present paper, but
also confirmed by the market that has established
simulator centres and facilities for more advanced
education and training. Also, development of
interdiciplinary research environments for introduction
and development of new technology and methods as well
as translation of new methods into practical medicine are
here to come.
References
[1] Population Ageing 2002, United Nations, Department
of Economic and Social Affairs, Presented at the Second
World Assembley on Ageing, Madrid.
[2] REINHART H., TRIPPEL B., WESTERMANN B.,
AND GRATZL (1999): “Computer aided surgery with
special focus on neuronavigation”, Computerized
Medical Imaging and Graphics, 23, pp 237-244
[3] NIMSKY C,. GANSLANDT O., VON KELLER B.,
ROMSTOCK J., FAHLBUSCH R. (2004):
IFMBE Proc. 2005;9: 2

Keynote
“Intraoperative high-field-strength MR imaging:
Implementation and experience in 200 patients”,
Radiology , 233: 67-78
[4] WIRTZ CR., KNAUTH M., STAUBERT M.,
BONSANTO MM., SARTOR K, KUNZE S.,
TRONNIER VM. (2000): ”Clinical evaluation and
follow-up results for intraoperative magnetic resonance
imaging in neurosurgey”. Neurosurgery, 46, pp 1112-
1122
[5] BONSANTO MM., STAUBERT A., WIRTZ CR.,
TRONNIER V., KUNZE S. (2001): ”Initial experience
with an Ultrasound-Integrated Single-Rack
neuronavigation System”, Acta Neurochir (Wien), 143:pp
1127-1132
[6] FENSTER A and DOWNEY DB (1996): ”3D
ultrasound Imaging: A Review.” IEEE Engineering in
Medicine and biology, Nov/Dec, pp 41-51
IFMBE Proc. 2005;9: 3

Keynote
WEB-BASED EHEALTH AND THE FUTURE ROLE OF MOLECULAR
BIOLOGY
R. Kitney 1
1 Imperial College, London, UK
r.kitney@imperial.ac.uk
Abstract
Information has been central to medical diagnosis and
treatment since time immemorial. However, with, for
example, the development of a wide range of imaging
techniques throughout the 20 th Century – and particularly
during the second half of the century – there has been an
increasing need to develop electronically based clinical
information systems. Throughout the early and mid
1990s there were a number of attempts to develop such
systems, almost all of which had very limited success.
Today the situation is very different. It is now possible to
develop clinical information systems (CIS) which are
reliable, affordable and accessible over the entire hospital
and beyond. This situation has been reached because of
the confluence of a number of factors; principal amongst
these are web-based technology, powerful pc technology,
international standards and broadband
telecommunications networks. Web-based technology
has proved to be an important vehicle for providing
clinical information on demand. However, this has only
been possible because of the development of broadband
telecommunications networks, which often, within the
hospital, can have a bandwidth of 1 GHz; the
development of powerful (Pentium) pc technology, which
has a performance as good as previous UNIX
workstations; and the development of two major
international standards, DICOM and HL7. These
international standards have enabled data from a wide
range of medical equipment to be connected into a
common information environment. The practical
manifestation of this technology now occurs in electronic
patient records (EPRs), PACS, RIS etc. However, current
clinical information systems, of which these are
examples, mainly operate at the systems, visceral and
tissue levels.
specialties. Hitherto, much of this information (for
example images) has been confined to a limited number
of specialties - a prime example being Radiology. This
situation is now about to change and will require the
development of a new type of CIS which can handle
information from across the BC; whilst, simultaneously,
integrating information which, at present, is held
separately in the EPR, PACS, RIS etc.
Key features of the new type of CIS will comprise the
ability to input data from a wide range of sources, and to
navigate seamlessly across the BC. Currently, at the
higher levels of the BC (namely, at the system, viscera
and tissue levels), data are frequently presented in
DICOM format; however, this does not apply at the
lower levels and it may, perhaps, be appropriate to extend
the international standards to these levels. The second
important aspect of the new type of CIS will be seamless
navigation across the BC. This will require the
development of visualisation and navigation techniques
which provide geometrical integrity, despite the fact that
the data at different levels may well be from different
modalities.
References
1. R.I. Kitney, "The Role of Engineering in the Post-
Genomic Age," The Royal Academy of Engineering, pp
1-31, ISBN 1-903496-09-8, 2003.
2. R. I. Kitney, “Clinical Information Systems Overview”
Keynote Address. Proc Medicon 2004 (IFMBE
Mediterranean Conference on Medical and Biological
Engineering) pp 56-62.
With the rapid developments which have occurred over
the last few decades in molecular and cellular biology, as
evidenced by the initial sequencing of the human
genome, it is now becoming important to include
information from other levels of the human organism[1].
To describe this range of information, we have coined the
term “The Biological Continuum” (BC)[2] – this
comprises the levels of the human organism from
systems to genes (ie, systems, viscera, tissues, cells,
proteins, genes). Another important trend is the need for
information from across the biological continuum to be
available, on demand, to a wide range of medical
IFMBE Proc. 2005;9: 4

Keynote
SCIENTIFIC ENTREPRENEURSHIP AND MARKET REALITIES:
G. Tesar 1
1 Umeå School of Marketing and International Buisness, Umeå University, Umeå, Sweden
Abstract
In the new technological age, an increasing number of
scientists are attempting to commercialize their
inventions, discoveries, and innovations. Especially in
the public sector, university-based scientists, academic
researchers, and clinical professors are examining their
commercial options for their ideas. Although scientific
research personnel in the private sector are frequently
constrained by an obligation to share their inventions,
discoveries, or innovations with their employer, they may
actively participate in the commercialization of their
ideas. In both cases, the introduction of scientifically
new ideas requires a great deal of creativity, patience, and
perseverance. The ability to bridge the scientific and
commercial environments requires a strong
entrepreneurial stamina. Although some scientists
succeed in bridging the two environments, many do not.
Those scientists who manage to commercialize their
ideas in the new technological age today are considered
scientific entrepreneurs.
Scientific entrepreneurship is becoming more complex.
In the new technological age, scientific entrepreneurs
need the ability to communicate with the outside world to
develop their entrepreneurial skills. The purpose of this
presentation is to examine the fundamental concepts of
entrepreneurship in the increasingly internationally
competitive age of new technology.
george.tesar@fek.umu.se
IFMBE Proc. 2005;9: 5

Keynote
FROM A SCIENTIFIC STUDY, TO A MANUSCRIPT,
TO ITS PUBLICATION
A. Murray
Editor in Chief, Medical & Biological Engineering & Computing, Newcastle University,
Freeman Hospital, Newcastle upon Tyne, UK
mbec@nuth.northy.nhs.uk
Abstract: Scientific research should be followed by a
publication detailing the results of the research, so
that others can learn from the work. Without
communication, research will be duplicated and this
will hamper further research. However, much
scientific research after much effort is never
published. It should be relatively easy to avoid such
disappointment. It is necessary to plan research
with vision, being clear about what other researchers
will want to know and to read. Potential authors
need to take time to consider what is needed by
others. There are many reasons why publication is
not always easy, and it is necessary to start from the
planning of the study. This paper explores the
practical and personal issues underlying scientific
success and its well-deserved publication.
Introduction
It seems logical that any scientific research will be
followed by a publication detailing the results of the
research, so that others can learn from the work. This is
essential, as without communication, research will be
duplicated and insights necessary for the next
breakthrough may not appear. However, much research
is never published, and many scientific papers are not
accepted for publication without many tedious
revisions, and sometimes additional work. This leads to
much disappointment, and in some cases to the waste of
the money which funded the research.
Researchers can lack the vision to plan their work so
that others will want to read about it. Is that because
they do not take time to consider what society needs or
what will excite other researchers. Or perhaps is it
simply because the research is approached with the
limited vision of a preferred technique, or is the thought
of publication left until it is too late, or does an
interesting personal goal lack the discipline needed for a
publication. Has the excitement of discovery and
communication been lost by some of us? Surely not.
Aim
The aim of this paper is to help introduce the skills
of writing a scientific paper to those who need to
communicate science. It also includes the necessary
preparation for this communication. Those with many
years experience in publishing may also benefit from
understanding the difficulties others have experienced.
The following comments are derived from my
experience as an editor of a scientific international
journal.
Why Write Scientific Papers?
It is very important to communicate the results of
scientific research work. It is necessary for the general
public to understand the impact of science and
participate in decision making. This elevates the role of
science, and hence the specific role of Bioengineers and
Medical Physicists. This can also bring with it increased
willingness to support science, and its funding.
Scientists and engineers need to be able to
communicate the research they are pursuing, and when
finished, need to be able to make the results available to
the wider community.
When to Start Preparing for Your Publication
It is never too early to start preparing your
publication. In fact, as soon as you have started to
design your study, you should ask the question “How
will I write this work up for publication?” That will
show whether you have a suitable structure or not.
Research studies tend to develop and change, and
sometimes the methods and data sets may alter several
time during a study for simple practical reasons. If that
happens you are unlikely to be able to write up a series
of short studies as a single study. It will appear as a
historical document rather than a scientific report.
As soon as a study diverts from the original plan, a
new plan will need to be produced, and perhaps the
study will need to start over again. It is better to decide
that early on rather than waiting to the end when it
might be discovered that the work just cannot be written
up as a scientific paper.
Any Editor will be able to relay many examples of
when authors try, usually without success, to produce a
scientific paper from a series of developing sub-studies.
Why Some Research is never Published
Communicating science is most successfully
achieved through journals, after formal review by our
scientific peers. Through this process there is a
permanent record for study, so that scientific experience
is handed on and those who follow can learn from those
who went before.
IFMBE Proc. 2005;9: 6

Keynote
Unfortunately there are a number of reasons why
publication is not always achieved.
Because many find the process of writing a paper
difficult, the result can be that some scientific work is
never communicated, which if it was worth doing in the
first place is almost unforgivable.
Sometimes work is prepared hurriedly or without
care and is rejected. Then pressure from other work in
the author’s laboratory can prevent the original research
being resubmitted for publication. Only rarely do
referees say that the research presented was a waste of
time. Referees give comments for improving papers.
These comments are there to help. They should not halt
the ambition to produce a good paper, but sadly they
often do.
Papers can be rejected for very simple reasons,
which careful preparation would have avoided. I
comment on some below. I have been involved on many
editorial boards, and have refereed for very many
journals, as well as being Editor in Chief of Medical &
Biological Engineering & Computing for many years. I
aim to use this experience to help others.
Successful Publication
The first thing to remember is that the most
important task is to communicate, and not to publish.
However, even though our means of communication is
by publishing, it is important not to loose sight of the
original goal. This is the greatest and most frequently
encountered mistake made by authors. Very few papers
are rejected because the work done was a waste of time.
That is not to say that it could not have been
substantially improved by planning it differently or by
clarifying the main aim or reading the published
literature more thoroughly. Many papers are rejected
because they communicate poorly.
Preparation for Writing
Before you write your first sentence you must have a
plan, or the paper will simply drift from one thing to
another. Producing this plan is a very important step in
successful publication. Remember that you should have
a written plan before the study starts, even if it needs to
be changed later.
Before writing your paper, set out the structure first,
so you are clear how the background, aims, methods
and results all fit together.
Be especially carefully about your aim. You need to
make sure that you can prove you have achieved your
aim, and if this is not by numerical data you need to
have a very good reason why not.
Clarity
Then as you write your paper make sure it is clear to
any reader. It has to pass referees and if they do not
understand your paper it will get no further. If your
native tongue is not English, do not worry, as there are
usually staff to help. At Medical & Biological
Engineering & Computing the editorial staff will make
any necessary amendments to the English for you to
check, but the paper needs to be clear first, or the staff
may be confused about what you are trying to say.
Structuring the Paper
My strong advice is that the paper should have a
traditional structure.
The introduction needs to explain why you needed
to do the work, setting it in context of what others have
done. The introduction makes the case for why the
research was worth doing, and this is always the first
question the referee will ask. This is often done poorly.
There needs to be an aim, and this is usually defined
clearly at the end of the introduction. All too often
referees say that they could not understand what the
authors were trying to achieve. The methods section
should be short and clear, describing what you did,
without any tutorial on the techniques or description of
what others have done. Referees are often confused
about what authors have actually done. Also, the
methods section needs to contain all the methods. No
other section of the paper should contain methods. Next,
the results section should contain all the results.
Although it may have some comments, keep your
discussion to the next section.
If you use such a traditional structure you will find
the paper easy to write. Also, you should keep papers
short as they are much easier to read.
For more information see an Editorial on this
subject. [1]
Ask Others to be Critical
After completing the first good draft of the paper, it
is very important to seek comments from your
colleagues. They should give you feedback on how well
your paper communicates, but beware that most
colleagues will want to please you. You need to ask for
honest criticism. At the same time you should try to see
your paper the way a referee would see it. [2]
Conclusion
Enjoy your research work and enjoy communicating
your science and engineering. Write your research
publications so that others will want to read what you
have achieved. [3] If you were excited by your work,
relay this excitement to others.
References
[1] MURRAY A. (2001): ‘On becoming a virtual
editor’, Med. Biol. Eng. Comput., 39, p. 1
[2] MURRAY A. (2003): ‘The other side of the
table’, Med. Biol. Eng. Comput., 41, p. 1
[3] MURRAY A. (2002): ‘A good read’, Med.
Biol. Eng. Comput., 40, p. 1
IFMBE Proc. 2005;9: 7

Keynote
COMPLEXITY OF RESPIRATORY NEURAL NETWORK DURING
MATURATION
M. Akay 1
1 Thayer School of Engineering Dartmouth College, NH, USA
Metin.Akay@Dartmouth.Edu
Abstract
Epidemiologic studies have identified numerous risk
factors for the Sudden Infant Death Syndrome (SIDS).
Prominent among these risk factors are the prone
sleeping position and overheating, either by excessive
clothing of the infant or by excessive heating of the room.
Recent studies also indicate that SIDS babies may have
significant neurotransmitter receptor defects. The number
and distribution of muscarinic, kainate and serotonergic
receptors in the brainstem were significantly less in
babies who died of the SIDS compared to babies who
died as a result of chronic illnesses or accidents.
Identification of these risk factors, some of which can be
modified, has led to recommendations that have
significantly reduced the risk of the SIDS. Unfortunately,
less progress has been made integrating these
epidemiologic and neuroanatomical findings into a
coherent pathophysiological explanation of the exact
mechanism whereby infants die in the SIDS. It is our
working hypothesis that the neurotransmitter defects
interfere with protective homeostatic responses to
potentially life-threatening, but often occurring events
during infant sleep. Thus, the epidemiologically defined
risk factors identify either a potentially life threatening
event (e.g., sleeping prone presumably increases the risk
of asphyxiation) or a factor that interferes with protective
homeostatic responses (e.g., overheating or the
neurotransmitter defects may interfere with cardiac,
respiratory or arousal responses to life threatening
events).
changes in the complexity of the respiratory neural
network that accompany maturation in an animal model,
the piglet. In a series of experiments, we characterize and
quantify the complexity of the output of the central
respiratory, monitored as the diaphragm EMG, to gain
insights into how respiration is generated during
wakefulness and sleep (REM and NREM); determine
the influence of complete inhibition of neurons in the
rostral ventral medulla (RVM) on the dynamics of
respiratory patterns during wakefulness and sleep (REM
and NREM) in unanesthetized, chronically instrumented,
intact piglets. We believe that the quantification of the
complexity of the respiratory pattern generator will be
useful for understanding the etiology of and developing
therapies for several clinically important disorders of
breathing patterns such as obstructive sleep apnea and
especially sudden infant death syndrome.
This research was supported in part by the NIH- HL
6573
Previous studies in various animal models have also
shown that respiratory premotor and motor neurons
undergo rapid changes in biochemical and bioelectric
properties during the first month of postnatal life. Early
in postnatal life, there is an increase in the complexity of
the dendritic tree of respiratory neurons as it changes
from a bipolar to a multipolar morphology. The
respiratory motor output including the phrenic neurogram
or diaphragm EMG depends on the integrated properties
of the respiratory pattern generator and has features that
reflect the dynamics of it; we use these features to
develop our model.
In this study, we propose novel nonlinear dynamical
analysis methods including the maximum likelihood
estimator (MLE) and the expectation-maximization MLE
method based fractal methods to define and quantify
IFMBE Proc. 2005;9: 8

Keynote
TISSUE ABLATION: DEVICES AND PROCEDURES
J. G. Webster
University of Wisconsin-Madison/Department of Biomedical Engineering,
1550 Engineering Drive, Madison WI 53706 USA webster@engr.wisc.edu
Abstract: Ablation is a method of delivering
physical, chemical, or energy treatment to tissue for
the purpose of removing, altering, creating scar
tissue or causing apoptosis (cell death). Cardiac
accessory pathways permit ventricular excitation to
escape to the atria and cause ventricular
tachycardia. Catheters permit mapping the location
of the accessory pathways. 2.6 mm diameter
catheter electrodes pass 450 kHz power of about 20
W for about 20 s to heat the pathway above 50 °C to
kill the pathway tissue. To kill hepatic cancer, the
radiologist inserts a probe percutaneously, expands
an umbrella-like array and applies power for about
8 min. Alternatively in an open procedure the
surgeon can heat using a variety of probes or freeze
using cryo-ablation. To avoid exceeding 100 °C,
which causes charring, steam and popping, saline
may circulate through the probe interior or perfuse
into the tissue. Microwave ablation provides
quicker heating to prevent hepatic vessels from
carrying heat away during ablation. Finite element
method electrical-thermal models help to develop
new methods including bipolar, multiple, phased
array, noncontact and needle electrodes.
Applications include ablation of prostate, brain,
gastrointestinal tract, capsule, breast, varicose
veins, blood clots, skin wrinkles, cornea, teeth, and
bone.
Introduction
Current technology has allowed radiofrequency
(RF) ablation to be a more established option compared
to other types of ablation for surgery alternatives,
especially in treating supraventricular cardiac
arrhythmias. However it has limitations that might be
answerable by other ablation methods.
A typical frequency range used for RF ablation is
between 300 to 1000 kHz. An exception for this usual
range should be made for the Thermage TM ThermaCool
Tissue Contraction which uses 6 MHz. Radiofrequency
could also be used to cut tissue (RF electrosurgery) and
to seal vessels (RF electrocautery).
The most popular cancer treatment applications of
RF ablation are in liver cancer. They are performed in
around 200 centers in the USA. Although it might not
cure, it has been found to improve quality of life and
lessen the symptoms of hormone secreting tumors [1].
Lung cancer patients might consider having RF
ablation for several reasons. First, RF ablation can
preserve more lung function compared to surgery.
Next, RF ablation in addition to chemotherapy has
been shown to be a more effective treatment than
chemotherapy alone [2]. Chemotherapy causes tissue
to become more sensitive to heat. RF ablation helps to
ablate tissue in a region with less blood flow, which is
difficult to reach by chemotherapy [2].
Impressive cancer pain relief caused by RF ablation
for bone metastases has received FDA approval. Not
only does the RF process ablate nerves, accounting for
pain of degree 7.5 on a scale of 10 (10 being the
worst), it also kills some of the cancer tissue [3]. After
12 weeks, the pain degree decreases continually to
reach an average of around 2.7.
Table 1 Several RF ablations or hyperthermia therapy
done in different organs or treatments showing similar
technique due to similarity in type of tissue to be
ablated or manipulated.
Types of target
ablation
Muscle cell
(goal: tissue
necrosis)
Collagenous
connective cell
(goal: collagen
shrinkage)
Nerve
Comparison of RF ablation
RF ablation applications
Cardiac arrhythmia; tumor in
liver, kidney, breast, prostate,
pancreas; menorrhagia
Ligament shrinkage; skin
tightening; varicose vein
treatment; corneal modification;
low back pain treatment
Low back pain treatment,
pallidotomy, thalamotomy
Compared to surgery, RF ablation (especially the
percutaneous RF ablation) and other minimally
invasive ablation techniques have the following
advantages: smaller morbidity, shorter stays in the
hospital, less pain (smaller incision) and reduced
sedation or even elimination of general anesthesia.
Weighed against microwave, laser, and ultrasound
system, RF ablation does not need an additional
transmitter such as fiber optic for laser, antenna for
microwave and transducer for ultrasound system. This
advantage allows RF ablation to be simpler and its
technology to be developed quicker.
Judged against microwave ablation, the main
advantage of RF ablation is its relatively controllable
process. In microwave ablation, the surgeon does not
know exactly how long the microwave signal needs to
be delivered. This can lead to indefinable hyperthermia
status leading to uncertainty effects. Heat-induced
IFMBE Proc. 2005;9: 9

Keynote
blebbing has been a reported effect; the effect of the
blebs being carried away by blood flow is still
unknown. The design for microwave generators has
been more complex due to the nature of the power loss
along the catheter by the possible presence of
unmatched transmission line. However, microwave
ablation has been found to have better ablation toward
tissue nearby blood vessels and thus reduces tumor
recurrence rate. It is also known to have deeper lesion
performance and lesion control through different
shapes of antennas.
Laser ablation outperforms RF ablation in its
precision. However, cost and safety might lead to RF
ablation as the selected choice. It has been shown that
Interstitial bipolar RF-thermotherapy (RFITT) has had
a comparable performance with those of Laser Induced
Thermotherapy (LITT) [4]. In addition there is no
carbonization occurring in RFITT as in LITT for bone
ablation [5]. Anther advantage of RF ablation is that
the fiber optic transmitter for lasers needs to be as
straight as possible to deliver maximum power. In RF
ablation a maneuverable catheter is able to reach
relatively difficult locations.
Finite Element Modeling
Temperature distribution in tissue surrounding the
electrode can be explained using the bioheat equation:
∂T
c = ∇ ⋅ k∇T
+ J ⋅ E − Q h
∂t
ρ .
The RF current flows through the tip of the cardiac
electrode and then to the ground pad placed usually on
the back of the patient’s body. The target tissue
becomes a heat source by Joule heating, sometimes it is
also called resistive or volume heating (represented by
J·E). J is current density and E is electric field. The
tissue provides resistance to the current generated from
the catheter tip, which then causes tissue heating. In the
bioheat equation, k is thermal conductivity and T is
temperature thus representing the conduction process.
ρ is mass density and c is specific heat. Liquid or gas
near a heat source can cause convection. Blood
perfusion acts like a heat sink carrying away heat loss
Q h .
The bioheat equation is also used modeling liver
ablation [6].
Summary and future research
Since wide application of RF ablation is relatively
new, studies that evaluate long-term effects of the
method still need to be investigated. In the case of pain
relief of bone metastatic cancer, for example,
researchers still have questions about how long the
pain relief will last.
More experiments also need to be done to compare
RF ablation and other ablation methods. It has been
said that since much of cryoablation has been done
earlier than RF ablation, researchers may have taken
into account more problems with cryoablation. With
this information, the comparison with liver ablation
could be inaccurate [7].
Other methods are being considered to increase RF
ablation efficacy. These include combining RF
therapies with chemotherapy and/or radiation. Both
saline and ferric oxide MRI contrast agent have been
shown to increase heat conduction within tissue [8].
In summary RF ablation has been shown to have
similar functions as surgical resection with the
potential of being minimally invasive. RF ablation
efficacy has been improved mainly by electrode
design. Several other methods have since been
developed to increase lesion size and depth of RF
ablation. Some of them are the Pringle maneuver,
pulsed signal, cool–wet technique and combined
therapies with chemicals or chemotherapy. The main
goal of improving RF ablation currently includes
increasing efficacy for ablating tissue nearby blood
vessels to reduce high recurrence rate. Complications
in RF ablation are usually very minimal. RF ablation is
relatively less complex and costly compared to other
ablation techniques. RF ablation in general has
performed well, although physiological process of
tissue necrosis is not yet clearly understood.
References
[1] CLEVELAND CLINIC (2003): ‘Liver Tumor Ablation
Program: Frequently Asked Question’. [Online]
www.clevelandclinic.org/general/rfa/faq.html
[2] CANCERABLATION.COM (2003): ‘RF Ablation’.
[Online] www.cancerablation.com/index.html
[3] LOWE, R (2003): ‘Procedure Offers Relief of
Intractable Bone Pain’. [Online]
www.cancerpage.com/cancernews/cancernews5097.ht
m
[4] DESINGER, K., STEIN, T., MULLER, G., MACK, M.,
VOGL, T. J. (1998): ‘Interstitial bipolar RFthermotherapy
(RFITT) Therapy Planning by computer
simulation and MRI-monitoring – A new concept for
minimally invasive procedures’. Proc. Surg. Appl.
Energy, 3249, pp. 147–9
[5] GROENMEYER, D. H. W., SCHIRP S., GEVARGEZ A.
(2002): ‘Image-guided percutaneous thermal ablation
of bone tumors’, Academic Radiol., 9, pp. 467–77
[6] HAEMMERICH, D., TUNGJITKUSOLMUN, S., STAELIN,
S. T., LEE, F. T. JR., MAHVI, D. M. AND WEBSTER, J. G.
(2002): ‘Finite element analysis of hepatic multiple
probe radio-frequency ablation’, IEEE Trans. Biomed.
Eng., 49, pp. 836–42
[7] MAHVI, D. M, LEE, F. T., JR. (1999):
‘Radiofrequency ablation of hepatic malignancies: Is
heat better than cold?’. Ann. Surg., 230, pp. 9–11
[8] GOLDBERG, S. N. (2001): ‘RF Tumor ablation:
improving therapeutic efficacy with combined
therapies’, Proc. SPIE Prog. Biomed. Optics Imaging,
4247, pp. 41–52
IFMBE Proc. 2005;9: 10

Keynote
HEART RATE VARIABILITY: MODELS, METHODS, AND APPLICATIONS IN
STRESS TESTING
P. Laguna 1
1 Inst. for Engineering Research, Zaragoza University, Zaragoza, Spain
Abstract
The study of heart rate variability (HRV) has become
increasingly popular because information on the state
of the autonomic nervous system (ANS) can be
noninvasively inferred by the use of relatively basic
signal processing techniques.
Despite the seeming simplicity of deriving the series of
RR intervals from the ECG signal and defining a
related measure of dispersion, it is essential to make
sure that the heart rate variability is accurately
characterized. Several definitions of signals for
representing the heart rhythm have been suggested
which characterize variability either in terms of
successive RR intervals or instantaneous heart rate
(HR). It is difficult to established criteria to judge
which of the HRV representations is better suited for
clinical purposes. However, adequate modelling of the
underlying physiological mechanisms could give us a
framework for this analysis, under the assumption the
modelling correctly represent the phenomena. In
particular, Integral Pulse Frequency Modulation
(IPFM) models are used to connect the ANS action
with the observed HR series, and the HRV
representations are reviewed under the view of this
modelling.
Spectral analysis of heart rhythm signals has received
considerable attention since oscillations embedded in
the rhythm, for example due to respiratory sinus
arrhythmia or the blood pressure control system, can
be quantified from their corresponding peaks in the
power spectrum. Such oscillations are characterized
by low frequency components, which typically are
located in the interval below 0.5 Hz. These techniques
are also reviewed.
HRV representations assume that the beating time
results from a sine atrial beat, otherwise no direct
relation with the ANS is included in the beating time.
This requires to be sure that ectopic and other
abnormal beats are not taken into consideration, but
even more, their action on neighboring sine atrial
beats need to be considered and corrected to prevent
modifications from the disturbances introduced by
the ectopic beats. Strategies to deal with this problem
are discussed.
Both, HRV modeling and spectral estimation, work
under stationary conditions, circumstances that are
not always satisfied in clinical practice like stress test
analysis, rest-to-tilt trials and others. For these
situations time-varying modelling and time-frequency
analysis will be better suited. An outline of this
tendency will be reviewed in stress test data.
HRV representations assume that the beating time
results from a sine atrial beat, otherwise no direct
relation with the ANS is included in the beating time.
This requires to be sure that ectopic and other
abnormal beats are not taken into consideration, but
even more, their action on neighboring sine atrial
beats need to be considered and corrected to prevent
modifications from the disturbances introduced by
the ectopic beats. Strategies to deal with this problem
are discussed.
Both, HRV modeling and spectral estimation, work
under stationary conditions, circumstances that are
not always satisfied in clinical practice like stress test
analysis, rest-to-tilt trials and others. For these
situations time-varying modelling and time-frequency
analysis will be better suited. An outline of this
tendency will be reviewed in stress test data.
HRV representations assume that the beating time
results from a sine atrial beat, otherwise no direct
relation with the ANS is included in the beating time.
This requires to be sure that ectopic and other
abnormal beats are not taken into consideration, but
even more, their action on neighboring sine atrial
beats need to be considered and corrected to prevent
modifications from the disturbances introduced by
the ectopic beats. Strategies to deal with this problem
are discussed.
Both, HRV modeling and spectral estimation, work
under stationary conditions, circumstances that are
not always satisfied in clinical practice like stress test
analysis, rest-to-tilt trials and others. For these
situations time-varying modelling and time-frequency
analysis will be better suited. An outline of this
tendency will be reviewed in stress test data.
IFMBE Proc. 2005;9: 11

Keynote
THE EUROPEAN HIGHER EDUCATION AREA IN BIOMEDICAL
ENGINEERING – ACHIEVEMENTS, TRENDS AND DEVELOPMENTS
Joachim H. Nagel
Department of Biomedical Engineering, University of Stuttgart, Stuttgart, Germany
jn@bmt.uni-stuttgart.de
The Bologna Process, aiming at setting up the
European Higher Education Area (EHEA), is dramatically
changing the national systems of higher
education in most of the participating 45 countries,
and has created a unique opportunity to promote
Medical and Biological Engineering and Sciences
(MBES). The movement has triggered an initiative
of the European MBES community to establish their
Higher Education Area by harmonizing the educational
programs, setting up best educational practices,
specifying minimum qualifications and establishing
criteria for an efficient quality control of
education, training and life-long learning. The main
objectives of the current initiatives within MBES are
to establish general European consensus on guidelines
for harmonized, but not standardized, high
quality MBES programs, their accreditation and for
certification and continuing education of professionals
working in the health care systems. Adoption and
adherence to these guidelines will ensure mobility in
education and employment throughout Europe and
still leave the necessary freedom for the European
countries and universities to maintain their unique
identities.
Starting in 1999, the International Federation for
Medical and Biological Engineering (IFMBE), its
European member societies and numerous European
universities with an interest in MBES, as well as the
European Alliance for Medical and Biological Engineering
and Sciences (EAMBES), an association of
European national and trans-national societies which
was founded in 2003 on the initiative of IFMBE,
have been engaged in projects aiming at creating a
comprehensive survey of the status of MBES education
and research in Europe, charting the MBES
community, developing recommendations on harmonized
MBES education and training, and establishing
criteria for the accreditation of MBES programs
in Europe.
In 2004, a Europe-wide participation project,
BIOMEDEA, has been launched, aiming at contributing
to the realization of the European Higher Education
Area in MBES. The objective of the project is
to establish Europe-wide consensus on guidelines for
the harmonization of high quality MBES programs,
their accreditation and for the certification or even
registration and continuing education of professionals
working in the health care systems. Improved
quality assurance of MBES education and training is
a vital component and is also directly related to the
issues of health care quality. It offers the advantages
of providing confidence for the employer that the
employee has the necessary education, training and
responsible experience, and the reassurance for the
user of the service, meaning the patients, that those
providing the service are effective and competent.
Adherence to these guidelines will insure mobility in
education and employment, and improved competitiveness
of the European biomedical industries.
The expected results of BIOMEDEA will be a
white paper on BME education, educational methods
and best practices in Europe, protocols for the formation,
training, certification and continuing education
of clinical engineers in Europe, and guidelines for
the accreditation of BME programs in Europe.
IFMBE, the main sponsor of BIOMEDEA, will, in
cooperation with WHO, as a part of the initiatives of
the World Alliance for Patient Safety, set up a global
registry of certified clinical engineers with the goal
of world wide mutual recognition of certification,
and strive towards making certification and registration
of clinical engineers mandatory everywhere in
the world, based on the same criteria.
Primary goal of BIOMEDEA remains, however,
to prepare the BME European Higher Education
Area and to find recognition by the national governments
throughout Europe, the European Union and
those European bodies that are the main players in
engineering education and accreditation.
A new project with broad support from the European
bodies entrusted by the Bologna countries with
the task of establishing European standards and procedures
for quality assurance and accreditation in
higher education, EUR-ACE (Accreditation of European
Engineering Programs and Graduates) aims at
setting up a European system for accreditation of
Engineering education with the following main
goals: provide an appropriate “European label” to the
graduates of the accredited educational programs,
improve the quality of educational programs in engineering,
facilitate trans-national recognition by the
label marking, facilitate recognition by the competent
authorities in accord with the EU Directives and
facilitate mutual recognition agreements.
BIOMEDEA, and thus the European biomedical
engineering community, has been accepted to represent
the specific issues and interests of MBES within
EUR-ACE, and to test its specific criteria for accreditation
within the project. For the BME community
this is a major step forward towards the establishment
of EHEA under due consideration of the
specific needs of Medical and Biological Engineering
and Sciences.
IFMBE Proc. 2005;9: 12

Keynote
EDUCATION AND TRAINING OF THE EUROPEAN MEDICAL PHYSICIST;
ROLES AND RESPONSIBILITIES IN RADIATION PROTECTION AND
SAFETY.
I. Lamm 1
1 Radiation Physics, Lund University Hospital, SE-221 85 LUND, SWEDEN
Abstract
European legislation has challenged many professional
organisations to propose harmonised professional
standards of high quality. The Directives of the Council
of the European Union concerning medical exposures and
basic safety standards have given a statutory requirement
for physicists to be involved in the medical uses of
ionising radiation, and have given impetus to the
discussions of education and training requirements in
medical physics. Whilst these Directives primarily deal
with medical radiation physics, their consequences will
also effectively set the standards for other branches of
medical physics. They will gradually affect every
European country, even though they are binding only on
EU Member States.
The medical physicist working in a clinical setting is a
member of the clinical team responsible for diagnosis and
treatment of patients. The qualified medical physicist has
a unique competence and carries a range of
responsibilities in his/her area of practice; for equipment,
techniques and methods used in the clinical routine, for
the introduction, adaptation and optimisation of new
methods, for calibration, accuracy, safety, quality
assurance and quality control, and generally also for
many areas of research and development.
In order to acquire and maintain sufficient knowledge and
an appropriate level of competence, both initial and
continuing education and training are necessary.
The European Federation of Organisations for Medical
Physics, EFOMP, is an umbrella organisation for
National Medical Physics Organisations, with one of its
main objectives to harmonise and promote the best
practice of Medical Physics within Europe. The EFOMP
approach to reach this objective is to encourage the
establishment of national education and training schemes
at all levels, in line with EFOMP recommendations. The
EFOMP efforts, resulting in recommendations on a
structured system for education and training, have been
recognised by the European Community.
inger-lena.lamm@skane.se
IFMBE Proc. 2005;9: 13

Healthcare assessment and clinical engineering
TRENDS IN HEALTH CARE ASSESSMENT
J. Persson 1
1 CMT, Linköping University, Linköping, Sweden
jan.persson@ihs.liu.se
Abstract
Trends in healthcare assessment – impact on adoption of
medical devices
Healthcare assessment, usually called ”health technology
assessment” (HTA), has developed rapidly as a tool for
policy making over the last decade. The gap between
what is affordable and what is technologically possible is
huge and increasing. There is, therefore, a need for
mechanisms to manage decision making and priority
setting in health care. The complexity of the system is
indicated in Fig.1. Various actors with different
incentives and often conflicting interests are involved. An
urgent issue is the role of HTA in the diffusion of
medical devices into healthcare. Are there losers and
winners in the game and where is biomedical engineering
and medical devices?
Fig. 1. The innovation and diffusion process for health
technology
A number of important achievements described below
contribute to the strengths and impact of HTA.
Health economy and outcomes analysis are presently well
established and recognised areas. Cost-effectiveness
(CEA) and cost-utility analysis (CUA) are widely used
and problems and pitfalls extensively reported and
debated. QALYs (Quality Adjusted Life Years) gained
through interventions are often used in decision making.
In policy making, classifications of cost-effective,
intermediate cost-effective and cost-ineffective are often
used when considering new technologies and their
possible adoption in healthcare. The concept of QALYs
has an increasing acceptance. In reports of applied studies
from 1975 – 2001, few studies (CEA, CUA, CBA (cost
Benefit Analyses) were reported up til 1985. From 1985
the number of CBA increased moderately, while number
of CEA increased dramatically and CUA even more
(OHE Health Economic Evaluations Database 2003).
Evidence based medicine (EBM) has developed in order
to address the above described problems. EBM means a
conscious and systematic use of best available
knowledge. Clinical experience should be combined with
best scientific evidence from external sources. A new
approach to the use of previous knowledge evolved,
much thanks to Archie Cochrane in the 70ties, leading to
the formation of the world-wide Cochrane Collaboration
with international data bases. Meta-analyses are used to
improve strength of evidence, to a high degree based on
randomized control studies. National agencies for HTA
have been established, several participating in the
INAHTA (International Network of Agencies for Health
Technology Assessment). The scientific society HTAi
(Health Technology Assessment International) was
established two decades ago, the International Society for
Priority Setting in Heath Care a few years ago. Also,
there is a network dealing with the adoption of new
methods, Euroscan – The European Information Network
on New and Changing Health Technologies.
What is the consequences of the progress described? It
seems that the demand for scientific evidence in policy
making means that some are winners (pharmaceuticals)
and some are losers (rehabilitation, health promotion)?
Where are medical devices?
Medical devices are important in effectiveness of
healthcare, but also a driver in increase of healthcare
expenditure. The global market for drugs was in 2002
USD 250 billion, while the market for medical devices
was 169 billion in 2001 and expected to be 260 billion in
2006 (US FDA estimates). It was also estimated that 10
000 new medical devices were introduced into the market
2000, i.e. many times more than the number of new
drugs. The issue of rational adoption of devices is
obviously important.
EBM for device-related interventions may be more
difficult to achieve than for drugs. Well-controlled
prospective clinical trials means design challenges.
Outcomes are influenced not only by the device but also
by the skill of the user (professional or patient) and by the
environment. Blinded treatments can rarely be obtained.
The manufacturer of devices is often a small company,
not having the same resources for sponsoring clinical
trials as the pharmaceutical companies.
Therefore, with a requirement for blinded RCTs to reach
the highest strength of evidence, it seems that it is often
more difficult than for drugs to enter the market.
Challenges are to proceed the work with strong studies on
effectiveness and cost-effectiveness of medical devices,
and to refine criteria for adoption of new non-drug
technologies.
IFMBE Proc. 2005;9: 14

Healthcare assessment and clinical engineering
BUSINESS DEVELOPMENT OF BIOMEDICAL ENGINEERING INVENTIONS
O. Lindahl 1
1 Centre for biomedical engineering and physic, c/o TFE, Umeå University, Umeå, Sweden
olof.lindahl@tfe.umu.se
Abstract
Biomedical engineering research is of potential interest
for industry. The business development of a company
based on a research result about eye pressure monitoring
is described in this paper. The business development
process is discussed. The importance of patents and
financial as well as market alliances are evident. The
teamwork between different competences are concluded
very important.
Introduction
Biomedical Engineering is a rapidly developing field in
Science and Industry. Umeå University aims to take a
leading part in this field. Therefore Umeå University has
decided to establish a Centre for biomedical engineering
& physics. The means are co-operation between research,
industry and health care and the goals are products and
methods for a better and more safe Health Care. Regional
goals of the centre is to, in the region: strengthen current
industry, establish new companies due to innovations,
locate mature companies due to cutting edge competence,
create new employments and develope infrastructure.
The aim of this study was to establish a company based
on a patent from a research project on eye pressure
monitoring and the paper describes how this was
achieved.
Methods
Intra ocular pressure (IOP) monitoring is important for
screening, diagnosing, and following up glaucoma
patients. The incidence of glaucoma is 5% of the human
population older than 70 years. There are several existing
methods today but Goldmanns Applanation Tonometry is
regarded by most physicians as the golden standard.
However the needs for an instrument which is more
easily handled for IOP measurements are widely
documented. In a PhD project within the Centre for
Biomedical Engineering and Physics at Umeå University
(CMTF), we used resonance sensor technology 1 for
measuring hardness of prostate tissue 2 . During the study
we found that resonance sensors could measure contact
area. We naturally discussed new application areas for
this new feature of the resonance sensor and found that
we could possibly develope a new IOP-instrument since
IOP is measured indirectly by measuring contact area (A)
and contact force (F). According to Imbert Ficks law the
IOP=F/A. Thus, we started construction and research on
IOP with resonance sensors 2 .
Results
A patent was filed in the year 1999. Negotiations
commenced with NUTEK about seed finances and an
agreement was reached with the condition that we found
as much fundings from other sources also. Through a
personal network we got in touch with an investment
company that was prepared to match the NUTEK
funding. In the year 2000 the company Bioresonator Co,.
Ltd was established and situated in Umeå. A business
plan, a market investigation and a design project for the
product were achieved. Several agreements between the
company and the researchers, Umeå university and the
county council, as well as a product plan was drafted. In
the period 2002-2004 the research continued to be
performed in parallel to that new finances for running the
company was chased. Patents and design-patents
were filed and payed in several countries and the CEmarking
procedure for the product was started up.
Furthermore, business alliances were searched for and
there was an intensive chase for a commercial partner
with market experience in the eye pressure monitoring
field.
Discussion
The process of starting a new company from a research
result is difficult but also rewarding. In our case we had
help from Uminova (Umeå university innovation centre)
with the business development and we had the luck to
find an investment company that believed in our product
ideas. The involvement of different people with different
competences like research, product development,
economy, law, business development, patent etc. has
been very important. This is of course since not any
single person has all this competence him or herself. One
difficult question is when the time is right to start the
commercialisation of a research result. In what stage
should the research be, in order to be enough mature for
productification. Usually the business development takes
more time than one usually think and so was the case for
us too. The entusiasm made our project survive though
(Fig 1). However, today we are signing a contract with a
large company in Europe and we plan to have the market
introduction in the autumn of 2005.
Conclusions
Commercialising a research result is not easy and usually
it takes much more time than one could expect and the
teamwork between key-persons are extremely
important. In the unusual commercialisation process for
IFMBE Proc. 2005;9: 15

Healthcare assessment and clinical engineering
the scientist, there are som important things to concider:
the network is very important, your patients have to be
strong and you have to strive on, a patent is a good
protection but also very costly, involve people with many
competences in your company and do not forget the
cumbersome CE-marking process.
References
1 Omata and Terunma, New tactile sensor like the human
hand and its applications, Sensors and Actuators A, 35,
9-15, 1992
2 Eklund, Resonator sensor technique for medical use,
Umeå university dissertation 801, 1992
IFMBE Proc. 2005;9: 16

Healthcare assessment and clinical engineering
TUBERCULOSIS-THE SILENT KILLER OF DEVELOPING WORLD AND
WHERE WE ARE?
M. Rahman 1
1 Land & Water Resources Engg., MSc. in EESI Program, Royal Institute of Technology, KTH,
Stockholm, Sweden
tauhid_cee@yahoo.com
Abstract
The misfortune of this current century is that in spite of
overwhelming technical advancement, it has to be a
witness of a lot of tragic deaths of the poor human beings
due to the epidemic of Tuberculosis (TB). Realizing the
gravity of out bursting of TB, the United Nations,
represented by World Health Organization (WHO),
declared the year 1993 as the global emergency year
which ultimately resulted in the announcement of the
Millennium Development Goals (MDGs) in 2001. The
target of the MDG s is to detect the case (70%) and to
turn around likelihood (85% success rate) of it within the
coming decade (WHO, 2004).
An evaluation, forecasted by WHO (2002), illustrates that
around two million of the TB infected population-which
is again more than one third of the total population-die
annually. The worst hit states- the 22 High Burden
Countries (HBCs) are from the Sub Saharan Africa,
South East Asia and East Europe regions. From 1997 to
1999, globally, 6% increment of new TB cases was
noticed. Again, in the last ten years, the males of
Botswana have lost their average life expectancy by 23.8
years due to the lethal combination of HIV and TB
(Paluzzi & Kim’2003, p.7).
TB combat program. DOTS’ success is dependent on
widening of identification and taking care of TB patient.
Data of notification and arising of smear positive cases in
a particular year are not always possible to collect as
uncertainty lies in arising of smear positive cases in
HBCs. One of the strongest barriers to combat TB is the
prolonged duration of treatment phase, which often make
patients frustrated.
In order to fulfill the MDG target, WHO should maintain
a worldwide database service to keep records of case
detection of patients and the places where acute
emergency of TB cases is found, immediate treatment
and logistics should be provided. High quality of first line
and second line anti TB drugs should be supplied,
uninterruptedly. Coordination of effective intersectoral
actions should be ensured in worst hit regions,
immediately.
In this context, a desk top study has been carried out to
assess the global tuberculosis situations and to evaluate
the status of the Millennium Development Goals (MDGs)
and further to discuss how far we are in reaching the goal.
In 1994, WHO recommended DOTS (Directly Observed
Treatment Short course, the effective tools for fighting
against TB. Around 180 of the 210 (WHO, 2004)
countries in the world are under the coverage of DOTS
program.
DOTS identified 37% of active TB cases within its total
coverage, in the year 2002 (The Global Fund, 2004). If
the present style of controlling of TB is expected to be
continued, then it needs to be waited up to 2013 to
observe the 70% success rate of case detection and
efficient treatment. The prime obstacles to reach the
MDGs’ target are inadequate political commitment, in
availability of funds, lack of skilled health workers, poor
health infrastructures and logistics, poor quality and
interrupted supply of anti TB medicines and lack of
collaboration and supervision between community and
IFMBE Proc. 2005;9: 17

Healthcare assessment and clinical engineering
ADOPTION OF MEDICAL DEVICES: THE NEONATAL INTENSIVE CARE
UNIT AS A CASE STUDY
K. Roback 1 , P. Gäddlin 2 , N. Nelson 3 , J. Persson 1
1 Center for Medical Technology Assessment, Linköpings universitet, Linköping, Sweden
2 Division of Pedriatics, County Hospital Ryhov, Jönköping, Sweden
3 Division of Pedriatics, Department of Molecular and Clinical Medicine, Linköping university,
Linköping, Sweden
kerstin.roback@ihs.liu.se
Abstract
Adoption and deployment of healthcare technologies
will be an important issue in the shaping of future
healthcare systems. This study investigates factors
affecting adoption of medical devices in the neonatal
intensive care setting. Interviews have been held with
medical professionals. Our results show that a major
inhibitor of diffusion is low compatibility with other
equipment and working routines, while the perceived
relative advantage is an important facilitator.
Futhermore the supply push is a strong driving force
in the diffusion process and at the ward unit available
communication channels may play an important role
in choice of new products.
Key words: medical devices, diffusion of innovation,
adoption, decision making
Introduction
Adoption of new medical devices often requires large
economic and educational resources. Therefore it is of
vital importance that adoption decisions are guided to
reach optimum benefits. To achieve this, developments
are needed in two main areas: (1) the availability and
diffusion of scientific evidence of benefits, risks, and
costs associated with the devices and (2) knowledge
about the healthcare innovation process, including the
large number of non-medical determinants involved in
adoption decisions. This study sets focus on the second
area. Adoption and deployment of medical devices in the
neonatal intensive care unit (NICU) is studied as a part of
a larger study of the market for medical devices, which
includes healthcare organizations and professionals,
manufacturers, distributors and political decision makers.
The aim is to identify important inhibitors and facilitators
in the diffusion process, to identify important actors in
management of technological change and to find methods
to achieve a more evidence based adoption of medical
devices.
Adoption decisions include both purchase of new devices
and replacement of technically obsolete or worn out
devices. Comprehensive understanding of the adoption
process is only achieved through a synthesis of different
perspectives, e.g. technical, economical, organizational,
ethical and political. The NICU study represents the
healthcare perspective on factors affecting the success or
failure of medical devices in the market.
Possible actors and determinants were identified in a
literature study [1] and research questions were
formulated. Questions of special interest in this case
study are:
• Which are the most important information sources
affecting adoption decisions? How does the information
reach the ward unit?
• What are the differences between judgments on devices
of different NICUs? What is the cause of these
differences?
Methods
Data collection is performed through interviews. An
iterative approach is applied where data collection and
analysis are alternately performed. This enables early
results to be included in the course of letting the set of
questions converge into those most apt to catch central
issues of the study. Respondents are healthcare
professionals and clinical engineers expected to work
with devices for neonatal intensive care in Swedish
hospitals. Devices at issue in the interviews have been
subject to an adoption decision and may either be in use
at the ward or rejected in the decision process.
Respondents can suggest devices that ought to be
included in the study. Devices included are e.g. infusion
devices and similar pumps, ventilators, incubators, blood
gas analyzers, patient monitoring equipment, glucometers
and pulse oximeters.
Results
The iterative character of the study enables presentation
of results at an early stage. The following preliminary
results can be drawn at present.
Explanatory factors of adoption rate
Rogers [2] suggests five characteristics of innovations as
main explanatory factors of adoption rates: Relative
advantage, compatibility, complexity, trialability, and
IFMBE Proc. 2005;9: 18

Healthcare assessment and clinical engineering
observability. According to Rogers, the rate of adoption
is positively related to perceived relative advantage,
compatibility, trialability, and observability, and is
negatively related to perceived complexity of the
innovation. Our study confirms that perceived relative
advantage is the most important factor, i.e. new devices
must be perceived as adding a value to the treatments,
the ward unit or the hospital. Complexity as a negative
force, on the contrary, could not be confirmed in this
setting. Neither are the rising costs perceived as a strong
inhibitor to adoption, but budgetary limits sometimes
acted to slow down the adoption rate.
Supply push vs. demand pull
Rosenberg [3] argue that the supply side variable acts as a
strong driving force, especially in medical innovation.
There is a growing stock of useful knowledge seeking its
applications. This is true for most new health
technologies, but also to some extent for replacement
devices. When devices are worn out demand for a
replacement device arises. But the new models available
on the market are often equipped with several new
functions and alterations, which are not initially asked for
by the potential buyer. These product developments are
not always perceived as advantages and do cause a lot of
uncertainty in the adoption decision.
Discussion
Introduction of new devices in healthcare may raise
ethical questions about healthcare need and risk, equity,
authority and control. Whose needs are to be met in a
future healthcare system? How are healthcare resources
allocated when technological advances enable us to
successfully treat increasingly severe conditions? We
expect that a thorough analysis of the results of this study
will contribute to these discussions by an identification of
areas where medical innovation can be made more
effective in a healthcare and a societal perspective.
References
[1] ROBACK, K., PERSSON, J. and HASS, U. (2003):
‘Diffusion and implementation of medical devices:
Background report’ [in Swedish], Centrum för
utvärdering av medicinsk teknologi, Linköpings
universitet, CMT-rapport 2003:1
[2] ROGERS, E. M. (1983): ‘Diffusion of Innovations’,
3rd edition, first ed. 1962 (New York: Free Press)
[3] ROSENBERG, N. (1974) Science, ‘Invention and
Economic Growth’, The Economic Journal, 84:333, pp.
90-108
Information sources and communication channels
Information about health technologies is often sought
from near colleagues, especially information about their
subjective evaluations of the products, while initial
information about a new technology is typically
presented by the manufacturer, either at an exhibition,
medical conference or by a sales representative visiting
the hospital. Inter hospital communication channels and
other external sources are primarily used by the doctors.
Rogers [2] found that early adopters of innovation have
more education and higher social status than late
adopters. A trend in our study, which could partly
explained by, is that the more contact medical staff have
with external sources, the higher value these sources are
ascribed.
Choice of equipment
Experience from the device in an actual caregiving
situation is crucial for the choice of replacement
products, while more innovative new devices could be
introduced without prior testing at the ward. The cause of
rejection of devices after the test period could be e.g. a
bad user interface, functional deficiencies or low quality
introductory education.
In some cases evaluations of the products differ between
hospitals. The most common explanation is that the new
device had low compatibility with other equipment at the
ward or with working routines. In a few cases is seems
likely that individual preferences and available
communication channels of high professionals have been
conclusive in the adoption decision.
IFMBE Proc. 2005;9: 19

Healthcare assessment and clinical engineering
THE NEED OF PROCEDURES COMPILED WITH MDD TO INVESTIGATE
MEDICAL DEVICE FAILURES INVOLVING PATIENT INJURY
H. Gilly*
* Department of Anaesthesia, Medical University of Vienna, A-1090 Vienna, Austria
hermann.gilly@meduniwien.ac.at
Abstract: The medical device directive [MDD; 1]
gives a straightforward procedure how to report and
what to do with a failing medical device involving
patient injury. Two cases with malfunction of
anaesthesia machine were examined and in both
cases the investigator was confronted with the fact of
a culture of messy documentation. Standard
operating procedures tailored to the specific needs of
the department involved and complying with the
MDD may help in overcoming present deficits.
Introduction
The medical device directive [MDD; 1] gives a
straightforward procedure how to respond and report
and what to do with a failing medical device involving
patient injury. We have noted, at least on two occasions,
that confronted with the medical device failure the staff
was not sufficiently trained to handle the serious
adverse events and its straightforward reporting in order
to provide objective and comprehensive information in
the course of actions taken after the equipment failure.
Materials and Methods
Case 1: Malfunction of an anaesthesia machine due
to an apparent breakdown of the inlet to a micro filter
protecting a pressure transducer used for monitoring.
Case 2: Missing gas flow from the y-piece.
Case3: Foreign body displaced into the lung.
All three incidents caused patient injury.
Results
The cases were reported in due time to the
manufacturer and the national body. However, when the
(legally authorized) investigator tried to fully trace the
device failure in retrospective he was surprised by
missing documentation (no photos of the equipment
status, none of the broken plastic parts). When the
service personnel of the manufacturer’s representative
checked the equipment they readily found the fault and
in case 1 repaired it on spot; in case 2 the anaesthesia
department staff performed a check and changed the
failing subunit. In case 3 only the detached part was
secured. The failing part (case 1) was sent to the
manufacturer for further investigation (no picture
archiving). Unfortunately the micro filter was sent to the
wrong OEM source, finally discarded and no more
available to the investigator.
A somewhat different approach in handling case 2
was noted, but as in case 1 the investigator was
confronted with the fact of a culture of messy
documentation. In case 3 yet missing but relevant
information could be collected retrospectively.
Discussion
With medicinal products the medical staff seems to
be experienced in post market reporting of serious side
effects of medicinal drugs whereas lacking experience is
to be suspected with medical devices.
According to a recent report [2] 1004 critical
incidents with a total of 20 deaths (12 due to device
failures) have been collected within a 1-year period.
33% of the incidents were related to ventilation equipment
(5 deaths) triggering equipment redesign at least in
2 cases. In Austria approx. 80 critical incidents are
reported to the authorities each year. However there is
no similar to the French [2] overview of national
registers on a European wide level even though there is
little doubt that post marketing vigilance is a most
useful way of improving the quality of medical devices.
Conclusions
Appropriate retrospective analysis of critical
incidents should be made accessible in due time and
anonymous form to the community on a European wide
level. Such a comprehensive data collection would
allow individual institutions to tailor their strategies and
adapt policies aimed at reducing the risks of critical
incidents in a pre-emptive approach.
In respect to the reporting system standard operating
procedures tailored to the specific needs of the
department involved and complying with the MDD may
help in overcoming present deficits.
References
[1] MDD. Council directive 93/42/EEC (1993). Art.10;
Information on incidents occurring following placing of
devices on the market.
[2] Beydon L, Conreux F, LeGall R et al. Analysis of
the French health ministry’s national register of
incidents involving medical devices in anaesthesia and
intensive care. Brit J Anaesth 2001; 86:382-7
IFMBE Proc. 2005;9: 20

Healthcare assessment and clinical engineering
IMPROVEMENT OF PATIENT SAFETY IN PRACTICE
- EXAMPLES OF PREVENTION OF ACCIDENTS
H. Teriö*
* Department of Biomedical Engineering, Karolinska University Hospital, Stockholm, Sweden
heikki.terio@karolinska.se
Abstract: Adequate routines to analyse the causes
of accidents or incidents in the health care work are
needed. The results from these analyses are
important in prevention of similar cases in the
future. The departments of biomedical and clinical
engineering should have a central role to work
together with the health care staff at the hospitals to
fulfil these tasks. At the Karolinska University
Hospital, Huddinge, two committees were set up to
work with improvement of patient safety. The
Product Safety Advisory Board with assignment to
support the clinics at the hospital with assessment of
medical technology, re-use of products for single-use
and in-house produced equipment. The Safety
Commission, that had the Swedish Accident
Investigation Board as an example, investigated
accidents and incidents at the hospital.
Introduction
Improper handling or usage, or technical failures of
medical equipment will always cause accidents and
incidents. To prevent these accidents the persons
working with the equipment must have a proper
education and training. Service of the equipment is a
natural part of the preventive work.
The discussion how to improve the patient safety
started in Sweden for almost 30 years ago. At that time
the electrical safety issues dominated, but even handling
of the medical gases and the radiation safety were dealt
with. The medical equipment, since those days have
been improved in many details because the apparatus
faults, accidents and incidents have been analysed.
In the beginning of 90’s quality assurance of the
biomedical and clinical engineering work was active in
a wide scale. This development has included both the
service and the organisation of the biomedical work at
the Swedish hospitals. The biomedical work includes of
course contacts with the health care staff; the doctors
and the nurses. Their education in technology and
training in handling the medical equipment have
become as a natural part of the engineering work at the
Swedish hospitals.
Since mid 90’s Sweden has a legislation that
requires that the health care organisations have quality
systems that assure the continuous development of, for
example, patient safety. Specific demands on the
responsibilities that the health care provider has when
using medical equipment or producing in-house
equipment were published at the same time. It became
also mandatory to analyse accidents and near accidents
and to report them to the authorities. How these tasks
are carried out and how the demands of the legislation
are met, vary from organisation to organisation and
from hospital to hospital. In any case there must be an
adequate routine to analyse the causes of the accident or
incident. The results from these analyses are important
in prevention of similar cases in the future.
The departments of biomedical and clinical
engineering should have a central role to work together
with the health care staff at the hospitals to fulfil these
tasks. At the Karolinska University Hospital, Huddinge,
the people from the department of Biomedical
Engineering have been natural members in the groups
that have analysed accident, conducted risk analyses and
safety assessments.
Materials and Methods
Two committees were set up to work with
improvement of patient safety. The first one was called
the Product Safety Advisory Board and it’s assignment
was to support the clinics at the hospital with issues
dealing with assessment of medical technology, re-use
of products for single-use and in-house produced
equipment. The members of this board represented
different areas within the hospital, for example the
hospital management, biomedical engineering, surgery,
radiology and orthopaedics.
The other board was the Safety Commission that had
the Swedish Accident Investigation Board as an
example. The overall aim of this commission was to
increase the knowledge of the basic reasons to the
incidents and accidents that has happened in order to
prevent occurrence of them. It was stressed that the aim
of this commission was not to act as an authority for
penalty. The members of this commission represented
different competences within the hospital and their
personnel suitability for the membership was critically
evaluated. The chairman of the commission was the
only person who had the right to give a statement about
an ongoing investigation. There was also a possibility to
call in experts for specific investigation.
The main task for the commission was to clarify the
course of events and analyse it to bring out the causes
and its effects. The investigations were carried out
methodologically considering the incident or accident
from different angles or areas of competence. In the first
interview with the persons involved the whole
commission was represented. In this way every member
IFMBE Proc. 2005;9: 21

Healthcare assessment and clinical engineering
of the commission got the same initial information, but
they could also notice different details in the given
statement. This interview is the first step in collection of
facts about the incident or accident and it includes all
persons that were involved. Their background,
education, competence, authorities and task were
mapped. The equipment involved was naturally
scrutinized. The adjustments, accessories, fittings,
checklists and logbooks were examined. Even the
organisation and management of the clinic/department
was investigated. The psychosocial climate at the
working place, instructions, routines and rules were
included in this step.
The commission must present its work in a written
report. It is important that the writing of the report must
involve every member of the commission who has taken
part in the particular investigation. It is also important
that no individual can be identified in the text. A
preliminary report should be written after only few days
after the work started. This report must describe the
course of events based on the interviews with the staff
involved. The draft of the report is also sent to the
persons who have been involved in the accident or
incident for comments. The report becomes public first
after the chair of the commission signs it. Every
member of the commission must after this point accept
the report without a possibility to make changes. If
some member has a different opinion in some detail this
should be written in the report before it is signed.
The last step for the commission in its work is the
follow up. The commission must acquire knowledge of
the measures taken to carry out the recommendations
that the commission had presented in its report.
Results
The Product Advisory Board had during the test
period 6 regular meetings where 10 issues were
discussed. 5 of the issues were dealing with resterilization
of medical products like orthopaedic
implants, pacemakers and handling of factory sterilized
products at the hospital are some examples. Other issues
that were linked to these were also discussed. For
example the ethical aspects of re-usage of pacemakers
or the economical benefits of re-sterilization of medical
products were very essential. The group worked out also
routines for in-house produced equipment and initiated
testing of digital blood-pressure instruments. After the
test period most of the tasks dealing with medical
equipment were transferred to the department of
Biomedical Engineering, where they are handled still.
The Safety Commission investigated two cases of
which one had medical equipment involved and that
caused a death of a patient. The equipment involved was
Telequard telemetry system for monitoring heart
patients. The investigation was carried out according to
the routine used by the Swedish Accident Investigation
Board and that has been used several times in
investigations of accidents and incident with aircrafts
involved. Personnel from the Swedish Accident
Investigation Board also supported the investigation.
In the beginning it was suspected that the medical
equipment used, did not work properly. The problems
were thought to be in the signal collection or signal
transmission. The investigation showed that the medical
equipment used worked as intended. The final report
showed that the main cause to the accident was that the
local routines for communication at the clinic did not
work and that the staff had too strong belief on what
was shown on the monitor display.
Discussion
The human factor is the cause of, or contributes to
90% of all accidents [1]. The shortcomings in education
of the staff contribute very often to the accidents, since
the education dose not follow the development of
technology. The demands of the tasks that have to be
carried out should match the competence of the
individual who is going to carry out the particular task;
otherwise there is a risk of an incident or accident. But,
the staff must also realize their own responsibility in
their work, provided that they have the qualification to
do so, i.e. right education and training.
The work of the Safety Commission exposed clearly
some shortcoming in the working routines of the
specific clinic. However, the case investigated was
perhaps an “easy” one and there certainly is more
complex situation that the commission will meet in the
future. The study shows also that the education and
training plays a central role in the safe health care. To
keep the staff’s competence on a necessary level
requires recourses that the management has to provide.
The management has to create a culture of safety that
must be accepted by all levels of the organisation.
Conclusion
In the hospitals the health care staff need support for
tasks that are not directly involved with the every day
patient care, but contribute very strongly to the patient
safety. The know-how for this kind of work can be
created among the hospital’s own staff with initial help
from outside expertise. Together with these experts
well-defined projects can be started up to give necessary
practice to carry on the work in the future.
References
[1] BOGNER, M. S. (1994) 'Medical Devices and Human
Error' in MOULOUA, M. and PARASURAMAN, R.
(Eds) 'Human Performance in Automated Systems:
Current Research and Trends', (Hillsdale, NJ,
Lawrence Erlbaum), pp 6467
IFMBE Proc. 2005;9: 22

Healthcare assessment and clinical engineering
EDUCATIONAL DEVELOPMENT AND CURRICULUM PLANNING WHEN
USING SIMULATORS.
-USEFUL TOOLS FOR TRAINING DOCTORS AND ENGINEERS.
K. Mäkinen*, L. Felländer-Tsai**, P. Ström**, A. Kjellin**, T. Wredmark** and L. Hedman**
* Department of Biomedical Engineering and Center for advanced medical Simulation, Karolinska
University Hospital, Stockholm, Sweden
** Department of Clinical Science, intervention and technology (CLINTEC), Karolinska Institute,
Stockholm, Sweden
www.simulatorcentrum.se
Abstract: The need for technical-safety-education
for clinical staff was identified during the work to
improve patient safety. Also the engineers and
technicians expressed their need to improve the
understanding of needs and clinical use of medical
equipment. A mandatory courseon technical safety
was started for Surgeons working with minimal
invasive surgery at the hospital and a course on
medical engineering applied on Anaesthesia and
surgery was started for medical engineers.
Background
While working with improved patient safety, needs
for technical-safety-education were identified. The
Department of Biomedical Engineering was
commissioned to, in cooperation with the Center for
advanced medical simulation and the different surgical
departments, create an educational programme.
Curriculum
As the curriculum proceeded, engineers and
technicians from the Department of Biomedical
Engineering, also wanted education, but more
appropriate for technicians. The idea was to reach a
profound understanding of clinical use and needs and
have a dialogue with the users of medical equipment.
Implementation
A basic accreditation course was started. This course
is mandatory for Surgeons, working with minimal
invasive surgery at the hospital. One full day is spent
with the Department of Biomedical Engineering and
half a day with simulator training.
The role of instructors at the Centre for advanced
medical Simulation
40 experienced surgeons were tested on the
simulators. The mean value of the performance of the
experience surgeons was calculated. To pass the test, the
participants has to perform this average performance
twice
The role of Engineers from the Department of
Biomedical Engineering
Main items of the course are
• Presentation of Biomedical Engineering and its
role in healthcare
• How to buy medical equipment
• The importance of regular preventive
maintenance
• Responsibility issues concerning medical
equipment
• Basic electronics
• Handling electrical equipment, what is the risk
with?
• Handling gas equipment, what is the risk with?
• Leaking currents, how do they occur?
• How to act when an accident has happened
• When to make a security check
• Diathermia, how to handle and what are the
risks?
• Risks with high frequent current.
• Image guided equipment, practical hands-ontraining.
Examination, 10 questions, 7 correct answers to
pass.
Engineering education
A engineering oriented course was started:
Medical engineering applied on Anaesthesia and
surgery. - For Medical Engineers
Employees at the Department of Biomedical
Engineering at Karolinska University Hospital, who
work with equipment used by anaesthesia doctors and /
or surgeons, are offered a one-day course.
The course is cooperation with the Centre for
advanced medical simulation and aims to teach how
IFMBE Proc. 2005;9: 23

Healthcare assessment and clinical engineering
medical equipment is clinically used, procedures,
different treatments and some anatomy.
General anaesthesia for engineers starts this course
and it teaches the unspoken checklist before anaesthesia
and the decision-loop. Each Engineer is given the
opportunity to act as an anaesthesiologist, assistant and
scrub nurse. Practical induction, intubations and
awakening hands-on-training.
Teacher: Carl-Johan Wallin, M.D, Ph.D DEAA,
senior consultant.
The afternoon has a surgeon approach. Experienced
surgeons use the simulators in order to teach anatomy
and surgical procedures. An opportunity to perform
surgery in a safe environment is given thanks to
simulators. Disadvantages and advantages in using
different kind of surgical equipment are discussed.
Teachers:
Li Tsai, M.D, Ph.D senior consultant and associate
professor. (Orthopaedics)
Lars Henningsohn, M.D, Ph.D (Urology)
Lars Enochsson, M.D, Ph.D (Endoscopes)
IFMBE Proc. 2005;9: 24

Healthcare assessment and clinical engineering
EARLY EXPOSURE TO HAPTIC FEEDBACK ENHANCES PERFORMANCE
IN IMAGE GUIDED SURGICAL SIMULATOR TRAINING
- A PROSPECTIVE RANDOMIZED CROSS OVER STUDY IN SURGICAL RESIDENTS
L. Felländer-Tsai*, K. Mäkinen**, P. Ström*, A. Kjellin*, T. Wredmark* and L. Hedman*
* Department of Clinical Science, intervention and technology (CLINTEC), Karolinska Institute,
Stockholm, Sweden
** Department of Biomedical Engineering and Center for advanced medical Simulation, Karolinska
University Hospital, Stockholm, Sweden
www.simulatorcentrum.se
Abstract: The rate of skill acquisition is often
assumed to be depended on what has been learned in
a similar context. The importance of haptic feedback
in early training phase of skill acquisition in an
image guided surgical simulator was studied in a
randomized cross over study. The results of the
study indicate that haptic feedback is important in
the early training phase of skill acquisition in image
guided simulator training.
Conclusion
Our findings indicate that haptic feedback is
important in the early training phase of skill acquisition
in image guided simulator training.
Background
In the literature on the transfer of skills it is often
assumed, that the rate of skill acquisition depends on
what has been learned in a similar context i.e. models
providing high level of fidelity, including haptic
feedback. The aim of this study was to study the
importance of haptic feedback in early training phase of
skill acquisition in an image guided surgical simulator.
Methods
We used a randomized cross over study design. 38
surgical residents were randomized to start a two hour
simulator training session with either haptic or nonhaptic
training following cross over after 1 hour. The
graphic context was a virtual upper abdomen. Two
diathermy tasks were used. We also used two validated
tests to control for differences in visual-spatial ability:
BasIQ general cognitive ability test and Mental Rotation
Test A (MRT-A).
Results
After two hours of training the group who had
started with haptic feedback performed significantly
better in the two diathermy tasks (unpaired t-test,
p

Healthcare assessment and clinical engineering
EXPOSURE TO MAGNETIC FIELDS OF THERAPEUTIC STAFF
DURING TMS/rTMS TREATMENTS
Ronnie Lundström 1 , Eduardo Figueroa Karlström 2 , Olle Stensson 2 , Kjell Hansson Mild 2,3
1 Ocupational Medicine, Umeå University, SE-901 85 Umeå, Sweden.
2 National Institute of Working Life, Box 7654, SE-907 13 Umeå, Sweden.
3 Department of Natural Sciences, Örebro University, SE-701 82 Örebro, Sweden
Ronnie.Lundstrom@vll.se
Abstract: Transcranial Magnetic Stimulation or
Repetitive Transcranial Magnetic Stimulation –
TMS/rTMS, is currently being used in
treatments of the central nervous system
diseases as for instance depressive states. The
principles of localised magnetic stimulation are
summarised and the risk and level of possible
field exposure of therapeutic staff is described.
Measurements and analysis of the potential
levels of exposure to magnetic fields of the staff
working with TMS/rTMS are presented.
Introduction
The application of localised magnetic fields for
non-invasive stimulus of the cortical region of
conscious patients are carried out using coils array,
among others, in the form of a figure-eight. Pulses
generated can be from single pulses to trains of up
to about 350 pulses per stimuli. When a transient
current flows through the coil system exterior to the
head, a strong time-varying magnetic field is
generated in the brain, which in turn induces eddy
currents that stimulate nerve fibres.
When a figure-eight coil is used, current flow
patterns results in two vortices that merge at the
point beneath the intersection of the figure-eight.
Coils are slightly tilted as to accomplish a
convergent focus of the effect at the cortical region.
The field resulting in the backward direction, i.e.
towards the therapeutic staff, is therefore somewhat
divergent. Although TMS is aimed to expose
patients, the nursing staff using this devices could
also be exposed to magnetic pulses, eventually even
to an extent surpassing the limits of occupational
levels of exposure given in the new EU directive
and the ICNIRP guideline. Motivated by concern
among the nursing staff, measurements were
performed at the equipment in current use at the
Norrland University Hospital of Umeå.
The magnetic field transients currently
generated by TMS equipments can be of the order
of 1 Tesla with a duration of about 0.05 – 0.2 ms.
The resulting time derivative of the field can be of
several tens of kT/s. This transient magnetic field
contribute to the depolarisation of nerve cells in the
brain, allowing stimulation of the cortex of the
brain within a volume of about 5 mm 3 . The
stimulus however has been found to extend to
functionally related sub-cortical regions of the
focused cortical centre. This provides a basis for
using TMS to treat the pathologic neural activity
that may underlie neuropsychiatric illness. The
method is known since the early 50 th [1] and its
value as a therapeutic option was reviewed in 1985
by Baker et al [2].
TMS System
The TMS/rTMS system used in this study was
a MegPro unit with a Magnetic coil transducer
model MC-B70 (Medtronic Synectics AB). The
system is designed to operate at an ambient
temperature of 22 ºC and is capable of generating a
maximum initial dB/dt of 35 kT/s. The active pulse
width is 72 µs, which is approximately ¼ of the
total sine wave period. The system has a mean
repetition rate of 5 pulses per second.
The coils in the transducer MC-B70 are two
partially overlapped coils (~30% overlap)
consisting of 10 turns of wire with a 10 mm inner
and 50 mm of outer radius and a winding height of
~6 mm. Coils have been encapsulated in PVC with
a minimum of 2 mm overall encapsulation and with
the coils symmetrically placed in the polymeric
matrix.
Measurements
The measuring system used consisted of a
measuring coil and a data acquisition unit. The coil
was a calibrated 10 turns electrically shielded
circular coil with 2.5 cm radius and the induced
voltage was registered with a Tektronix TDS 1012
two channels digital store oscilloscope.
Measurements were performed at several of the
most commonly used pulse settings of the
equipment.
The dB/dt data recording measurements were
performed along a vertical axis, perpendicular to
the flat surface of the double coil arrangement of
the Medtronic Dantec MC-B70 Magnetic
IFMBE Proc. 2005;9: 26

Healthcare assessment and clinical engineering
Transducer. The dB/dt scans were taken at 0.1-
0.5 m distances from the transducer with ~10 cm
interval, along the axis of the transducer and away
from the patient, thus to asses the field who could
affect the nursing staff. Measurements were done
even along the axis of one of the coils of the
arrangement of the MC-70B transducer.
The measurements were done at environmental
conditions similar to those corresponding to normal
therapeutic treatments (temperature, humidity, etc).
The transducer however was just resting in a pillow
at the treatment bed.
charge enough to accomplish cortical depolarisation
effects [5]. Pulses are applied at a mean repetition
rate of 5 pulses per second. The effective pulse
width according to the manufacturer of 72 µs could
be confirmed within the accuracy range of the
experimental set up.
Results
The first finding is a confirmation of intensity
decay of the field gradient proportional to 1/r 3 (r:
distance from the coil), see Fig. 1.
Figure 3: The B-field pulses as integrated from
measured dB/dt values, see Fig. 2.
Discussion
Figure 1: Log-log diagram of peak dB/dt vs.
distance giving a calculated slope of 3 ± 7,5%.
The pulse shape of dB/dt signal was recorded
as shown in Fig. 2.
Figure 2: Measured dB/dt at 10 cm, central axis.
The signal was numerically integrated to obtain
the B-field, in Fig. 3. There are rather isolated
stimulus although applied in a sequence of several
pulses, thus, allowing for the stimuli to accumulate
ICNIRP guidelines [4] and the newly published
EU directive [3] provide limitations for the workers
level of exposure to the risks arising from
electromagnetic fields as well as instructions for
precautionary actions such as training and
information. These regulations set limits aimed to
avoid excitation of the central nervous system of
workers, among others, while TMS/rTMS is aimed
just to accomplish high level of local exposure to
produce cortical excitations in patients under
controlled forms. For the pulse trains in use we
found a pulse period T~0.3 ms and about 72 µs
active pulse width, which gives an equivalent
frequency of about 3.5 kHz. For this frequency the
limit value for dB/dt is about 1 T/s (see Fig. 3 in [4]
for recommended levels of exposure to pulsed wave
forms).
We could verify that the limits for the magnetic
field pulses are transgressed at distances of about
0.7 m from the surface of the transducer’s coils
during normal treatment conditions. In spite of this
finding, further studies, especially of different
designs of TMS devices, should be made to bring
deeper insight in the issue of weather the settled
limits in terms of current density also are
transgressed. Until those studies are pursued it is
suggested that the clinical staff should not work
while their body or parts of it, are at distances
closer to 0.7 m from the transducer to avoid risks of
over exposure to magnetic pulses.
Until further studies of the levels of exposure
of the clinical staff at all possible instruments and
settings are performed it is recommended that the
IFMBE Proc. 2005;9: 27

Healthcare assessment and clinical engineering
body or parts of it for the nursing staff should not
be exposed to fields at shorter distances than 0.7 m
from the back side of the coils.
The Dantec MC-B70 equipment could be used
with a mechanical arm holding the transducer in the
right position for the patient. This device should
always be used in order to avoid exposures that can
arise while handholding the probe during treatment
sessions. If similar devices are available for other
TMS/rTMS products, they should be used instead
of hand holding transducers during treatments, thus
allowing the nursing staff to step away from the
zone of high level of exposure (at least 0,7 m apart
from the transducer and its cable).
Conclusions
The staff working with patient treatments with
TMS/rTMS can become exposed to magnetic fields
levels exciding both EU directive and ICNIRP
guidelines, therefore, it is recommended that
procedures are develop to avoid unnecessary
exposure of nursing staff, along side with
instructions as recommended in the referred
documents.
Acknowledgments
The fruitful cooperation between the
participating institutions is deeply acknowledged.
Furthermore, the technical instrumental description
provided by the manufacturer is much appreciated.
References
[1] Penfield W, Jasper H. Epilepsy and the
functional anatomy of the human brain.
Boston, Mass: Little, Brown & Co; 1954.
[2] Barker A T, Jalinous R, Freeston I L. Noninvasive
magnetic stimulation of human motor
cortex, The Lancet. 1985; 1:1106-1107.
[3] European Parliament and the Council Directive
2004/40/EC on the minimum health and safety
requirements regarding the exposure of
workers to the risks arising from physical
agents (electromagnetic fields)(18 th individual
directive within the meaning of article 16.1 of
Directive 89/391/EEG. PE-CONS 3655/04,
SOC 190, CODEC 586, 2004:04.
[4] Guidance on determining compliance of
exposure to pulsed and complex non-sinusoidal
waveforms below 100 kHz with ICNIRP
guidelines, Health Physics. 2003; 84(3): 383-
387.
[5] J E Randall, Elements of Biophysics, 2 nd Ed.,
pg. 264-267 Year Book Medical Publishers,
Inc. (1962).
IFMBE Proc. 2005;9: 28
3

Healthcare assessment and clinical engineering
PRELIMINARY REFERENCE LEVELS FOR DIAGNOSTIC RADIOLOGY IN
ESTONIA
K. Kepler 1 , A. Servomaa 2 , I. Filippova 3
1 Training Centre of Medical Physics and Biomedical Engineering, University of Tartu, Tartu, Estonia
2 University of Oulu, Oulu, Finland
3 University of Tartu, Tartu, Estonia
kalle.kepler@ut.ee
Abstract
The present survey of adult patient doses in x-ray
diagnostics was carried out in 24 hospitals, covering thus
56% of hospitals and about 20% of conventional X-ray
units in Estonia. Entrance surface dose (ESD) of the
patient was assessed by indirect method, using data of
radiation yield of x-ray tubes and examination technique.
Data were collected for 1050 radiographs of adult
patients. Average entrance surface doses varied by a
factor of up to 5,5-12 between hospitals. The average
doses in chest, lumbar spine and pelvis examinations do
not exceed the European diagnostic reference levels
(DRL), but are higher than the dose reference levels in
Nordic countries. In this study the preliminary DRLs,
based on the 3rd quartile of the dose distribution, have
been suggested for chest PA examinations – 0,3 mGy, for
chest LAT – 1 mGy, for lumbar spine AP – 7 mGy, for
lumbar spine LAT – 12 mGy, for pelvis AP – 6 mGy.
The recommended DRLs are preliminary, until the data
collected through more comprehensive dosimetric
surveys will be available. Likewise, it presumes
implementing all requirements of the European Medical
Exposure Directive (97/43/Euratom) concerning patient
dosimetry into the national legislation in Estonia.
Introduction
About one million medical X-ray examinations are
carried out each year in Estonia, corresponding on
average to 0.7 examinations per head of population [1].
In 1997 the essentially voluntary system of patient dose
management, developed by the International Commission
on Radiological Protection, became mandatory in
European Union [2]. By 1999 European Diagnostic
Reference Levels (DRLs) were available in three sets of
European Guidelines on quality criteria for radiographic
examinations in adults or children, and for computed
tomography examinations [3]. For now patient dose
surveys are carried out in most of the Member States, and
quite often it is done in the framework of international
collaboration (e.g. Nordic survey [4]).
In Estonia doses in pediatric radiology were studied from
1999 to 2002 [5]. Unfortunately there is no legal
regulation for establishment of patient dose assessment
system in Estonian radiology departments yet. A new
regulation for use of radiation in medical radiology,
following all requirements of Medical Exposure Directive
(MED) [2], is in preparation in the Ministry of Social
Affairs.
The present study is the first attempt to evaluate patient
doses in the majority of the Estonian health care
institutions equipped with x-ray departments. The aim of
this study was to measure average patient doses and to
estimate preliminary reference doses for most typical
examinations in radiology departments in Estonia.
Methods
The present survey of adult patient doses in X-ray
diagnostics was carried out during the years of 2002-
2003. Dose measurements were carried out in 24
hospitals. The data were collected for 1050 radiographs
of adult patients. The sample of patients was chosen so
that the weight of the patients is between 60-80 kg and
the average of the weight 70 ± 3 kg. Five typical x-ray
examinations were chosen for the study: chest PA, chest
LAT, lumbar spine AP, lumbar spine LAT, pelvis AP.
The exposure factors (focus size, filtration, film-screen,
grid, examination projection, tube potential, MAS, FFD,
field size on film) were recorded along the details of
patient age, gender, height, weight and focus-to-skin
distance (FSD). The surface dose of the patient was
assessed by an indirect method, using the radiation yield
of the x-ray tube and examination techniques. The
radiation output of the X-ray tube has been measured
beforehand at the relevant tube voltage, focal spot and
filtration. Using the measurements of the absorbed dose
to the air, ESD was calculated by applying the inverse
square law to obtain the dose at the FSD and by
multiplying by the mean backscatter factor (1.35).
Results
The doses in chest PA examinations (Figure 1) vary by a
factor of 12 between hospitals. The minimum value of
the ESD in chest PA examinations is 0,05 mGy, the
maximum value is 0,6 mGy.
The average doses in lumbar spine AP examinations vary
by a factor of about 4 between hospitals. The minimum
value of the ESD is 1,8 mGy, the maximum value of ESD
is 10 mGy.
IFMBE Proc. 2005;9: 29

Healthcare assessment and clinical engineering
Results in all different examinations are shown in Table 1
and the third quartile of the dose distribution and
recommended DRLs are given in Table 2.
Table 1: Number of radiographs, ESD minimum and
maximum values, average ESD and European ESD
reference values
in Estonia. Cooperation between the professional
societies and the national authorities responsible for
creation of the national legislation and sustainable system
of quality assurance including patient dose optimisation
in Estonia is necessary.
References
[1] Thomson H., Ruuge M., Rätsep M., Teemusk L.
(Editors) (2003): ‘Estonian health statistics 2000-2002’,
(Ministry of Social Affairs, Tallinn)
[2] Council of European Union (1997): ‘Council
Directive 97/43/EURATOM of 30 June 1997 on health
protection of individuals against the dangers of ionizing
radiation in relation to medical exposure, and repealing
Directive 84/466/EURATOM’, Official Journal of the
European Communities, 40 (L 180), pp. 22-27
[3] European Commission (1999): ‘Radiation Protection
109: Guidance on diagnostic reference levels (DRLs) for
medical exposures’, (Office for Official Publications of
the European Communities, Luxembourg)
Table 2: Average ESD, estimated third quartile ESD and
recommended dose reference levels (DRL) in Estonia
[4] Gron B., Olerud H., Einarsson G., Leitz W.,
Servomaa A., Schoultz B. and Hjardemaal O. (2000): ‘A
Nordic survey of patient doses in diagnostic radiology’,
Eur. Radiol., 10, pp.1988-1992
[5] Kepler K., Lintrop M., Servomaa A., Filippova I.,
Parviainen T. and Eek V.(2003): ‘Radiation dose
measurement of paediatric patients in Estonia’, in:
‘STUK-A195: Radiation Protection in the 2000s –
Theory and Practice’, (Finnish Radiation and Nuclear
Safety Authority, Helsinki), pp. 287-292
Discussion
The average doses vary by a factor of up to 4-12 between
hospitals, which can be explained by differences in the
radiology equipment and techniques. Average ESD for
chest, lumbar spine and pelvis examinations in the most
of the hospitals in Estonia does not exceed the European
diagnostic reference levels [3]. Average ESD in chest PA
examination is of the same value as the European
reference level. But it can be also estimated, that the
estimated doses are higher than in Nordic countries [4].
Conclusions
Quantitative methods for assessment of patient doses
should be implemented in all radiology departments.
The recommended DRLs are preliminary, until the data
collected through more comprehensive dosimetric
surveys are available. It presumes implementing of all
requirements of the MED [2] into the national legislation
IFMBE Proc. 2005;9: 30

Healthcare assessment and clinical engineering
EXPERIENCE FROM TEACHING BIOMEDICAL
ENGINEERING TO HEALTH CARE PERSONAL
M. Folke* and A. Jonsson*
* Department of Computer Science and Electronics, Mälardalen University, Västerås,
Sweden
mia.folke@mdh.se
Abstract: Technical equipments are becoming more
and more common in clinical environment. In Sweden
there is still a lack of undergraduate education in
biomedical engineering adapted to clinical personal. A
course in biomedical engineering aimed for health
care personal at Mälardalen University, Sweden, have
been given in two different structures and evaluated.
The course content has not been changed with the new
structure. We have found that it is of great importance
to motivate the students and to show the use of the
biomedical engineering knowledge in clinical practice
to get them interested in the subject. The subject must
also be taught at an appropriate level.
Introduction
The use of technical support for diagnostics,
monitoring and treatment in health care has been
increased over the past years. Because of the increasing
number of elderly in the future, the use of technical
support will be more useful in elderly care and home
health care.
To get a safe health care environment, for both
patients and personal, it is important that the personal
using the technical equipment have the relevant
knowledge. Except specific apparatus knowledge it is
important to have knowledge about risks and how to
interpret the result so that the plausibility of the result can
be verified and quick detection of artefacts can be made.
Another study have shown that personal, after education
and training, have experienced less anxiety in the
utilisation of medical equipment and less uncertainty in its
use [1].
In Sweden there is lack of relevant undergraduate
education in biomedical engineering for health care
personal.
Since the autumn of 2003 a course in biomedical
engineering for health care personal is given at
Mälardalen University, Sweden [2]. The course has been
given in two different structures to evaluate and find a
better way to teach biomedical engineering to health care
personal.
Materials and Methods
The course in biomedical engineering (7.5 ECTS) at
Mälardalen University, Sweden is optional for students
during their three years of nursing studies, but is also
accessible for practicing health care personal. The
students attending the course have different experience
and technical knowledge. When designing the course we
assume that the students do not have any higher level of
technical education than that given in compulsory school.
The main emphasis of the course is the principles of
the techniques behind, for non-specialised nurse, relevant
equipment and not the specific apparatus knowledge. The
course also includes regulations about medical equipment
and safety aspects of electricity, gas and radiation. The
course includes lectures, laboratory classes and a written
examination.
The first time the course was held the structure of the
teaching packages were divided in diagnostics and
treatment. Since this was a new kind of student group for
us to teach, we misjudged their previous medical
knowledge and thought that all of them already knew the
medical reason for using different equipments.
The following times the course has been given, the
structure of the subjects has been related to the different
technical principles i.e. all equipments with optical
principles. When changing the structure of the course we
also expanded the number of laboratory classes from three
to four, so that more of the content of the course was
included in the laboratory classes. We also introduced
three minor elective written examination dispersed over
the course. A pass on such an examination will give
bonus for the final examination, but the reason is
primarily to encourage the students to work during the
whole course. No difference in course content has been
made between the two different structures, but after the
redesign, the education starts at a more basic level even in
medical knowledge.
Results
The first time the course was held 22 students attended
and 64 percent of them passed the final examination,
table 1. Unfortunately only 50 percent answered that they
would recommend the course to another student, table 1.
This lead to the conclusion that 50 percent of the students
were not satisfied, table 1. One can also see that most of
IFMBE Proc. 2005;9: 31

Healthcare assessment and clinical engineering
the unsatisfied students were the same students that did
not pass the final examination.
A bad rumour about the course did result in fewer
students for the next occasion, table 1. Redesign of the
course and encouraging the students to attend the lectures
and start studying at the beginning of the course gave a
positive result and also a positive rumour after the second
time the course was held.
After the second course occasion we can see a positive
trend in number of applicants and response from the
students who have finished the course, table 1. In addition
to, an increased total number of applicants, the number of
practicing health care personals attending the course, have
increased.
Only few students have had any use of the bonus from
the minor written examinations, since they passed the
final written examination without the bonus. However,
most students say that the minor written examinations are
very important during their studies and encourage them to
start studying earlier in the course.
The students also mean that the laboratory lessons are
important, since the practical use of the theoretical
knowledge makes it easier to understand.
Table 1: Number of student and their success.
Course
occasion
1
2
Number of
students
Passed final
examination
(%)
22 64 50
5 100 100
3 13 92 100
4
36
preliminary
Satisfied
Although the lectures starts at a lower level now the
students reach at least the same level at the end of the
course as when the course first was given. The change in
structure of the course has not lead to an increased burden
of work for the teachers.
Discussion
The need of relevant education in biomedical
engineering for health care personal is obvious. However,
some nurse students still believe that they will have a
work which will not require any technical knowledge.
The nurse students’ interest and understanding for the
need of technical knowledge is greatly dependent on
where they have done their practical training. In general,
the students that have done their practical training at i.e.
the emergency department or the intensive care realise the
need of technical knowledge and their interest in
biomedical engineering increase. Further, they are not to
the same extent afraid to and do not feel so insecure with
technical equipment.
Because of the differences in the students’ background
and interest it is important to start on a low level. It is also
important to make all students interested in the subject,
since the interest is necessary in learning [3]. The subject
must therefore be taught from a clinical point of view, i.e.
stress the medical reason and how to use the equipments
in diagnostics, monitoring and treatment.
The importance of laboratory lessons for learning is
confirmed by the fact that the combination of theory and
practice has been proved to strengthen the knowledge [4].
A positive rumour about the course has made the
students more interested in the subject and since the sharp
drop of students after the first course occasion, the
number of applicant have increased.
The minor written examinations is of large importance
for the students to realise what we expect from them
during the final written examination and helps the
students to focus on the more important content in the
course.
Conclusions
To get the nurse students and health care personal
more interested in biomedical engineering, the teacher has
to teach the subject at a for them appropriate level. The
subject has to be made less dramatic since many of these
students have a preconceived notion that engineering and
physics are very difficult to learn. It is also important to
relate all theories to clinical practice.
References
[1] PERSON, J., EKBERG, K. and LINDÉN, M. (1993):
'Work Organisation, Work Environment and the Use
of Medical Equipment: A Survey Study of the Impact
on Quality and Satety', Med. & Biol. Eng. & Comput.,
31, pp. HTA20-HTA24
[2] MÄLARDALEN UNIVERSITY, 'Course Syllabus':
http://www.mdh.se/studieinformation/VisaKursplan?k
urskod=LA1730&sprak=en
[3] SJØBERG, S. (2000): 'Naturvetenskap
som allmänbildning', Studentlitteratur, pp. 348-377
[4] RYEGÅRD, Å. (2004): 'INTERAKTION MELLAN
TEORETISKT OCH PRAKTISKT LÄRANDE - En
Studie i Elektronikundervisning på Högskolenivå'.
Department of Electronics. Västerås, Mälardalen
University, Mälardalen University Press, ISBN: 91-
88834-72-7.
IFMBE Proc. 2005;9: 32

Medical informatics
IMPLEMENTATION OF GLUCOSE-INSULIN CONTROL IN H 2 /H ∞ SPACE
USING MATHEMATICA
Levente Kovács*, Béla Paláncz**, Zsuzsa Almássy*** and Zoltán Benyó*
* Dept. of Control Engineering and Information Techn., Budapest Univ. of Techn. and Economics,
H-1117 Budapest, Magyar Tudósok krt. 2, Hungary
** Dept. of Photogrammetry, Budapest Univ.of Techn. and Economics, Budapest, Hungary
*** Heim Pál Children Hospital, Pediatric Department, Budapest, Hungary
lkovacs@seeger.iit.bme.hu, klevi77@yahoo.com
Abstract: In this case study, an optimal control in
H 2 /H ∞ space is presented for the glucose-insulin
system in case of diabetic patients under intensive
care. The analysis is based on a modified twocompartment
Bergman model. To design the optimal
controller, the disturbance rejection LQ method has
been applied. The critical value of the scaling
parameter γ crit is determined by a symbolic solution
of the modified Ricatti equation. The numeric
evaluation of the symbolic computation with γ > γ crit
leads to two different solutions for the gain matrix.
The numerical results are in close match with that of
the µ-Toolbox of MATLAB, however this latest gives
only one of the two solutions.
Introduction
From an engineering point of view, treatment of
diabetes mellitus can be represented by an outer control
loop to replace the partially or totally failing bloodglucose
control system of the human body. Nowadays,
maintaining the glucose level in a diabetic patient under
intensive care is an actively researched topic.
Most of the glucose-insulin models were realized for
"artificial pancreas" function, in conditions where the
patient’s blood glucose level is monitored and insulin
injection is performed continuously during surgery. We
have chosen a modified two-compartment model [1],
considering it as the best appropriate model, [2]. The
main advantage of the selected model in comparison
with the others is its on-line adaptive nature, based on
strong theoretical foundations, but also describing the
physiological system appropriately.
Improving the control strategy, an optimal glucoseinsulin
control in H 2 /H ∞ space has been designed and
simulation of food (sugar) intake was carried out. The
symbolic and numerical computations were determined
with Mathematica 5, as well as with MATLAB 6.5.
Materials and Methods
To simulate the insulin-glucose interaction in human
body the following two-compartment model was
employed, [1]:
X&
1(t)
= p1X1(t)
+ p 2h(t)
(1)
X&
(t) = (p − X (t))X (t) + i(t) + p
2
3
1
2
4
h(t) and i(t) represent the input variables: the
exogenous insulin and glucose. X 1 (t) and X 2 (t) (being
the state as well as the outputs variables) stand for the
concentration of glucose in the plasma and insulin
remote from plasma. Parameters p i (i=1…4) are: p 1 =-
0.02115, p 2 =0.09255, p 3 =-0.01418, p 4 =0.07794, [3].
To design the H 2 /H ∞ control for this system, the
nonlinear system has been linearized in the vicinity of
steady state, namely at (X 1 0, X 2 0, h0, i0). Using a
classical LQ method, an optimal control is obtained by
minimizing the following quadratic cost functional, [4]:
∞
1 T
T
J (u) = ∫[y
(t)Qy(t) + u (t)Ru(t)]dt (2)
2
0
The disturbance rejection LQ method, represents a
generalization of the classical LQ method and is based
on the minimax criteria in H 2 /H ∞ space. However, now
the input variable u(t) is divided into two parts, control
input u (t)
and disturbance d(t). Therefore, the
quadratic cost functional is, [5]:
∞
1 T
T
2 T
J (u,d) = ∫[y
(t)y(t) + u (t)u(t) − γ d (t)d(t)]dt (3)
2
0
It was demonstrated that the solution of this
differential-game exists (“worst-case” design) [5], it is
unique and satisfies the saddle point condition:
*
J(u , d) ≤ J(u, d) ≤ J(u, d )
(4)
*
is the worst-
where u is the optimal control and
case disturbance.
Results
*
*
d
The H 2 /H ∞ controls were designed both in
MATLAB and Mathematica, [4]. To test our controllers
via simulation, we considered the situation of food
intake (disturbance) simulating the sugar absorption in
the body. According to our clinical experiments we used
the following function (for a duration of 20 minutes):
2
( t−10)
−
h(t) = 0.05 ⋅ e 45
(5)
Using the classical LQ and disturbance rejection LQ
methods, results are presented in Figure 1 and Figure 2.
Employing symbolic computation for H 2 /H ∞ control,
in order to solve the modified Ricatti equation, it was
possible to determine the critical value of the scaling
parameter γ as function of the model parameters:
IFMBE Proc. 2005;9: 33

Medical informatics
Figure 1: The performances of LQ (continuous) and
disturbance rejection LQ (dashed) control considering
insulin concentrations.
Figure 3: Bifurcation of the gain matrix element K 11 (γ)
from the singular point, γ = γ crit .
However, these controls are not implemented yet,
but after the necessary further verifications they could
provide a useful help in control of blood glucose level,
and in the optimization process of diabetic
administration.
Acknowledgements
Figure 2: The performances of LQ (continuous) and
disturbance rejection LQ (dashed) control considering
glucose concentrations.
2 2
p 2 p 3 + p 4
crit =
p1p
3
γ (6)
It was found that for γ > γ crit there are more than one
positive definite solutions for the gain matrix (Table 1).
The values are in good agreement with the result of µ-
Toolbox of MATLAB, however, MATLAB gives only
one of the solutions (Figure 3).
Conclusion
Nowadays scientists are trying to obtain on-line
adaptive control laws using compartment models, [6],
but results are still in an initial phase. Until now, the
existing adaptive control models worked in a flip-flop
manner, selecting the control command between two
values, but they have neither physiological nor control
theoretical background. Using the presented control
methods, a continuous control input can be achieved.
Moreover, the model is not complicated (it uses only
two differential equations), so it could give a great
advantage in case of practical implementation.
According to this case-study, the disturbance
rejection LQ control proved to be superior to the
classical LQ and its robustness can moderate the
eventual modeling errors of this simplified model.
Table 1. Bifurcation of the gain matrix element K 11 (γ).
This research has been supported by Hungarian
National Research Fund, Grants No. OTKA T029830,
T042990 and by Hungarian Ministry of Education Grant
No. FKFP 200/2001 and Pro Progressio Foundation.
References
[1] BERGMAN B.N, IDER Y.Z., BOWDEN C.R., and
COBELLI C. (1979): ‘Quantitive estimation of
insulin sensitivity’. Am. Journal of Physiology, 236,
pp. 667-677.
[2] JUHÁSZ CS (1993), ‘Medical Application of
Adaptive control supporting insulin therapy in case
of diabetes mellitus’, (Ph.D. dissertation, Budapest).
[3] KOVÁCS L., BENYÓ B., PALÁNCZ B. and
BENYÓ Z. (2004): ‘A Fully Symbolic Design and
Modelling of Nonlinear Glucose Control with
Control System Professional Suite (CSPS) of
Mathematica’, Acta Physiologica Hungarica, 91
(2), pp. 149-158.
[4] KOVÁCS L., PALÁNCZ B., ALMÁSSY Zs.,
BENYÓ Z. (2004): ‘Optimal Glucose-Insulin
Control in H 2 space’, Proc. of 26th Ann. Int. Conf.
of IEEE Eng. in Biomedicine Soc., San Francisco,
CA, p. 762-765.
[5] ZHOU K. (1996): ‘Robust and optimal control,
(Prentice Hall, New Jersey), pp. 376–412.
[6] BAURA G.D. (2002): ‘System Theory and
Practical Applications of Biomedical Signals’,
(IEEE in Biomedical Eng, New York), pp. 340-383.
γ / γ crit 1.1 1.2 1.3 1.4 1.5 2 3 5
a
K 11
b
K 11
0.2284 0.1021 0.0824 0.0685 0.0580 0.0305 0.0130 0.0046
0.3237 0.3548 0.3745 0.3884 0.3988 0.4264 0.4439 0.4524
IFMBE Proc. 2005;9: 34

Medical informatics
A PILOT STUDY OF DEVELOPING THE LABORATORY INFORMATION
SYSTEM OF A COMMERCIAL LABORATORY
S. Tang 1 , J. Lin 2 , P. Li 2 , Y. Huang 3 , S. Young 2
1 Department of Biomedical Engineering, Yuanpei Institute of Science and Technology, Hsin-Chu,
Taiwan
2 Institute of Biomedical Engineering, National Yang-Ming University, Taipei, Taiwan
3 Department of Nursing, Chang Gung Institute of Technology, Tao-Yuan, Taiwan
sttang@ms1.hinet.net
Abstract
In Taiwan, the numbers of commercial medical
laboratories play a very important role in the healthcare
environment. These laboratories require the information
technology to facilitate the huge data processing daily.
But they are difficult to develop their own system for
their scale. Consequently, the modern IT is not beneficial
to the laboratories, and results in problems in efficiency
and quality. This study adopted the Rapid Prototyping
(RP) method for solving the problem. The method
successfully cooperates with the users to develop a
feasible system.
Figure 1 illustrates the general workflow of the RP
method, which involves intensive interaction that closely
links system developers and users throughout the life
cycle of system development. A prototype system is
generally created in the primary development phase, and
is used as a platform for cooperation between the
developers and users. During the prototype system is
continuously tested and refined. This gradually integrates
the users’ requirements and system design, and the
prototype system evolves toward the final operational
system.
Introduction
Nowadays, the information system is becoming a
survival key to the most commercial organizations [1].
This situation is same to the healthcare institutes.
Primarily, The healthcare institute fully committed IT
(Information Technology) company to develop the
information system. Because of the special
characteristics, the develop strategy of a healthcare
information system is quite differ to the general
commercial firm. And then the immedicable defects
usually happen to the system [2-4]. As a result, the
modern healthcare institute becomes to develop the IT
department by themselves.
In Taiwan, the numbers of medial- or small-scale
commercial medical laboratories play a very important
role in the healthcare environment. These laboratories
also require the IT to facilitate the huge daily data
processing. But they are difficult to develop their own
system for their scale. Consequently, the modern IT is not
beneficial to the laboratories, and results in problems in
efficiency and quality [5-7].
This study adopted the Rapid Prototyping (RP) method
[8] for solving the abovementioned problem.
Methods
Contrast to conventional Waterfall method, RP is an
alternative method for system development, which is
characterized by producing a working prototype during
the primary progress of the development life cycle.
Figure 1: The workflow of the RP method
Results
The initial system prototype was as the basis for system
development and as the platform for collaboration of
developers and users. A series of cooperation refinements
were followed, as shown in figure 2.
The interested data were identified firstly. We referred
the current working processes, and then induced a
workflow to integrate the progression of full data
acquisition. This flow acts as a checklist to ensure data
completeness.
The acquisition, storage, and manipulation of actual
data were the next result. The system architecture of the
system was derived from the identified data. That is a
IFMBE Proc. 2005;9: 35

Medical informatics
multitiered architecture that separates the GUI
applications, data service, data-preparation, and back-end
data sources into distinct layers. The data-preparation is
the initial process and “cleansing” system for data that
are moving toward the data server (here “cleansing”
involves correction, filtering, detection, and reporting).
Finally the GUI was implemented and used as the main
platform for analyzing end-user response and
acceptability. The user interface comprised an
instrument-like panel. The arrangement is similar to the
control panel of many medical instruments.
Figure 2: The evolution progress and the results
Discussion
Many medical information systems fail due to
opposition from caregivers, and so the successful
introduction of a new information system into healthcare
institutes requires an effective blend of technical and
organizational skills. The proposed RP strategy is
focusing on intensive user involvement, which can easy
achieve user support. The RP strategy has also proven
beneficial to other researchers. Dugas et al. adopted an
RP strategy to satisfy users’ expectations and
successfully developed a decision support system for
hepatic surgery [9]. Kinzie et al. deployed an RP strategy
to assess needs, to develop user interface, and finally to
create a user-centered model for a personal health-history
web site [10]. The technological key issues of this project
are to solve the problems include streamline workflow,
historical data migration, and reporting system
construction. The commercial laboratory would possess
the ability to handle its information system development
after this project. Then it is effectively to decrease the
manpower, increase efficiency, quality assurance.
[1] HIMSS (2000): ‘Final report: Trends in Healthcare
Information and Technology’, the eleventh annual
HIMSS leadership survey sponsored by IBM, 2000.
[2] Young S. T., Chang J. S. (1997): ‘Implementation of
a patient-centred and physician-oriented healthcare
information system’, Med. Inform., 22(3), pp. 207–214.
[3] Klein C. S. (2003): ‘LIMS user acceptance testing’,
Qual. Assur., 10(2), pp. 91–106.
[4] Stead W. W., Miller R. A., Musen M. A., and Hersh
W. R. (2000): ‘Integration and beyond: linking
information from disparate sources and into workflow’, J.
Am. Med. Inform. Assoc., 7(2), pp. 135–145.
[5] Sanchez-Villeda H., Schroeder S., Polacco M., et al.
(2003): ‘Development of an integrated laboratory
information management system for the maize mapping
project’, Bioinformatics, 19(16), pp. 2022–2030.
[6] Chae Y. M., Lim H. S., Lee J. H., et al. (2001):
‘Development of an intelligent laboratory information
system for community health promotion center’,
Medinfo., 10(Pt. 1), pp. 425–428.
[7] Min W. K., Lee W., Park H. (2002): ‘The
development of systematic quality control method using
laboratory information system and unity program’,
Southeast Asian J. Trop. Med. Public Health, 33(Suppl.
2), pp. 74–78.
[8] American Society for Testing and Materials (1996):
‘E1340-96 Standard guide for rapid prototyping of
computerized systems’, West Conshohocken, PA, pp. 1–
11.
[9] Dugas M, Schauer R, Volk A, Rau H. (2002):
‘Interactive decision support in hepatic surgery’, BMC
Med Inform Decis Mak, 2(1), p5.
[10] Kinzie M. B., Cohn W. F., Julian M. F., Knaus W.
A. (2002): ‘A user-centered model for web site design:
needs assessment, user interface design, and rapid
prototyping’, J Am Med Inform Assoc, 9(4), pp. 320–
330.
Acknowledgements
We gratefully acknowledge the support of the National
Science Council, Taiwan, under grant NSC 93-2622-E-
264-003-CC3.
Conclusions
This informatics project has applied an RP strategy to
the development of a commercial laboratory information
system. The framework is a solution to other healthcare
information system, which can be modified and expanded
to provide new services or to support new application
domains.
References
IFMBE Proc. 2005;9: 36

Telemedicine and patient data management
BIOMEDICAL SENSOR NETWORK ARCHITECTURE BASED ON
TCP/IP
C. López 1 , J. C. Tejero-Calado 2 , A. Bernal 3 , M. A. López 1 , G. Quesada 4 , J. Lorca 5
1 R&D Department, Andalusian ICT Centre (CITIC), Málaga, Spain
2 Electronic Department, University of Málaga, Málaga, Spain
3 R&D Department, IMABIS, Málaga, Spain
4 Critic Care Unit, Carlos Haya Hospital Complex, Málaga, Spain
5 Directorship, revistaesalud.com journal, Málaga, Spain
mclopez@citic.es
Abstract: Most of the patients who are in hospitals
and, increasingly, patients controlled remotely from
their homes, at-home monitoring, are continuously
monitored in order to control their evolution. The
medical devices used up to now, force the sanitary
staff to go to the patients’ room to control the
biosignals that are being monitored, although, in
many cases, patients are in perfect conditions. If
patient is at home, it is he or she who has to go to the
hospital to take the record of the monitored signal.
New wireless technologies, such as BlueTooth and
WLAN, make possible the deployment of systems
that allow the display and storage of those signals in
any place where the hospital intranet is accessible. In
that way, unnecessary displacements are avoided.
Introduction
There are signals which must be monitored
continuously or periodically in patients. The most
important ones are: pulse-oximetry, non-invasive blood
pressure, electrocardiography... Traditionally, patients
have been connected to large devices, what reduces
their mobility and wellbeing during the hospital stay. It
also forces the physicians to go to the patient’s room to
check the evolution of these signals.
New wireless technologies, such as IEEE
802.11[1,2] or BlueTooth[3], make possible the design
of high level integration devices for bioengineering
applications, as well as, the development of new
systems which make easier the at-home patients’
control. Thus, biomedical signals can be sent to other
devices (screen, PDA, PC...) or processing centres,
without restricting the patients’ mobility when they are
at hospital; or increasing the patients’ wellbeing when
they are at home.
In the last few decades, patients supervision
systems[4-5] have been developed, but they have not
been widely diffused and implanted. The principal
reason of this fail is that they were designed to solve
only a partial area within monitoring problem or that the
used technologies were not scalable.
The designed system is based on TCP/IP, a really
used and extended architecture that allows the
development of a global, scalable and innovative
solution to patients’ supervision. The architecture is
based on two TCP/IP servers which manage the
acquisition devices, the information they transmit and
handle the representation devices (PC, PDA...). The
acquisition devices (AD) that have been used are the
ones presented in the next section. Their main
characteristics are the use of wireless technologies to
transmit the data and that they are treated as any other
IP node of the hospital intranet.
In the following sections the system structure is
presented, as well as the interaction between the
devices. To conclude, a project discussion is exposed.
Methods
The system is based on a star topology where both
servers are placed in the centre and coordinate the
communication between the devices (Figure 1). Thus,
every device, both acquisition and representation, will
have only one open connection once the server accepts
the connection request. This distribution makes the
device communication software less complex and
makes the servers have the biggest part of the
complexity. These handle a huge number of
connections, meanwhile the ADs only need an open
connection whatever the number of representation
devices that demand their data.
In the following paragraphs the functions of the
elements that form part of the structure are detailed,
including their interaction.
Monitoring Server- Biosignal Acquisition Devices:
The Monitoring Server (MS) handles the acquisition
device connections. Its principal functions are:
management of the biosignal acquisition devices and
reception and monitoring information management.
The ADs are high level integration wireless devices
for biomedical signal acquiring. These devices will
implement the following functionalities: biosignal
sensing, signal conditioning, A/D conversion and signal
processing in order to reduce the noise level and to
IFMBE Proc. 2005;9: 37

Telemedicine and patient data management
stabilize the signal. Once the signal has been processed,
it is sent to the server through the wireless capabilities
integrated into the device .These devices act as network
clients and send connection requests to the MS once
they are active.
Rep.
Device
Data
Connection request and
ADs information
Representation
Server
Chosen AD
Data Monitoring
Server
AD List
Acq.
Device
Rep.
Device
Monitoring
Server
Representation
Server
Figure 1: System topology.
Acq.
Device
Rep.
Device
Figure 3: Server-Representation Devices Interaction.
The MS sends to the RS the AD list every time a
new AD is connected to the network or when one of
them is disconnected from it. When RS receive it, it is
diffused to every RD.
The MS is always waiting for AD connection
requests. When a connection request is received, the
device identification and its network registration are
made. The device identifies itself indicating the device
group it forms part: electrocardiograph, pulseoximetry...
This kind of identification is not singleminded
because several devices from the same group
can be connected. To avoid this ambiguity each
connection is identified with both the device group and
the device IP address.
Once the identification has finished, the MS has to
update the AD list. This list contains the identification
of all the ADs which have an open connection with the
MS and will be used by the Representation Server to
inform its clients about the AD connected to the
network.
From this moment and while the AD is active, it
starts the transmission of the biosignal it is acquiring.
When the AD is inactive, the MS eliminate it from the
AD list (Figure 2).
Acq.
Device
Data
Connection request
and identification
Monitoring
Server
Device list
update
Figure 2: Server and Acquisition Devices Interaction.
Representation Server- Representation Devices: The
Representation Server (RS) manages the connection
requests produced by the Representation Devices (RD)
and sends to them the data they request.
In the same way as the MS does, the RS is always
waiting for connection requests made by the RDs. Once
the connection is established, the RS sends to the RD
the AD list. This information is showed to the user and
this one chooses the AD to display. When the choice
has been realized, the AD identification is sent to the RS
and it communicates with the MS in order to receive the
data from the selected AD. When the data is in the RS,
it sends it to the RD that had already demanded it.
At any moment the user can change the AD to view
and the process to receive the new data is the same as
the one explained above (Figure 3).
Discussion
The main emphasis of this project has been the
development of a network architecture where every
device, both AD and RD, is treated as an IP node, and
where the interaction between devices is hold without
sanitary staff intervention.
The IP identity of each node allows the reception
and display of the acquired information of any device
connected to the network, both inside the hospital and
from any other place with access to the hospital intranet.
Thus, physicist can receive and display the signals
wherever the patients are, improving the wellbeing and
making easier the reintegration of patients to their
quotidian life when at home monitoring is acceptable. If
hospital monitoring is required, sanitary staff is not
needed to go to the patients’ room each time a control
must be made.
This project has been supported by the Andalusian Health Service
(SAS204/03), Andalusian ICT Centre (CITIC) and The Mediterranean
Institute for Advance in Biotechnology and Sanitary Investigation
(IMABIS).
References
[1] GAST M. S., LOUKIDES M. (2002): ‘802.11
Wireless Networks: The Definitive Guide’,
(O'Reilly & Associates, Sebastopol)
[2] ROAHAN P., LEARY J. (2003): ‘802.11 Wireless
Local-Area Network Fundamentals’, (Cisco Press,
Indianapolis)
[3] BRAY J., SENESE B. (2001): ‘Bluetooth
Application Developer´s Guide’, (Syngress
Publishing, Rockland)
[4] BAI J., HU B., ZHANG Y., YE D. (1997): ‘A
Communication Server for Telemedicine
Applications’. IEEE Transactions on Information
Technology, Vol. 1, No. 3, pp. 295-209
[5] Priddy B.; Jovanov E. (2002): ‘Wireless distributed
data acquisition system’, Proc. of the Thirty-Fourth
Southeastern Symposium on System Theory.
Huntsville, Alabama, USA, 2002, p.463 – 466
IFMBE Proc. 2005;9: 38

Telemedicine and patient data management
GAP ANALYSIS BUSINESS IMPACT OF MODEL DRIVEN ARCHITECTURE
(MDA) ON TELEMEDICINE HEALTHCARE SOLUTION
R. Nabiev 1 , N. Tariq 2 , S. Jonsson 1 , H. Teriö 1 , D. Andersson 3
1 Biomedical Engineering Department, Karolinska University Hospital, Stockholm, Sweden
2 Department of Computer and System Science, Royal Institute of Technology, Stockholm, Sweden
3 Department of Innovation and Medical Informatics, Karolinska University Hospital, Stockholm,
Sweden
rustam.nabiev@karolinska.se, emis-nat@dsv.su.se, sven.jonsson@karolinska.se,
heikki.terio@karolinska.se, daniel.andersson@karolinska.se, naveed_eltaf@yahoo.com,
naveed_eltaf@yahoo.com
Abstract
The need for improvement of existing business
processes within the organization leads often to the
construction of a new IT system. Such a system needs
to be analyzed with respect to its feasibility, business
impact, and economical impact on the organization. A
vertical gap analysis provides such a comparison
focusing on the differences between the two systems.
When the functionality of the system has been
defined, several implementation approaches need to
be considered and compared to each other in order to
find the best one. A horizontal gap analysis can be
used in this aim. The application of both vertical and
horizontal gap analysis within the framework of
Model Driven Architecture (MDA) is demonstrated
within the Telemed HC System (TMHC) in the
Biomedical Engineering Department (MTA) at
Karolinska University Hospital, Stockholm, Sweden.
TMHC is telemedicine solution for remote monitoring
of biomedical equipments located in patients’ homes.
Introduction
New systems are usually built because a need has been
identified for improving existing business procedures.
When such needs are identified, automation of certain
procedures is a common approach to improvement;
however, such a new system should be analyzed with
respect to feasibility, business and economic justification,
among others.
In order to be able to compare any future solution to the
solution in place at the time being, a benchmark needs to
be established. This benchmark needs to describe the
business procedures and workflows in place at the time
being. Then, the envisioned solution needs to be sketched
out and defined in terms of business procedures and
workflows. Having defined these two benchmarks, a
Vertical Gap Analysis (VGA) can be established. A VGA
investigates if it is more profitable to do nothing (keep
the status) or to do something (build a new system). If the
VGA shows that building a new system is worth the
investment, different possible solutions should be
compared to each other in order to find the one that is
best suited. In order to be able to do so, a Platform
Independent Model (PIM) describing the system to be
built in a platform-independent manner needs to be
established. This PIM will be mapped to several Platform
Specific Models (PSM) describing different possible
technical implementations. These PSM can then be
compared to each other investigating their costs,
strengths, weaknesses, problems, etc. Such a comparison
is called a Horizontal Gap Analysis (HGA). The concepts
outlined above have been applied in a project called
‘Telemedicine Home Care’, shortly ‘TeleMed HC
(TMHC)’, in the Biomedical Engineering Department
(MTA) at Karolinska University Hospital, Stockholm,
Sweden. TMHC is an Enterprise Application Integration
(EAI) solution for remote monitoring of biomedical
equipments located at patients’ homes.
Methods
Within the framework of MDA [1] gap analyses can be
applied at different stages of the development process.
The use of gap analyses is presented in this research for
making architectural decisions regarding the usage of
Message Oriented Middleware (MOM) [2] versus Virtual
Private Network (VPN) [3] architecture.
Figure 1 and 2 visualizes the overall conceptual
architecture of PIM with regards to TMHC using
Message Queuing (MQ) service MOM architecture (see
Figure 1) versus the VPN architecture (see Figure 2).
Please, see discussion section for details of both
architectures.
Results
A comparison of MOM architecture using MQ services
verses VPN architecture gave following results. For
details on these results please refer to [4]
- From a business point of view, building an EAI using
MQ is preferable, since it allows integration with existing
and new systems.
- From a technical point of view, building a system using
MQ is clearly preferable to a VPN solution, since it is
IFMBE Proc. 2005;9: 39

Telemedicine and patient data management
easier to manage and most critical aspects that is securing
is built in or provided by MQ services.
- MQ is easier to build a scalable, secure, flexible and
less complex to maintain.
- MQ or MOM needed some extra training on how to use
MQ services such as WebSphereMQ, JBossMQ, etc
when compare to VPN solution.
- From economical point of view although MOM solution
can be higher in cost than VPN solution such as building
a MQ system would cost about SEK 18,000 more than
building a VPN system. But, when considering the above
mentioned results such integration capability, scalability,
security, etc this cost would be less when the developed
solution such as TMHC solution will be running is
production environment and millions of transactions will
be handled.
Discussion
Following is the short description of both MOM and
VPN architecture being compared for this research.
In Figure 1, MQ architecture is shown for TMHC
solution. A message is sent to MQ server using some MQ
service such as WebSphereMQ, JBossMQ, etc from
patient’s premises where medical device is connected to
one client side application named TMHC patient
application. This message is received at MQ server which
forwards this message to TMHC platform or server
application. This server application is further connected
to repository and also supports existing application at
Karolinska for viewing and managing the different
information such as medical device information, patient
information, etc.
Figure 2: TMHC VPN Architecture
Conclusions
MDA provides a sound approach to system development
by providing models at different levels, supporting a
decision taking process at several stages and points in
time. Vertical and horizontal gap analyses support this
approach by providing the required information to base a
decision on different project reviews. These may suggest
adopting a different solution than the conclusion after
consideration of only the short-term goals would. A
comparison of MOM and VPN architecture is presented
in this research for measuring the gap analysis business
impact of MDA on TMHC solution. One of the main
advantages gained from the use of MDA in TMHC
solution is that it made it easier to act proactive; it
allowed identifying the different puzzle bits the system
consisted of and estimating their complexity.
References
(Electronically Publications)
[1] MDA Guide, Internet site address:
http://www.omg.org/docs/omg/03-06-01.pdf
[2] Minnesota State University Mankato, Internet site
address:
http://krypton.mnsu.edu/~spiral/eta/glossary/indxGlossO
Oxml.html
[3] Chris De Herrera’s Windows CE web site, Internet
site address:
http://www.cewindows.net/pocketpc/glossary.htm
(Karolinska University Hospital Internal Document)
Figure 1: TMHC MOM Architecture
Using MQ
[4] Andress N., Secure Transaction of Patient Data –
Comparison of Different Available Technologies and
Solutions. Karolinska University Hospital, 2003
In Figure 2, VPN architecture is shown for TMHC
solution. A VPN connection is established between
patient’s computer which is connected to medical device
and server. This server is further transforming the
information or data received from patient’s premises to
some database such as Oracle.
IFMBE Proc. 2005;9: 40

Telemedicine and patient data management
WEB-BASED SUPPORT TO ENHANCE SELF-MANAGED DIABETES CARE
M. Psaros 1 , A. Ekbom-Schnell 2 , A. Vidmark 2 , S. Koch 3
1 Medical Informatics and Engineering, Uppsala University Hospital, Uppsala, Sweden
2 Internal Medicine, Endocrinology and Diabetology, Uppsala University Hospital, Uppsala, Sweden
3 Medical Science, Biomedical Informatics and Engineering, Uppsala University, Uppsala, Sweden
magdalena.psaros@akademiska.se
Abstract
The wide spread use of Internet technology in our
society can be used to enhance patient information
and patient engagement in order to improve
treatment results. Patients with chronic diseases, such
as diabetes type 1, usually do have good knowledge
about their diseases. However, to enhance their
involvement in their health care process, they have to
have access to certain health care and medical
information and be able to communicate with their
nurse and physician in an adequate way. Based on a
user needs analysis, we implemented a personalised
information and communication space for patients
and health care personnel.
Introduction
The Swedish Board of Health and Welfare’s national
guidelines for diabetes mellitus [1] confirm that the
individual health care covenant between patient and
caregivers is not established in an adequate way. This
complicates the process of formulating clear treatment
goals and teaching patients to master and control their
conditions.
In the health care sector, stress is a common factor. This
forces caregivers to prioritise some of their day-to-day
work. As a result, patients do not have the opportunity to
discuss their treatment which leads to less qualitative
self-care. It is important, that the health care personnel
have time to maintain a continuous contact with the
patient. [2]
Today, there are few web pages that support the
communication between the patients with diabetes.
Through these forums patients can discuss different
questions and thoughts related to their disease. The lack
of useful communication places for both patients and
health care personnel, is however striking.
The aim of this project is therefore to develop a web
service for both patients and health care providers that
will support the participation of patients with diabetes in
their medical care.
Methods
The web service was developed applying a user-centred
system design (UCSD) process. [3] Representative users
continuously participated from the beginning to the end
of the implementation in an iterative way. A feasibility
study in terms of a user needs analysis was an important
part of this project. The purpose was to analyse
conditions and demands for the use of specific web
services in diabetic care, mainly from the patients’ points
of view. The user needs analysis was made through deep
interviews with 12 patients with Diabetes Type 1 who use
insulin pump. The patients are registered at the Unit for
Endocrinology and Diabetology, Uppsala University
Hospital. In parallel, interviews and questionnaires
acquired the user needs from the staff perspective with
two nurses and one physician.
Iterative prototyping was used for development of
different design solutions. Four test patients were
selected for deep evaluation of the resulting prototype in
three iterations. Their viewpoints were of crucial
importance for further developments of the web
application.
System design
The web service consists of two different parts, a general
part and a more personal one. The general part includes a
forum where the patients can discuss and exchange their
experiences of Diabetes Mellitus, type 1. There is also an
extensive information portal, with information links to
news and different subjects related to the disease. The
topics included in the portal and in the forum were
defined by the users. Furthermore, in the personal part an
e-mail system was implemented where the patients can e-
mail their nurse or physician for care-related subjects
they do not want to share in the discussion forum. There
are also different service functions that make it possible
to order medical recipes and certificates in an electronic
way.
The information, which mediates between the patient and
the physician or the nurse in the application, contributes
to the care process and should therefore be stored in the
IFMBE Proc. 2005;9: 41

Telemedicine and patient data management
patient record. Moreover, this information is often
confidential. With that in mind the technical and secure
solutions were implemented at the same level as today’s
Internet banks. This requires that the user is authorised
with a username, a personal code and finally with a one
time password. If the initial identification is correct, the
one-time password is generated. The password is
automatically received by the user’s cell phone as an
SMS - message, every time the user wishes to log in.
• An improved HbA1c – value. The web service
increases the accessibility of the caregivers
which improves patient’s trust (the medical
personnel is communicative and the patient can
expect an answer).
• An improved self-care and empowerment – the
possibility to receive knowledge of the disease
by the information portal and by active learning.
• An improvement of the patients’ co-operation –
the patients themselves can communicate
through the discussion forum.
In an administrative way this test period aims to give:
• A more efficient way to communicate. The
patient and the caregiver do not have to keep a
specific time to communicate.
• Less interruptions in the caregiver’s daily work
– fewer phone calls.
• A more efficient way to handle the
administrative work, for example renewal of
medical recipes and certificates.
Figure 1: The user receives a one-time password through
his/her cell phone as an SMS-message.
Results
The web service is in operation over a test period during
5 months, started in the end of November 2004 to the last
of April 2005 with a user group of 20 patients and 5
caregivers from the Unit of Endocrinology and
Diabetology, Uppsala University Hospital. The overall
impressions and reactions have been positive so far. In
the end of April the user group will receive a user’s
questionnaire. Follow-up in terms of analysing the
questionnaires and, if needed, using complementary
interviews will be done in May 2005. The results thereof
will be presented at the conference.
References
[1] Socialstyrelsen, Nationella riktlinjer för vård och
behandling vid diabetes mellitus:
http://www.sos.se/fulltext/9900-061/9900-061.htm
[2] Vårdriktlinjer vid diabetes, Örebro läns landsting:
http://www.orebroll.se/prim/page____11175.aspx
[3] Gulliksen, J. & Göransson, B. (2002):
’Användarcentrerad systemdesign’, Studentlitteratur.
Discussion
The way we applied a user-centred system design has
turned out very well. The users themselves have been
involved in the entire process and have had the chance to
affect the development. The user questionnaires,
complemented by interviews, will be used as a basis for
analysing if this IT-tool has a positive effect on the
patients’ treatment and whether it supports patient
empowerment.
In medical meaning this test period aims to give:
IFMBE Proc. 2005;9: 42

Telemedicine and patient data management
COMPARISON OF MODEL DRIVEN ARCHITECTURE (MDA) BASED TOOLS
USING TELEMEDICINE HEALTHCARE SYSTEM
N. Tariq 1 , S. Jonsson 2 , R. Nabiev 2 , H. Teriö 2 , D. Andersson 3
1 Department of Computer and System Science, Royal Institute of Technology, Stockholm, Sweden
2 Biomedical Engineering Department, Karolinska University Hospital, Stockholm, Sweden
3 Department of Innovation and Medical Informatics, Karolinska University Hospital, Stockholm,
Sweden
emis-nat@dsv.su.se, sven.jonsson@karolinska.se, rustam.nabiev@karolinska.se,
heikki.terio@karolinska.se, daniel.andersson@karolinska.se, naveed_eltaf@yahoo.com,
naveed_eltaf@yahoo.com
Abstract
Success of any software system is dependent on the
selection of specific tools like design, development
tools etc, technologies like Java, .Net, CORBA etc and
methodologies like Water Fall Model, Spiral Model,
Incremental Model etc. It is difficult to choose
between which tools, technologies and methodologies
are appropriate for desired software system. This
difficulty is fairly simplified by one of the standard
development organization named Object
Management Group (OMG) by providing a new
methodology Model Driven Architecture (MDA).
MDA’s main target is to construct software from
graphically viewable, models. It raises the
development of software system from level of
abstraction to level of requirement gathering and
designing. In the end of year 2004 there are more than
40 MDA based tools available in market. Need was
felt by one of the research department named MTA at
Karolinska University Hospital (Karolinska) to
evaluate and compare specific MDA based tools for
Telemedicine and patient management systems which
may include information systems, mission critical
systems, embedded systems etc. This research
evaluates and compares selected MDA based tool
against Evaluation Criteria (EC) by using one of
Telemedicine healthcare and patient management
system name Telemed HC System (TMHC) developed
at MTA.
Introduction
In year 2001 Object Management Group (OMG) [1]
adopted Model Driven Architecture (MDA) [2]. MDA is
based on ideas of raising the level of abstraction,
automated software generation and platform
independence. Till the end of year 2004 there are more
than 40 MDA based tools available in the market [3].
MDA based tool vendors such as Compuware, Borland,
IO-Software etc are claiming that their products are fully
complaint of MDA.
One can’t just rely on the marketing statement of the
vendors while choosing between numbers of MDA based
tools. Different aspects are required for specific needs
like no compromise on integration capability, iterative
development and UML. Cost is also one of main factor
means why to pay for additional features which are not
required.
This research helps in choosing specific MDA based tool.
Since, there should be no further need to spend time in
evaluating the tools for certain features and match tools
to particular requirements.
Methods
Inductive research approach[5] is followed to conduct
this research. Five significant steps are followed to
achieve the goals of this research, as given below.
1. Information Gathering
This step involves questionnaire, formal specifications
from OMG for MDA [6] and study of related literature.
Questioner is developed for prioritizing and gathering
aspects & factors which are important for MDA based
tools. Experts who solved this questioner are from OMG,
industry and academia. Formal specifications of MDA
from OMG are used to formalize this research work
according to the standards and guidelines given by OMG.
2. Development of Evaluation Criteria (EC)
Development of EC is an important aspect of this
research since, selected MDA tools are measured
according to this EC. EC is break down in two broad
categories MDA specifications and General Factors (GF)
such as reusability, visual interface, etc. MDA
specifications are further categorized as Pervasive
Services, Domain Facilities and other OMG standards
which are constituents of MDA which are UML, MOF,
XMI, CWM, etc.
3. Evaluation of MDA based tools
In the third step MDA based tools are evaluated
according to developed EC. This evaluation of MDA
based tools are done practically by executing the software
provided by selected MDA tool vendors which includes
Aonix’s Ameos, Artisan’s Real Time Studio, BitPlan’s
IFMBE Proc. 2005;9: 43

Telemedicine and patient data management
UML2PHP, Borland’s Together, Compuware’s OptimalJ,
DomianSolutions’s Codegenie, IO-Software’s Arcstyler,
Mentor Graphic’s Nucleus Bridgepoint and Xactium’s
XMF-Mosaic.
EC are applied using one of Telemedicine healthcare and
patient management system name Telemed HC System
(TMHC) [7] developed at MTA one of the research
departments at Karolinska, Stockholm, Sweden.
4. Comparison of MDA based tools
This step contains the comparison of MDA tools based
on the evaluation of MDA tools done is last step.
5. Conclusions
Finally the results, conclusions and future work are
drawn on the bases of comparison of MDA based tools
done in the previous step.
Results
It is difficult to say which MDA based tool is better than
the other selected MDA based tool. Borland Together is
an advanced tool for modeling which can be used for
designing, modeling and code generation. Compuware’s
OptimalJ and IO-Software’s Arcstyler are quite mature
tools for MDA they also support designing and modeling
capabilities, one of the main advantages of these tools are
they take your PIM or some initial level model like class
model and generates the PSM and then generates the
code for you which reduce development time enormously
especially for J2EE platform. Bitplan UML2PHP
generates ready to use PHP5 based application from
UML models. Aonix Ameos supports the UML 2.0
profiles. Domain Solution’s CodeGenie provide support
Executable UML and generates C++ and/or Java code.
Xactium’s XMF-Mosaic is also an advance tool for MDA
based development for Java. Artisan’s Real time studio
and Mentor Graphic’s Nucleus Bridgepoint tools are best
for real time and embedded systems where results or code
generation is critical and should be reliable.
Discussion
During the evaluation it has been noticed that selected
MDA tools can be used for specific need. Artisan’s Real
time studio and Mentor Graphics’ Nucleus Bridgepoint
have power full compilers which generates reliable code
for mission critical and embedded software.
Compuware’s OptimalJ, IO-Software’s Arcstyler,
Xactium’s XMF-Mosaic, Borland’s Together and
Domain Solution’s CodeGenie can be used for
information management system for Java and C/C++
platforms. Bitplan’s UML2PHP, Compuware’s OptimalJ
and Borland’s Together can be used for web based
software development.
Conclusions
MDA is based on ideas of raising the level of abstraction,
automated software generation and platform
independence. MDA allows definition of machine
readable application and models that can allow long term
flexibility in terms of implementation, integration,
maintenance, testing and simulation. There are more than
40 MDA based tools available in the market till the end
of year 2004. Selected MDA tools for this research have
different features for different need or type of system.
Borland’ Together provide advance designing, modeling
and code generation facilities. Compuware’s OptimalJ,
IO-Software’s Arcstyler and Xactium’s XMF-Mosaic are
good for Java and iterative development. Bitplan’s
UML2PHP generates PHP5 deployable code. Domain
Solution’s CodeGenie supports Executable UML and
generates C++ and Java code. Artisan’s Real time studio
supports code generation for C, Java and Ada and can be
used for mission critical software. Mentor Graphic’s
Nucleus Bridgepoint generates C and Java code for
embedded systems.
References
(Electronically Publications)
[1] OMG, Internet site address:
http://www.omg.org/
[2] MDA Guide, Internet site address:
http://www.omg.org/docs/omg/03-06-01.pdf
[3] ADTmag, Internet site address:
http://www.adtmag.com/print.asp?id=7850
[4] Compuware, Internet site address:
http://www.compuware.com/products/
optimalj/default.htm
[5] Barbara Kitchenham, University of Keele, Internet
site address:
http://www.keele.ac.uk/depts/cs/se/e&m/tr9609.pdf
[6] MDA Specifications, Internet site address:
http://www.omg.org/mda/specs.htm
[7] KTH, Internet site address
http://web.it.kth.se/~iw01_nru/exjobb/documents/
deliverables/final_report/TS_021223_D04_V02.pdf
Acknowledgements
We acknowledge the support of companies which
includes Aonix, Artisan, Bolrand, Bitplan, Compuware,
Domain Solutions, IO-Software, Mentor Graphics and
Xactium who were involved in this research and provided
technical trainings, their MDA based tools, maintenance
and technical supports. We would also like to thank
people from OMG, Karolinska, KTH, Kings
College(London), Cephas Consulting, David Frankel
Consulting, EDS, IBM, OCI, InfoTech Consulting,
Mathworks and MID, who were involved in this research
and provided us their valuable suggestions and advices.
IFMBE Proc. 2005;9: 44

Telemedicine and patient data management
A LOW COST ECG MONITORING SYSTEM EMPLOYING
TELEMEDICINE
Mamun Bin Ibne Reaz
Faculty of Engineering, Multimedia University, 63100 Cyberjaya, Selangor, Malaysia
mamun.reaz@mmu.edu.my
Abstract: This paper describes an implementation of
a low cost real time remote ECG monitoring system.
The system is capable of acquiring and storing
patient’s ECG data and transfers it in real time to
the remote terminal. It comprises of an acquisition
and a networking part. The acquisition part acquires
ECG signal through electrodes and then amplifies
the weak ECG signal by an amplifier and filter out
noises. ADC modules carry out A/D conversion of
the ECG signal and fed into the interface circuit for
level conversion. Serial port program enables data to
be stored in a PC from acquisition device. The
networking part is in the form of a Java based
client/server pair application and installed in local
and remote terminal respectively via TCP/IP to
provide transfer of data and enable chat session.
This system is utilizing Internet protocols,
commercial software and low cost component to
transmit ECG data to physicians for monitoring,
diagnosis and patients care at a significantly low
cost, regardless of patient’s location.
Introduction
The use of electronic and communications
technologies to provide and support health care when
distance separates the participants is well known
definition of telemedicine which was published in 1996
by the Institute of medicine [1]. Reports indicating that
telemedicine has been a great concern for physicians
with a passion for technology, and barriers still remain
for a low cost, comprehensive and integrated use in the
daily operations [2]. Telemedicine reduce costs by
enabling in-home monitoring of patients, eliminating
the need for utilization of expensive facilities, and
reducing the need for transportation of patients to
physicians and medical centers [3].
One application of telemedicine is to rapid
transmission of electrocardiogram (ECG) data to
physicians so as to improve patient care and conserve
healthcare resources in managed care environment.
ECG transmission has been particularly useful for
pacemaker follow-up [4] and other patient monitoring
applications [5]. The use of ECG transmission for
emergency settings has been emphasized in order to
reduce response time in infarct size control or
resuscitation of sudden cardiac-death victims [5].
Real-time ECG transmission via the Internet has
been previously reported elsewhere [6], in order to
provide direct access to physicians in remote locations
to coronary-care-unit patient-monitoring data and to
check patients being monitored at their homes. This
paper describes a complete low cost real-time ECG
monitoring system, including the signal acquisition
hardware, client, and server applications, the required
transmission protocols, and the consultancy features.
The system is user friendly and does not require any
particular training aside from knowledge of widespread
and standard Internet tools. Due to the interactive
approach of the proposed system, the physician is also
able to make online consultation directly from the server
software provided.
Design Overview
This project comprises of two main parts - namely
ECG acquisition system and networking application.
Electrodes are placed on human body to capture small
electrical voltage produced by contracting muscle due to
each heartbeat. The ECG signal obtained by the
electrodes is in the range of 1 to 5mV. Due to the weak
voltage level, the signal is fed into an instrumentation
amplifier to amplify and filter the acquired signal. The
amplified signal is then fed into the ADC circuit for
A/D conversion. Digital output of the ADC is sent to
local terminal (patient’s terminal) via an RS232
interface circuit for level conversion. The digital data is
then read and stored by a serial port program running in
the local terminal and transferred to a remote terminal
via TCP/IP and sets up a full duplex communication.
There are various features such as plot graph, chat, play
of sound clip and patient’s database incorporated into
the networking as depicted in Fig. 1.
Figure 1: Structure of the project
Hardware Design
The first stage of hardware design for ECG data
acquisition system is a simple ECG sensing, using
electrode. The output from ECG sensor is fed into the
IFMBE Proc. 2005;9: 45

Telemedicine and patient data management
second stage for signal amplification and filtering
purposes. Next, the analog output from the second stage
is fed into the third stage for analog to digital
conversion. Finally, the digital output from ADC is sent
to a PC via an RS232 interface circuit. Fig. 2 shows the
schematic diagram of the circuit.
Results and Discussion
A test is carried out on each part of the developed
hardware to view how one affects the other. The AD620
has the common mode noise rejection ratio feature.
Therefore, the 50Hz or 60 Hz line interface is reduced.
However, there is still noise overridden on the ECG
waveform. These noises are due to the electrical activity
of the active muscle and white noise.
The low noise frequency that exists while measuring
the ECG signal from human’s body is greatly reduced
through the low frequency filtering. It is noticed that the
peak R-wave is clearly shown.
To test the networking program, both LPMS and
RPMS is installed in respective computer and
successfully execute all the modules.
Conclusions
Figure 2: Schematic diagram of amplification and
filtering stage.
Software Development
The software is developed using Java programming
language and the networking program is manifested in
the form of a client/server pair application. The client
application, known as a Local Patient Monitoring
System (LPMS), is installed in the patient’s terminal
(PC) while the server application, known as Remote
Patient Monitoring System (RPMS), is installed in the
physician’s terminal. A connection is established when
a user in LPMS chooses a desired location to be
connected. The desired locations are restricted to the
predefined hospitals or clinics. Once the network has
been successfully set up, user sends ECG data file as
well as sound file to RPMS. Physicians at the remote
location display the ECG graph and carry out analysis
on the patient’s heart condition. The physician may also
listen to the heart beat sounds using a standalone audio
player, which enables the sound clip to be played and
looped, as desired. In addition, users from both ends can
communicate with each other via a simple chat program
that is incorporated in both LPMS and RPMS. Finally,
RPMS has a patient’s medical information database that
enables patient’s data entry to be created, updated and
retrieved. The interface for both LPMS and RPMS
provides similar features except for the database feature,
which is only available at RPMS.
Networking module, Send audio file module, Chat
module, Plot ECG-waveform module, and Audio player
module make up the network program. The process of
algorithm design and code writing is based on a
modular architecture. Each of these modules is
constructed separately using Java software JDK1.2.2.
This programs are then incorporated into the Graphical
User Interface (GUI) using Java IDE tool JBUILDER
which coordinate and manage all these functional
modules together.
The ECG data acquisition device is successfully
developed. The device can get ECG data from human
body and send the ECG data in digital form to a PC or
patient’s terminal. In addition, the network application
successfully sets up a full duplex and point-to-point
connection. The networking system enables transfer of
ECG and heart beat sound files, online chat, ECG
display and also retrieve, update and create patient’s
medical record database. Thus, the objective for this
project, which is real time implementation of
Electrocardiogram (ECG) data transfer is achieved.
References
[1] Institute of Medicine (U.S.). Committee on
Evaluating Clinical Applications of Telemedicine.
Telemedicine: a guide to assessing
telecommunications in health care. National
Academy Press, Washington DC, 1996.
[2] Swedish federation of county councils, What are the
barriers facing telemedicine?, Stockholm, 2000.
[3] Jannett, T. C., Prashanth, S., Mishra, S., Ved, V.,
Mangalvedhekar, A., Deshpande, J., Proc. of the
34th Southeastern Symposium on System Theory,
2002, pp53-56.
[4] H. Hutten, G. Schreier, P. Kastner, and M.
Schaldach, Cardiac telemonitoring by integrating
pacemaker telemetry within worldwide data
communication systems, Proc. XIX Annu. IEEE
EMBS Conf., Chicago, 1997, pp 974-976.
[5] S. Pavlopoulos, E. Kyriacou, A. Berler, S.
Dembeyiotis, D. Koutsouris, A novel emergency
telemedicine system based on wireless
communication technology—AMBULANCE, IEEE
Trans. Inform. Technol. Biomed., vol. 2, 1998, pp
261–267.
[6] S. H. Park, J. H. Park, S. H. Ryu, T. Jeong, H. H.
Lee, C. H. Yim, Real-time monitoring of patients on
remote sites, Proc. of the 20th Annual International
Conf. of the IEEE Eng. in Med. and Bio. Society, vol.
3, 1998, pp 1321-1325.
IFMBE Proc. 2005;9: 46

Telemedicine and patient data management
A TECHNICAL PLATFORM FOR REMOTE AUSCULTATION AND
REAL-TIME MONITORING OF PHYSIOLOGICAL PARAMETERS
J. Skönevik*, P. Hallberg*, A. Müller*, R. Lundström*, U. Wiklund*,**, J.S. Karlsson*,**
*Department of Biomedical Engineering & Informatics, University Hospital, Umeå, Sweden
**Department of Radiation Sciences, Umeå University, Umeå, Sweden
E-mail: stefan.karlsson@vll.se
Abstract
New technical solutions change the way health
care providers communicate and how the patient
is diagnosed. Our aim was to develop a platform
for remote consultations and monitoring of
physiological parameters in real-time over the
Internet. The system worked satisfying and
preliminary tests with pediatric cardiologists
were promising for further development.
Introduction
Demographic changes, with the aging problem,
and a search of cost effectiveness in healthcare
motivate the development of new diagnostic
methods and tools that will reduce the need for
transportation of patients and specialists. Current
trends, supported by the EU eHealth program, aim
towards better information and communication
technology solutions; telediagnosis, home health
care and smart monitoring solutions [1]. Such
systems are implemented in balance between
optimistic redesign and the traditions of the
profession; many new telemedicine solutions are
ready to take off, especially with the Swedish
healthcare network Sjunet and evolving wireless
broadband techniques.
We aim to build a generic platform that fit
different types of monitoring needs. A system of
PC software, computers, sensors and headphones
that will be a useful tool for remote consultation
based on auscultation, electrocardiogram (ECG)
and other physiological parameters.
Platform Design and Operation
Developing the platform and then specialize it
to specific monitoring needs was an iterative
process in cooperation with users in the medical
profession. User studies gave us an understanding
of how the users work; frequency and sequence of
actions, tools and method of clinical examination.
This guided us in designing a system that was easy
to use and solved the right problems. The first user
group identified and targeted for evaluation of the
platform was pediatric cardiologists at Norrland’s
University Hospital (NUS), Umeå, with a need to
remotely monitor auscultations of children at
Skellefteå Hospital, 130 km away. These children
were frequently referred to specialist cardiologists
in Umeå since the common healthy functional heart
murmurs could not be definitely distinguished from
the pathological ones. Remote auscultation would
save patients and specialists from unnecessary
travels and reduce the cost of the examination
(figure 1). Also the patient would not have to
anxiously wait for the delayed diagnosis.
Figure 1. Real-time monitoring over the Internet
enables the specialist to remotely diagnose the
patient.
Persona and scenario building resulted in a low
fidelity prototype and an interactive prototype.
Software implementation focused on modularity
and reusability. The design was object oriented and
multithreaded to handle software complexity.
Scenarios, use cases, object interaction diagrams
and class diagrams lead to an implementation that
was modularized into separate software packages
for handling graphics, sound, sensor digitalization
and network communication. Specific features were
implemented for the pediatric cardiologists, such as
registering the stethoscopes position on the body
during auscultation, customized filtering and high
fidelity headphones for listening.
IFMBE Proc. 2005;9: 47

Telemedicine and patient data management
A platform for remote monitoring of physiological
parameters in real-time over the Internet has
been implemented and tested. The core of the
software is techniques for reliable network
streaming and presentation of the physiological
parameters as graphs and sound. The main concept
is that you share your sensors with colleagues over
the internet. Any or all of the physiological
parameters currently monitored on your own
computer such as files, remote sources or locally
measured signals, can be shared with others in realtime.
Sharing sensors and connecting to others is
easy since a central hub keeps track of online users
and their signals. Navigation controls are used to
make selections in the graph and play the sounds
repeatedly. Filters applied in real-time enhance the
sound quality and selected frequency bands.
Platform Implementation
Streaming the continuously sampled physiological
parameters over the Internet uses the
Transmission Control Protocol (TCP). Data is prebuffered
at the receiver to handle variability in
network packet delays. Overall latency depends on
network bandwidth, load and congestion, but
usually it is between 200-500 ms. To reduce the
required bandwidth, a lossless compression is used.
Sound is typically reduced to 1/3 of the original
size, without loosing any information. Heart sounds
are digitized with a Meditron M30 electronic
stethoscope sampled with a soundcard. Sensors for
ECG and other parameters are connected to our
wireless system [2].
Discussion
This software sends physiological parameters in
real-time over the Internet. But, as opposed to the
technology typically used for live Internet feeds of
music, radio or videoconference it uses more
reliable transmission channels; no medical data is
lost or distorted. Another major concern is to create
an interface and features that fit the profession and
their method of examination. Therefore we do not
base our software on any proprietary components
that would effectively prevent us from customizing
and making it work exactly the way it needs to.
TCP is reliable; no network data packets are
lost, reordered or duplicated during transmission.
We considered defining mechanisms for recovering
from packet loss on top of User Datagram Protocol
(UDP) or Real-time Transport Protocol (RTP). This
way, reliability could be implemented with better
real-time performance, without enforcing the
congestion control mechanisms that decrease the
congestion window when packet losses are
detected. We decided against it though, since it is
quite an effort for little gain; monitoring is not very
delay sensitive, and voice communication can be
solved separately if necessary, with lossy high ratio
compression and UDP or RTP streaming. The next
generation of Internet protocols, IPv6, will give a
better quality of service for real-time streaming and
new possibilities for this platform implementation.
Meditron M30 electronic stethoscope, used to
perform auscultations, amplifies the sound and is
perceived differently than traditional acoustic
stethoscopes. The doctors need to practice to listen
and analyze sound from electronic stethoscopes to
feel comfortable with the technique. Soundcards
can sample with rates up to 96 kHz and 24 bits/s
but this is not necessary for auscultation; the
frequency range of acoustic stethoscopes is
typically 20-2000 Hz and higher frequencies are
effectively damped by the body tissue [3].
Psychoacoustic modeling methods like Ogg Vorbis
or Advanced Audio Coding (AAC) were not used
for a number of reasons. The fact that they are lossy
may ruin future analysis of the data. They only
work for sound data and also they are quite
computationally expensive. Our design was built on
the principle that no medical data should be lost or
altered, and not be saved in any closed standards
format.
Summary
This paper explains the design and software
implementation of a platform for real-time remote
auscultation and monitoring of physiological
parameters, specialized to the needs of pediatric
cardiologists at NUS. Acceptance testing and
further development is next.
Acknowledgements
Thanks to Annika Rydberg and Dag Teien,
pediatric cardiologists at NUS, for evaluating the
system during the development process.
References
[1] LYMBERIS A, DE ROSSI D. (2004): ‘Wearable
eHealth Systems for Personalized Health Management:
State of the Art and Future Challenges’, in
Studies in Health Technology and Informatics.
[2] KARLSSON JS, BÄCKLUND T, EDSTRÖM U
(2003): ’A new wireless multi-channel data system
for acquisition and analysis of physiological
signals’, proc. 17 th International Symposium on
Biotelemetry, September, Brisbane, Australia.
[3] LUKKARINEN S, NOPONEN A, ANGERLA A,
SIKIÖ K, SEPPONEN R. (1996): ‘Frequency
Response Measurements on Commercially
Available Stethoscopes’, in Medical and Biological
Engineering and Computing.
IFMBE Proc. 2005;9: 48

Telemedicine and patient data management
WEARABLE MOBILE TECHNOLOGY FOR HEALTHCARE:
O. Atzmon 1 , D. Shklarski 1 , S. Jonsson 2 , H. Teriö 2
1 Tadiran Spectralink Ltd., Tel Aviv, Isreal
2 Biomedical Engineering, Karolinska University Hospital, Stockholm, Sweden
oatzmon@tadspec.com
Abstract
TeleMedIS is a new e-Health system for remote
monitoring of patients, which will be evaluated in
Karolinska University Hospital, Huddinge. The system
consists of: TMW (Telemedicine Monitoring
Wristwatch), a personal wireless wrist-wearable remote
monitoring device, developed by Tadiran Spectralink
(Israel); Telemedicine Home-Care Platform (TeleMed
HC Platform), which is a platform for remote
management and support of biomedical equipments
located in patients' homes, and established in the
Biomedical Engineering Department at Karolinska
University Hospital, and Wireless Network infrastructure
provided by TeliaSonera Sweden AB. The project will
evaluate and investigate the possible deployment and
commercialization of TeleMedIS solution in the swedish
healthcare sector:
Introduction
The healthcare industry is undergoing significant
changes. e-Health is becoming increasingly important
and is being considered as an answer to the rising costs of
healthcare, the ageing of the population and the
increasing demand for high quality healthcare.
Results
TeleMedIS enables continuous monitoring of vital
signs of patients, such as pulse rate, 1-lead ECG, and
blood oxygen saturation level (SpO2) with real-time
cellular communications, using its built-in embedded
cellular engine, transmitting medical date, receiving
medical advice and conducting two-way voice-based
consultation with a remote medical professional. Patients
can be monitored from anywhere, in their homes, at work
or during traveling.
Discussion
The presentation will discuss the systems’ features, the
trials program, the expected outcomes and the potential
benefits for patients, physicians, administrators and other
stakeholders, as well as future research which will be
needed for exploiting the systems benefits and its
influence on future health delivery methods.
TeleMedIS is a new e-Health system for remote
monitoring of patients , which will be evaluated in
Karolinska, Stockholm. The system consists of: TMW
(Telemedicine Monitoring Wristwatch), a personal
wireless wrist-wearable remote monitoring device,
developed by Tadiran Spectralink (Israel); Telemedicine
Home-Care Platform (TeleMed HC Platform), which is a
software platform established in the Biomedical
Engineering Department at Karolinska University
Hospital, and Wireless Network infrastructure provided
by TeliaSonera Sweden AB.
Methods
TeleMedIS enables continuous monitoring of vital
signs of patients, such as pulse rate, 1-lead ECG, and
blood oxygen saturation level (SpO2) with real-time
cellular communications, using its built-in embedded
cellular engine, transmitting medical date, receiving
medical advice and conducting two-way voice-based
consultation with a remote medical professional. Patients
can be monitored from anywhere, in their homes, at work
or during traveling.
IFMBE Proc. 2005;9: 49

Telemedicine and patient data management
IMPLEMENTATION OF A TECHNICAL SYSTEM IN
DISTRIBUTED CARE -
ATTITUDES AND POSSIBILITIES
L. Rattfält 1 , C. Hagström 2 , M. Lindén 3 and P. Hult 1
1 Department of Biomedical Engineering, Linköping University, Linköping, Sweden
2 Biomedical Engineering, Örebro University Hospital, Örebro, Sweden
3 Computer Science and Electronics, Mälardalen University, Västerås, Sweden
linra@imt.liu.se
Abstract: Demographic conditions are changing and
the age distribution is showing an increase of elderly
people. Today's health care needs to reassess its
possibilities to maintain a high quality care despite
these changes. A reborn concept of distributed care
has been seen as a solution. In this study, a
hypothesis of a future system for distributed care is
presented. A questionnaire study has been
performed to investigate the attitudes among nurse
assistants towards working with high technology
solutions according to the presented model. The
results show an overall interest in learning and using
new technologies on the premises that the care is
improved. Based on the condition that increased
distributed care will generate a large amount of
data, a discussion is made of the possibilities that
technology will provide. It can for example help to
distinguish, interpret and visualize sensory signals in
such a system.
Introduction
Several studies have been made on the subject of
technology in home health care. Ongoing projects in
telehealth are for example pain estimation, diabetes
counseling and support for informal caregivers [1]. The
need of biosensors has been discussed in for example
[2] and [3]. The organization of future health care
systems is an open question, but current research on
distributed care will be an important part of it. Even
though it is hard to evaluate the total societal cost for
different types of cares, studies have shown that present
home care programs cost as much as regular care when
the informal caregiver’s work is included [4]. However,
the prospect is that along with a more adjusted
organization, the home care will be more effective
compared to hospital care.
In Fig. 1, a hypothesis of what a future health care
system could look like is illustrated. There are two types
of locations; centrally at the hospital or distributed.
(Distributed is where the caretaker is; at home, on
vacation etc.) At home, relevant parameters such as
ECG and blood pressure are measured and a simple
analysis is performed. The measurements are then
transmitted via Internet, GPRS or alike to a centrally
placed database. If needed, a physician can access the
database to suggest an action.
Distributed
Sensor
signals
Analyis
database
Centrally
Analyis
Decision
support
Figure 1: A schematic view of a distributed
health care system.
An example of the system's utility would be if an
elderly caretaker is ill and the arriving nurse wants a
consultation about appropriate treatments. The doctor
can order an ECG, which the nurse executes and
transmits to the database. The doctor accesses the ECG
via the database with little time delay and can then
suggest what to do. Other advantages with this system
are for example that it allows follow up measurements
after diseases, long time monitoring and that the
caretaker does not need to travel.
The aim of this study was to perform a questionnaire
and interview study among personnel in the health care
sector to investigate their attitudes towards integrating
technology needed for distributed care in their work. A
complementary practical trial was performed with
personnel unaccustomed to technological equipment.
Materials and Methods
The questionnaire study comprised 77 nurses and
nurse assistants and focused on four different types of
technologies; computers, Internet, mobile phones and
digital cameras. For each one of those, the informants
stated their present ability, frequency used, the need for
basic training and future prospect of using the
technology in their work. 96% of the informants were
female.
The results from this study were used to develop
questions and a strategy for the interview series which
included 25 persons from different professions within
the distributed care sector. The participants were nurse
assistants, nurses, physicians and administrative
personnel. Due to different perspectives, needs and
knowledge the interviews were held either in groups or
individually and had different foci. For the medical staff
the questions reflected their present working situation
IFMBE Proc. 2005;9: 50

Telemedicine and patient data management
and future visions and needs concerning retrieval of
objective bio parameters. For the nurses and nurse
assistants focus was usability and attitude towards
technical devices in their work and for the
administrative staff, organizational questions were
highlighted. For the practical trial an ECG-amplifier and
a graphical user interface were developed.
Combining this parameter with the measured weight, a
graph indicating safe and dangerous regions can be
created, see Fig. 2.
Results
The results from both the interviews and the
questionnaire showed a general positive attitude towards
using more technology in home health care, but it has to
be motivated by for example better diagnostic
capabilities, better care or improved security for the
caretaker. However, before technical devices are taken
into use, the staff expressed their wish to be adequately
educated in order to feel comfortable. It was a
consensus among the informants that increased use of
technology under no circumstances was to replace or
decrease the current personal care.
Nurses and physicians stated that one of their main
tasks is to assess the caretakers’ condition. Normally
this is done by the general impression the care provider
gets by watching, listening and by palpation. Having
access to more objective parameters such as ECG and
heart sound would improve the physicians’ possibilities
to suggest better treatments. In the nurse group, they
stressed that their role should be restricted to being
information providers for the doctors when it comes to
ECG and heart sound interpretation. In the trial study,
the personnel were able to handle the equipment
satisfactorily.
Discussion
The positive results of this questionnaire and trial
study are in accordance with similar studies, indicating
a future increase of technical devices in distributed care
[1]. However it is important to reflect over the
consequences distributed care may have in the care
takers' homes, for example to rearrange furniture in
favour for measurement equipment [5].
The use of biosensors will increase the amount of
data rapidly. It is not possible for physicians to survey
all of this information, and to reduce it to a reasonable
level is crucial. This motivates an extension of the
previous model (Fig. 1) with a signal processing part
connected to the database. Its main task would be to
extract interesting parameters and help presenting them
in an efficient way. Having these parameters, diagnosis
support can also be performed. An example is the
Minnesota code, which is a register of parameter values
for detecting pathologies in ECG:s [7].
To demonstrate the possibilities of signal processing
in this context an example is given. For persons with
congestive heart failure, heart rate variability and weight
are interesting parameters to monitor over time. The
heart rate variability can for example be derived from
analysis of the ECG by finding the QRS complex.
Figure 2: A visualization of how weight and heart rate
variability change over time.
In the figure, two manifolds represent the regions of
healthy and unhealthy states. If the measurement falls
into the unhealthy manifold, an alarm should be raised
and a doctor or nurse contacted. In this example the
dotted curve changes over time towards the healthy
state, meanwhile the dashed curve causes an alarm
when it intersects the unhealthy manifold.
Conclusions
This study shows that the attitude towards
technology in the distribute care is positive as long as
the quality of care is maintained. A pilot study shows
that personnel unfamiliar with technology can handle
technological equipment such as ECG registration.
References
[1] UTBULT, M. (2004): ‘Vård nära dig’, Teldok
rapport 152
[2] ESSÉN, A. (2003): ‘Kvarboende och äldrevård i
hemmet med modern teknik’, Institute for Futures
Studies. ISBN: 91-89655-34-6
[3] ROSS P. E. (2004): ‘Managing care through the air’,
IEEE Spectrum, Dec, 14-19
[4] ANDERSSON A. (2002): ‘Health economic studies
on advanced home care’, Dissertation, ISBN: 91-
7373-445-4
[5] TAMM, M. (1999): ‘What does a home mean and
when does it cease to be a home? Home as a setting
for rehabilitation and care’, Disability and
Rehabilitation, 21, pp 49-55
[7] KORS, J. A., CROW, R. S., HANNAN, P. J.,
RAUTAHARJU, P. M., FOLSOM, A. R. (2000):
‘Comparison of computer-assigned Minnesota
codes with the visual Standard method for coronary
heart disease events’, Am. J. of Ep., 151, pp 790-
797
IFMBE Proc. 2005;9: 51

Telemedicine and patient data management
SPEX-SPREADING EXCELLENCE IN HEALTHCARE-A TELEMEDICINE
PROJECT
E. Fridén 1 , B. Eklund 2 , B. Gerdin 3 , P. Gustafsson 4 , K. Eriksson 1
1 Department of Medical Physics and Biomedical Engineering, Mälarsjukhuset, Eskilstuna, Sweden
2 County Council, County Council of Uppsala, Uppsala, Sweden
3 Department of Plastic Surgery, University Hospital, Uppsala, Sweden
4 Department of Surgery and Urology, Mälarsjukhuset, Eskilstuna, Sweden
erik.friden@dll.se
Abstract
The focus of the EU funded telemedicine project
SPEX- “SPreading EXcellence in Healthcare” is to
market validate a healthcare organizational model
based on healthcare units using tele-medicine. The
model relies on a network between a highly
specialized Centre of Excellence and one or several
peripheral Points of Care. By spreading clinical
knowledge from a university hospital to a county
hospital, overall resources are used more efficient.
Hospitals in Italy, Spain and Sweden participate in
SPEX. The Swedish group is formed by the University
Hospital of Uppsala and Mälarsjukhuset in Eskilstuna
with SYSteam Udac as a technical partner. Using teleconsultations,
doctors in Eskilstuna can discuss
diagnosis and treatment with colleagues in Uppsala.
The procedure is so far used specifically for wound
patients in plastic surgery but can be adapted to other
medical situations. Efficient decisions due to added
knowledge can shorten treatment time for this group
of patients at a lower cost, which is a major advantage
to both patients and society. Many patients with
chronic ulcers and severe burns demand health care
resources during several years or even for life.
Technology used in the project is mainly consumer
electronic products, which have proven to perform
well and to be user-friendly at low costs.
Introduction
In this paper the Swedish part of a project called SPEX -
SPreading EXcellence in Healthcare will be described
concerning the organizational model together with the
technical platform that is being used. Some preliminary
results from the first half of the project’s field trials will
also be presented.
SPEX is a telemedicine project funded by the European
Union. The objective of SPEX is to market validate an
organizational model used when knowledge in healthcare
are spread using telemedicine. The market validation is
done through the project’s field trials. The field trials are
successful if the technical platform and the organisational
model work together in a way that medical knowledge is
spread. Knowledge is spread from a highly specialist
Centre of Excellence (e.g. a University Hospital) to a
smaller healthcare unit called a Point of Care unit (e.g.
County hospitals).
In SPEX three European countries participate; Italy,
Spain and Sweden. A pilot site for each country is set up
with some differences from each other regarding national
healthcare systems, medical specialty and structures of
telecommunications.
The Swedish group is formed by the Department of
Plastic Surgery at the University Hospital of Uppsala in
County Council of Uppsala, which is collaborating with
the Department of Surgery and Urology at
Mälarsjukhuset in Eskilstuna, County Council of
Södermanland. SYSteam Udac is a technical partner for
the project. This group is focused on plastic surgery,
specifically burns, chronic ulcers and pressure sores.
Many patients with chronic ulcers and burns demand
resources from health care during several years or even
for life.
Both hospitals have prime responsibility for surgical care
in their respective counties, but the Department of Plastic
Surgery at the University Hospital of Uppsala also get
referrals concerning more complicated cases from other
counties.
Methods
Organizational model: Complicated cases in the County
Council of Södermanland that are possible referrals to the
University Hospital of Uppsala will be selected for
treatment within the SPEX project. Tele-consultations
between doctors are scheduled every two weeks, but can
also take place at other occasions.
Overview, technical platform: All data transfer and
communication takes place on Sjunet, the Swedish
Hospital Network. The technical platform will be tested
by the doctors during the project’s field trials. The
technical platform contains five different tools for teleconsultations
that can be used depending on the situation.
Most of the technical products used are consumer
electronic products. The five tools are:
Video-streaming: Tele-consultations are performed using
video-streaming equipment in the outpatient clinic in
Eskilstuna. The patient’s wounds are being filmed by a
DV camera operated by a nurse. The camera’s output
video signal is connected to the video streaming system
IFMBE Proc. 2005;9: 52

Telemedicine and patient data management
and sent live to Uppsala. In Uppsala one or two
physicians can comment and take part in the patient visit.
Video traffic is only one-way but speakerphones in both
ends are used by the doctors to communicate.
Videoconference equipment: The use of a PC- based
video conferencing system makes it possible for more
than two physicians to participate in a tele-consultation.
The patient is not necessarily present in this situation.
The videoconferencing software is installed in a standard
PC workstation together with a web camera and a
microphone headset. Doctors can take part from their
office or even from home while on call using a VPN
connection.
Video calls using 3G mobile phones: Video calls with
standard 3G mobile phones can be used in urgent
situations that demand tele-consultations. The camera in
the mobile phone captures video images of patients in
Eskilstuna and sends video images to the doctor in
Uppsala. The video quality is low but good enough for a
less advanced consultation.
Shared patient records: Doctors participating in the
SPEX project at both hospitals have access to patient
records with notes and digital images from the SPEX
tele-consultations. These notes can be discussed using
email in a more asynchronous way.
See and share PC desktops and documents: Software that
enable sharing PC desktops between two or more
participants give doctors the opportunity to discuss digital
medical images before and after surgery. The software
also has a tool for drawing and annotations on screen.
Table 1: Products used in the SPEX project
Manufacturer Product
Video streaming AXIS
Video server
equipment Communications, 250S
Sweden
Speaker phone Konftel, Sweden Konftel 50
Videoconference VCON, Israel vPoint HD
software
PC desktop Tandberg, See&Share
sharing software Norway
Web Camera Creative, Ultra FX
Singapore
DV camera Panasonic, Japan NV-GS 200
Results
Field trials in the Swedish part of the SPEX project show
so far that twelve (12) patients have been participating in
fourteen (14) tele-consultations. Eleven (11) of these
patients (92 %) have been treated in the County Council
of Södermanland instead of being sent to Uppsala.
Patients are treated locally due to the added knowledge
that is spread between doctors during the consultations,
which results in fewer transports of patients. Costs are
lower for the County Council of Södermanland however
incomes are decreasing to the University Hospital of
Uppsala caused by less wound patient referrals from the
County of Södermanland. Preliminary results show that
patients and doctors are satisfied with the teleconsultations
in SPEX. Among the five tools in the
technical platform, the video streaming equipment has
been used in most of the consultations. The system has
proven to produce high video quality for diagnosis and is
also easy to use.
Discussion
Spreading new medical research results and knowledge
from university hospitals to peripheral hospitals is a
driving force for continuous development in medicine.
By spreading knowledge from a university hospital to a
peripheral hospital, the overall resources can be used in a
more efficient way. Using tele-consultation is one way to
achieve that. The use of excellent specialist resources in
the handling of the most complex clinical cases is a
natural approach. If the care of ordinary clinical cases
interferes with this, resources are not used in the most
efficient way.
Recent years’ fast expansion of broadband-networks and
especially Sjunet, the Swedish Hospital Network, makes
it possible to send large amounts of data between
hospitals in a secure way. The combination of increased
data network capacity and improved consumer
electronics products makes tele-medicine available at
lower costs.
Conclusions
More efficient decisions due to added knowledge from
tele-consultations in SPEX give the possibility to shorten
treatment time for patients with problematic wounds,
which is also a major advantage to society. Using the
SPEX service, a patient in the County of Södermanland
can be treated in the local hospital or in primary care with
fewer costs than before.
In SPEX mostly consumer electronic products have
successfully been used with relatively low investment
costs.
A solid organizational model also requires an agreement
between hospitals on how the Centre of Excellence
covers their costs for the consultations. This agreement
together with further testing of the technical platform will
show if SPEX can survive the project phase and be used
in other counties and medical specialties.
References
(Journals)
[1] LOULA P, RAUHALA E, ERKIJUNTTI M, RATY
E, HIRVONEN K and HAKKINEN V. (1997):
‘Distributed clinical neurophysiology’, J of Telemed
Telecare 1997;3:89-95.
[2] MANSON N. ‘Telemedicine and the New Children’s
Hospital (Royal Alexandra Hospital for Children)’, J
Telemed Telecare 1997;3 Suppl 1:46-48.
IFMBE Proc. 2005;9: 53

Telemedicine and patient data management
TECHNICAL CONSIDERATIONS AND WORKAROUNDS INTEGRATING
FUSION IMAGE DATA INTO A TRADITIONAL PACS ENVIRONMENT
J. Leal 1
1 Radiology, Johns Hopkins University, Baltimore, United States
jpleal@jhu.edu
Abstract
We evaluated the integration of fusion image data into a
traditional PACS environment. Inadequacies of existing
systems and methodologies were learned; workarounds
and alternative strategies were developed and are
presented here.
Introduction
The emerging area of fusion imaging, where functional
images are either co-acquired or co-registered in space
with anatomical image data sets, has resulted in the
acquisition and manipulation of data sets with processing
and data handling requirements much different than the
typical Radiological image data set. In many situations,
this has forced the adoption of PACS technologies in
areas of diagnostic imaging which heretofore have been
able to do without. In this work, we share our experiences
attempting to do so and strategies learned to make such
an endeavor a successful one.
Methods
A variety of imaging devices were used as sources for the
data in this evaluation. These included a General Electric
Medical System’s (GEMS) Discovery LS PET/CT
scanner, GEMS Hawkeye and Infinia SPECT/CT
scanners, Siemens E-CAM SPECT scanner, Philips
Skylite SPECT scanner, Siemens CT scanner and GEMS
MRI scanner. The archival PACS systems used in this
evaluation were the Siemens MagicStore PACS system
(with MagicView workstations) and a GEMS Centricity
v2 PACS (with Radworks workstations). Fusion
workstations used in this work were the GEMS Xeleris
workstations and the CTI-Mirada Reveal-MVS
workstation.
Results
While the PACS archive systems could adequately store
and move the functional image data, the associated
workstations were generally unable to properly interpret
and display the image data. The primary error in this case
was the misinterpretation of image data and/or image
header data, specifically the DICOM fields Rescale
Intercept (0028,1052) and Rescale Slope
(0028,1053).[1] Within vendor interoperability was better
but still less than optimal.
Network transmission of fusion data sets by traditional
PACS was frustratingly slow to the physicians reading
the cases. This was determined to be for two reasons: 1)
the primary methodology chosen by most vendors for the
transport of image data utilizes the image-by-image
method as specified by the DICOM standard. As fusion
data sets incorporate two or more times as many images
than typical diagnostic data sets, the number of network
transfers is similarly greater; and 2) depending on the
manner in which the images were acquired, the
segmentation of a study by series or study can also
significantly impact the queuing of studies for network
transmission at the source.
Discussion
The existing PACS environment failed us on several
counts. Both the network transmission of image data
sets, as well as the interpretation of components of the
data sets, was less than optimal. In both cases, the failure
was due to the lack of understanding on the part of the
PACS environment of a fusion data set. To compensate
for this, we incorporated a PACS within a PACS. Using a
GE Centricity 2 archive and our functional imaging
workstations, we built our own PACS environment for
our functional imaging group. We then connected this
‘fusion’ PACS to the larger, enterprise PACS. This
provided our functional imaging group with the
performance and tools required to read and analyze the
fusion data sets while also providing image data access to
users hosted on either PACS system. While the
workstations on the enterprise PACS can not perform
fusion analysis, they do have access to the image data
generated by the fusion scanners and can capture result
sets of fusion, e.g. screenshots as secondary captures, and
display them locally. Using query-spanning and moveforwarding
technologies, the ‘fusion’ PACS is able to
access and retrieve data sets for software fusion which
have been acquired outside the realm of the functional
imaging group, i.e. CT, MR scans.
Conclusions
Traditional PACS systems can be used in the
management of multi-modality data sets, though the
typical PACS workstation falls short of providing the
clinician the tools they require to properly read these data
sets. By creating a ‘fusion’ PACS within the larger
enterprise PACS environment, we found that we were
able to leverage the best of both environments and
provide the clinician with the data and tools required to
properly display and interpret the fusion data sets. Next
generation PACS architectures will need to these lessons
IFMBE Proc. 2005;9: 54

Telemedicine and patient data management
learned into account as both fusion scanners and the
fusion of multi-modality image data through software
becomes an increasingly standard practice in the review
and interpretation of diagnostic medical images.
References
LEAL, J.P., WAHL, R.L. (2003): 'Technical
considerations integrating PET/CT into a Traditional
PACS environment', Society of Nuclear Medicine Annual
Meeting, New Orleans U.S., 2003.
IFMBE Proc. 2005;9: 55

Neural engineering
BOLD SIGNAL ANALYSIS DURING ACOUSTIC STIMULATION OF THE
BRAIN AT 3 TESLA
S. Casciaro 1,2 , D. Zacà 2 , R. Bianco 2 , S. Neglia 2 , F. Esposito 4 , F. Di Salle 3 , A. Distante 1,2
1 Consiglio Nazionale delle Ricerche/Istituto di Fisiologia Clinica, Sezione di Lecce, Lecce, Italy
2 Istituto Scientifico Biomedico Euromediterraneo/Divisione di Bioingegneria, Brindisi, Italy
3 Dept. of Neurological Science, Pisa University, Italy
4 Division of Neurology, Second University of Naples, Italy
casciaro@ifc.cnr.it
Abstract: In this work a functional study of the
brain was carried out by means of MRI; in
particular, the response to auditory stimuli was
detected using the Blood Oxygen Level Dependent
(BOLD) contrast technique. A normal volunteer
underwent a series of short acoustic stimuli of
different duration, repeated at a long enough
interstimulus interval (ISI) to avoid signal
overlapping of two consecutive stimulations [1]. The
raw data were processed in order to reduce the
presence of noise in the signal, to draw brain
activation maps and to extract a temporal activation
function. The analysis of the latter showed that brain
activation grows with increasing stimulus duration
though in a nonlinear way.
Introduction
of each voxel with an ideal response function (Fig. 1),
obtained through convolution of the stimulus with an
impulse response function. As to the latter, the
functions hypothesized namely by Cohen [5] and Cox
[6] were considered; the results given by these models
were then compared. The signal of thirty voxels from
both brain emispheres showing highest correlation
(r>0.785) was averaged, giving rise to a new function
taken as index of the temporal brain activation.
Two parameters of this new curve were correlated
with the stimulus duration, i.e. the peak value and the
area underlying the curve itself [7]. Besides, the relation
between BOLD signal and stimulation was assessed in
terms of system linearity.
Finally, the brain activation pattern was analyzed
and compared with the ones foreseen by using the linear
models proposed by Cohen and Cox [5,6].
BOLD functional Magnetic Resonance Imaging
(fMRI) is a well-established non-invasive technique
able to detect the neuronal activation sites with an
optimum spatial resolution (1 mm) and a good temporal
resolution (1 sec) [2,3]. The aim of our experiment was
to investigate the brain response to acoustic stimulation
storing raw data obtained by MRI and analysing them
through fMRI signal and image processing techniques.
Materials and Methods
Real-time Echo Planar Imaging (EPI) scannings
were performed on a healthy volunteer with a GE Signa
3T MR scanner, while the acoustic stimulation was
provided to the subject by means of the commercial
software Presentation (Neurobehavioral Systems,
Inc). The subject wore headphones to allow the auditory
system to get the stimulation and at the same time to
reduce the gradient coil noise effects.
The auditory stimulation experiment was repeated
four times and the output signal was averaged to
improve the overall data signal-to-noise (SNR) ratio.
Data and images were then analyzed and processed with
AFNI © [4] suite software. The activation of the brain
regions was assessed by correlating the temporal signal
Fig. 1: Measured and simulated (Cox) responses
Results
The activation maps computed according to Cox and
Cohen’s functions of do not have any substantial
difference, demonstrating that these two models are
equally effective in revealing brain activation.
The activation patterns show a nonlinear behaviour
of the response peak and area versus the duration of the
stimulus. As shown in Fig. 2, brain activation intensity
for short stimuli duration is higher than if the response
was exactly linear, or, conversely, the response intensity
IFMBE Proc. 2005;9: 56

Neural engineering
grows less than linearly as the stimulus duration
increases.
they are an easy and useful means to make approximate
predictions in BOLD functional analyses. Thus, other
impulse response functions could be provided to better
reproduce the real experimental data.
Fig. 3: Time to signal curve for several stimulus
durations. The stimulus is intended to start at time t=0
and to last as long as indicated in the box.
References
Fig. 2: Plot of stimulus duration vs. curve area (top) and
peak intensity (bottom). The non-zero intercept for zero
stimulus duration indicates system nonlinearity
The nonlinearity of the system is revealed by the
different trends of the simulated and measured signals
(Fig. 1). In this case Cohen’s and Cox’ linear models
give only a rough indication about the signal temporal
course, failing, particularly, in the response to long
stimuli (4-5 sec), as the measured signal is up to 2,5
times less intense than expected with a linear model.
Discussion
Fig. 3 shows the activation curves for different
stimulus duration. It can easily be shown that these
output signals do not have the typical properties of
linear systems signal, e.g., scaling and superposition
[8,9].
This behaviour can be explained by considering that
the phenomena which give rise to the BOLD signal
change may have an inherent nonlinear nature, such as
neuronal activation, metabolism modifications, blood
flow changes and others [10].
Moreover, it has been shown that the nonlinear
character of BOLD signal depends on the kind of tissue
involved in the signal change, with neuronal activation
sites showing a higher degree of nonlinearity [10].
Conclusions
As already evidenced for other brain activation
regions such as primary visual [8] and motor [9] cortex,
linear models fail in giving a precise description of the
BOLD signal due to auditory cortex activation. Still,
[1] Hall DA, Haggard MP, Akeroyd MA, Palmer AR,
Summerfield AQ, Elliott MR,Gurney EM, Bowtell
RW (1999),“‘Sparse’ temporal sampling in auditory
fMRI”, Hum Brain Mapp, 7,pp 213–23
[2] KIM S.-G., LEE S.P., GOODYEAR B., SILVA A.C.
(2000), ‘Spatial Resolution of BOLD and Other fMRI
Techniques’ in MOONEN C.T.W. and BANDETTINI
P.A. (Eds.): ‘Functional MRI’ (Springer Verlag,
Berlin), pp. 195-203
[3] BANDETTINI P.A. (2000), ‘The Temporal Resolution
of Functional MRI’ in MOONEN C.T.W. and
BANDETTINI P.A. (Eds.): ‘Functional MRI’ (Springer
Verlag, Berlin), pp. 205-219
[4] COX R.W. (1996): ‘AFNI: Software for Analysis
and Visualization of Functional Magnetic Resonance
Neuroimages’, Comput. Biomed. Res. 29, pp. 162–
173
[5] COHEN M.S. (1997): ‘Parametric analysis of fMRI
data using linear systems methods’, Neuroimage, 6,
pp. 93-103
[6] AFNI, Internet site address: http://afni.nimh.nih.gov
[7] LIU H.L., GAO J.H. (2000): ‘An investigation of the
impulse functions for the nonlinear BOLD response
in functional MRI’, Magn. Reson. Imaging, 18, pp.
931–938
[8] VAZQUEZ A.L., NOLL D.C. (1998), ‘Nonlinear
aspects of the BOLD Response in Functional MRI’,
Neuroimage, 7, pp. 108-118
[9] BIRN R.M., SAAD Z.S., BANDETTINI P.A. (2001),
‘Spatial heterogeneity of the nonlinear dynamics in
the fMRI BOLD response’, Neuroimage, 14, pp. 817-
826
[10] PFEUFFER J., MCCULLOUGH J.C., VAN DE
MOORTELE P.F., UGURBIL K., XIAOPING H. (2003),
‘Spatial dependence of the nonlinear BOLD response
at short stimulus duration’, Neuroimage, 18, pp. 990-
1000
IFMBE Proc. 2005;9: 57

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THE MAGNITUDE-SQUARED COHERENCE IN THE DETECTION OF
STIMULATION/NO-STIMULATION TRANSITIONS
D. B. Melges, A.F.C. Infantosi, M. Cagy and A.M.F.L. Miranda de Sá
Biomedical Engineering Program, Federal University of Rio de Janeiro, Rio de Janeiro, Brazil
PO Box 68510, Zip Code 21941-972
dmelges@peb.ufrj.br, afci@peb.ufrj.br, mcagy@peb.ufrj.br, amflms@peb.ufrj.br
Abstract: The present work proposes the
Magnitude-Squared Coherence (MSC) application to
electroencephalographic (EEG) signals during
somatosensory stimulation (SEP) at the motor
threshold. The maximum response band was
identified (gamma band). Temporal evolution of
MSC at selected frequencies was evaluated to reflect
the rest-stimulation-rest transitions. The promising
results indicate this technique to be a useful tool to
monitor the patient neurophysiologic responsiveness.
Introduction
The Somatosensory Evoked Potential (SEP) has
been used in the clinical diagnostics to assess
neuropathologies and intra-operatively to monitor the
patient status in order to prevent against neurological
impairment. It has been applied during spine surgery,
thalamotomy, brachial plexus and pelvic fracture
surgery [1], and other clinical and intraoperative
applications. Its usefulness in the post-operative period
has also been pointed out in [2].
In spite of the optimistic results achieved with SEP,
morphological-based analysis, accomplished in surgery
setup, is subjective and highly dependent on the
specialist expertise. Hence, the development of
techniques that could objectively detect stimulus
response is necessary. The Magnitude-Squared
Coherence (MSC), an Objective Response Detection
(ORD) technique, has been applied to sensory potentials
evoked by photic [3] and somatosensory [4] stimulation.
Among others statistical tests, the MSC has achieved
prominent results in stimulus response detection. The
aim of this work is to verify the MSC efficacy in
identifying stimulation to no-stimulation (and
conversely) transitions.
Materials and Methods
SEP record was obtained (sample rate = 3 kHz) with
gold electrodes from Cz´-Fpz´ (Cz´: 2 cm posterior to
Cz; Fpz´: mid-point between Fpz and Fz) and C3´-C4´
(2 cm posterior to C3 and C4, respectively) from ten
adult volunteers. Stimulation was presented at the
nominal rate of 5 Hz (4.91 Hz to avoid stimuli rate
multiples coincident with 60 Hz and harmonics),
through 200 µs width pulses. Stimuli were applied to
the right tibial nerve at the motor threshold – intensity
that produces feet interior muscle involuntary
contraction.
Considering the linear model of Figure 1, the MSC
estimate was calculated as [5]:
M
∑
Yi
( f )
2
i=
1
ˆ κ ( f ) = (1)
M
2
M Y ( f )
∑
i=
1
where Y i (f) is i th window Discrete Fourier Transform
and M, the number of epochs used in the estimation. In
order to avoid stimulus artefact that is broad band,
stimuli-synchronized and follow the first 10 ms after
stimulus presentation, windows of 190 ms (spectral
resolution = 5.26 Hz) were used. Automatic artefact
rejection algorithm was used as described in [4] to avoid
high variance (low signal-to-noise ratio) windows.
Establishing the Null Hypothesis H 0 , Response
Absence, it can be shown that ˆ κ 2 ( f ) is related to the F
distribution with 2,2M-2 degrees of freedom. Hence, it
is possible to obtain the critical value for the estimate
for a given significance level (α) and the number of
epochs (M) as [4]:
ˆ κ
F
i
2
2,2M
−2,
α
crit =
(2)
M −1+
F2,2
M −2,
α
where F 2,2M-2 , α is the critical value of the above F
distribution for an α significance level. False positives
are expected at the rate α in all frequencies in the nostimulation
condition.
The maximum response band was visually evaluated
in order to select suitable frequencies for analysing and
the temporal evolution of some
Figure 1: Linear Model: x[n] is an impulse train, h[n] is
the system transfer function, s[n] is the stimulus
response, r[n] is the background EEG and y[n] is the
measured signal.
2
IFMBE Proc. 2005;9: 58

Neural engineering
frequencies was obtained as
exemplified in Figure 2. Horizontal
line indicates the critical value
2
ˆ κ crit = 0.0298, determined by
M = 100 and α = 0.05. When this
value is exceeded by ˆ κ 2 ( f ) , one
can reject H 0 and accept the
Alternative Hypothesis (Response
Presence).
Results
Figure 2: ˆ κ 2 ( f ) temporal evolution at f = 36.8 Hz for Cz´-Fpz´ of volunteer #1.
The horizontal line is the ˆ κ 2 crit = 0.0298 determined by M = 100 and α = 0.05.
In the no-stimulation condition, the detection rate
with ˆ κ 2 ( f ) was bounded to 5%, which reflects the
significance level adopted in this constant false-alarm
rate (CFAR) detector.
The number of volunteers for whom it was possible
to detect stimulus response is shown in Table 1. It was
often impossible to detect potential in both derivations
for the frequency range considered (19.64 - 54.01 Hz)
and no detection was obtained in 49.10 and 54.01 Hz.
Considering the 29.46 - 44.19 Hz range, stimulus
response was detected in 90% (at C3´-C4´) and 60% (at
Cz´-Fpz´) of the volunteers, and in 100% if both
derivations are considered together.
As exemplified in Figure 2, the temporal evolution
of ˆ κ 2 ( f ) at f = 36.8 Hz (volunteer #1) accompanies the
no-stimulus to stimulus transition (t = 881 epoch) and
conversely (t = 2346 epochs).
Table 1: Number of Response Detection for ten
volunteers, per frequency (Hz), in both derivations
19.64 24.55 29.46 34.37 39.28 44.19
Cz´-Fpz´ 1 1 3 4 3 3
C3´-C4´ 1 4 7 6 4 5
Discussion
The MSC application to the tibial nerve SEP was
demonstrated to be suitable to represent the reststimulation-rest
transitions. However, it is very
important to be aware of the derivation selection
problem, by considering the natural interindividual
variability. In fact, concern about the derivation
selection has been reported in [6].
During stimulation, when response detection is
verified, the maximum response band is consistently
identified in the gamma band for all volunteers.
Especially frequencies between 29.46 and 44.19 Hz
were capable of reflecting the patient condition
(Table 1). This finding agrees with [4], in which the
frequency range 30-50 Hz was established.
In this work, only response to the motor threshold
was investigated, although lower stimulus intensities
can be used as evidenced in [4].
Conclusions
The application of the MSC to SEP can be a useful
tool to assess, with a known false-positive rate (false
detection), the current patient status. Although it was
possible to detect the response in 100% of the
volunteers in the 20.46 - 44.19 Hz frequency range, by
considering both derivations, it would be interesting to
verify whether it is possible to improve individual
frequency detection. As suggested by [3], multiple
coherence could emphasize the stimuli-synchronized
activity and improve detection rate. Further, a large
casuistry should be considered to confirm the results.
Acknowledgement
To CAPES and CNPq for financial support.
References
[1] LINDEN, R.D., ZAPPULA, R., SHIELDS, C.B. (1997):
‘Intraoperative Evoked Potential Monitoring’, in:
CHIAPPA, K.H. (Ed.): ‘Evoked Potentials in Clinical
Medicine’, (Lippincott-Raven, NY), pp. 601-638.
[2] DONG C.C., MACDONALD D.B., JANUSZ M.T.,
(2002); ‘Intraoperative Spinal Cord Monitoring
During Descending Thoracic and Thoracoabdominal
Aneurysm Surgery’, Annals of Thoracic
Surgery, 74, pp. S1873-S1876.
[3] MIRANDA DE SÁ A.M.F.L., FELIX L.B., INFANTOSI
A.F.C. (2004): ‘A Matrix-based Algorithm for
Estimating Multiple Coherence of Periodic Signal
and its Application to the Multi-channel EEG
During Sensory Stimulation’, IEEE Trans. Biom.
Eng., 51, pp. 1140-1146.
[4] TIERRA-CRIOLLO C.J., INFANTOSI A.F.C. (2002):
‘Determinação da Banda de Máxima Resposta para
Estimulação do Nervo Tibial’, Anais do XVIII
Cong. Brasileiro de Eng. Biomédica, v. 5. São José
dos Campos, Brasil, 2002, pp. 476-479.
[5] DOBIE R. A., WILSON M. J. (1989): ‘Analysis of
Auditory Evoked Potentials by Magnitude-Squared
Coherence’, Ear and Hearing, 10, pp. 2-13.
[6] MACDONALD D.B., STIGSBY B., ZAYED Z.A. (2004):
‘A Comparison between Derivation Optimization
and Cz´-FPz for Posterior Tibial P37 Somatosensory
Evoked Potential Intraoperative Monitoring’,
Clinical Neurophysiology, 115, pp. 1925-1930.
IFMBE Proc. 2005;9: 59

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MULTIVARIATE SPECTRAL ANALYSIS APPLIED TO THE EEG DURING
RHYTHMIC STIMULATION – A COHERENCE-BASED APPROACH
A.M.F.L. Miranda de Sá, M. Cagy, D. B. Melges and A.F.C. Infantosi
Biomedical Engineering Program, Federal University of Rio de Janeiro, Rio de Janeiro, Brazil
PO Box 68510, Zip Code 21941-972
amflms@peb.ufrj.br, mcagy@peb.ufrj.br, dmelges@peb.ufrj.br, afci@peb.ufrj.br
Abstract: The present work proposes two
multivariate, coherence-based techniques for
evaluating the inter-relationship in the EEG during
sensory, rhythmic stimulation. The first is the
coherence between two signals that is due to the
stimulation signal and the latter, the partial
coherence after removing the stimuli contribution.
Such techniques have been developed based on the
fact that the stimulation signal is periodic. They were
tested through Monte Carlo simulations and an
example is provided in electroencephalographic
signals during stroboscopic photic stimulation.
Introduction
The coherence function is a frequency-domain
technique, which is analogous to the correlation
coefficient. Its squared-modulus (which is often called
just coherence) is not affected by a time delay between
the signals under investigation. This aspect, in addition
to frequency selectivity, turns coherence very useful for
evaluating the relationship in electroencephalographic
(EEG) signals, where the inter-relationship is known to
occur within certain frequency bands [1].
The extension of the concepts of coherence to the
multivariate case leads to multiple and partial
coherences. The first is the fraction of the power
accounted for in a given signal via linear relationships
with a group of other ones, and the latter, the coherence
between two signals after removing the linear
contribution from a set of other signals. Multiple and
partial coherences would thus provide a more complete
evaluation of the interrelationship in the multichannel
EEG [1].
However, in quantifying synchronization of EEG
signals due to rhythmic stimulation, the strong
coherence of background EEG from neighbouring areas
is a confounding influence. Thus high coherence
estimates at the frequency of stimulation does not
necessarily indicate a synchronized response to
stimulation, but could largely reflect correlation of the
spontaneous activity. A modified coherence estimate
has been proposed in [2] for measuring the degree of
synchronization due to stimulation. On the other hand,
in event-related synchronization/desynchronization
(ERD/ERS) studies, one aims at evaluating spectral
changes caused by stimulation but that are not
synchronized with it, i.e. that are time-locked but not
phase-locked to the stimuli.
In the present work, two multivariate, coherencebased
techniques are suggested for evaluating the interrelationship
in the EEG during sensory, rhythmic
stimulation. An example of these techniques is provided
with EEG data during stroboscopic photic stimulation.
Materials and Methods
Considering the two-input-one-output linear model
of Fig. 1, the modified coherence estimate for
measuring the degree of synchronism between the
output signals y 1 [k] and y 2 [k] accounted for by the
periodic input signal x[k] is given as [2]:
M
∑
∑
Y ( f ) Y ( f )
2
ˆκ
1,2
( f )
(1)
2
( f )
1i
2i
i=
1
i=
1
= ⋅
M
M
2
M∑
Y1
i
( f ) M∑
Y2i
i=
1
i=
1
2
where Y ji (f) (j=1,2) is i th window Fourier Transform of
y j [k], and M, the number of segments used in the
estimation. For the same model, the partial coherence
estimate expression between the output signals
removing the contribution from the input signal, can be
obtained by simplifying the general matrix expression
provided in [3] for the particular case when x[k] is
periodic. Thus, it may be expressed as:
ˆκ
⎡
⎢
⎢⎣
2
y1y2•
( f ) =
M
*
1
Y1
i
( f ) Y2i
( f ) −
M i
2
M
1 ⎤ ⎡
− ∑Y1
i(
f ) ⎥ ⋅ ⎢
M i 1 ⎥⎦
⎢⎣
*
∑ ∑Y1
i
( f ) ∑
i= 1 = 1 i=
1
M
2
∑ Y1
i(
f )
i= 1
=
M
M
M
2
Y ( f )
2i
1
−
M
M
M
2
∑ Y2i
( f ) ∑Y2i
i= 1
i=
1
2
( f )
where “*” superscript denotes complex conjugate. It is
interesting to note that both coherence estimates are
independent of the input signal. Critical values for
ˆ 2
1,2
f
ˆ 2
y 1y2•
f
2
⎤
⎥
⎥⎦
(2)
κ ( ) are provided in [2] and for κ ( ) can be
obtained with the multivariate extension of the
invariance of coherence statistics when one signal is
Gaussian and coherence is zero [4] as:
ˆ κ 2 crit = beta (1, 2)
(3)
y1y2• crit
M −
where beta crit (1,M–2) is the critical value of the beta
distribution with parameters p=1 and q=M–2 for a given
significance level.
The application of the techniques is illustrated on the
EEG signal (derivations O 1 and O 2 with ipsilateral
IFMBE Proc. 2005;9: 60

Neural engineering
earlobe reference) of a normal subject recorded over a
period of 24 seconds during stroboscopic flash
stimulation at 6 Hz. The signals were digitised at
ˆ 2
1,2
f
ˆ 2
y 1y2•
f
256 Hz. Next, κ ( ) and κ ( ) were obtained
according to expressions (1) and (2), respectively.
Simple coherence was also estimated. In all these
spectral estimates, the window length was set equal to
2 s, leading to M=12 data segments in each estimation.
ˆ 2
1,2
f
This led to a critical value (α=5%) of 0.057 for κ ( )
ˆ 2
y 1y2•
f
(from [2]) and of 0.259 for κ ( ) , according to
expression (3). Simple coherence critical value was
found to be 0.283 (from [5]).
n 1 [k]
x[k]
H 1
(f)
H 2 (f)
v 1 [k]
v 2 [k]
+
+
n 2 [k]
y 1 [k]
y 2[k]
ˆ 2
1,2
f
Figure 1: Linear model used in deriving κ ( ) and
ˆ 2
y 1y2•
( f
κ
) . x[k] is the stimulus, v 1 [k] and v 2 [k] are the
responses [output of the filters H 1 (f) and H 2 (f)], and
n 1 [k] and n 2 [k] are the contributions of background activity
to the measured EEG signals y 1 [k] and y 2 [k].
Results
ˆ 2
1,2
f
The simple coherence is higher than κ ( ) , but
the latter shows clearer peaks at the stimulus frequency
and its harmonics (Fig. 2). The former also shows
statistically significant coherence at almost all
frequencies displayed, whereas for the latter, significant
Coherence
1.0
0.8
0.6
0.4
0.2
0.0
0 6 12 18 24 30
Frequency
ˆ 2
1,2
f
Figure 2: Illustration with EEG. κ ( ) (continuous
ˆ 2
y 1y2•
f
line), κ ( ) (bold) and simple coherence (dotted line).
Critical values indicated accordingly in horizontal lines.
synchronisation is only observed at the harmonics of
stimulation (α=5%).
The major differences between partial coherence
ˆ 2
y 1y2•
f
estimate κ ( ) and simple coherence occur in the
stimulation frequency and its harmonics. The first
clearly reduces the peaks due to the stimuli but keeps
the remaining frequencies almost unchanged.
Discussion
ˆ 2
1,2
f
κ ( ) seems not to be affected by background
EEG activity, emphasizing the response at the stimulus
frequency and harmonics [2]. On the other hand,
ˆ 2
y 1y2•
f
κ ( ) informs about the similarity in two EEG
derivations during stimulation that is not due to the
stimuli. It is close to simple coherence in the remaining
frequencies. Thus, it has a high specificity.
Conclusions
ˆ 2
1,2
f
ˆ 2
y 1y2•
f
κ ( ) and κ ( ) provide complementary
information about the stimulation effect to EEG signals.
While the first quantifies the degree of synchronism due
to stimulation, the latter informs about the relationship
due to the background, non-phase locked activities.
Thus, they may be useful in the ERD/ERS studies.
Both coherence estimates presented are independent
of the periodic signal. This result is relevant, especially
for a transient-periodic stimulation, when the stimuli
amplitude may decrease very fast. Thus, they could be
useful for reducing random error in data acquisition.
Acknowledgement
To CNPq for financial support.
References
[1] NUNEZ, P.L., SRINIVASAN, R. et al. (1997): ‘EEG
coherency: I statistics, reference electrode, volume
conduction, Laplacians, cortical imaging, and
interpretation at multiple scales’, Electroenceph.
Clin. Neurophysiol., 103, pp. 499-515.
[2] MIRANDA DE SÁ, A.M.F.L., INFANTOSI, A.F.C.,
SIMPSON, D. M. (2001): ‘A statistical technique for
measuring synchronism between cortical regions in
the EEG during rhythmic stimulation’, IEEE Trans.
Biom. Eng., 48, pp. 1211-1215.
[3] OTNES, R.A., ENOCHSON, L. (1978): ‘Applied Time
Series Analysis, volume 1 – Basic Techniques’,
(John Wiley and Sons, New York), pp. 374-376.
[4] NUTTALL, A.H. (1981): ‘Invariance of distribution of
coherence estimate to second-channel statistics’,
IEEE Trans. Acoust. Speech, Signal Processing,
ASSP-29, pp. 120-122.
[5] MIRANDA DE SÁ, A.M.F.L. (2004): ‘A note on the
sampling distribution of coherence estimate for the
detection of periodic signals’, IEEE Signal
Processing Letters, 11, pp.323-225.
IFMBE Proc. 2005;9: 61

Neural engineering
SIMULATIONS OF RADIO-FREQUENCY LESIONS WITH VARYING
BRAIN ELECTRODE DIMENSIONS
J. D. Johansson 1 , J. Wren 2 , O. Eriksson 1/3 , D. Loyd 2 and K. Wårdell 1
1 Department of Biomedical Engineering, Linköping University, Linköping, Sweden
2 Department of Mechanical Engineering, Linköping University, Linköping, Sweden
3 Elekta Instrument AB, Sweden
E-mail: johjo@imt.liu.se
Abstract: Radio-frequency (RF) lesioning in the
brain was simulated using the finite element method
(FEM). Heating for 60 s with temperature control in
order to keep the tip at 80 °C was simulated. Length,
L, (2 – 4 mm) and diameter, D, (0.5 – 2.5 mm) of the
electrode tip were varied and the resulting lesion
volumes were used to calculate a regression model:
Lesion Volume = – 13.1D + 15.7L⋅D + 13.1D 2 mm 3 .
The results can be useful for electrode design and
prediction of lesion size.
Keywords: Radio-frequency surgery, Brain, Lesion size,
Electrode dimensions, Finite Element Method (FEM)
Introduction
Radio Frequency (RF) lesioning is an electrosurgical
method in which a radio frequent current from a suitable
electrode is used to thermally coagulate malfunctioning
tissue. It can among others be used for treatment of
severe chronic pain and symptoms from Parkinson’s
disease [1]. A thermocouple in the tip of the RFelectrode
is used to control the current in order to keep
the tip at a preset temperature, usually between 70 and
90 °C. It is important that the size of the resulting lesion
is correct so that the desired effect is obtained without
destroying too much tissue. In this study the influence
of the brain electrode dimensions on the lesion size was
investigated with simulations using the finite element
method (FEM).
Materials and Methods
Monopolar brain electrodes with varying active tip
length (L = 2, 3 and 4 mm) and diameter (D = 0.5, 1,
1.5, 2 and 2.5 mm), see figure 1, were modelled using
the finite element program FEMLAB 3.0 (Comsol AB,
Sweden). The material parameters are given in Table 1.
An electric potential, V, of 25 V was applied to the
electrode tip and a boundary sufficiently far away from
the electrode tip (30 mm) in order not to interfere with
the lesioning process was set to ground and to a constant
temperature, T, of 37 °C.
The thermal field was simulated using the heat
conduction equation [2] with resistive heating, Q, & [3] as
heat source:
Upper electrode
Electric Insulation
Electrode tip
Axis of Symmetry
2 mm
Diameter, D
0.5 - 2.5 mm
Length, L
2 - 4 mm
Figure 1: Electrode. Length, L, and diameter, D, of the
active electrode tip were varied in the models.
∂T
ρc = ∇ ⋅ ( k∇T
) + Q&
∂t
(W/m 3 ) (1)
Q & = J⋅∇V = σ(T)(∇V) 2 (W/m 3 ) (2)
ρ denotes mass density (kg/m 3 ), c specific heat capacity
(J/(kg⋅K)), t time (s), k thermal conductivity (W/(m⋅K)),
Q & power density (W/m
3 ), J current density (A/m 2 ) and
σ(T) temperature dependent electric conductivity. The
electric potential field was calculated using the equation
for steady current density [3]:
∇⋅J = -∇⋅(σ(T)∇V) = 0 (A/m 3 ) (3)
Table 1: Material Parameters Used, Mean Values From:
a) [4] b) [5] c) [6]
Grey matter
σ (S/m) ρ
(512 kHz) (kg/m 3 )
(T-20 °C) ⋅ 1.04⋅10 3
0.175 a) b)
c
k
(J/(kg⋅K)) (W/(m⋅K))
3.6⋅10 3 b) 0.35 c)
Upper - 4.75⋅10 3 0.70⋅10 3 9.0
electrode a)
Electric - 2.37⋅10 3 1.3⋅10 3 3.7
Insulation a)
Electrode
tip a) - 6.0⋅10 3 0.62⋅10 3 11.5
The heating was controlled so that a point in the
electrode tip corresponding to the position of the
thermocouple reached and held a preset temperature of
80 °C.
IFMBE Proc. 2005;9: 62

Neural engineering
The resulting volume of the produced lesion was
calculated as the volume of all tissue reaching at least
the temperature of 60 °C after 60 s of heating. The
width of the lesion perpendicular to the electrode was
also calculated. Finally, a regression model [7] with
lesion volume as a function of length and diameter of
the electrode tip was made.
Results
2 mm
60 °C
80
(°C)
Residuals (mm 3 )
37
Volume (mm 3 )
The results, which can be seen in table 2, yielded the
following regression model with 95 % confidence
intervals for the regression coefficients in parenthesis:
Lesion Volume = – 13.1(±1.9)D +
15.7(±0.4)L⋅D + 13.1(±0.7)D 2 (mm 3 )
The regression model is visualised in figure 2. An
example of a simulation and the residuals for the
regression model can be seen in figure 3. Our group has
obtained similar results with real electrodes in
gelatinous albumin test solution [8].
Table 2: Resulting Lesion Volume and Width as a
Function of Electrode Tip Diameter and Length.
Tip
Diameter
(mm)
Tip
Length
(mm)
Lesion
Volume
(mm 3 )
Lesion
Width
(mm)
0.5 2 12.8 2.7
1 2 31.2 3.8
1.5 2 58.1 4.9
2 2 89.7 5.8
2.5 2 127 6.6
0.5 3 20.9 3.0
1 3 45.9 4.2
1.5 3 80.7 5.2
2 3 121 6.2
2.5 3 167 7.1
0.5 4 30.1 3.2
1 4 62.4 4.4
1.5 4 103 5.4
2 4 153 6.5
2.5 4 207 7.3
Lesion Volume (mm 3 )
Tip diameter, D (mm)
Tip length, L (mm)
Figure 2: Surface plot of the regression model:
Lesion Volume = – 13.1D + 15.7L⋅D + 13.1D 2 mm 3 .
Figure 3: (left) An example of a simulation result
zoomed around the electrode tip (L = 2 mm, D = 1 mm).
(right) Residuals for the regression model.
Discussion
The diameter of the tip has clearly a much larger
impact on the lesion volume than the length has. This is
expected since the diameter affects the size of the
electrode in two dimensions while the length affects it
in only one. Increasing the diameter quickly increases
the volume of the lesion. However, a large diameter
electrode will cause more tissue damage on the way to
the target site. A thin electrode, on the other hand, tends
to be flimsier which can make precise targeting more
difficult. The results can be useful for lesion size
prediction when designing new electrodes.
References
[1] COSMAN, E. (1996), 'Radiofrequency Lesions', in
GILDENBERG, P. L. and TASKER, R., (Ed):
'Textbook of Stereotactic and Functional
Neurosurgery'. (Quebecor Printing, Kingsport) pp.
973-85
[2] CARSLAW, H. S. and JAEGER, J. C. (1959):
'Conduction of Heat in Solids', 2 ed. (Oxford
University Press, Oxford)
[3] CHENG, D. K. (1989): 'Field and Wave
Electromagnetics'. (Addison-Wesley Publishing
Company, Inc., USA)
[4] JOHANSSON, J. D., WREN, J., LOYD, D., and
WÅRDELL, K. (2004), 'Comparison between a
Detailed and a Simplified Finite Element Model of
Radio-Frequency Lesioning in the Brain', 26th Ann.
Int. Conf. of the IEEE Eng. in Med. and Biol. Soc.,
San Fransisco, USA. pp. 2510-13
[5] DUCK, A. F. (1990): 'Physical Properties of Tissue'.
(The University Press, Cambridge)
[6] CHATO, J. C. (1985), 'Selected Thermophysical
Properties of Biological Materials', in SHITZER A
and RC, E., (Ed): 'Heat Transfer in Medicine and
Biology, Analysis and Applications', vol. 2.
(Plenum Press, New York and London) pp. 413-18
[7] MONTGOMERY, D. C. (1996): 'Design and Analysis
of Experiments'. (John Wiley & Sons, Inc., USA)
[8] ERIKSSON, O., BACKLUND, E.-O., LUNDBERG, P.,
LINDSTAM, H., LINDSTRÖM, S., and WÅRDELL, K.
(2002): 'Experimental Radiofrequency Brain
Lesions: A Volumetric Study', Neurosurgery, 51,
pp. 781-88
IFMBE Proc. 2005;9: 63

Neural engineering
IMPROVING COHEN’S MODEL FOR BOLD FUNCTIONAL RESPONSE
OF THE AUDITORY CORTEX
S. Casciaro 1 , S. Neglia 2 , G. Palma 2 , D. Zacà 2 , R. Bianco 2 , E. Casciaro 1,2 , A. Distante 1,2
1 Consiglio Nazionale delle Ricerche/Istituto di Fisiologia Clinica, Sezione di Lecce, Lecce, Italy
2 Istituto Scientifico Biomedico Euromediterraneo/Divisione di Bioingegneria, Brindisi, Italy
casciaro@ifc.cnr.it
Abstract: Previous studies in our labs had
investigated the response of the auditory cortex to
stimuli with a duration in the range from 0.5 to 5
seconds. The peak values were chosen to detect the
dependance of the response from the duration of the
stimuli. As no expressions have been suggested in
literature, in this study, an analytical expression for
the trend of the data was developed and a
supralinear dependence was found. This observation
was used to improve Cohen’s model [1] in the sense
of attenuating the increase in the growth of the peak
values with the duration of the stimuli as to better
represent the trend of the real response.
The signal of 11 voxels showing highest correlation
(r > 0.785) was averaged, giving a new function taken
as reference curve of the brain activation. By comparing
the brain activation pattern with the ones foreseen by
using the linear models proposed by Cohen and Cox
Fig. 1 is obtained (it is evident that noise has already
been filtered from the measured response).
Introduction
Previous investigations [2, 3] have demonstrated a
non-linear dependence between the BOLD response
curves and the duration of the stimuli, but no other
model has been suggested to describe this dependence.
Our measurements were first compared with the
simulated answers calculated using the linear models by
Cohen and Cox [4]. It was shown that linear models are
inadequate for short duration stimuli (less than 6
seconds). This work focuses on the definition of an
analytical expression for the trend of the signal response
and on the definition of a new model in the non-linear
region.
Materials and Methods
The data were collected by scanning a healthy
volounteer with a GE Signa 3T MR while the acoustic
stimulation was provided to the subject by means of the
commercial software Presentation (Neurobehavioral
Systems, Inc) and through headphones. The latter also
allowed to reduce the gradient coil noise effects. The
activation of the brain regions was calculated by
correlating the temporal signal of each voxel with an
ideal response function (Fig. 1), obtained through
convolution of the stimulus with an impulse response
function. As to the latter, the functions used are those
hypothesized by Cohen and Cox; the resulting activated
regions given by using these models were though almost
identical.
Fig. 1: Comparison between the measured response and
the simulated response with Cohen’s model.
The modeling was then carried out with MATLAB®
(MathWorks, MA). First, for the peak values, different
kinds of curves were tested and then the one which gave
the better fitting parameters was chosen. In particular,
four values indicate the goodness of a fitting: Sum of
Squares due to Error (SSE), R 2 , Adjusted R 2 and Root
Mean Squared Error (RMSE). The one of greatest
interest was for us the R 2 , as this statistic measures how
successful the fit is in explaining the variation of the
data. A value closer to 1 indicates a better fit. This
means focusing attention on the behaviour of the
system, and therefore on the physical meaning of the
trend. Then, by carrying out some considerations
relating to Cohen’s model and the analytical funcion
obtained for the peaks, a new model was suggested.
Results
The fitting of the data, as described in the previous
paragraph, led to a power law for the peak values:
( x)
b
f = a ⋅ x + c;
a = −12.4,
b = −1.351,
c = 43.05
IFMBE Proc. 2005;9: 64

Neural engineering
The results are represented graphically in Fig. 2.
Fig. 2: Values of the peaks plotted against the duration
of the stimuli and fitted with MATLAB.
Very good statistics were found for this curve. Next,
the new model was defined. The result is given in Fig.
3:
these are originated by metabolical phenomena which
arise as a consequence of the reaction to stimuli.
Now, longer the stimuli, higher the consumption of
oxygen. Obviously, for longer stimulus durations, the
increase of the oxygenated blood flow will be higher, so
this must lead to an increase in the signal registered. On
the other hand, when the stimulus is very short,
probabily the “oxygenated blood excess” will be higher
than for longer stimuli, and this leads to a supralinear
dependence of the peak values from the duration of the
stimuli. When the stimulus lasts for long periods
(>6seconds, [5]), the blood flow is not so much higher
than the real necessity, so maybe for this reason the
BOLD signal increases almost linearly with the duration
itself. It must be also considered that a kind of
“saturation” of the vessels will take place which will
limit the maximum level of the BOLD response.
The second step in this study is the suggestion of a
new model for the response, which tries to better
represent the physical behaviour described and
confirmed by the trend of the peak values. In fact, the
greatest limit of Cohen’s model was the fact that it
didn’t show adequately the supralinearity of the growth
of the values of the peaks for very short stmuli. Our
model tries to limit this failure.
Conclusions
Fig. 3:Our model compared with Cohen’s and the
measured data.
The expression used to obtain this curve is given by:
hx ()= α ⋅ x β −γ ⋅x ⋅e − x δ
The most evident change compared to Cohen’s
model is the introduction of the term γx. This term
allows us to smooth out the increase in the values of the
peaks in the simulated response. The other parameters
allow us to better adapt the model to represent the
measured response.
Discussion
It is interesting to observe how consistent the trend of
the power law is with the nature of the physical
phenomenon that is being described. The BOLD
technique is based on the variations of the signal
obtained during the MR scan with the changes of the
oxygenated over deoxygenated haemoglobin ratio, and
A previous work [6] had suggested a neural
adaptation in the visual cortex for stimuli of duration
shorter than 3s. For the auditory cortex the limit may be
that of 6s, and the adaptation may take longer in this
region of the brain. Nevertheless, a linear model is very
simple to use, and this can be a great advantage to
exploit. In this direction, we have worked as if the
system was linear, trying to find an expression that
describes more accurately than Cohen’s the measured
curves.
As previously shown and discussed the method is
simple and succesfull.
References
[1] COHEN M.S., (1997), “Parametric analysis of fMRI
data using linear systems methods”, NeuroImage, 6,
pp 93-103;
[2] FRISTON KJ, JOSEPHS O, REES G, TURNER R, (1998),
“Nonlinear event-related responses in fMRI”, Magn
Reson Med.;39, pp 41-52.
[3] TALAVAGE TM, EDMISTER WB, (2004),
“Nonlinearity of FMRI responses in human auditory
cortex”, Hum Brain Mapp; 22, pp:216-28;
[4] AFNI, Internet site address: http://afni.nimh.nih.gov
[5] ROBSON MD, DOROSZ JL, GORE JC, (1998),
“Measurement of the temporal fMRI response of the
human auditory cortex to trains of tones”,
NeuroImage, 7, pp 185-98;
[6] VAZQUEZ AL, NOLL DC, (1998);”Nonlinear aspects
of the BOLD response in functional MRI”,
Neuroimage; 7, pp 108-118.
IFMBE Proc. 2005;9: 65

Neural engineering
PSYCHO-ACOUSTIC EXPERIMENTS OF THE PERCEPTION OF THE
MISSING FUNDAMENTAL
T. Matsuoka 1 , D. Konno 1
1 Information and Control Systems Science, Graduate School of Engineering, Utsunomiya University,
Utsunomiya, Japan
matsuoka@cc.utsunomiya-u.ac.jp
Abstract
The frequency band of one channel of a telephone
line is from 300Hz to 3400Hz. Although the pitch
frequency of speech is not in this frequency band, the
pitch frequency can be perceived over the telephone.
When we listen to a complex tone of f 1 =nf 0 and
f 2 =(n+k)f 0 , it is considered that the pitch ( f 0 ) is
produced in the auditory center. f 0 is known as the
missing fundamental. We made cochlear models -
Anteroventral Cochlear Nucleus models (AVCN
models), compared the output data, from the part of
an AVCN model, to physiological data, and
investigated the mechanism generating the missing
fundamental by using the models. The frequency
information of the missing fundamental of input
signals lower than 900Hz explicitly appeared in the
interspike-interval histogram of the aggregated
autocorrelogram of the output pulse trains from
AVCN models. To confirm it by the perceptual
experiments, we have carried out psycho-acoustic
experiments of the perception of the missing
fundamental. Close agreement between model
experimental results and psycho-acoustic
experimental results has been obtained.
Introduction
Several phenomena, which made clear by psychoacoustic
experiments, have no evidence of electric
physiological data. The missing fundamental
phenomenon is one of those phenomena. The missing
fundamental is not in the frequency band of one channel
of a telephone line but we can perceived it over the
telephone. It is considered that the missing fundamental f 0
is produced in the auditory center when we listen to a
complex tone of f 1 =nf 0 and f 2 =(n+k)f 0 . f 0 is known as the
missing fundamental.
Methods
We made a cochlear model and an Anteroventral
Cochlear Nucleus model[1],[2] as shown in Fig.1. Each
autocorrelogram of each output pulse train from the
AVCN model of right ear and the AVCN model of left
ear was made. The frequency information of the missing
fundamental of input signals lower than 900Hz explicitly
appeared in the interspike-interval histogram of the
aggregated autocorrelogram of the output pulse trains
from AVCN models as shown in Fig.2. In this report, the
psycho-acoustic experiments for confirming the
perception of the missing fundamental are carrid out. We
discuss the psycho-acoustic experimental results by
comparing to the model experimental results. The
procedure of the psycho-acoustic experiments is in Fig.3.
A stimulus tone series consists of stimulus #1 and
stimulus #2. At stimulus #1, a subject listens to tones of f 0
at both ears through a head phone. At stimulus #2, a
subject simultaneously listens to a tone of f 1 at his/her one
ear and a tone of f 2 at his/her another ear. It is for
avoiding that a subject perceives a combination tone.
Each subject listens to a stimulus tone series and says
whether stimulus #2 includes the same frequency
component to stimulus #1 or not.
Results
The results of psycho-acoustic experiments are shown in
Table 1. Each stimulus tone series is presented to a
subject five times in random order. Subjects are five
males and four females. The maximum number of replies
in Table 1 is 45. 16, 8, and 15 at bottom 3 lines ( f 1 and f 2
are higher than 900Hz) are significantly smaller than 31
at the top line ( f 1 and f 2 are lower than 900Hz). It means
that the perception of the missing fundamental of input
signal higher than 900Hz is difficult. These results (in
Table 1) of psycho-acoustic experiments agree with
experimental results (in Figure 2) of cochlear models -
Anteroventral Cochlear Nucleus models.
IFMBE Proc. 2005;9: 66

Neural engineering
Discussion
The synchronization index and the entrainment index of
output pulse train from ‘cochlear models and
Anteroventral Cochlear Nucleus models’ keep 0.9 until
1000Hz and 700Hz, respectively. These properties agree
with those from physiological data [2].
The frequency information of the missing fundamental of
input signals lower than 900Hz explicitly appeared in the
interspike-interval histogram of the aggregated
autocorrelogram of output pulse trains from ‘cochlear
models and Anteroventral Cochlear Nucleus models’. To
compare it to the perception of the missing fundamental,
we have carried out psycho-acoustic experiments. The
result of psycho-acoustic experiments agrees with that of
model experiments.
References
(Books)
[1] Greenberg S., Rhode W. S., Yost W. A., Watson
C. S., editors (1987) : “Auditory Processing of Complex
Sound”, Lawrence Erblaum Associates, pp.225-236
(Journals)
[2] Joris P. X., Smith L. H., Yin T. C. (1994) :
“Enhancement of Neural Synchronization in the
Anteroventral Cochlear Nucleus. I. Responses to Tones at
the Characteristic Frequency”, J. Neurophysiol., Vol.71,
pp.1022-1036
IFMBE Proc. 2005;9: 67

Biomedia
BIOMEDIA
A. Anani 1
1 Umeå University, TFE, Umeå, Sweden
adi.anani@tfe.umu.se
Abstract
Biomedia is a new discipline that deals with the
interaction between the human being and it’s
surrounding IT environment. In other words, how do
the computers and computerised equipment
surrounding the human body in the daily life
understand the human being and his behaviour?
physiology and behaviour is the principal source of
information to be picked up, analysed and understood by
the surrounding communicational and computational
means.
Introduction
Biomedia is about the interaction between human and its
environment with focus on how to understand human
beings from computing, communications, and interaction
points of view.
Within the discipline of Biomedia, one can tackle one or
more of a wide range of topics, such as facial expression,
body gesture, eye gaze, biometrics and biosignals.
The main issues are how to get biomedia signals, how to
characterize these signals, how much information can we
get, what kind of information and last but not least how
can we use it?
Biomedia means understanding human emotions. It is the
answer of daily social questions such as, how do you do?
Or how are you today? It means thus understanding the
human being in his daily life; when she works, when she
studies and when at home. It can thus be used in
healthcare, distance learning and when driving to
guarantee better life conditions. It can also be used for
security purposes to guarantee a secure life or, in the
worst case, the other way round. The applications are left
to the imagination.
Information Theory is needed to estimate the amount of
information obtained from the face, fingerprints,
biosignals etc. Moreover it is of great importance to study
how this information is transmitted to computers and how
it is processed by it. So, signal processing, image
processing, pattern recognition and communications are
inevitable pillars of this discipline.and communications
are inevitable pillars of this discipline.
Conclusions
Biomedia is a new discipline with a lot of interesting
applications in which the humanbeing itself in his
IFMBE Proc. 2005;9: 68

Biomedia
Figure Checker
Noor Anani* and Haibo Li
Digital Media Lab, Department of Applied Physics and Electronics,
Umeå University, Umeå, Sweden
*norani02@student.umu.se
Abstract: A simple and smart system to check
differences in a human body’s shape is designed and
implemented with out the need to use neither a scale or
a tape measure nor a skinfold. It can be used in
medical applications and also in physical fitness and
diet programs comparing changes in human figure.
This system, consisting of a web camera and a worked
up software based on image processing and dynamic
programming, is applied on a person with suitable
clothes showing the bodylines clearly. Two pictures
are taken at different points of time and then compared
to notice any difference taking place in the waist area.
This will help people who exercise or are on a diet to
keep fit while keeping an eye on their own figure.
Introduction
Biomedia systems are more common in medical
applications. The system suggested in this paper is a
live example on Biomedia where a surrounding
computing device interpretes any change in the human
figure as a result of loosing or adding fat.
A lot of people suffer from obesity and try various
ways to lose weight and get slimmer. The most
common way is to go on a diet, another is to do more
exercise and the most expensive way is to do a fat
suction.
We all know that exercising is the healthiest way to
lose weight. It’s known that by exercising we burn fat
but at the same time we gain in muscles which can be
very confusing for a lot of people. They get
disappointed when they see that their weight is either
the same or even higher than before although they
have been exercising for a while. Of course it doesn’t
have to mean that they gained more fat, they probably
gained in muscles. In reality they have become thinner,
but they lack an effective control method to realize
that. Unfortunately this can be a reason, good enough
for them to give up the process.
An easy example is that two people can have the same
BMI (Body Mass Index, is said to measure the fat in
the body) but a different percent body fat. A body
builder with a large muscle mass and a low percent
body fat may have the same BMI as a person who has
more body fat because BMI is calculated using weight
and height only [1]. This means that BMI doesn’t
really measure body fat.
Other way to check the body fat is with skinfold. It can
be done to measure decreases in body fat. The problem
with skinfold measurement is that they must be done by
somebody who is trained and experienced in this type of
measuring. Nevertheless, it is an accurate way to
measure progress in excess fat reduction [2]. The
purpose of this study is to develop an easy system for
people to have control over their figure.
Figure 1: These men have the same height, weight, and BMI.
Materials and Methods
In this project a testing person with very tight, thin
black clothes, a white squared piece of paper placed on
the stomach area stands against a white background and
is facing with a webcam. Matlab is used in the whole
experienment.
The testing person stands a distance from the camera
and behind her is the white background. The propose of
dressing such clothes in front of the white background
is to get a more exact and cleaner black-white image
after getting binarised from the programming. The
image is saved and after that the program measures the
first and the last black pixel in every row and saves that
information for later use. After a period of time the
exact same process is made on a new picture on the
same person and is saved for comparison later.
Now if the distance between the testing person and the
camera is longer in the second picture than in the first, it
may look as if she is smaller or slimmer. To avoid this
problem the white paper as a rigid reference is weared.
The images can be normalized by comparing the size of
the white square in the two pictures. In other words if
the size of the square is for example smaller the one in
the new picture then following normalization can be
done:
Square area in the first image = A 1 , square area in the
second image = A 2 , then zoomfactor s= A 2 /A 1. after that
the new image is zoomed in a factor of s into the same
size as the old image.
IFMBE Proc. 2005;9: 69

Biomedia
The comparison part will be done by the powerful
dynamic programming, which will try to fit the points
in the first picture with the points in the second picture
as good as possible. After that the distance between the
points in the left edge and the ones at the right edge
can be calculated and compared: Did she lose or gain
some fat?
Discussion
We believe that this system is more appropriate for
those who are on a diet or follow a fitness program, just
so it can motivate them. In other words when the person
thinks that he/she has become thinner/fatter, and only
wants to check. If the differences are too small it should
be neglected, because too small changes aren’t giving so
much of information.
The system developed here acn be used to measure
other parts of our body. We will focus on some parts
like: legs, chest, arms etc, where one can find more fat.
Conclusions
Figure 2: The first picture to the left, second to the right
Results
Here some experimental results are shown here. From
the right picture one can see that her body shapes have
changed: the distance between the first black pixels on
the left edge to the right edge is shorter than the one in
the new picture. One could almost see that she is
smaller or thinner now. Of course a more systematic
way to check by dynamic programming is under
development. We will report it during conference. The
goal of this project is to develop a functional system
that will show the temporal varying profile of human’s
waist figure.
The system can be very useful, for people who want to
get slimmer. It can motivate them in an easy way, when
they see their figure changes, instead of always having
to deal with the scales. This system is of course more
suitable in cases where the testing person hasn’t
dramatically changed in figure from the first picture to
the second. Besides, It is of great help when
measurements on sore patients are necessary.
References
(Electronical Publications)
[1] http://www.cdc.gov/nccdphp/dnpa/bmi/calc-bmi.htm
[2]http://www.mines.edu/~skimpel/measuringprogress.pdf
IFMBE Proc. 2005;9: 70

Biomedia
Abstract
A Hand Geometry Based Personal Verification System
Sani Anani and Haibo Li
Digital Media Lab, Department of applied Physics and Electronics,
Umeå University, Umeå, Sweden
mr_sani@hotmail.com
Biometrics is being used more and more in
security systems for identifying or verifying
persons especially after the September 11
event. In this study a low cost personal
verification System is developed based on
Hand Geometry. It is intended for civil
applications where a reasonably high
security level is sought.
Introduction
It is said that the 9/11 attacks has changed the
world, at least that’s the case in biometrics
security applications. Biometric security
systems use components common to machine
vision systems (cameras, A/D, edge
detection/thresholding, etc.) to compare
features of a person's anatomy to stored,
authenticated transform of the features in order
to confirm their identity [1]. This method of
identification requires the person’s physical
presence at the point-of-identification.
Identification based on biometric techniques
obviates the need to remember a password or
carry a token [2]. In this work the hand
geometry is used which involves the
measurement and analysis of the shape of one's
hand [3]. The aim of the study is to build a
low cost, reliable personal verification system
based on the hands geometry. Hand geometry
as a biometric identity is considered to be
acceptable by users, not associated with
criminology as fingerprints do and easy to
perform. Still, in spite of the fact that it is not
listed as a high security biometric identity, it is
covetable in civil applications where a
reasonable level of security is satisfactory. If a
false acceptance takes place, no catastrophe
will take place, but possibly a minor economic
loss. In case of a higher security requirement,
the method can be combined with other
parameters like passwords and PIN numbers as
in the normal case of a key card and a code,
only in this case the hand would be a key card
that you never forget at home because it
follows you where ever you go. It also requires
of the person to be physically present at the
point-of-verification which guaranties that only
you can get access.
Methods
The system described in this paper consists of
two main parts; a hardware part and a software
part. The first part is the physical part where
the user interacts with the system. It is a box
with a Plexiglas on which the user places his
palm. Under the glass there is a light source
where an effort has been made to make the
illumination as equally distributed as possible.
The other component in the physical part is a
web camera that captures an image of the
hand. The software part is where the image
processing takes place. After extracting the
features of the captured picture of the palm, the
Image is then processed in MATLAB and
thereafter saved in a database.
The interesting part of the captured image is
the edge, which is easily detected after some
image processing. The image goes through
three stages of processing. The first step is
noise reduction, which is achieved by using a
low pass filter. After the noise reduction the
processed image is binarized. Thus, it changes
from a colour image to a black and white
image (ones and zeros).
After binarizing the image it is easier to detect
the edge of the palm and get the profile, which
is the feature of interest. The edge is then
followed pixel by pixel from the beginning of
the edge point to its end saving the stages
between the pixels as different directions. This
data is then normalised so that the system
would recognise a person’s hand even if it
were turned some degrees to right or left. The
information is saved as a profile vector in a
database.
When a user is trying to access the system an
image of his hand would be scanned and going
through the stages of processing above so that
its vector can be obtained and compared with
the vectors already existing in the database to
IFMBE Proc. 2005;9: 71

Biomedia
see if this person’s hand exists in the database
or not. To compare a vector with the rest of the
vectors, dynamic programming is used.
Dynamic Programming is an approach
developed to solve sequential, or multi-stage,
decision problems [4] in this case to ease the
measurement of the vectors similarity in the
database.
Results
The system is a low cost functional one that
can verify a person’s identity by finding the
most similar hand image within a database.
hand in this system has replaced the card. It
does not require the person to put the hand for
scanning in the exact same position every time
as it is done in other hand verification systems
where pegs are used to insure the right
placement of the palm for scanning. The
system is user friendly not only in a practical
sense but even in a psychological sense
meaning that it is easier for the user to accept it
than a fingerprint based system, which usually
is associated with criminology.
Conclusions
The system presented in this paper is a
personal verification system that can be used in
the daily life without any problem. It can be
used as an access method to fitness centres,
school restaurants, buildings, clubs and other
places where only members of a defined group
are allowed to have access.
Figure 1: hand profile after binarization
Tests to evaluate the quality and the property
of the system will be performed. False
acceptance rate (FAR) and false recognition
rate (FRR) are among the main things to be
looked at.
References
[1]
http://www.machinevisiononline.org/p
ublic/articles/archivedetails.cfm?id=12
39
[2]
http://biometrics.cse.msu.edu/info.html
[3]
http://ctl.ncsc.dni.us/biomet%20web/B
MHand.html
[4]
http://benli.bcc.bilkent.edu.tr/~omer/re
search/dynprog.html
[5]
http://www.sbc.su.se/~per/molbioinfo2
001/dynprog/dynamic.html
Figure 2: Image after edge detection
Discussion
This system is a low cost verifying one
compared to card-based systems because the
IFMBE Proc. 2005;9: 72

Biomedia
A Physiological signal Based Personal Verification System
A pilot Study
Haris Begic* and Adi Anani
Digital Media Lab, Department of applied Physics and Electronics,
Umeå University, Umeå, Sweden
el00hbc@hotmail.com
Abstract: This paper presents an ongoing project
that investigates the possibilities of using the EMG
signals of the zygomaticus major muscle, when
activated by a laugh or a smile, as a biometric
identity. A possibility to verify or identify a user
through this area of skin contact in the cheeks may
arise in some applications where helmets are used.
obtained have no predefined pattern; it is therefore the
choice of using the Markov model is the most
appropriate. Matlab is used as a platform for
implementation.
Introduction
Human facial expressions are of interest to both
psychologist and psycho-physiologists. Facial
electromyography (EMG) is one of the ways to
measure emotional expressions such as fear, surprise,
happiness, sadness and anger. EMG stands for
electromyogram and is a measurement of the muscle
activity.
Facial EMG can be distinguished based on the activity
across specific facial muscles. Activities in the
zygomaticus major muscle tend to correspond with
positive emotions (happiness, surprise) and the
corrigator supercilii muscle tends to correspond with
negative emotions (sadness, anger).
The main objective of this study is to read human
physiological signals, EMG signals in this occasion, in
order to check the possibility of using it as a biometric
identity. To be more specific, the objective is to
investigate whether the EMG signal of the zygomatic
major can be used for personal identification or not.
Materials and Methods
The EMG signals of the zygomatic major are picked
up using surface electrodes applied on the skin area of
the face covering the muscle. These signals are the
input of this hypothetical biometric system. The
signals are then filtered to reduce the noise, amplified,
normalized and then stored in a database as identified
records. The database includes all the records of the
people who are members of the test group. When a
person having a record in the database would like to
check if she is verified or identified a new EMG record
is taken and compared with already existing records.
The comparison has to be simple and reliable. The
algorithm used in this work is based on the Hidden
Markov Model, HMM. The EMG signal patterns
Figure 1: The picture to the left is when the examined
person is in a passive state and the picture to the right is
when the same person is in a smile-laugh state.
Results
The preliminary results show that they are sensitive to
where the surface electrodes are placed. It can also be
added that the sizes of the zygomatic major may play a
decisive role. If so then the measurements between
different individuals may make a distinction. However,
the preliminary results based on the HMM will be
reported later.
Discussion
The method of using EMG signals is quite different
from other identification methods that are commercially
available; it may be somehow abstract but still very
interesting. Physiological signals, as biometric identity
is natural, personal and hard to fake.
One issue with this method is that it takes a lot of
collaboration from the person, which is going to be
identified. The method is individually based and
position dependendent. Misplacing the electrodes may
result in different signals, which can make the
identification more difficult
IFMBE Proc. 2005;9: 73

Biomedia
Conclusions
It is worthwhile to investigate the utilisation of
physiological signals as biometric identities. Like
fingerprints, hand geometry and iris pattern the
physiological signals are available all the time with the
humanbeing, with the only difference that they are
activated as long as there is a sign of life, they cannot
be cut, removed and probably impossible to fake.
References
http://biopac.com/
http://www.jor.se/sv/front.html
http://face-andemotion.com/dataface/general/homepage.jsp
http://face-and
emotion.com/dataface/facs/guide/FACSIV1.html#21971
1
IFMBE Proc. 2005;9: 74

Biomedia
IMPROVEMENT IN BRAILLE READING USING A FINGER COVER
Kouki DOI*, Satoko SHINOHARA**, Hiroshi FUJIMOTO***
* Graduate School of Human Science, Waseda University, Tokorozawa, Japan
** School of Human Science, Waseda University, Tokorozawa, Japan
*** Faculty of Human Science, Waseda University, Tokorozawa, Japan
*e-mail: koukidoi@akane.waseda.jp
Abstract: Recently, transparent-resinous-ultravioletcuring-type
(TRUCT) Braille signs have been
introduced for printing Braille together with visual
characters. However, it has been pointed out that
friction between the forefinger and the base material
affects TRUCT Braille reading ability. To reduce
this friction, we have used a polyester nonwoven
fabric finger cover and have examined its effect on
Braille reading ability. The subjects used were 12
Braille learners with acquired visual impairment,
who were asked to read randomly selected
characters with and without wearing the finger
cover. Most of the subjects could read TRUCT
Braille significantly faster when they wore the finger
cover than when they did not. This result suggests
that wearing a finger cover enables Braille learners
to read TRUCT Braille more efficiently, and that it
can be used as an assisting tool for Braille reading.
Introduction
Braille is an important communication medium for
the visually impaired, and even though the reading
speed for Braille is about a quarter that for visual
characters [1], highly experienced Braille readers can
read up to 200 letters per minute [2]. However, the
number of patients with acquired visual impairment has
increased in recent years, and this is mainly due to the
pigmentary degeneration of the retina and diabetic
retinopathy. Most of these patients cannot read Braille,
and intensive training is required to achieve even a
modest level of reading ability.
On the other hand, transparent-resinous-ultravioletcuring-type
(TRUCT) Braille signs are becoming
popular in Japan, especially when they are printed
together with visual characters. These signs are made by
screen-printing a resinous ink that is cured using
ultraviolet radiation. The screen-printing technique can
be applied to various base materials, such as paper,
metals, and plastics, on which the Braille dots are
printed, and TRUCT Braille signs have begun to be
used in public facilities, such as on tactile maps, ticket
vending machines, and on handrails. Naturally, it is
expected that beginners in Braille reading will utilize
these signs. However, it has been pointed out that the
friction between the forefinger and the base material
may affect TRUCT Braille reading.
We have carried out a study to investigate the effect
of using a finger cover to reduce friction during Braille
reading.
Materials and Methods
We conducted an experiment to compare the
readability of TRUCT Braille both when wearing a
polyester nonwoven fabric finger cover and when not
wearing a finger cover (Fig. 1). We chose polyester
nonwoven fabric as a finger cover because of its
durability, and when it absorbs moisture, the friction is
decreased. Twelve Braille learners with acquired visual
impairment participated as subjects in our experiments.
Their average age was 54.8 years (SD = 9.34 y). They
were asked to read randomly arranged Japanese TRUCT
Braille characters both when wearing and when not
wearing a finger cover. The characters were printed on a
laminate film, which is generally used as a base material,
because it protects the surface of the visual characters. It
is known that laminate film is base material on which
the forefinger cannot slide easily. The Braille used in
this study consisted of two different dot heights: 0.15
and 0.40 mm, with a Japanese standard distance
between the dots of 2.3 mm. The subjects were asked to
read the characters verbally for one minute, both with
and without wearing the finger cover, twelve times. The
reading speed and error rate were recorded, and the
reading speed was determined as the number of
characters read per minute. The error rate was displayed
as the percentage of reading mistakes. In each case, the
results were analysed statistically using two factors
(finger cover and dot-height) between the group
employing the analysis of the variance (ANOVA)
technique with Bonferroni multiple comparison test
corrections.
Results
The average reading speed and error rate were
measured, and the results are shown in Fig. 2. A
significant improvement in the reading speed and error
rate was observed when participants wore the finger
cover.
IFMBE Proc. 2005;9: 75

Biomedia
Reading speed
In the case of the 0.15 mm dot height characters, the
reading speed when the subjects wore the finger cover
was about twice that when they did not. In the case of
the 0.4 mm dot height characters, the reading speed was
1.37 times faster when they wore the finger cover than
when they did not. When the subjects did not wear the
finger cover, the 0.4 mm dot height character-reading
speed was 1.48 times faster than the 0.15 mm dot height
character-reading speed. In contrast, no significant
difference in reading speed between the two dot height
characters was observed when the subjects wore the
finger cover. These results show that most subjects
could read TRUCT Braille significantly faster when
they wore the finger cover than when they did not.
Error rate
The error rate for the 0.15 mm dot height characters
with the finger cover on was about one-third that for
readings with no finger cover on. The error rate for the
0.4 mm dot height characters with the finger cover on
was two-fifths that for readings with no finger cover on.
In the cases where no finger cover was worn, the error
rate of readings using the 0.4 mm dot height characters
was about half that for readings using the 0.15 mm dot
height characters. In the cases where the finger cover
was worn, the error rate of readings using the 0.4 mm
dot height characters was three-fifths that for readings
using the 0.15 mm dot height characters. These results
show that most subjects could read TRUCT Braille
significantly more correctly when they wore the finger
cover than when they did not.
Discussion
As the forefinger has an extremely sensitive tactile
sense, especially at the fingertip, it could be supposed
that covering the fingertip will disturb such a sense, and
surgeons and dentists, who must wear gloves to prevent
infection during surgery, have commented that the use
of gloves decreases their touch sensitivity [3].
In contrast, our results suggest that interposing a
finger cover between a Braille object and the forefinger
enhances the tactile sense. An interpretation of our
results could be that the signal for identifying each dot
is transmitted through the interface material, while the
noise due to the friction between the base material and
the skin is reduced, so that the effect is that the signal to
noise ratio is increased. In practice, most of the subjects
claimed that it was difficult to read TRUCT Braille
without wearing the finger cover because their
forefingers could not slide easily due to the friction.
Therefore, friction may obstruct the movement of the
forefinger, and so the forefinger may fail to receive
sufficient stimuli from the dots to be able to read them.
On the other hand, most subjects could read TRUCT
Braille significantly faster and more correctly when
wearing a finger cover than when not wearing a finger
cover, regardless of the character dot height. In
particular, the reading speed when wearing a finger
cover was about twice as fast as when not wearing a
finger cover for a character dot height of 0.15 mm. Our
results show that wearing a finger cover enhances the
readability of TRUCT Braille. The wearing of a finger
cover, therefore, may enable persons with acquired
visual impairment who are learning Braille to be able to
read TRUCT Braille more efficiently.
Conclusions
We have conducted an experiment to compare the
readability of TRUCT Braille both when wearing a
forefinger cover and when not wearing a forefinger
cover. Twelve persons with acquired visual impairment
who were learning to read Braille participated in our
experiment. The results show that most participants
could read TRUCT Braille significantly faster and more
correctly with a finger cover than without it. This result
suggests that wearing a finger cover may enable the
learners of Braille to read TRUCT Braille more
efficiently, and that it can be used as a Braille-reading
assistance tool for learners.
Acknowledgements
This research was partly supported by a Grant-in-Aid
for Scientific Research from the Ministry of Education,
Culture, Sports, Science and Technology of Japan
(16300187), and by the 21st Century Center of
Excellence Program,Japan Society for the Promotion
of Science.
References
[1] KIZUKA, Y., ODA, K. (1989): ‘A program for
teaching Braille based on a new theory of Braille
reading’, National Institute of Special Education
Bulletin., 3, pp. 49–56
[2] FOULKE, E. (1995): ‘Braille’, in MORTON, A.H.,
WILLIAM, S. (Ed): ‘The psychology of touch’,
(Lawrence Erlbraum Associates, Publishers, New
Jersey), pp. 219–233
[3] CUNNINGHAM, J.L, DELARGY, S.M., WARNOCK,
C.M. (1992): ‘Glove wearing in Northern Ireland and
an assessment of the loss of tactile perception’, Journal
of the Irish Dental Association., 39, pp. 12–14
IFMBE Proc. 2005;9: 76

Biomedia
HAPTIC VIBROTACTILE: DRIVER ASSISTANT SYSTEM
L. Liu 1
1 TFE, Umeå, Sweden
Abstract
Driving on a busy road is a critical job. Drivers need to
combine all senses to handle upcomimg events and
situations. The capability of observing road situations
through visual sensors is a strong requirement for future
driver assistance systems. The duty of driver assistance
systems is dedicated to reduce the number of fatalities
and severities of traffic accidents. Major challenge is
detecting and classifying pedestrians, and displaying
warning signals.
This paper presents a driver assistant system providing
tactual alert on detection of pedestrians. A tactile display
is embedded on the steer as a warning indicator for the
drivers. The system consists of three unites: infrared
camera, image analysis system, and tactile display. A
pedestrian detection system from infrared video is
performed, and it involves three steps: target area
selection, feature extraction, and pedestrian detection.
li.liu@tfe.umu.se
IFMBE Proc. 2005;9: 77

Biomedia
TACTILE VIDEO
L. Liu 1
1 TFE, Umeå University, Umeå, Sweden
Abstract
It is possible to substitute one type of malfunctioned
sense with another. For the blind people they can not see
just because of the lacking of vision. One of the ways to
utilize a range of perception capabilities of their body, is
to sent information by means of different senses through
touch, auditory and thermal feeling
Our prime objective is to enhance the quality of life for
the visually impaired people. A tactile video system is
therefore developed by taking advantage of the fingertip.
The system contains three units: imaging system, image
analysis system, and tactile display system. Tactile video
system allows the interfering and interpretation of visual
information via a video camera. The visual information is
then sent to the blind as tactual patterns at the fingertip.
The work will result in a specific system providing
navigational aid and conquering the visual recognition
problem faced by visually impaired. This device will use
piezoelectric actuators affecting local nervous tension
patterns.
li.liu@tfe.umu.se
IFMBE Proc. 2005;9: 78

Cardiovascular engineering
THE INTELLIGENT STETHOCOPE AS A TOOL
IN MODERN HEALTH CARE
P. Hult*, C. Ahlstrom*, L. Rattfält*, C. Hagström**, N-E. Pettersson** and P. Ask*
* Department of Biomedical Engineering, Linköping University, Linköping, Sweden
** Biomedical Engineering, Örebro University Hospital, Örebro, Sweden
peras@imt.liu.se
Abstract: We are developing the intelligent
stethoscope as a powerful instrument in distributed
health care. Taking advantage of the rapid
development in computer technology we intend to
improve the nearly two hundred year old
stethoscope. Advanced signal processing is used to
relate the sounds to specific pathologies of the heart
and lungs. Implemented preferably in a PDA we can
store and process acoustic recordings from the
patients and, if necessary, communicate the signal
for external expert advice.
Introduction
We envision the intelligent stethoscope as a device
combining the classic stethoscope with modern
information technology. The bioacoustic technique is
basically simple and robust, and therefore appropriate in
a distributed health care system. The intelligent
stethoscope is very suitable for telemedicine
applications. The bioacoustic signal is recorded on site,
no matter where the patient is. Decision support via
advanced signal processing methods could be consulted
locally using portable computers, and the collected data
could also be uploaded to an electronic health record via
internet, GPRS or alike. The uploaded sounds could
then be interpreted by an expert at the specialist clinic at
the hospital [1].
Mechanical processes in the body produce sounds
which indicate the health status of the individual. The
most important body sounds are heart sounds and lung
sounds, but sounds from swallowing, micturition,
muscles and arteries also have clinical relevance.
Previous works in bioacoustic signal processing
mainly focus on presenting information about the
signals rather than classification and interpretation.
However, various methods and algorithms for lung
sound analysis [2] and heart sound investigations [3]
have been presented.
As auscultation skills amongst today’s physicians
tend to get worse [4], the need for simple and robust
investigative techniques grows stronger. The intelligent
stethoscope is such a technique, providing extended use
of auscultation amongst medical personnel. A
prerequisite for this development is the increasing
availability of cheap, compact and portable computers,
powerful enough to execute the decision support
algorithms and flexible enough to provide the necessary
communication tasks. In a future where home nursing is
likely to be a big part of health care, the intelligent
stethoscope will be a valuable tool. Our aims are to:
1. Develop tailored signal processing methods to
extract clinically relevant information.
2. Develop telecommunication solutions to transfer
obtained data to a centrally placed database or
electronic health record.
Methods and Results
Most of our bioacoustic research has been
concentrated in two fields: lung sound analysis and
heart sound analysis. This is a short review of our
previous and ongoing research within the scope of the
intelligent stethoscope.
Automatic detection of the third heart sound: The 3 rd
heart sound is clinically important due to its relation to
heart failure/myocardial insufficiency. Our automatic
method for detection of the 3 rd heart sound was based on
a tailored wavelet approach [5], capable of utilising both
time and frequency information from the signal at the
same time, see Fig. 1. The method investigated four
frequency bands of the phonocardiogram, 17, 35, 60 and
160 Hz. These frequency bands were compared
according to rules compiled from knowledge about
signals. The method was verified in a study on heart
failure patients and proved capable of detecting a
majority of the 3 rd heart sounds.
Automatic timing of respiration phases: The aim of
this work was to develop a method for respiration
monitoring, where the start and stop of the respiration
phases could be timed accurately [6]. A microphone
applied over trachea measured sounds induced by
turbulent airflow, and the method is hence based on a
Figure 1: The wavelet transform of one heart cycle in a
phonocardiogram.
IFMBE Proc. 2005;9: 79

Cardiovascular engineering
direct measurement of the flow. The analytical method
used was a moving FFT summation capturing features
of the frequency content of the respiration sounds. The
accuracy of the respiration phase detections was 50 ms
and the algorithm could accurately detect and separate
inspiration and expiration.
Characterisation of heart murmurs: Heart murmurs
are generated by turbulent blood flow and occur when
the blood flow exceeds the Reynolds number. The main
reasons for generation of murmurs are high rates of flow
through the valves, flow through constricted valves
(stenosis) or backward flow through incompetent valves
(insufficiency). A more turbulent blood flow induces a
more chaotic sound, which is reflected in the recorded
acoustic signal. The degree of self-similarity, and hence
chaos, is represented by the object’s fractal dimension -
a measure of a signals complexity in terms of
morphology, entropy, spectra or variance in the time
domain [7]. We have used features from a time
dependent variance fractal dimension to extract features
vectors. Preliminary classification of various heart
murmurs has been conducted using cluster analysis, and
the results are indicated in Fig. 2.
Analysis of adventitious lung sounds: Diagnosis
based on lung sounds is a complex task. Differentiating
features depend on characteristics of the patient as well
as on the environment, and some signs are lost with, for
instance, a cough. This project investigates the
nonlinear properties of adventitious lung sounds (sounds
superimposed on normal breath sounds) in comparison
with normal lung sounds. The analysis is restricted to
two abnormal sounds, wheezes and crackles. Wheezes
are musical lung sounds common in patients with
obstructive airway diseases such as asthma or COPD.
Crackles are discontinuous, explosive and transient in
character. Typical diseases where crackles are present
are alveolitis, emphysema and COPD. Our
investigations indicate that different adventitious lung
sounds have different appearance in phase space. These
differences could be visualised in a recurrence plot
where ordinary breath sounds are represented by a flat
recurrence plot, wheezes by diagonal lines and crackles
by vertical lines. The intention of the project is to find
AS
Normal
Figure 2: Results from cluster analysis of the feature
vectors provided by the fractal dimension. Squares
indicate aortic stenosis, circles indicate mitral
insufficiency and filled circles indicate normals.
MI
relevant discriminating features to classify different
diseases such as heart failure and COPD based on their
lung sounds.
Discussion
The stethoscope has been one of the most important
clinical instruments for a long time, and auscultation is
probably the most common method for diagnosis in
primary health care. Further development of the
stethoscope has not made any considerable progress, but
bioacoustic methods have now reached a stage where it
is possible to develop a new and modern instrument.
This new instrument has the potential to include both
advanced signal analysis, decision support and
telecommunication abilities. By adding communication
technology such as blue-tooth or GPRS, wireless
transfer from the stethoscope is possible. A promising
scenario is where medical personnel caring a patient in
its home can transfer a signal instead of an individual. A
physician at the hospital makes the diagnosis and
decides about the treatment which is carried out in the
patients’ home.
Conclusions
Our new signal processing algorithms gives us a
powerful tool for relating bioacoustic signals to specific
cardiac and pulmonary pathologies. The data from an
intelligent stethoscope can be stored as an acoustic
patient record, for diagnostic support and for external
expert advice in a distributed and home heath care
system.
References
[1] DAHL L., HASVOLD P., ARILD E., HASVOLD T.
(2003): 'Kan hjertebilyder evalueres med
telemedisin?' Medisin og vitenskap, 123 pp. 3021-3.
[2] PASTERKAMP H., KRAMAN S. S., WODICKA G. R.
(1997): 'Respiratory sounds. Advances beyond the
stethoscope', Am J Respir Crit Care Med, 156 pp.
974-87.
[3] OLMEZ T., DOKUR Z. (2003): 'Classification of
heart sounds using an artificial neural network',
Pattern Recognition Letters, 24 pp. 617-29.
[4] MANGIONE S., NIEMAN L. Z. (1997): 'Cardiac
auscultatory skills of internal medicine and family
practice trainees. A comparison of diagnostic
proficiency', Jama, 278 pp. 717-22.
[5] HULT P., FJALLBRANT T., WRANNE B., ASK P.
(2004): 'Detection of the third heart sound using a
tailored wavelet approach', Med Biol Eng Comput,
42 pp. 253-8.
[6] HULT P., FJALLBRANT T., WRANNE B., ENGDAHL
O., ASK P. (2004): 'An improved bioacoustic
method for monitoring of respiration', Technol
Health Care, 12 pp. 323-32.
[7] KINSNER W. (1994): 'Batch and real-time
computation of a fractal dimension based on
variance of a time series', Dept. of Electrical &
Computer Eng., University of Manitoba, pp. 1-22.
IFMBE Proc. 2005;9: 80

Cardiovascular engineering
A DSP SYSTEM FOR REAL-TIME ANALYSIS OF PERIPHERAL
VESSELS FROM SEQUENCES OF ECHOGRAPHIC IMAGES
F. Faita (*) , V. Gemignani (*) , M. Demi
(*) (**)
(*)
CNR Institute of Clinical Physiology, Pisa, Italy
(**)
ESAOTE SpA, Florence, Italy
{f.faita, gemi, demi }@ifc.cnr.it
Abstract: Computerized measurement systems are
currently used to monitor the diameter of arterial
vessels over time. These systems provide accurate
and robust measurements but they require a tedious
off-line analysis. The system we present in this paper
is a video processing board which performs realtime
measurements of an artery diameter, thus
providing the physician with an immediate response
while the examination is still in progress. The board
is based on the digital signal processor
TMS320C6415 and the algorithm is based on a
mathematical operator derived from the
generalization of the first absolute central moment.
Introduction
The endothelium is the tissue that lines the lumen of
all blood vessels but overall it is an organ that
synthesizes and releases vasoactive substances which
regulate vascular functions. A dysfunction of this organ
is an early step in the development of atherosclerosis
and is a useful indicator for the prediction of cardiac
events. For these reasons, the characterization of the
endothelial function is an attractive research topic in
modern vascular medicine. The examination consists in:
i) the application of a mechanical stimulus and ii) the
use of ultrasound equipment to measure the variation of
the vascular diameter.
Despite its widespread use, this technique has some
limitations due to the difficulties in obtaining an
accurate measurement of such a small vessel (3 to 5
mm) by using ultrasounds. Manual measurements on
the images using calipers proved to be difficult and too
time consuming for the analysis of long duration studies
(typically 5-10 minutes). To overcome these problems,
computerized systems have been developed which
automatically compute the diameter of the vessel [1].
They acquire the video signal from the ultrasound
equipment and subsequently analyze the sequence of
images. These systems provide accurate and robust
measurements but they require a tedious off-line
analysis.
The system we present in this paper can measure the
diameter of an artery in real-time thus providing the
physician with an immediate response while the
examination is still in progress. The main part of the
system is a video processing board based on a high
performance DSP. For every image, that is, at a rate of
25 frames/sec, the DSP automatically locates the two
borders of the vessel, computes the diameter and
displays the results on a user interface.
A new mathematical operator derived from the
generalization of the first absolute central moment is
used to automatically locate the two borders of the
vessel. The first absolute central moment belongs to the
wide class of moments of n order, which includes
variance, skewness and kurtosis. However, the first
absolute central moment has not been analyzed in depth
in the past and, in particular, its properties have never
been exploited in image processing. It is common
opinion that the first absolute central moment has not
been investigated because of the mathematical
difficulties introduced by the presence of the absolute
value which makes theorem proving difficult [2].
Materials and Methods
Given a gray level map f(p) of an image, the
generalized first absolute central moment e(p) and its
mass center b(p) can be written as follows:
where g(q,σ i ) are discrete Gaussian functions which are
normalized over circular domains Θ i with radius r i =3σ i .
Let us consider a straight gray level discontinuity with a
step profile. Vector b(p) always indicates the direction
of the path that joins point p to the nearest point of the
discontinuity and, when configurations with σ 1 >σ 2 /π
are chosen, b(p) indicates a point which is always closer
to the discontinuity than point p [3]. Consequently,
when the points p i of two approximate starting borders
IFMBE Proc. 2005;9: 81

Cardiovascular engineering
are given, the respective points of the vessel borders can
be localized by iteratively computing vectors b(p i ).
Once the borders of the vessel are determined on the
first frame of the sequence the latter can be used as
approximate starting borders to localize the borders of
the vessel on the second frame of the sequence and so
on. Therefore, when two approximate starting borders
are given on the first frame, the vessel borders along
with its diameter are automatically computed trough the
sequence.
Results
In Fig.1 a direct comparison between the
measurements of the diameter of a brachial artery on
1100 frames is shown. The two sets of measures were
obtained with our system and with a gold standard
method [1], respectively. The degree of correlation is
high and the slope is not significantly different from 1.
The intercept is different from zero since the two
methods compute the diameter of the vessel in different
ways. Our method finds the edge of the vessel at the
lumen-intima interface while the second finds the edge
of the vessel in proximity of the adventitia. In Fig.2 the
flow-mediated response (FMD) of a brachial artery to a
mechanical stimulus is shown.
compact and affordable apparatus which can be used
with any ultrasound system provided with a standard
video output.
Conclusions
A validation of the results with respect to both
manual measurements and automatic measurements is
still in progress. However, preliminary tests in three
clinical centers show that the system provides very
accurate measurements and that it is a remarkable step
forward towards a more systematic evaluation of FMD.
It is worth noting that the availability of the results in
real-time allows the sonographer to keep the exam
under control by correcting the patient movements
which could invalidate the final result. Indeed, due to
the reduced diameter of the artery, even small
movements of the patient’s neck can give rise to serious
artifacts.
References
Fig.2: Brachial artery FMD response
Fig.1: Gold standard method versus automatic system
Discussion
The video processing system we used to implement
the real-time measurement of the diameter is a
standalone video processing board [4]. The main
component is the Texas Instruments’ TMS32C6415, a
high performance DSP which is particular suited for
video processing applications. Its CPU, which has a
Very-Long-Instruction-Word (VLIW) architecture, can
carry out eight 32-bit instructions/cycle at 600MHz
clock rate, that is 4.8 billion instructions/second. The
board is equipped with an analog video decoder which
can acquire the most common video standards, a multichannel
analog I/O module, a USB interface, an RS232
serial interface and a considerable amount of memory:
512 Kbytes flash memory; 4 Mbytes synchronous
SRAM and 512 Mbytes DRAM. The hardware solution
proved to be a very powerful device in carrying out the
algorithm which is necessary to measure the diameter of
the vessel. Furthermore, the board proved to be a very
[1] F. Beux, S. Carmassi, M.V. Salvetti, L. Ghiadoni, Y.
Huang, S. Taddei, A. Salvetti. “Automatic
evaluation of Artery diameter variation from
vascular echographics images” Ultrasound in Med.
& Biol., vol 27, no. 12, pp. 1621-1629, 2001.
[2] Press W.H., Flannery B.P., Teukolsky S.A.,
Vetterling W.T.: Numerical Recipes, Cambridge
University Press, 1991.
[3] Demi M.: “On the Gray-Level Central and Absolute
Central Moments and the Mass Center of the Gray-
Level Variability in Low Level Image Processing”,
Computer Vision and Image Understanding, vol.97,
issue 2, pp.180-208, 2005.
[4] F. Faita, V. Gemignani, M. Giannoni, A. Benassi.
"A fully customizable DSP based system for realtime
imaging", Proc. of International Signal
Processing Conference – GSPx, Dallas – Texas,
pp.1-5, 2003.
IFMBE Proc. 2005;9: 82

Cardiovascular engineering
MYOCARDIAL PERFUSION ASSESSMENT USING AN ECG TRIGGERED
LASER DOPPLER TECHNIQUE
C. Fors 1 , M.G.D. Karlsson 1 , H. Casimir-Ahn 2 and K. Wårdell 1
1 Department of Biomedical Engineering, Linköping University, Linköping, Sweden
2 Linköping Heart Centre, University Hospital, Linköping, Sweden
carfo@imt.liu.se
Abstract: A new method to assess myocardial
perfusion during and after heart surgery has been
developed. Laser Doppler perfusion monitoring is
used in combination with ECG triggering to
minimize movement artifacts in the recorded signal.
The method has been evaluated during coronary
artery bypass surgery and interesting findings from
the measurements are presented. The evaluation
proved the method to be feasible for measuring
myocardial perfusion on the beating heart, provided
that the ECG is sufficient for triggering.
Introduction
Myocardial perfusion monitoring of patients that
undergo coronary artery bypass graft (CABG) surgery
could give valuable information about the heart’s
condition and, in the postoperative phase, enable early
detection of graft failure. A method for this purpose has
recently been developed and evaluated during surgery
[1]. It is based on laser Doppler perfusion monitoring
(LDPM), which is an established technique to assess
microvascular blood perfusion [2]. A drawback with
LDPM is the high sensitivity to tissue motion not
related to blood cells. In order to facilitate beating heart
measurements, the ECG is simultaneously acquired and
analyzed to select intervals in the cardiac cycle with
minimum tissue movements.
The overall project aim is to enable postoperative
monitoring of myocardial perfusion by a small probe
inserted into the myocardium through the chest wall. In
this paper the method is presented and some examples
from the measurements are shown.
Materials and Methods
A small and lightweight intramuscular probe (∅ =
0.6 mm) is inserted 3–5 mm into the myocardium. Two
optical fibres guide low power He-Ne laser light
bidirectionally between the probe tip and an LDPM
device. The perfusion signal generated by the LDPM
device and the ECG are continuously acquired by a
computer and processed by software developed in
LabVIEW (National Instruments Inc., USA).
Low myocardial tissue velocity has been found in
late systole and late diastole [3]. In order to determine
the perfusion signal in these intervals the ECG is
processed. The software identifies the T and P wave
peaks and calculates the perfusion signal as an average
over an interval of 10 ms starting 20 ms before each
peak. The signals are denoted Perf LateSyst and Perf LateDias ,
see Figure 1. The system is further described in [4].
Measurements have been performed during CABG
surgery (n = 13), both on the normal beating heart as
well as on the arrested heart during extracorporeal
circulation (ECC) and in the transitions between these
two states.
Results
A typical perfusion signal is shown in Figure 1.
Rapid variations and high amplitudes are found around
the QRS complex and in early diastole. In late systole
and late diastole, where Perf LateSyst and Perf LateDias are
calculated, the signal is low and steady.
Perfusion signal (a.u.)
12
10
8
6
4
2
0
Perf LateSyst
Perfusion signal
ECG
Perf LateDias
0 0.2 0.4 0.6 0.8 1.0
Time (s)
Figure 1: Two perfusion values are calculated every
heart-beat as averages over 10 ms.
Time traces of the perfusion signal in late systole
and late diastole are shown in Figure 2. These signals
are from a baseline measurement in the end of the
surgery, on the normal beating heart. The periodical
variations in the signals are in accordance with the
breathing frequency of the ventilator (f ≈ 16 min -1 ).
IFMBE Proc. 2005;9: 83

Cardiovascular engineering
Figure 2: Time traces of the perfusion signals measured
on the beating heart in the end of the surgery.
Figure 3 illustrates the transition to extracorporeal
circulation. The perfusion signals decrease after the
onset of ECC. Aortic cross-clamping and injection of
cardioplegia cause cardiac arrest and blood-emptying of
the myocardium. Perfusion values are calculated once a
second when the heart has stopped (from heartbeat no.
84).
Perfusion signal (a.u.)
Perfusion signal (a.u.)
5
4.5
4
3.5
3
2.5
2
1.5
1
0.5
6
5
4
3
2
1
0
0 5 10 15 20
Heartbeat
Figure 3: Perf LateSyst and Perf LateDias during the transition
to extracorporeal circulation.
Discussion
Onset of ECC
Perf LateSyst
Perf LateDias
Perf LateSyst
Perf LateDias
Cardioplegia, aortic
cross−clamping
0
0 10 20 30 40 50 60 70 80 90 100
Heartbeat
The method described aims to facilitate trend
monitoring of myocardial perfusion. Some examples
from measurements during CABG surgery have been
shown to illustrate the potential.
The rapid variations and high amplitudes in the
perfusion signal in early systole and early diastole are
assumed to be related to movements caused by the
heart’s contraction and relaxation, respectively. By
calculating perfusion signal values in late systole and
late diastole, the movement artifact contributions are
assumed to be at a minimum.
Figure 3 shows the difference in perfusion signal
level when measuring on the beating heart compared to
the blood-empty, arrested heart. The decrease of the
perfusion signals after the onset of ECC may be
explained by a reduced workload of the heart. The
transition from pulsatile to continuous blood flow may
also have an influence on the perfusion signals.
The method described requires that the ECG T and P
waves can be identified. Some CABG patients may,
however, have pathological ECGs where the T and/or P
waves cannot be detected, even by the naked eye.
Besides, a pathological ECG may imply abnormal
myocardial tissue motion. Therefore, other ways of
identifying intervals with low tissue velocity will be
investigated.
It is known that breathing has hemodynamic effects
[5], and influence from the ventilator on the perfusion
signal has been found. Exactly how the breathing is
related to the perfusion signal is not yet elucidated, but
it is an interesting finding that will be analyzed in detail
later on.
The evaluation during CABG surgery has shown that
the method is feasible for measuring myocardial
perfusion of the human heart, provided that the T and P
wave peaks can be identified. The next study will be
performed on CABG patients in postoperative care. The
probe will be left in the myocardium for up to 24 h after
the surgery and the perfusion signals will be monitored
online during this time.
References
[1] KARLSSON M. G. D., FORS C., WÅRDELL K. and
CASIMIR-AHN H. (2005): ‘Myocardial Perfusion
Monitoring during Coronary Artery Bypass using
an ECG-triggered Laser Doppler Technique’,
submitted.
[2] NILSSON G. E., SALERUD E. G., STRÖMBERG N. O. T.
and WÅRDELL K. (2003): ‘Laser Doppler Perfusion
Monitoring and Imaging’ in VO-DINH T. (Ed):
‘Biomedical photonics handbook’, (CRC Press,
Boca Raton, FL), pp. 15:1-24.
[3] KARLSSON M. G. D., HÜBBERT L., LÖNN U.,
JANEROT-SJÖBERG B., CASIMIR-AHN H. and
WÅRDELL K. (2005): ‘Myocardial Tissue Motion
Influence on Laser Doppler Perfusion Monitoring
using Tissue Doppler Imaging’, Med Biol Eng
Comput 42(6), pp. 770-776.
[4] FORS C., KARLSSON M. G. D., AHN H. C. and
WÅRDELL K. (2004): ‘A System for On-Line Laser
Doppler Monitoring of ECG-Traced Myocardial
Perfusion’, Proc. of the 26 th Annual International
Conference of the IEEE EMBS, San Francisco, CA,
USA, pp. 3796-3799.
[5] STEINGRUB J. S., TIDSWELL M. and HIGGINS T. L.
(2003): ‘Hemodynamic consequences of heart-lung
interactions’, J Intensive Care Med, 18(2), pp. 92-
99.
IFMBE Proc. 2005;9: 84

Cardiovascular engineering
WALL BACK FLOW IN HUMAN AORTA:
INFLUENCE OF GEOMETRY
J. Svensson*, R. Gårdhagen*, D. Loyd*, and M. Karlsson**
*Department of Mechanical Engineering, Linköping University, Linköping, SWEDEN
**Department of Biomedical Engineering, Linköping University, Linköping, SWEDEN
johsv@ikp.liu.se
Abstract: In the human arteries the development
of atherosclerosis can lead to serious
lesions that influence the blood distribution
in the body. By combining MRI and CFD
the flow field in the human aorta can be simulated.
Due to indications that recirculation
can be a influencing factor for atherosclerosis
development a backflow parameter is derived
from the CFD results, and the influence of
model geometry is investigated.
Introduction
The arterial blood flow in the human body affects
the arterial wall. With computational fluid dynamics
(CFD) it is possible to estimate forces on the arterial
wall induced by the blood flow. Geometries of the
human aorta used for the simulations are gained from
magnetic resonance imaging (MRI) measurements.
Due to MRI image resolution, interpretation of MRI
images and segmentation procedures the geometry
is not entirely reproducible. It is known that the
geometries from segmentation become different and
that it is influencing some flow parameters more then
others for steady state CFD simulations [2]. The parameter
investigated for this study is wall back flow
(WBF) which is a parameter that describes if there
is a back or forward flow near the wall. WBF can be
used to show areas with recirculation or flow reversal
which is believed to be a factor in atherosclerosis
development [3]. Knowledge of the development of
atherosclerosis are highly interesting in the clinical
use e.g. diagnosis, intervention planning and followup.
The geometrical influence is in focus in the
evaluation of the WBF parameter.
Material
Image data (MRI) used here come from a juvenile
with an aorta with both a coarctation and an
aneurysm. The MRI images used are obtained with
a GE Signa Horizon MRI scanner at the University
Hospital, Linköping University. Slice thickness in
the scans is 2.6 mm and the distance between center
of slices is 3.9 mm and 35 slices were used to create
the models. The in-plane resolution is 512×512
pixels and each pixel has a size of 0.56×0.56 mm.
Method
Manual segmentation is performed by making delineation
of the arterial wall in the MRI images.
The segmentation procedure is performed by different
individuals resulting in three different geometries
which are called A, B and C model. The models are
describing the human aorta from the end of the aortic
arch to the mid thoracic section, see figure 1 left.
Steady state simulations are performed for each of
the three geometries with varying different inflow velocities
and inlet rotations. There are three different
inflow velocities and three different rotations, resulting
in 27 different simulations, see figure 1. The different
inflow rotations are set to be 50% of the axial
velocity component.
Figure 1: Left: Part of the human aorta that the models describe.
Right: Steady state simulations performed.
WBF is calculated from the WSS vectors at
the perimeter on cross-sectional planes through the
artery. These planes are defined perpendicular to a
Stokes flow streamline. The normal of the plane is
set to be the direction of the streamline. WBF have
two signs, + or -, which means forward flow or backward
flow, at a location in the present plane, figure 2.
WSS vectors that have a positive component in the
normal vector ⃗n direction are set to + which means
forward flow and vice versa for the back flow.
Figure 2: Illustration of WBF definition, -=backflow
and +=forward flow.
To make it possible to compare different geometries,
the macro variations in geometry of the models
are used to define a similar region where a valid
comparison can be performed. The cross-sectional
area of the models are used and the minimum and
maximum area are located. The region for comparison
is located between the coarctation (minimum
cross-sectional area) and the middle of the aneurysm
(maximum cross-sectional area).
Results
The geometrically defined region discussed previously
is used to derive a WBF ratio. For each simulation,
there is a ratio of WBF in %, see table 1.
IFMBE Proc. 2005;9: 85

Cardiovascular engineering
Table 1: Wall backflow ratio [%], between A min and A max
A B C
cw 0 ccw cw 0 ccw cw 0 ccw
0.2 m/s 28 28 28 27 27 27 29 29 30
0.4 m/s 25 25 25 23 23 23 28 29 34
0.6 m/s 24 28 24 19 20 20 27 24 34
This ratio is only one value and therefore the distribution
is also of interest in order to see where the
WBF is located at the arterial wall, see figure 3.
Here one map describes results from each model, this
map view of the wall was introduced and used by [1]
and [2], where streamline coordinate is on the x-axis
and the y-axis describes a circumferential coordinate.
Figure 3: WBF maps from left A,B and C model, between
A min and A max, black areas show WBF.
Another view is obtained by calculating a circumferential
WBF ratio for each streamline coordinate,
which leads to a curve describing changes of WBF
ratio between A min and A max , see figure 4.
larger then for the rotation with exception for the
C model. At higher flow rates the WBF decreases.
Variations in the rotation of the inflow influences the
WBF ratio very little especially for the lower velocities
there WBF ratio is almost identical. At higher
velocities the ratio is different with different rotations,
at least for the A and C model. The B model
is not influenced by the rotation. Finally, looking at
variations due to the difference between the models
it is seen that the B model is different, from the other
two.
The WBF areas are located in the aneurysm area.
Which is not surprising, due to the coarctation just
before the aneurysm creates a jet with counter rotating
vortices around it in the aneurysm. There are
two main areas of WBF which show that there are
two main vortices. Between these regions there is
forward flow due to the jet is wide enough to hit the
wall at these locations.
In the WBF ratio variations the behavior is similar
for the three models and the main difference here
is that the C model that has the peak WBF ratio
before the other two. This is probably due to a more
rapid change in cross-sectional area earlier in the C
model then in the other two.
Conclusions
The WBF is not so influenced by the geometrical
differences when looking at the ratio over the whole
region and along the streamline direction. The distribution
of WBF are similar in the different models
with two distinct areas with WBF in the aneurysm
part. But when looking at a more local level the
differences in geometry are influencing the WBF distribution
significantly.
References
Figure 4: Circumferential WBF ratio between A min
and A max.
Discussion
WBF is an interesting parameter which shows recirculation
areas at the arterial wall which is believed to
be a factor influencing development of atherosclerosis.
This definition of WBF gives a rough value due
to the fact that if the WSS vector is pointing in some
way downstream it is considered forward flow and if
the vector has a negative component in the normal
streamline direction it is considered to be backflow.
The advantage is that it is rather easy to interpret
WBF.
In the analyzed region the ratio of WBF (table 1)
is varying with flow, rotation and model geometry.
The variations with flow seem generally to be slightly
[1] R. Gårdhagen, J. Svensson, and M. Karlsson.
CFD studies of rotating blood flows in human
aorta - a parameter estimation. Proceedings of
ICCFD3, Toronto, Canada, 2004.
[2] J. Svensson, R. Gårdhagen, and M. Karlsson.
Geometrical considerations in patient specific
models of a human aorta with stenosis and
aneurysm. Proceedings of ICCFD3, Toronto,
Canada, 2004.
[3] H.M. Honda, T. Hsiai, C.M. Charles
M. Wortham, M. Chen, H. Lin, M. Navab,
and L.L. Demer. A complex flow pattern of low
shear stress and flow reversal promotes monocyte
binding to endothelial cells. Atherosclerosis,
(158):385–390, 2001.
IFMBE Proc. 2005;9: 86

Cardiovascular engineering
MODELING OF ACTION POTENTIAL PROPAGATION IN
CARDIAC CELLS DISPERSED IN HETEROGENEOUS MEDIA
Izharul Haq 1 , Lina Al-Kury 2 , Ziad Hani 1 , and Tala Musallam 1
Ajman University, 1 Faculty of Computer Science and Computer Engineering
2 Faculty of Pharmacy, Abu Dhabi, United Arab Emirates.
ihaq27@hotmail.com, lina_kury@yahoo.com, tmusallam@hotamail.com
Abstract: In this paper a cardiac cell capacitor
model (CCCM) has been proposed.
Myocardial cells are represented as capacitors
dispersed in an amorphous material
representing the extracellular fluid. Using
discretized Laplace equation we calculate the
distribution of the electric field (EF) within the
intercellular space. Our results show that the
EF induces spontaneous propagation of action
potential (AP) that cascades across all
myocardial cell membranes leading to
excitation. We also show the effect of varying
the number of gap junctions (GJs) on AP
propagation. Finally, we propose a new
bionano switch based on the GJ
characteristics.
Introduction
Several theories have discussed the precise
mechanism by which the action potential (AP)
propagates from cell to cell during heart
contraction. Furthermore, many of these have
been translated to computer simulation models,
which showed close correlation with laboratory
work. Recently, prototyping of AP propagation
has been performed using electronic engineering
tools such as PSpice [1].
GJs are protein structures that traverse the
cell membranes of two adjacent cells. The pore
joining the two cells is approximately 2nm in
diameter. The GJs perform intercellular
regulatory functions. It is often regarded that GJs
are the sites for transmission of AP and involved
in coordination of synchronized heart beating.
Studies have shown that the AP transmission
is carried intercellularly through the membranes
[2]. Stimulation of a cell membrane causes a local
depolarization resulting in the traveling of AP
across the cell membrane surface. This is a result
of ionic exchange (sodium and potassium ions)
between ICF and ECF through ion channels. The
AP across the cell membrane and the spontaneous
release of calcium ions may induce the closure of
GJs according to some recent studies [2].
It must be noted that there are various
constrains on using living cells in studying the
behavior of cell membranes. In view of this, an
alternative technology has recently emerged
using nanotechnology. Nanotubes can be
fabricated with characteristics very close to living
cell membranes [3].
Hypotheses
The myocardial cell is typically cylindrical
with a length of 150µm and a diameter of about
16µm. The intercellular space is about 3.5nm [4].
The ICF and the ECF are heterogeneous and can
be considered as amorphous media. At the
cellular and molecular scale the environment can
be considered as highly vibrant and dynamic.
Based on the above facts, we have proposed a
simplified cardiac cell capacitor model (CCCM)
in which we considered the myocardial cells as
capacitors of various values. The shape, size and
orientation of the capacitors are arranged in semirandom
manner conforming to the way observed
in real structure found in myocardial fibers.
Although, there is a very high level of
uniformity in cell arrangement, there are fine and
subtle variations, which may influence the
behavior of cellular functions. One of the
objectives of CCCM is to investigate the effects
caused by such morphological differences. We
used the discretized Laplace equation to
determine the EF distribution within the ECF.
From such study it may be possible to determine
the mechanism by which the depolarization
propagates. Further investigation is carried out to
determine the role of GJs in propagation of AP.
Using CCCM, it may be possible to determine
how myocardial cells repolarize.
Results
Figure 2(a) shows computer simulation
results of excitation from the first to the sixth
myocardial cells. As the stimulation reaches its
threshold in the first cell, the consecutive five
cells become partially depolarized. The
magnitude of depolarization decreases with
distance as expected. Figure 2(b) shows the
simulation at a later time at which the first two
cells begin to go through repolarization. Whilst,
the remaining cells are still in the process of
depolarization.
IFMBE Proc. 2005;9: 87

Cardiovascular engineering
Excitatio
Excitation
n
10
5
0
1 2 3 4
Number of Cells 5 6
10
8
6
4
2
0
1 2 3 4
Number of Cells
Figure 2: AP Propagation along myocardial cells
(a) at stimulation (b) subsequent interval.
The AP values are given in Table [1].When
cell C 00 was stimulated, this induced cells C 01 and
C 10 to reach the threshold value. Cells at the far
edges of the matrix (e.g., C 12,8 ) were not affected.
Stimulation opens Na + channels thus the cells
begin to self excite reaching the threshold values.
Consequently, this leads to the propagation of AP
across the whole structure.
Table 1: AP of myocardial cells induced by
stimulation of cell C 00 .
Figure 3 shows the speed of propagation of
AP with respect to the number of GJs. As the
number of GJs increases, this causes an initial
nonlinear rise in the propagation speed and
eventually reaching a plateau.
10
S (arbitrary units)
1
1 2 3 4 5 6 7 8 9
Gap Junctions (arbitrary units)
Figure 3: Speed of propagation Vs. no. of GJs.
5
6
Figure 2(a)
Figure 2(b)
Discussion
Results from this work have shown that
immediately after excitation, an EF is developed
between the excited and adjacent cell. If the value
of EF is sufficiently large, it can surpass the
threshold for excitation. At this point AP wave
propagates to the adjacent cell. Likewise, this
process continues to cascade until all the
myocardial cells along the path are depolarized. It
is found that there is a delay in transmission at
the transjunctional membranes that agrees with
other findings [2]. From our CCCM it is found
that the EF distribution in the ICF is not uniform
but causes no significant disruption to the AP. As
each myocardial cell fires, the EF diminishes
between the adjacent cells and the process of
repolarization begins.
The repolarization occurs as a consequent
event leading to the resting membrane potential.
This can be explained as follows. As the
excitation potential reaches its threshold, sodium
channels close followed by the opening of
potassium channels. Since the process is dictated
by diffusion, repolarization is slower. This result
is reflected in the asymmetrical AP curve seen
experimentally [5].
Conclusions
CCCM has verified that AP can travel
without the presence of GJs although the
presence of these junctions increases the rate of
propagation. Based on our findings, we suggest
further investigation to determine the function of
GJs using nanotubes. We propose using
nanotubes to encapsulate the connexin proteins.
These connexin-nanotubes may display similar
characteristics to the GJs and can be
characterized for their physical and electrical
properties. Such structures can also behave as
bionanoGJ switches under certain conditions.
Currently we are investigating how bionanoGJ switch
can be fabricated.
Reference
[1] Sperelakis,N. et al. (2005): Propagated
repolarization of simulated action potentials
in cardiac muscle and smooth muscle. Theor.
Boil. Med. Model., 2.
[2] Sperelakis,N. and Ramasamy,L. (2005): Gapjunction
channels inhibit transverse
propagation in cardiac muscle. BioMed. Eng.
Online, 4.
[3] Pirio, G. et al. (2002): Fabrication and
electrical characteristics of carbon nanotubes
field emission micro cathode with integrated
gate electrodes. IoP Elec. J. Nanotech. 13.
[4] Severs,N.J. (2000): The cardiac muscle cell.
BioEssays, 22, 188 – 199.
[5] Guyton,A.C. and Hall,J.E. (1996): Text Book
of Medical Physiology. W. B. Saunders
Company, Pennsylvania.
IFMBE Proc. 2005;9: 88

Cardiovascular engineering
A MICROSYSTEM FOR MONITORING HEART MOTION
L. Hoff * , O.J. Elle ** , M.J. Grimnes * , S. Halvorsen ** , H.J. Alker * , E. Fosse **
*
Vestfold University College, Faculty of Science and Engineering, Horten, Norway
*
Rikshospitalet University Hospital, The Interventional Centre, Oslo, Norway
lars.hoff@hive.no
Abstract: The Micro Heart project is a collaboration
between Vestfold University College and
Rikshospitalet University Hospital. The aim is to
construct a miniaturized motion sensor for use
during heart surgery. The first feasibility studies
have used commercial dual-axis accelerometers to
monitor the heart motion in anesthetized pigs. Heart
motion is due to both respiration and heart beating,
and respiration movements were filtered out for a
meaningful interpretation of the data. Velocity and
position of the heart wall were calculated by
numerically integrating high-pass filtered
acceleration traces. The results show that the
acceleration sensors can measure heart motion in
great detail. Abnormal events, e.g. arrhythmias, that
occurred during the measurement were recognized,
and confirmed by comparison with synchronously
recorded ECG data.
Introduction
The Micro Heart project is a collaboration between
The Interventional Centre at Rikshospitalet University
Hospital and the Institute for Microsystems Technology
at Vestfold University College. The project is based on
an idea that originated at Rikshospitalet, and it is funded
by a grant from the Norwegian Research Council.
Single-axis accelerometers have previously been
used on the heart to better understand the sources of
heart sounds [1]. The goal of our work is to measure the
heart motion in detail using triple-axis accelerometers,
during and after surgery [2][3][4]. This monitoring may
give an early warning of complications, e.g. ischemia,
after coronary bypass surgery. Constructing a dedicated
triple-axis sensor is the major part of our project, but it
also includes systems for acquisition, processing and
interpretation of the sensor data.
This paper presents results of our first animal
studies. They were done using sensor prototypes made
from commercially available accelerometers. The
purpose of these studies was to investigate whether the
concept is feasible, and to obtain results needed for the
specification of a final sensor.
Measurements were done on anesthetized pigs. The
chest wall was opened, and two acceleration sensors
were sutured to the heart wall. The first sensor was
fastened near the apex, receiving its blood supply from
the LAD coronary artery. The second sensor was
fastened higher on the left ventricle, on a region
supplied by the CX coronary artery.
The data acquisition setup is illustrated in Figure 1.
Analog acceleration data from each sensor was sampled
at 250 samples/s, using a 12-bit AD-converter (National
Instruments DAQPad 6020E). The ECG signal was
sampled synchronously as a time-reference. Several
hours of heart beating were recorded. Results were
stored in a computer, and processed using Matlab.
We are interested in the motion from the heart's own
beating. The heart also moves due to respiration, and
movements of the whole animal may further complicate
the interpretation of the signals. The pig heart beats at a
rate of about 1.5 Hz, or 90 beats per minute, while the
respiration rate is typically less than 0.4 Hz. Hence, the
heart beating could be isolated by high pass filtering:
DC components, caused by gravitation, were removed
by subtracting the mean value. The resulting traces were
run through a digital 4th order high-pass Butterworth
filter with cut-off frequency 1.0 Hz. This filter was
applied both forwards and backwards, to obtain zero
phase.
Results
The power spectrum of a received acceleration trace
is plotted in Figure 2. Two series of harmonic peaks are
identified, corresponding to two distinct periodic
motions: The respiration at 0.3 Hz and the heart beating
at 1.7 Hz. The position spectrum is obtained by
weighting this acceleration spectrum by 1/f 2 , where f is
the frequency. Hence, the position spectrum is
dominated by the peak at 0.3 Hz, i.e., the heart's
absolute position is mainly determined by the
respiration. Another interesting feature in the spectrum
is the increased height of the respiration peaks around
Methods
Triple-axis acceleration sensors were made by
mounting two commercial dual-axis accelerometers at
90° angles (ADXL311, Analog Devices, Norwood, MA,
USA). The bandwidth from the sensors was reduced to
50 Hz, selected to preserve the finer structures in the
heart motion, while minimizing noise.
Figure 1. Data acquisition setup. Data was collected
from two sensors attached to a pig’s heart.
IFMBE Proc. 2005;9: 89

Cardiovascular engineering
Figure 2. Power spectrum of the heart acceleration.
Two series of harmonic peaks are recognized: One from
respiration, at 0.3 Hz, and one from the heart beating, at
1.7 Hz.
the heart motion peaks. This is interpreted as nonlinear
mixing between the respiration and heart motion peaks:
The heart motion changes slightly through the
respiration cycle.
Figure 3shows the heart position calculated from the
acceleration traces. For simplicity, only one of the three
axes is plotted. The ECG signal is included for
reference. The curve displays a periodic motion with
amplitude slightly less than 1 cm, perfectly synchronized
with the ECG. An arrhythmia seen in the ECG signal at
638 s is easily recognized in the motion graph. The
motion curve shows in detail how the arrhythmia briefly
disturbs the heart motion, and how the motion is
restored to normal after a few heartbeats.
The heart wall velocity is of particular interest, as
novel ultrasound methods, i.e. Tissue Doppler, or Tissue
Velocity Imaging, TVI, have provided new information
about the heart velocity pattern. An example of a heart
wall velocity calculation is shown in Figure 4. Structures
in the velocity pattern can be recognized, and the shape
of these structures have been shown to be indicators of
heart failure [4].
Discussion
This study has shown that accelerometers can
measure the heart motion in great detail. Abnormalities
in the motion pattern are recognized.
The prototype sensors used in this study are too big
for practical use on humans. However, they have given
us confidence that the concept is feasible, and have
provided the experimental data needed to get
specifications for a final miniaturized sensor. Work on a
miniaturized triple-axis sensor with size and
specifications optimized for the application is in
progress, and the first test sensor designs are in
production at the MultiMEMS foundry service
(SensoNor AS, Horten, Norway).
Conclusions
Triple-axis accelerometers can measure heart motion
with good resolution. The measured motion traces can
reveal patterns that may be an indication of heart
circulation failure.
Figure 3. Heart position as function of time, measured
with the prototype sensor. Only one of the three axes is
shown. An arrhythmia at 638 s is captured by the sensor.
Figure 4. Heart wall velocity as function of time,
calculated from the acceleration measurements.
References
[1] WOOD, J. C., BUDA, A. J., and BARRY, D. T., (1992):
'Time-Frequency Transforms: A New Approach to
First Heart Sound Frequency Dynamics', IEEE
Trans. Biomed. Eng. 39, pp. 730-740.
[2] ELLE, O.J, HALVORSEN, S., GULBRANDSEN, M.G.,
AURDAL, L., BAKKEN, A., SAMSET, E., DUGSTAD, H., and
FOSSE, E. (2005): 'Early recognition of regional
cardiac ischemia using a 3-axis accelerometer
sensor', Physiol. Meas. 26. In press.
[3] HOFF, L., ELLE, O. J., GRIMNES, M. J., HALVORSEN, S.,
ALKER, H. J., and FOSSE, E. (2004): 'Measurements of
Heart Motion using Accelerometers', Proc. of 26th
Ann. Internat. Conf., IEEE Eng. in Med. and Biol.
Society, San Fransisco, CA, USA, pp. 2049-2051.
[4] GRIMNES, M., HOFF, L., HALVORSEN, S., ELLE, O. J.,
ALKER, H. J., and FOSSE, E. (2004): 'Velocity and
Position Approximations from Left Ventricular 3D
Accelerometer Data', Proc. of IEEE 30th Ann.
Northeast Bioeng. Conf., Springfield, MA, USA.
[5] EDVARDSEN, T., URHEIM, S., SKULSTAD, H., STEINE, K.,
IHLEN, H., SMISETH, O. A. (2002): 'Quantification of
left ventricular systolic function by tissue Doppler
echocardiography: added value of measuring preand
postejection velocities in ischemic myocardium',
Circulation, 105, pp. 2071- 2077.
IFMBE Proc. 2005;9: 90

MITRAL VALVE OPENING IN THE FAILING HEART
Cardiovascular engineering
K. Kindberg and M. Karlsson
Division of Biomedical Modelling and Simulation, Department of Biomedical Engineering,
Linköping University, Linköping, SWEDEN
katki@imt.liu.se
Abstract: The opening mechanics of the mitral
valve in seven failing ovine hearts has been
quantified. This was done by using data from
a research protocol with a dense radiopaque
marker array on the mitral valve of seven ovine
hearts, as well as an array of markers implanted
in the left ventricular wall.
as t = 0. The distance between the markers at the
edges of the anterior and posterior leaflets is used as a
measure of the opening of the mitral valve.
Introduction
The normal mitral valve opening was studied in detail
in [1]. The same animals were here used to investigate
the behaviour of the mitral valve in the failing heart,
a heart that is pumping a diminishing amount of blood.
Materials and Methods
Radiopaque markers were surgically implanted in the
left atrial and ventricular walls of seven ovine hearts.
In addition, an array of 14 markers was placed on the
mitral annular ring and on the anterior and posterior
leaflets of the mitral valve. The surgery and the data
acquisitionwerereportedindetailin[1],andhence
only a brief review will be done here. Seven healthy
sheep were anesthetized and an array of radiopaque
markers was implanted to silhouette the left ventricular,
LV, chamber. Four markers were sutured to the
left atrial epicardium. Six markers, 2.5 mm diameter,
were then sutured equidistant from each other along
the central meridian of each mitral leaflet, four on the
anterior leaflet and two on the posterior leaflet. Eight
additional markers were sutured to the atrial side of the
mitral annulus at equal distances around its circumference,
one near each commissure and three along the
perimeters of the anterior and posterior leaflets. After
eight weeks of recovery, hemodynamic and biplane
videofluoroscopic data were obtained with hearts in
normal sinus rhythm. Two biplane fluoroscopic images
were recorded simultaneously at 60 Hz and an analog
LV pressure, LVP, signal was recorded in digital format
on each individual video image. The two-dimensional
coordinates of each marker in each projection were digitized
frame-by-frame. Data from the two views were
merged to yield the three-dimensional coordinates of
the centroid of each marker every 16.7 ms.
During data acquisition, vena cava was occluded
during a few cardiac cycles to generate a series of
shrinking pressure-volume loops, shown in Figure 1, in
order to study the end-systolic pressure-volume relationship.
The vena cava occlusion, VCO, is here used
to study the mitral valve opening, MVO, in the failing
heart. Four phases during VCO will be studied. The
first phase, before occlusion, consists of three consecutive
beats. Phase two, three and four consist each of
one beat after some time of occlusion, see Figure 1.
The different beats were time aligned by choosing the
time frame during rapid pressure decay immediately
after LVP is 10% larger than the minimum LVP value
Figure 1: Pressure-volume loops of the left ventricle
during vena cava occlusion, VCO. The four loops correspond
to four phases during VCO.
At each sample time, the angle, θ Ant , between an
annular septal-lateral reference vector and the vector
from the septal annular point to the edge of the anterior
leaflet was computed. Similarly, the angle, θ Post ,
between the annular reference vector and the vector
from the lateral annular marker to the edge of the
posterior leaflet was computed.
Statistics: The three beats of each animal during
the first phase were treated as one beat, using the
mean value of the three beats. All values are given
as mean±SE for n = 7 animals (phase one) or n = 6
animals (phase two–four) unless otherwise specified.
The significance of the difference between two values
of a given variable was assessed using paired t-test.
Significance was accepted at P

Cardiovascular engineering
Throughout the whole cardiac cycle, both the
septal-lateral distance of the annulus and the
commissure-commissure distance decrease during
VCO.
Discussion
In the failing heart, the mitral valve opening starts
later in the cardiac cycle and the maximum distance
between the opened anterior and posterior leaflets decreases.
This behaviour is achieved with a smaller
mitral annulus, since the annular septal-lateral and
commissure-commissure distances decrease. The maximum
opening angles of the mitral leaflets do not
change, but the leaflet angles of the closed valve increase.
Figure 2: MVO during the four phases of VCO. LVP
for the first phase.
The maximum anterior and posterior opening angles
do not change during VCO, see Figure 3. The
maximum values of the angles are reached later during
VCO, which is consistent with the longer opening
intervals. The angles of the closed leaflets get larger
mean values during VCO.
References
[1] M. O. Karlsson, J. R. Glasson, A. F. Bolger, G.
T. Daughters, M. Komeda, L. E. Foppiano, D. C.
Miller, N. B. Ingels Jr. Mitral valve opening in
the ovine heart. Am J Physiol Heart Circ Physiol,
274(43):552-63, 1998.
Figure 3: The leaflet angles before and during vena
cava occlusion. Above: The anterior leaflet angle θ Ant .
Below: the posterior leaflet angle θ Post .
IFMBE Proc. 2005;9: 92

Cardiovascular engineering
INSTRUMENTATION FOR REPRODUCING THE POSITIONING OF A
PERSON IN BALLISTOCARDIOGRAPHIC MEASUREMENT
L. Leppäkorpi*, R. Sepponen**
* Helsinki Univ. of Technology/Applied Electronics Lab., P.O.Box 3000, 02015 Espoo, Finland
** Helsinki Univ. of Technology/Applied Electronics Lab., P.O.Box 3000, 02015 Espoo, Finland
lasse.leppakorpi@hut.fi
Abstract: A new type of instrumentation for
enhancing the reproducibility of
ballistocardiographic (BCG) measurements is
introduced. A specially designed vest transfers the
mechanical excitation to the sternum of a person.
While propagating through the body, this excitation
is superposed with all internally originated signals.
The signal is then measured with a BCG measuring
device and the signal component corresponding to
the excitation is detected and analysed. The results
indicate that the reproducibility of the positioning of
a person may be achieved.
Materials and Methods
The calibration instrumentation consists of three
basic parts: mechanical excitator, electronics and a
computer with a data analysing program. A block
diagram of the system is shown in Figure 1.
Introduction
In ballistocardiography (BCG), the reproducibility
of measurements has been questioned.
In BCG, the mechanics of a body is about the
vibrations caused by the cardiac function and how they
are transferred through the body to the supporting
measurement device. A posture of a person on a
measuring device partially defines the internal
transmission bath and affects the body-device coupling;
it therefore has a remarkable effect on the related
measured ballistocardiogram [1].
Despite the distortion caused by the coupling, it is
better to take the measurements from the supporting
device than directly from the body. This is because the
platform averages all simultaneous incidents in the
whole body, whereas measurement directly from the
body emphasizes local events near the sensor [2].
The basic procedure for investigating the effects of
whole-body vibration is to vibrate subjects with
platform and measure the responses directly from their
bodies [3]. This procedure is obviously reversed in
actual ballistocardiography.
The objective of this study is to introduce and design
improved instrumentation for the calibration of
ballistocardiographic measurement [4]. In this new
concept, the excitation signal is transformed into
mechanical excitation and coupled directly to the
sternum of a person under measurement. The
significance comes from the fact that the method
follows the actual transmission direction of
ballistocardiographic vibration, and the signal is
measured using the BCG measuring device.
Figure 1: Block chart of the calibration instrumentation.
A microcontroller and digital-to-analog converter
are used to create two reference signals with ninetydegree
phase shift. The excitation signal is fed to an
electromagnetic mechanical excitator. This can be
attached to a special vest that a person can wear. The
vest and the coupling of the vest to the body of a person
are illustrated in Figure 2.
Figure 2: The vest is designed to fit persons with
different body sizes. In Figure A, the pectoralis major is
shown, B shows the body of a sternum, C a piece of
viscoelastic foam, D a mechanical excitator and E a
tightening strap. F illustrates an additional strap that can
be attached between legs.
The detection electronics uses the signal from the
amplifier of a BCG measurement system. The signal is a
superposed signal of the calibration and the
ballistocardiographic signal components. The
IFMBE Proc. 2005;9: 93

Cardiovascular engineering
calibration signal component is detected with a
quadrature detector and analysed with a computer. This
method gives the ability to sense the prevailing setup
including the posture of the person.
To demonstrate the effect of posture on the detected
signal, a set of measurements with documentation of
changes in posture has been assembled. A posture of a
person is captured with a digital camera before and after
the posture changes. These digital images are used to
calculate a difference image, which can be compared; it
should correlate with the changes in the measured
signals.
From the point of view of reproducibility, we can
assume that the posture is reproduced when measured
signal amplitudes are the same as in the previous
measurement. This is demonstrated in figure 5.
Results
Coupling of the excitation to the measurement
system is illustrated in the Figure 3. In this case, the
frequency of 12 Hz is used.
Figure 5: Changes in the posture of a person when there
is an attempt to get signal amplitudes at the same levels
as in a previous measurement.
Discussion
The measurements and results demonstrate the
importance of calibration and a careful positioning of
the person on the BCG instrument. This implies that in
designing a BCG instrument one should consider means
to improve the reproducibility of the positioning of the
person.
Conclusions
Figure 3: The excitation reference signal and the BCGsignal
measured separately with a BCG-measurement
system and the measured superposition of these two
when the vest is on a person.
The positioning of a person: The results demonstrate
that changes in the amplitude and the phase of the
calibration signal are correlating with the changes in the
posture of a person. For example, bending the upper
body forward causes a remarkable change in measured
output. This change is shown in Figure 4. On the left
side of the screen is an illustration of the traces in the
starting posture and on the right side after the posture
has changed.
We have designed an improved instrumentation for
the calibration of the BCG measurements. The new type
of calibrator can be used to study mechanical properties
of the whole system including the body of a person
under study. Information provided by this calibration
system enables reproducible positioning of a person on
a BCG instrument.
References
[1]WERDOUW P. D., WESTERHOF N. and NORDERGRAAF
A. (1969): ‘Analysis of the Effect of Posture on the
BCG’, Bibl. Cardiol., 24
[2] CUNNINGHAM D. M. and GRISWOLD H. E. (1974):
‘Acceleration-Frequency Response of the Isolated
Human Body Over the Ballistocardiograph
Spectrum’, Bibl. Cardiol. 34
[3] GOLDMAN D. E., and VON GIERKE H. E. (1960): ‘The
Effects of Shock and Vibration on Man’, Lecture and
Review Series, Naval Medical Research Institute,
Bethesda, Maryland, vol. 60-3
Figure 4: Oscilloscope traces represent both signals (I
and Q) in quadrature detector and a calculated sum of
squared signal amplitudes (S). The change in signal S
corresponds to the amplitude difference; the change in
ratio of the signals I and Q corresponds to the phase
difference. The difference in posture is represented by
the black shadow in the image of the person.
[4] LEPPÄKORPI L. and SEPPONEN R. (2005):
‘Instrumentation for the Calibration of
Ballistocardiographic Measurement’, submitted for
publishing
IFMBE Proc. 2005;9: 94

Cardiovascular engineering
STUDY OF THE CAUSES OF BLOOD PRESSURE VARIABILITY DURING
MANUAL SPHYGMOMANOMETER MEASUREMENT
F.E. Smith, A. Haigh, J. Wild, E.J. Bowers, S.T. King, J. Allen, D. Zheng, P. Langley, A. Murray
Cardiovascular Physics and Engineering Research Group, Medical Physics, Newcastle University,
Freeman Hospital, Newcastle upon Tyne, UK
alan.murray@ncl.ac.uk
Abstract: Manual blood pressure measurement is an
important clinical measurement. However, it is
highly variable and often inaccurately performed.
Overestimating or underestimating blood pressure
by even 5 mmHg can seriously compromise accurate
diagnosis. Blood pressure is influenced by factors
that are not always well controlled during the
measurement procedure. However, little data exists
on the magnitude of their effect. Twelve subjects
were studied, repeating blood pressure measurement
8 times in each. Each group of four measurements
was made during resting, deep breathing, talking
and head movement. The order of these
measurements was randomised. During deep
breathing diastolic pressure fell by mean (SD) 5.5
(2.5) mmHg (p

Cardiovascular engineering
Head movement: The subject rotated their head from
side to side during the measurement.
Data Analysis
Using variance analysis, the effect of patient factors
was studied. This also took into account sequential
changes which might arise during the four
measurements.
Results
Variance analysis showed that there were no
differences in blood pressure measured as a result of the
measurement’s time position in the sequence of four
measurements, but that there were very significant
differences in both systolic (p

Cardiovascular engineering
PRELIMINARY CLINICAL RESULTS FROM A COMPUTER-ASSISTED
CORONARY ANGIOPLASTY BALLOON INFLATION DEVICE
T. Olbrich and A. Murray
Cardiovascular Physics and Engineering Research Group, Medical Physics, Newcastle University,
Freeman Hospital, Newcastle upon Tyne, UK
alan.murray@ncl.ac.uk
Abstract: A computer-assisted balloon inflation
device with pressure-volume (PV) feedback
capabilities has been developed for use in the clinical
environment. Preliminary results from 3 patients
have been studied. They showed different PV
characteristics between manual stent deployment
and computer assisted stent deployment, and
illustrated the effect of angioplasty predilation and
stent deployment.
Introduction
Coronary artery disease is the cause of most
premature deaths in the western world and, despite
considerable mechanical and pharmacological
developments in therapy, restenosis still affects a
significant number of patients [1]. Also, up to 30
percent of all conventional balloon angioplasty
procedures result in angiographically significant
coronary artery dissections [2]. Since the introduction of
stents almost a decade ago, stenting has now become the
most commonly used treatment of percutaneous
coronary interventions (PCI) for patients with coronary
artery disease (CAD).
In a standard PCI procedure, a small angioplasty
balloon is positioned inside the stenosed segment and is
inflated with an incompressible fluid to open up the
narrowing large enough for the following stent to pass
through. The stent is mounted on the tip of a balloon
catheter and deployed by inflating that balloon to a
predefined pressure.
Manual balloon inflation and stent deployment may
induce more trauma to the vessel wall than necessary.
This could result in immediate vessel closure in the
short term or in restenosis in the long term. Also, apart
from a low-resolution angiogram that is produced after
each inflation, the operator has no other information
about lesion properties, lumen improvement or the
quality of the stent deployment.
Aim
The aim was to develop a computer-assisted balloon
inflation (CABI) device for clinical use in Percutaneous
Coronary Intervention (angioplasty and stent
deployment) that was also capable of assisting the
cardiologist with feedback from pressure-volume data
recorded during the procedure in real time.
Methods
Computer assisted inflation device
A laboratory version of our automatic measurement
system was redesigned for use in the clinical
environment. The system has been described in depth
previously [3,4] and therefore only a very brief
description is given.
The inflation device consisted of a high pressure
glass syringe (Hamilton Gasthigth®), a syringe holder, a
stepper motor and two pressure transducers (RS249-
3858). The second pressure transducer was used to
check for any faults of the first pressure transducer and
ensure additional safety when in clinical use. The
syringe plunger was attached to the stepper motor using
a threaded shaft that converted the rotational movement
of the motor into a pushing / pulling movement of the
syringe plunger. A computer controled the motor via the
control unit and ran the measurement sequence
automatically. The control unit generated a pulse train
that drove the stepper motor and monitored the pressure
signals from both pressure transducers. The measured
inflation pressure P was constantly compared to a preset
maximum inflation pressure and the inflation was
stopped immediately when this threshold P max was
reached. After a predefined “wait-time”, the balloon
was deflated to the starting point. The current inflation
volume V was determined from the number of pulses
driving the stepper motor and the pressure-volume (PV)
data was displayed on the monitor in real time (Figure
1).
The system was fully equipped with safety features
to avoid any potential danger to the patient through
malfunction. The whole system was subjected to a full
function test and the high pressure syringe, the pressure
transducer manifold and the syringe housing were
disinfected (Gigasept®) before clinical use.
Clinical application
The CABI research system was set-up and operated
by a physicist following the instructions of the clinician
in charge. Patients were studied in the Freeman Hospital
Catheterization Laboratory during clinical angioplasty
and stent deployment. Preliminary results have been
obtained from 3 patients illustrating the PV data of the
existing manual technique, an automatic predilation and
an automatic stent deployment.
IFMBE Proc. 2005;9: 97

Cardiovascular engineering
P
Pressure
Transducer
V
Automatic
Balloon Inflator
Balloon Catheter
Figure 1: Control circuit of automatic balloon inflation
with feedback cababilities.
Results
The measurement system was successfully
integrated into the catheterization laboratory of the
Freeman Hospital. Figure 2 shows the P curve of a
manual stent deployment. It is noticeable that the
inflation pressure drops considerably after each increase
due to the interrupted twisting motion of the
Indeflator handle. Figure 3 shows the PV curve (1 st
inflation and 2 nd inflation) during automatic stent
deployment. The inflation was very smooth due to a
constant inflation speed and illustrates the mechanical
interaction between the balloon catheter, the stent and
the arterial wall.
1st
2nd
Figure 4: PV curve (1 st & 2 nd inflation) of an automatic
predilation.
Figure 4 shows the PV curve of the 1 st and 2 nd
inflation during angioplasty predilation. The difference
between 1 st and 2 nd inflation illustrates the work needed
to open the stenosed section suggesting an improved
vessel lumen.
Discussion and Conclusions
This preliminary clinical study has given us
confidence that the system can be used in the clinical
environment. It has been shown that there is a difference
between manual and automatic inflation technique and
that the lumen improvement achieved during
angioplasty predilation can be illustrated with pressurevolume
curves.
The system is now undergoing an extensive clinical
study to investigate mechanical characteristics during
balloon angioplasty and stent deployment in vivo and in
situ.
References
Figure 2: P curve (inflation) of a manual stent
deployment.
[1] DOBESH PP, STACY ZA, ANSARA AJ,
ENDERS JM. (2004): 'Drug eluting stents: a
mechanical and pharmacologic approach to
coronary artery disease', Pharmacotherapy, 24,
pp. 1554-1577
[2] ROGERS JH, LASALA JM. (2004): 'Coronary
artery dissection and perforation complicating
percutaneous coronary intervention', J Invasive
Cardiol., 16, pp. 493-499
[3] OLBRICH T, MURRAY A. (2001): 'Assessment
of computer-controlled inflation/deflation for
determining the properties of PTCA balloon
catheters with pressure-volume curves', Physiol.
Meas., 22, pp. 1-10
1st
2nd
Figure 3: PV curve (1 st & 2 nd inflation) of an automatic
stent deployment.
[4] OLBRICH T, MURRAY A. (2004): 'Assessment
of a technique to determine the mechanical
properties of coronary arteries using mock
arteries', Physiol. Meas., 25, pp. 1-15
IFMBE Proc. 2005;9: 98

Cardiovascular engineering
Development of Implantable Cardiac Measurement Device
-Modeling Approach
J. Väisänen, J. Hyttinen, J. Malmivuo
Ragnar Granit Institute, Tampere University of Technology, Tampere, Finland
E-mail: juho.vaisanen@tut.fi
Abstract: The designing methods should be applied
to the designing processes of new implantable ECG
device in a way that the effects of implantation on
measurement could be predicted and the need of
expensive clinical trials could be minimized.
Especially when effects of implantation on received
signal are studied these methods could be useful.
Modeling could offer an effective means of studying
effects of implantation on the ECG measurement
compared to standard body surface measurements.
This paper introduces a project where a modeling
method was applied to study measurement
sensitivities of implantable ECG devices. The study
was based on 3D Finite Difference Method (FDM)
applying lead field and reciprocity theorems. The
results of the study indicate that the modeling could
be used as an effective tool when implants are
designed.
Introduction
The 12-lead ECG system is commonly used in
monitoring the electrical activity of the heart. The use of
the system needs skilled personel and hospital
environment, thus producing costs. The use of wireless
and especially implantable measurements could offer a
stable and long term monitoring possibility with lower
costs. Modeling serves an effective means of studying
e.g. the measurement of the heart’s electrical activity
without expensive and time consuming clinical
measurements. [1, 2] Especially the use of the modeling
for designing of the implants reduces the need of
expensive testing and iteration rounds because different
characteristics of the implant, could be tested with
models during the designing process. Furthermore,
modeling can be used for estimating the signals
measured from the implants and how they correlate with
the ECG measured from the standard surface leads. The
past modeling studies of the implanted devices have
been concentrating on modeling the current distributions
arrised by stimulative devices like cardiac defibrillators
[3]. Thus there is a lack of modeling studies where the
measurement sensitivities of implanted ECG monitors
had been studied.
The objective of this paper is to introduce the
modeling and especially the lead field and reciprocity
approach [4] in the development of the implantable
bioelectric measurement solutions by demonstrating the
effects of implant's size and implantation on
measurement sensitivity of implantable ECG monitor.
Materials and Methods
A. Lead field and reciprocity theory
The relationship between the measured signal V LE in
the lead and the cardiac current sources J i in the
volume conductor can be formulated as follows:
V
LE
1
= ∫ J σ
LE
i
• J dv
where J LE is the lead field and σ is the conductivity of
the volume conductor [4]. If the lead field and the
properties of the volume conductor are known, the
strength of the measured signal can be estimated.
The lead field can be obtained based on reciprocity
theorem which states that the locations of source and
detector can be swapped without any change in the
amplitude of the measured signal. For example the
result is same if a current source is located to the
volume conductor and the sensitivity is measured on the
surface of the conductor or other way around. The
current field in the volume conductor is equivalent with
the lead field. [4]
(1)
B. Models
The model used in this study was a 3D model of
thorax based on the Visible Human Man model (VHM)
[5, 6]. It represents data of 60 segmented slices of a
male thorax and contains 26 different tissue types. The
calculations are based on the finite difference method
(FDM), which is composed of cubic elements forming a
resistor network [2].
C. Calculations
To demonstrate the general effects of implantation
the lead field of small implant was compared with the
lead field of surface electrodes. The surface electrodes
were located on the body surface on the same locations
as the electrodes of the small implant. The implantation
depths for electrodes of the small implant were 15 mm.
To demonstrate the effects of the implant’s size to the
measurement sensitivity the lead field of large implant
(23.3 mm x 6.7 mm x 60 mm) was compared with the
lead field of a small implant (3.3 mm x 3.3 mm x 32
mm). Both implants had electrodes covering both ends.
The size of the small implant is the target size of
implant in our project (www.ele.tut.fi/tule). The outlook
of the 3D torso model with large and small implants is
illustrated in Figure 1.
IFMBE Proc. 2005;9: 99

Cardiovascular engineering
Table 1: Effects of Implant’s Size and Implantation to
the Measurement Sensitivity
Figure 1: 3D models of torsos containing nonconducting
large(A) and small(B) implants with white.
Results
In Table 1 is presented the average changes of the
lead vector magnitude and direction when the small
implant case was compared to surface electrode and
large implant cases. Standard deviations of lead field
magnitude and direction changes were also depicted.
The results show that implantation increases the average
lead field magnitude as could be expected. It can be also
seen that reducing the implants size the average lead
field magnitude decreases, as could be expected.
The standard deviation of surface electrode case
shows that the magnitude change due to the
implantation is not homogeneously distributed. For
clinicians it could be more interesting to know how
much the sensitivity distribution is changed due to the
implantation. Thus it could be used to estimate the
source area of the measured signal. The distribution of
magnitude change of lead field in small implant lead
field versus large implant is illustrated in Figure 2.
Conclusions
The study shows that the modeling and especially
the lead field and reciprocity theories are effective and
applicable tools when measurement sensitivities of new
implantable applications are modeled. It shows that
these methods could be used e.g. when new implantable
ECG devices are designed. With modeling the
appearing changes in signals can be estimated
beforehand without expensive clinical studies.
Small implant vs.
Surface
electrodes
Large
implant
∆Direction (°) 13.253 3.863
Std ± 0.138 0.056
∆Magnitude (%) 17.7 -47.3
Std ± 20.9 5.9
In the future the modeling could be applied as a
part of system designing for defining e.g. the implant’s
purpose of use or as testing tool when different
characteristics of the implant are tested.
Acknowledgements
We would like to thank PhD Noriyuki Takano for
providing the finite different method software and MSc
Tuukka Arola for providing software for model
illustrations. This work has been supported by the
Academy of Finland (202758), by a grant from The
Finnish Cultural Foundation, and by the Ragnar Granit
Foundation.
References
[1] Hyttinen J. (1994): 'Development of Regional
Aimed ECG Leads Especially for Myocardial
Ischemia Diagnosis', Ragnar Granit Institute,
Tampere University of Technology
[2] Takano N. (2002): 'Reduction of ECG Leads
and Equivalent Sources Using
Orthogonalization and Clustering Techniques',
Ragnar Granit Institute, Tampere University of
Technology
[3] Papazov S., Kostov Z. and Daskalov I. (2002):
'Electrical current distribution under
transthoracic defibrillation and pacing
electrodes', J Med Eng Technol, 26, pp. 22-7
[4] Malmivuo J. and Plonsey R. (1995):
'Bioelectromagnetism: Principles and
Applications of Bioelectric and Biomagnetic
Fields', Oxford University Press
[5] Kauppinen P., Hyttinen J., Heinonen T. and
Malmivuo J. (1998): 'Detailed model of the
thorax as a volume conductor based on the
visible human man data', J Med Eng Technol,
22, pp. 126-33
[6] Ackerman M. J. (1991): 'The Visible Human
Project', J Biocommun, 18, pp. 14
Figure 2: The distribution of the lead vector magnitude
changes inside the heart in one slice. The changes of the
lead vector magnitude when small implant was
compared with the large implant.
IFMBE Proc. 2005;9: 100

Cardiovascular engineering
A NOVEL ISOLATED MEASUREMENT SYSTEM FOR BIO-POTENTIALS
K. Piipponen 1 , R. Sepponen 1
1 Applied Electronics Laboratory, Helsinki University of Technology, Espoo, Finland
kari.piipponen@tkk.fi
Abstract
The common mode interference induced from
variety of sources is a major problem in bio-potential
instrumentation. Requirements
for patient safety and high signal to noise ratio raise a
demand for effective
isolation devices. We describe a new method and
means to implement a compact isolated
instrumentation network. The new concept applies
microwaves for communications
and power feed, allowing very low isolation
capacitance and bi-directional
communications over isolation. The system is suitable
and safe enough for any
line-powered bio-potential measurement and
significantly reduces the common
mode interference.
Introduction
Bio-potential signals are often recorded in the presence of
electromagnetic interference. The patient, the electrode
leads and the connected electronics form a system into
which interference couples several ways. Focus of this
paper is in the common mode interference. Figure 1
illustrates a typical, isolated three electrode Bio-potential
measurement setup. An isolation device is applied to
reduce the leakage current (I L ) and to increase the
common mode voltage range of the system. The
displacement current (I D ) induces a ground referred
voltage (V B ) to the body. V B decomposes to two voltages
inside the measurement system: the voltage between the
body and isolated amplifier common referred as common
mode voltage (V CM ) and the voltage between the
amplifier common and earth ground referred as isolation
mode voltage (V IM ). They are both interfering voltages
and can be treated separately. The resultant interference
caused by V B at the system output is [1] [2].
Eq. (1.1)
The reduction of interference due to V CM is accomplished
in two ways. 1) The effective CMMR of the pre-amplifier
is improved. 2) The V CM is actively or passively reduced
before it transforms into interference. The second
approach is done by decreasing the isolation capacitance
(C ISO ) of the isolation device and using a passively
coupled or actively driven third electrode to equalize the
voltage between the patient and the amplifier common.
[3][4]. The procedures decreasing the magnitude of V CM
give birth to V IM between the amplifier common and
earth ground. According to (1.1) interference caused by
V IM can only be reduced by increasing the IMRR of the
isolation device. [4]
Figure 1. A typical isolated three electrode bio-potential
measurement system. Potential divider effect divides V B
partly over Z E0 and partly over C ISO .
In this paper we will design a line powered
instrumentation system utilizing a method that allows
multi-channel signal transfer, control of the isolated
electronics and power transfer, using only capacitive
coupling over isolation. In addition, we will evaluate its
suitability for bio-potential measurements and capability
to reduce common mode interference.
Methods
In this new solution we exploit the fast development of
wireless technologies. The solution described below is
using Bluetooth as the communication technology,
because the Bluetooth concept is already at advanced
level of development. It must be pointed out that any
other suitable technology could be used. Figure 2
describes the block diagram of the system including the
central unit and a transducer unit.
As a typical configuration the new system includes a
grounded central unit and an isolated transducer unit
connected to each other by a coaxial cable (fig. 2). The
active transducer unit includes bio-potential amplifiers,
ADC’s, an MCU and the Bluetooth module acting as
slave. Bi-directional Bluetooth communication between
the transducer unit and the central unit allows transferring
IFMBE Proc. 2005;9: 101

Cardiovascular engineering
multi-channel measurement data and real-time control of
the transducer parameters during the measurement. In
addition to Bluetooth master the central unit includes a
serial interface for connecting data display device.
Figure 2. The block diagram of the novel instrumentation
system. Note that there is only one coaxial cable between
the central unit ant the isolated unit for communications
and power feed.
The operating power for the transducer unit is also
generated in the central unit. The 2 GHz PCS band
microwave power is fed to signal conditioning unit over
the isolation barrier using the same coaxial cable with the
2.4 GHz ISM band Bluetooth signal. The power and data
signals are separated from each other by diplexers acting
as band division filters (Fig. 2). At the isolated signal
conditioning module the PCS band microwave power is
rectified by an RF/DC-converter and the resulting DC
power is used as operating power for the isolated
electronics. Because the signal and power transfer take
place at very high frequency, C ISO can be very small.
Results
A prototype of the system was built to evaluate the
suitability of the method for bio-instrumentation. The
microwave power was efficiently transferred over the
isolation barrier with C ISO as low as 1.6 pF. At the
moment, the data rate between the Bluetooth modules is
115200 bits/s. The common mode response of the system
was evaluated with an analog circuit shown in figure 3.
C P was increased from its typical real value (~3 pF) to
150 nF to get a constant V B over the whole frequency
band. C B was 180 pF and Z E was 100 kW according to
worst case condition. The frequency response of the ratio
of V CM and V B was measured with three different values
of C ISO . The results are shown in figure 4.
Figure 3. Simplified schematic of the common mode test
circuit.
Figure 4. The ratio of V CM and V B as function of
frequency. Decreasing the C ISO by a decade decreases
V CM by 10-20 dB.
Discussion
The data rate of the system is limited by the current
Bluetooth Serial Port Profile (SPP) implementation. Still,
around 10 ksamples/s at 12 bit ADC resolution is
possible and that is enough for a variety of multi-channel
bio-potential applications. In figure 4 the experimental
results match with the simulations with higher values of
C ISO . When C ISO is minimized (1.6 pF) the distributed
capacitance between the isolated common and ground
becomes dominant. Further reducing the V CM demands
radical increase of the integration level of the isolated
electronics in the next stage of the project.
Conclusions
The novel instrumentation system described in this paper
has several benefits. The system is safe and has very
good common-mode performance due to very low
isolation capacitance, number of cables is small and realtime
control of the isolated electronics is possible.
References
[1] M.F.Chimene, R. Pallas-Areny, “A comprehensive
model for power line interference in biopotential
measurements”, IEEE Trans. Biomed. Eng., Vol. BME-
49, pp. 535-540, 2000.
[2] B. B. Winter and J. G. Webster, “Reduction of
interference due to common mode voltage in biopotential
amplifiers”, IEEE Trans. Biomed. Eng., vol. BME-30,
pp. 58–62, 1983.
[3] R. Pallás-Areny, “Interference-rejection
characteristics of biopotential amplifiers: A comparative
analysis”, IEEE Trans. Biomed. Eng., vol. 35, pp. 953–
959, 1988.
[4] D. E. Wood, D. J. Ewins, and W. Balachandran,
“Comparative analysis of power line interference
between two- or three-electrode biopotential amplifiers”,
Med. & Biol. Eng. & Comput., vol. 33, pp. 63–68, 1995.
IFMBE Proc. 2005;9: 102

Cardiovascular engineering
PARAMETRIC ASSESSMENT OF HRV IN CONGENITAL CENTRAL
HYPOVENTILATION SYNDROME: A CASE REPORT
T. Princi*, A. Accardo** and D. Peterec°
* Department of Physiology. University of Trieste, via Fleming 22, 34127 Trieste, Italy
** D.E.E.I., University of Trieste, via Valerio 10, 34100 Trieste, Italy
° Inst. of Physiology, Faculty of Medicine, University of Ljubljana, Zaloska 4, 1000 Ljubljana,Slovenia
princi@dfp.units.it
Abstract: Congenital central hypoventilation
syndrome (CCHS) is characterized by autonomic
nervous system dysfunction and decreased heart
rate variability (HRV). In the present study the
assessment of linear (FFT spectrum, Poincaré plot)
and non-linear (fractal dimension, β coefficient)
parameters of HRV indicated not only a
sympathetic and vagal cardiac dysfunction but
probably also an alteration of other autonomic
nervous mechanisms, which cooperate for the
functional integration between the cardiac and
respiratory functions.
Introduction
CCHS is a rare disorder of respiratory control [1]
that occurs in association with tumors of neural crest
origin and with symptoms of diffuse autonomic
nervous system dysregulation and/or dysfunction as
decreased HRV [2, 3] and reduced heart rate response
to exercise [4]. However, Woo et al. [2], by using
power spectral analysis, reported similar LF/HF ratio
values during wakefulness in both patients with CCHS
and controls. Congenital lack of central
chemosensitivity with absence of the ventilatory
response to hypercapnia is the main characteristic of
CCHS [1], but the pathophysiological mechanisms of
CCHS are still unknown. Failure of mechanisms that
integrate chemoreceptor inputs to the respiratory
centers is the current hypothesis [5].
The aim of the present study was to better delineate
the autonomic nervous control on the cardiac function
in one CCHS patient in comparison to a control by the
assessment of linear (FFT spectra, Poincaré plot) and
non-linear (fractal dimension, β coefficient)
parameters of HRV.
Methods
The study was performed in one female, 16 years
old, presenting a form of CCHS requiring night
respiratory assistance. This patient was compared to
one healthy volunteer of the same sex and age. The
heart rate (HR) was recorded (Polar S 810 HR
monitor) continuously for 20 min during daytime (4.00
p.m.), at rest in supine position while both subjects were
breathing spontaneously. The series of consecutive R-R
interval (tachogram) in function of beat numbers was
extracted and linearly interpolated in order to resample
the series at regular time intervals (500ms) for further
processing.
From the tachograms, FFT spectra and PSD were
calculated using the Hanning window on intervals of
1024 points. Low (LF: 0.04 - 0.15 Hz) and high (HF:
0.15 – 0.80 Hz) spectral bands were evaluated and the
LF/HF ratio was derived.
The Poincaré plot, able to measure the differences
among R-R intervals due to changes in vagal and
sympathetic modulation without the requirement for the
stationarity of the data [6], was also quantitatively
analysed. In this analysis, SD 1 parameter is used as a
marker of vagal influence, whereas SD 2 parameter
represents the more delayed R-R interval changes
correlated to sympathetic activity, and SD 1 /SD 2 ratio
indicates the vago/sympathetic balance. SD 1 and SD 2
represent the two axes of the best-fit ellipse that
contains the Poincaré points.
Linear regression analysis between log(Power) and
log(Frequency) was performed on the power spectrum
included between 0.004 Hz and 0.2 Hz, and the 1/f
spectral exponent, i.e. the β coefficient, was estimated.
The fractal dimension (FD) of HRV was
investigated as a possible indicator of the complex
interaction that might reflect the number of inputs to HR
controllers [7]. FD was evaluated by means of the
Higuchi’s algorithm on tracts of 120 consecutive RR
interval samples. The mean value of FD +/- SD was
considered in the analysis
Results
Figure 1 shows the tachograms, PSD behaviours and
Poincaré plots, evaluated in a patient with CCHS (top)
and in a control (bottom), during wakefulness at rest in
supine position at the same daytime (4.00 p.m).
In Table 1 R-R intervals as well as FFT spectral
parameters, Poincaré plot parameters (SD 1 , SD 2 ), FD
and β values, related to the patient (CCHS) and control,
are showed.
IFMBE Proc. 2005;9: 103

Cardiovascular engineering
Figure 1. Tachograms, β coefficient (linear slope) and Poincaré plots in the CCHS patient (top) and in a control (bottom).
Table 1. Linear and non-linear parameter values in
CCHS and Control subject. For R-R intervals and FD
values the mean +/-SD are reported.
CCHS subject Control
RR interval (ms) 752+/-35 1026+/-58
LF (ms 2 ) 156.3 725.5
HF (ms 2 ) 101.4 695.9
LF/HF (-) 1.54 1.04
Total Power (ms 2 ) 467.6 2141.9
SD 1 (ms) 13.2 46.8
SD 2 (ms) 47.0 67.3
SD 1 /SD 2 (-) 0.28 0.70
β coefficient 1.18 0.75
FD (-) 1.49+/-0.08 1.79+/-0.09
Discussion
This study, confirming the results of other Authors
[3], reports higher HR levels and reduced HR overall
variability, at rest in supine position, in CCHS patient
in comparison with a healthy volunteer. The CCHS
subject was characterized by lower values of HF
spectral component and SD 1 parameter, indicating a
decrease of vagal cardiac influence. Furthermore, the
LF component and SD 2 parameter were reduced even
less than HF and SD 1 parameters, while the LF/HF
ratio presented a higher value as expression of
sympatho-vagal dysregulation in this syndrome with
dominant sympathetic activity and prevalent vagal
withdrawal. The FD as well as β values suggest lower
cardiac complexity [8] and therefore a lack of
functional integrity of autonomic control mechanisms in
CCHS subject in comparison to the healthy volunteer.
In conclusion, linear parameters of HRV as well as nonlinear
dynamics of cardiac activity indicate not only a
sympathetic and vagal cardiac dysfunction in CCHS but
probably also an alteration of other autonomic nervous
mechanisms, which cooperate for the functional integration
between the cardiac and respiratory functions.
References
[1] WEESE-MAYER D. E., SHANNON D. C., KEENS T. G., SILVESTRI J. M.
(1999): ‘American Thoracic Society Statement. Idiopathic congenital
central hypoventilation syndrome. Diagnosis and management’, Am. J.
Respir. Crit. Care Med., 160, pp. 368-373
[2] WOO M. S., WOO M. A., GOZAL D., JANSEN M. T., KEENS T. G., HARPER
R. M. (1992): ‘Heart rate variability in congenital central hypoventilation
syndrome’, Pediatr. Res., 31, pp. 291-296
[3] TRANG. H., GIRARD A., LAUDE D., ELGHOZI J. L. (2005): ‘Short-term
blood pressure and heart rate variability in congenital central
hypoventilation syndrome (Ondine’s curse)’, Clin. Sci. (Lond.), 108(3),
pp. 225-230
[4] SILVESTRI J. M., WEESE-MAYER D. E., FLANAGAN E. A. (1995):
‘Congenital central hypoventilation syndrome: Cardiorespiratory
responses to moderate exercise, simulating daily activity’. Pediatr.
Pulmonol., 20, pp. 89-93
[5] SPLENGER M. C., GOZAL D., SHEA S. A. (2001): ‘Chemoreceptive
mechanisms elucidated by studies of congenital central hypoventilation
syndrome’, Respir. Physiol., 129, pp. 247-255
[6] TULPPO M. P., MÄKIKALLIO T. H., TAKALA T. E. S., SEPPÄNEN T., and
HUIKURI H. V., (1996): ‘Quantitative beat-to-beat analysis of heart rate
dynamics during exercise’, Am. J. Physiol., 271, pp. H244-H252
[7] NAKAMURA Y., YAMAMOTO Y., MURAOKA I. (1993): ‘Autonomic
control of heart rate during physical exercise and fractal dimension of
heart rate variability’. J. Appl. Physiol., 74(2), pp. 875-881
[8] GOLDBERGER A. L. (1996): ‘Non-linear dynamics for clinicians: Chaos
theory, fractals, and complexity at the bed-side’, The Lancet, 347, pp.
1312-1314
IFMBE Proc. 2005;9: 104

Cardiovascular engineering
DIFFERENT APPROACHES OF MEASUREMENT OF HEMODYNAMIC BY
ELECTRICAL IMPEDANCE PLETHYSMOGRAPHY METHOD
A. Stankus 1
1 Department of psychosomatic disturbances and sleep research, Institute of Psychophysiology and
Rehabilitation, KMU, Palanga, Lithuania
albstan@ktl.mii.lt
Abstract
The goal of the presented study is an elaboration of
methodology for a direct measurement of circulation
by means of volume units. With this aim in mind
particular hardware has been elaborated. A new
principle in measurement of relative changes of the
pulse wave amplitude allows registering of a
electroplethysmogram. It enabled to measure changes
directly in the volume of tissues and to assess
hemodynamics of various body parts by relative
(ml/100ml) volume units, and to make this method
more precise. This study enabled to improve this
method for measurement of the blood circulation in
legs. This method can be used for diagnostic purposes
in vascular pathology.
Introduction
Popularity of the impedance plethysmography method is
based on its simplicity. However, some biophysical and
methodological problems occur in the impedance
cardiography, performing measurements of the stroke
volume. During the last 30 years, the main attention of
the investigators was focused on the comparison of the
absolute stroke volume values, measured by this method
and defined by other methods. Correlation values
obtained in investigating persons in 23 studies were
various within the range from 0.49 to 0.97. Because of
physical reasons, results from earlier mentioned
dependences could not be absolutely exact. Use of this
method for investigations of the blood circulation in
lower limbs is not doubtful. The values used in
calculation of limbs volume are as follows: specific tissue
impedance (ρ, Ω*cm), distance between electrodes (l,
cm) and constant impedance (Z 0 Ω).
Specific impedance of tissues is considered constant and
equals to 135-150 Ω*cm. However, experiment shows
that it depends on blood hematocryte changes from 22 to
66%. Use of the specific impedance in the particular
cases makes this method more difficult. In real conditions
it is very difficult to hold parallelism of the placed band
electrodes to make a precise measurement. The constant
impedance is measured in many cases incorrectly, as it is
read from the indicator. Thus, there are too many
approximate calculations and not enough precise
measurements. Among three parameters mentioned
above, the measurement of the specific impedance,
depending from tissues and properties of blood (in the
other words – homogeneity of the object to electric
conductivity), is the most difficult one.
Methods
New measurement of approaches to hemodynamic is
based on the precondition, that we can directly measure
the ratio by tetrapolar means, as it is shown on equation:
∆V/V 0 =- k o (∆Z/Z 0 ). It is easy to make it using a
computer or/and electronic method [1]. Assuming, that
the specific impedance of the measured area is constant, I
measured a direct ratio between the alternating volume
changes and the total volume measured by the electrical
method [1]. Pulsated filling of the measured segment
with blood is measured in the following way: ∆V/V 0 =
ml/ml *100 = %. Thus, I got units of measurement used
in physiology – ml/100ml. The device is calibrated by
change to 0.1% of the main impedance by a parallel
connection of the resistor. Using this method, the
maximum defined values of amplitude indicate how
many ml volume increases in each 100 ml segment
located between electrodes. The amplitude changes
without additional calculations showed relative changes
of the volume in time. Circumferential electrodes were
placed over the segments of a thigh, a knee-joint, a shin
and a foot.
Results
I investigated 43 healthy subjects (HSs) at the age of 30 -
69 and 42 patients (Pts) with differently stage of
endarteritis obliterans, at the age of 40 - 69.
Table 1: Distribution of ratio resistance in leg
Healthy
Patients
Region Left, % Right,% Left, % Right,%
Thigh 0.115
±0.006
0.117
±0,007
0.099
±0.007*
0.095
±0.007*
Kneejoint
0.165
±0.010
0.169
±0.011
0.121
±0.008*
0.124
±0.008*
Shin 0.115 0.116 0.079 0.086
IFMBE Proc. 2005;9: 105

Cardiovascular engineering
Foot 0.103
±0.010
±0.073 ±0.083 ±0.006* ±0.007*
0.106
±0.080
0.073
±0.006*
0.074
±0.008*
Table 1 shows the results of measurement of magnitude
of the ratio of resistances. The magnitude of maximum
ratio between of amplitude pulse wave and constant
impedance was maximum in a knee-joint. In other region
this ratio was less. The statistical analysis of the main
parameters showed very distinct and significant
differences between the HSs and the Pts (p

Medical ultrasound
Fig. 1. Impedance characteristics of piezoelectric disk
D/t=10 before and after fitting.
Search for the optimal geometry of piezoelement
Mode coupling can also be used for rising k eff and
effectiveness of piezoelement, since some coupled
modes are more effective with regard to pure ones.
Having individual set of constants of pjezomaterial,
accurate modeling and simulation by FEM was possible
with the goal to find the best geometry parameters
regarding to k eff . The results of search for the best widthto-thickness
ratio for 2D pjezo-plates and 3D bars of
quadratic cross section are presented in Fig. 2.
Comparison of results will show that k eff for 3D bar can
be reached 4-7% better than in case of 2D plate. The
method proposed had allowed the answer the practical
question: which type of pjezo-ceramic is most suitable
for coupled mode operation. For example PZT-5J
ceramics k eff seems to be highest. It can be explained
that for that material bar mode is interacting with
transversal mode with involvement (what is typical to
that kind of pjezo-ceramic) of thickness mode. Such
complex mode is more effective but at the same time
more sensitive to the W/t ratio (Fig. 2.).
Conclusions
The success and adequacy of modeling of
pjezoelements by FEM depend on the accuracy of the
full set of elastic, pjezoelectric and dielectric constants
used for calculations. The adequate set of individual
parameters can be found by fitting of experimental and
simulated frequency dependencies of electrical
impedance of pjezo-sample at coupling vibration mode.
The maximum of electromechanical coupling factor
depends on aspect ratio ant it is different for different
type of piezoceramic. The method described potentially
can contribute to the individual computer assisted
design of highly effective multi-element pjezotransducers,
including composite ones to be applied for
medical scanners.
References
b)
Figure 2. Dependence of electromechanical coupling
factor k eff in piezoelements of 2D plate (a), and 3D bar
(b) shape (W- width, t- thickness)
a)
[1] SATO J., KAWABUCHI M., FUKUMOTO A.
(1979): ‘Dependence of the electromechanical
coupling coefficient on the with-to-thickness ratio of
plank-shaped transducers used for electronically
scanned ultrasound diagnostic systems’, J. Acoust.
Soc. Am.. 79 (6). pp.1609-1611.
[2] WOJCIK G.L., VAUGHAN D.K., ABBOUD N.N.,
MOULD J. (1993): ‘Electromechanical modeling
using explicit time-domain finite elements‘, IEEE
1993 Ultr. Symp. Proc., 93 (2). pp. 1107-1112.
[3] IEEE Standard on Piezoelectricity. IEEE-Std 176-
1987.
[4] CARCIONE L., MOULD J., PEREYRA V.,
POWELL D., WOJCIK G. (2001): 'Nonlinear
inversion of piezoelectrical tranducer impedance
data', J. Comp. Acoust.. 2001 (3).
[5] KYBARTAS D., LUKOSEVICIUS A. (2002):
‘Determination of piezoceramics parameters by the
use of mode interaction and fitting of impedance
characteristics’. Ultragarsas. Kaunas: Technologija.,
2002 4(45), pp. 22-28.
IFMBE Proc. 2005;9: 108

Medical ultrasound
ULTRASOUND SIGNAL ENHANCEMENT VARYING MICROBUBBLE
CONCENTRATION AT VERY LOW MECHANICAL INDICES
S. Casciaro 1,2 , R. Palmizio Errico 2 , F. Conversano 2 , A. Maffezzoli 3 , A. Sannino 3 , A. Distante 1,2
1
Institute of Clinical Physiology, National Council of Research, Lecce, Italy
2
Bioengineering Division, Euro Mediterranean Scientific Biomedical Institute, Brindisi, Italy
3 Innovation Engineering Department, Lecce University, Lecce, Italy
casciaro@ifc.cnr.it
Abstract: In this study we analyze the differences in
contrast enhancement produced varying the
concentration of phospholipidic ultrasound (US)
contrast agent (CA) microbubbles at low frequency
(2.5 MHz) and low mechanical index (MI=0.08). Five
CA concentrations (range 0.013-0.100 µL/mL) were
tested in a new hydrogel-based phantom and
insonified by a linear transducer. Backscattered
radiofrequency signals were acquired employing an
acquisition system able to get the full raw signal. We
have observed that the highest signal intensity was
found for the lowest tested concentration; signal
intensity also presented a strong linear decreasing
correlation (r=0.995) with CA concentration in the
range 0.013-0.033 µL/mL. This is a preliminary step
toward a complete modelling of USCA signal in
various experimental conditions.
Introduction
In recent years the potential biomedical applications
of microbubble ultrasound (US) contrast agents (CA)
have considerably increased [1-2]. On the other hand,
many discussions have been raised concerning safety in
the use of such CA [3], even after huge multicenter
studies aimed to investigate the safety of these agents in
several body districts [4].
As we will show in this work, the last technological
innovations in the field of signal processing and
radiofrequency (RF) spectrum analysis revealed that a
strategy to possibly improve the safety of microbubbles
is both the reduction of CA injection dose and of
mechanical indices.
In this work we present the preliminary results of an
in vitro system that can be developed and modulated in
order to be able to study the microbubble acoustical
behaviour in almost all kind of human tissues, being
also able to cover all vessel sizes and to reproduce
different vascular system conditions in terms of
geometrical configurations and flow velocity ranges.
In this study we assessed the differences in signal
contrast enhancement obtained by varying the
concentration of a phospholipidic US contrast agent
(supplied by Bracco Research SA, Geneva, Switzerland)
at low frequency (f) and very low mechanical index
(MI).
Materials and Methods
A new hydrogel based phantom was developed
having a sound propagation velocity similar to that of
the human liver: two hydrogels were used for the matrix
tissue and for the vessel wall. The phantom was 6 cm
deep, 5 cm long, 8 cm wide and it had two 1-mm
diameter vessels, both placed at 2 cm from the upper
surface. A saline-diluted microbubble infusion at room
temperature (25 °C) was pumped through the phantom
vessels by a peristaltic pump (Peri-Star Model 500304,
WPI Inc., FL, USA) at a constant flow rate of 8
mL/min. Five different CA concentrations were tested:
0.013, 0.025, 0.033, 0.050, 0.100 µL/mL. These
concentration values belong to the high concentration
range indicated by the manufacturer (0.002-0.100
µL/mL).
An US transducer (LA 532, Esaote Spa, Florence,
Italy) was positioned on the top of the phantom, so that
the imaging plane resulted perpendicular to the vessels.
The transducer was connected to a digital ecograph
(Megas GPX, Esaote Spa, Florence, Italy) linked to a
platform for RF spectrum acquisition and analysis
(FEMMINA, developed by Florence University), able to
get the full raw signal of the probe.
The phantom was insonified with 2.5 MHz US
pulses (MI=0.08). RF signals were sampled at 40 MHz
and a sequence of 180 data frames was acquired for
each concentration.
Off-line analyses were performed using the
prototype software (Fortezza, supplied by Florence
University). Raw data were not actually filtered. By
means of an ad hoc implemented Fortezza algorithm,
we selected a ROI approximately equivalent to a square
of 0.5 mm, covering exclusively the microbubble flow.
Since each data frame was 4.5 cm wide and composed
of 180 scan lines, 3 lines passed through a 0.5-mm wide
box. Assuming also a speed of sound of 1560 m/s, the
width of the time-gate equivalent to a 0.5-mm capture
gate was equal to 0.64 µs, resulting in 25 data points
along each of the selected scan lines. Consequently, the
considered ROI was composed of 75 data points.
The average intensity of the ROI was calculated for
each acquired frame and recorded in a file
corresponding to the considered concentration. Data
IFMBE Proc. 2005;9: 109

Medical ultrasound
were then analyzed, averaged and plotted versus CA
concentration.
For RF spectrum analysis, another Fortezza
algorithm was used to calculate backscatter values of
single components extracted from the Fast Fourier
Transform (FFT) curve. Before FFT calculation, the
selected raw data were zero-padded to 4096 points. FFT
was calculated for the 3 tracks of the ROI and averaged
to obtain the mean FFT curve of the considered ROI.
Then subharmonic, fundamental, second harmonic and
third harmonic backscatter values were extracted and
averaged over the corresponding frame sequence. These
values were plotted in an histogram versus CA
concentration.
Results
As expected, CA average signal intensity is always
higher than pure saline solution and the highest value
was found for the lowest concentration (Figure 1).
The fundamental component shows the highest
value for every dilution (Figure 2). The maximum is at
concentration 0.013 µL/mL and then intensity decreases
till concentration 0.050 µL/mL. Finally it rises again for
concentration 0.100 µL/mL. This trend confirms what
showed in Figure 1.
confirmed by backscatter values extracted from the FFT
curve (Figure 2).
Conclusions
In the concentration range 0.013-0.100 µL/mL,
employing US pulses with f=2.5 MHz and MI=0.08, the
tested CA shows the highest intensity for the lowest
concentration. Average signal intensity also presents a
strong linear decreasing correlation (r = 0.995) with CA
concentration in the range 0.013-0.033 µL/mL, while
for higher microbubble concentrations this linear
relationship disappears.
Further investigations are needed to obtain the
microbubble signal response at lower concentrations, in
order to cover all possible dilution ranges and develop
new analytical models of microbubble acoustic
behaviour.
Figure 2: Histogram of average backscatter intensity
versus CA concentration.
References
Figure 1: Plot of ROI average intensity versus CA
concentration (p

Medical ultrasound
FREQUENCY EFFECTS ON ECOCONTRAST AGENTS SIGNAL
BEHAVIOUR AT LOW MECHANICAL INDICES
R. Palmizio Errico 2 , S. Casciaro 1,2 , F. Conversano 2 , E. Casciaro 1,2 , G. Palma 2 , A. Distante 1,2
1
Institute of Clinical Physiology, National Council of Research, Lecce, Italy
2
Bioengineering Division, Euro Mediterranean Scientific Biomedical Institute, Brindisi, Italy
palmizio@isbem.it
Abstract: An in vitro system based on an apparatus
for the data analysis that provides a wide range of
settings was developed to investigate the signal
backscatter of contrast agents (CA). A commercial
echograph (MEGAS GPX) was used to insonate
microbubbles through a linear array probe (LA 532)
and a prototypal acquisition system was used to store
radiofrequency backscatter signals using no filters.
A new phospholipidic ultrasound CA of the last
generation was studied pumping contrast agent
stirred solution through the vessels of a hydrogel
phantom simulating human liver and capable of
reproducing several vessel sizes in order to
investigate insonification frequency effect. The
transmit frequencies were 2.5, 5, 7.5 MHz. Each
frequency effect was studied at four low mechanical
indices (MIs, 0.08, 0.1, 0.2, 0.3). The backscatter
intensity of the subharmonic, fundamental and
second harmonic frequency was calculated and
extracted from a ROI located inside to vessel cavity
through a just developed algorithm. The agent
provided the best backscatter intensity in the
fundamental component of 2.5 MHz transmit
frequency and in the subharmonic component of 5
and 7.5 MHz transmit frequencies.
Introduction
Over the past years, in vitro experiments [1-4] have
improved the understanding of the interaction between
contrast microbubbles and the US beam, but the effort
to accomplish a full understanding of the phenomena is
by no means complete. However an experimental setup
that can cover a wide range of settings (e.g. flow rate,
transmit frequency, mechanical index) similar to that
used in diagnostic ultrasound applications is yet to be
reported. The interaction of gas-filled microbubbles
with an ultrasound (US) beam has become, in recent
years, a major area of research. This is largely due to the
diversity of the reactions between the microbubbles and
the US beam and the subsequent implications on the
implementation of US diagnostic techniques.
In this work we present preliminary results obtained
with an in vitro system developed to study microbubble
acoustical behaviour using a phantom that can
reproduce all kind of human tissues and modulate the
vessel sizes and using an apparatus for the data analysis
that provides a full range of settings. The purpose of this
study was to determine differences in contrast
enhancement given by the experimental intravenous
contrast agent (supplied by Bracco Research SA,
Geneva, Switzerland), at constant dilution, employing
different transmit frequencies and at vary low
mechanical indices (MI).
Materials and Methods
The phantom (developed in collaboration with the
Material Engineering section of Lecce University) was a
custom-designed tissue-mimicking phantom, made of
hydrogel having a sound propagation velocity very
close to the human liver value. It contained two 1-mm
diameter vessels placed at the same stand-off distance
within the phantom, both at 2 cm from the upper
surface.
A stirred microbubble dilution (1:80000) was
pumped, at a constant flow rate, into phantom vessels
and insonated through a linear array probe (LA 532,
Esaote, Florence, Italy), positioned on the top of the
phantom so that the imaging plane resulted
perpendicular to the vessels, and connected to a digital
ecograph (Megas GPX, Esaote, Florence, Italy), in a
research configuration suitable for insonification
parameters adjustment. The ecograph was externally
linked to a prototype for radiofrequency (RF) analysis
(FEMMINA, developed by Florence University), able to
get the full raw signal of the probe [5]. The received
signals were digitized at 40 MHz. Each frequency (2.5,
5, 7.5 MHz) was tested at four MIs (0.08, 0.1, 0.2, 0.3)
and the raw data were acquired in sequences of 100
frames and they were stored in FEMMINA for further
off-line analysis.
Stored raw data were used to reconstruct images
data through the Fortezza software (supplied by
Florence University). This software was also utilized to
implement a new algorithm to properly choose the
Region Of Interest (ROI) inside the vessel cavity. We
selected a ROI approximately equivalent to a square of
0.5 mm. Since each data frame was 4.5 cm wide and
composed of 180 scan lines, 3 lines passed through a
0.5-mm wide box. Assuming also a speed of sound of
1560 m/s, the width of the time-gate equivalent to a 0.5-
mm capture gate was equal to 0.64 µs, resulting in 25
data points along each of the selected scan lines.
IFMBE Proc. 2005;9: 111

Medical ultrasound
Consequently, the considered ROI was composed of 75
data points. In order to evaluate the combined effect of
frequency and MI, another Fortezza algorithm was
developed to measure backscatter intensity by FFT
calculation over the defined ROI.
Before FFT calculations were carried out, the raw
data corresponding to the defined ROI and selected by
means of a 25-point “Rect” window were zero-padded
to 4096 points to increase the frequency resolution of
the spectra. FFT curves were averaged to obtain and
visualize the FFT averaged curve over the ROI. This
calculation has been repeated three times for every
acquired frame sequence, corresponding to a specific
combination of frequency and MI values. Each time the
single amplitude value of a different FFT component
(subharmonic, fundamental and second harmonic) was
extracted, visualized and recorded in a Fortezza
proprietary format file.
Results
In figure 1, the relationship between single FFT
components and the transmit frequency at the various
tested mechanical indices is displayed.
Av. Back. Int. (dB)
Av. Back. Int. (dB)
105
100
95
90
85
80
75
70
65
60
105
100
95
90
85
80
75
70
65
MI = 0.08
2,5 5 7,5
Frequency (MHz)
MI = 0.2
2,5 5 7,5
Frequency (MHz)
Av. Back. Int. (dB)
105
100
95
90
85
80
75
70
65
105
100
95
90
85
80
75
70
65
Av. Back. Int. (dB)
MI = 0.1
2,5 5 7,5
Frequency (MHz)
MI = 0,3
2,5 5 7,5
Frequency (MHz)
Fig. 1: Histogram of Average Backscatter Intensity
versus transmit frequency (MHz) ( subharmonic
fundamental second harmonic components)
In every histogram the highest intensity is observed
at 2.5 MHz FFT component for 2.5 MHz and 5 MHz
transmit frequencies, while at 7.5 MHz the highest value
is observed at 3.75 MHz FFT component.
In fig. 2-a and 2-b fundamental FFT component
corresponding to 2.5 MHz transmit frequency is the
highest, while, increasing MI, it is gradually overtaken
by the subharmonic component corresponding to 7.5
MHz transmit frequency
Discussion
This preliminary study was meant to determine the
frequency effect of a new contrast agent that flowed at
constant dilution through phantom vessels based on
hydrogel simulating human liver ecocontrast. The
transmit frequency influences acoustic microbubble
behaviour in fact the fundamental and second harmonic
backscatter intensities decrease with increasing
frequency, while subharmonic components increase.
Then fundamental component and second harmonic
values don’t increase in a remarkable way arising MI
but second harmonics values increase arising MI until
0.2 and then remain approximately constant till MI
equal to 0.3.
Conclusion
The last generation phospholipidic contrast agent
shows a significant contrast enhancement for each
tested combination of transmit frequency and MI at
1:80000 dilution. In particular, at 2.5 MHz transmit
frequency this contrast agent gives the best backscatter
intensity in the fundamental component and its value
doesn’t increase in a remarkable way arising MI, so it
should be possible to successfully use this microbubble
suspension in conventional fundamental B-mode
imaging already at the lower MI values. On the other
hand, for applications that require 5 MHz or 7.5 MHz
transmit frequencies, the best contrast enhancement
achievable with the tested contrast agent is observed in
correspondence of subharmonic components, for which
the values increase arising MI until 0.2. Finally we can
state that, to effectively employ this contrast agent
dilution at 5 MHz or 7.5 MHz transmit frequencies, it
should be necessary to work in subharmonic imaging
modality with MI lower than 0.3.
References
[1] FRINKING PJA, DE JONG N AND CESPEDES EI.
(1999): ‘Scattering properties of encapsulated gas
bubbles at high ultrasound pressures’ J. Acoust.
Soc. Am., 105, pp. 1989-1986
[2] SBOROS V, MORAN CM, PYE SD AND MC DICKEN
WN (2003): ‘The behaviour of individual contrast
agent microbubbles’, Ultrasound Med. Biol, 29, pp.
687-94
[3] SBOROS V, MORAN CM, PYE SD AND MC DICKEN
WN (2004): ‘An vitro study of a microbubble
contrast agent using a clinical ultrasound imaging
system’, Phys. Med. Biol., 49, pp 154-173
[4] PORTER TR, OBERDORFER J, RAFTER P, LOF J AND
XIE F. (2003): ‘Microbubble responses to a similar
mechanical index with different real-time perfusion
imaging techcniques’, Ultrasound Med. Biol., 29,
1187-92
[5] SCABIA M., BIAGI E., MASOTTI L. (2002): ‘Hardware
and Software Platform for Real-Time Processing
and Visualization of Echographic Radiofrequency
Signals’, IEEE Trans. UFFC, 49, pp. 1444-1452
Further references are available on request.
IFMBE Proc. 2005;9: 112

Medical ultrasound
Ultrasound contrast for perfusion studied.
Marcus Ressner a , Adriana Kvikliene b , Lars Hoff c , Rytis Jurkonis b , Tomas
Jansson d , Birgitta Janerot-Sjöberg e , Arunas Lukosevicius b , Per Ask a
a Departments of Biomedical Engineering and e Department of Clinical Physiology Linköping University, Linköping,
Sweden, b Institute of Biomedical Engineering, Kaunas University of Technology, Kaunas, Lithuania,
c Faculty of Science and Engineering, Vestfold University College, Horten, Norway,
d Department of Electrical Measurements, Lund University, Lund, Sweden.
Abstract: Our model driven approach is used to gain
better knowledge of the different processes involved in
the generation of the backscattered contrast echo. It
can be divided into three separable stages: linear and
non-linear wave propagation in tissue, the resulting
echo from the pulse interaction with the contrast
microbubble, and the propagation of the scattered
echo. Our tools are computer simulation and model
experiment.
Introductrion
Functional images with combined information of
blood flow and motion are applicable within a wide range
of basal science and physiology.
In a long term perspective the goal of our ultrasound
contrast research is to find a new model driven approach
for estimation of myocardial perfusion. An aim is to
develop nonlinear simulation of the contrast bubble and
ultrasound wave interaction as well as wave propagation
and to design an in vitro model including a perfusion
phantom for ultrasound contrast measurements. We plan
to use the simulation and in vitro model to evaluate and
optimize the wash-in technique after bubble destruction.
Methods and Results
Modelling of plane wave propagation was carried out
in successive forward steps applying the operators of
nonlinear distortion, attenuation and speed dispersion.
The operator of nonlinear distortion is based on the time
domain relation developed by Remenieras, [1] and the
operator of attenuation and speed dispersion is based on
spectral decomposition and modification methods
developed by He, [2].
Modelling of the signal field in a nonlinear medium
was carried out under the assumptions that the ultrasound
intensity is weak because of safety reasons related to high
mechanical index (MI) and ultrasound contrast agents,
and to increase contrast bubble survival in acoustic beam.
Only nonlinearity effects originating from nonlinear terms
in tissue elasticity that relates the pressure and the
material compression (expansion) are taken into account.
We further assume the medium is a nonlinear
homogeneous liquid with power law frequency function
of attenuation and sound speed dispersion.
Using a method developed by Jurkonis and
Lukosevicius [3] based on the spatial impulse response of
an aperture (SIRA) as well as spectral decomposition and
modification methods we calculated the acoustic pressure
pulse field in a nonlinear media. An algorithm of Spatial
Superposition of Attenuated Waves (SSAW) method was
adopted for field simulation in nonlinear medium. The
acoustic pressure waveform in a field point is calculated
by adding contributions from elementary waves and is
modified in steps accounting for nonlinear propagation,
attenuation and speed dispersion.
Simulations of the contrast microbubble response to
the incident pressure pulse were based on the Rayleigh-
Plesset equation of motion for the surrounding liquid with
the addition of a radiation damping factor [4]. The
pressure at the bubble surface was calculated for bubbles
encapsulated in a very thin viscoelastic shell, assuming an
exponential stress-strain relationship [5]. The acoustic
bubble response was calculated at a normalized distance.
The shell material parameters used in the simulations
are based on the properties of the contrast agent Sonazoid
(GE-Amersham Health, Oslo, Norway).
Interaction with nonlinearly distorted pulses was
studied in water for series of nonlinearly distorted pulses
and was measured by a needle hydrophone in an
experimental setup. Simulations of the interaction
between contrast bubbles and the sampled ultrasound
pressure pulse were performed to yield bubble echoes that
correspond to in vitro measures.
The simulations of nonlinear pressure pulses
correspond well to the in vitro hydrophone measurements
and shows that the attenuation will reduce the effects of
pulse distortion due to nonlinear wave propagation. As a
consequence, the nonlinear distorted pulse will have a
reduced energy content compared to the non distorted
pulse as more energy have been shifted to higher
frequencies and therefore suffered from a stronger
attenuation. The result of bubble response simulations is
presented in Figure 3. The simulations are performed with
distorted pulses and with theoretically generated nondistorted
pulse that interacted with the bubble model. The
increase of the second harmonic frequency amplitude for
the nonlinear distorted pulse is about 3 dB. The difference
IFMBE Proc. 2005;9: 113

Medical ultrasound
of the second harmonic amplitudes of the backscattered
pulses will increase with higher acoustic pressures but is
not detectable at pressure levels below 200 kPa.
Conclusion
The presented results of the pulse wave model in a
nonlinear, attenuating and speed-dispersive media look
reliable enough for comparative estimation of particular
effects. Assumptions taken make simulation quite simple
and suitable for model based processing of echographic
signals obtained with contrast agents. Quantitative
theoretical analysis as well as in-vitro experiments in soft
tissue mimicking medium show that the absorption
strongly reduces the nonlinear distortion originated in
tissue, as the higher frequency components are more
absorbed than the fundamental ones. Also theoretical
simulations show that contrast bubbles interaction with
excitation pulses is the main cause of nonlinear
distortions, and a 2-3 dB increase of second harmonic
amplitude depends on nonlinear distortions of incident
pulse.
Acknowledgement
This study was supported by the Swedish National
Research Council, Swedish National Foundation for
Strategic Research by the program Cortech and by the
Vinnova Competence Center Nimed.
REFERENCES
[1] J.P. Remenieras, O.B. Matar, V. Labat, F.
Patat, Time-domain modeling of nonlinear
distortion of pulsed finite amplitude sound
beams, Ultrasonics 38: 305–311, 2000.
[2] P. He, Simulation of ultrasound pulse
propagation in lossy media obeying a
frequency power law, IEEE Trans. Ultrason.
Ferroelect. Freq. Contr. 45 (1) 114–125, 1998.
[3] R. Jurkonis, A. Lukosevicius, New method of
spatial superposition of attenuated waves for
ultrasound field modelling, Ultrasonics
40: 823–827, 2002.
[4] L. Hoff, Acoustic characterization of contrast
agents for medical ultrasound imaging, Kluwer
Academic Publishers, Dordrecht, The
Netherlands, 2001.
[5] L. Hoff, Nonlinear response of Sonazoid,
Numerical simulations of pulse-inversion and
subharmonics, in: Proc. IEEE Ultrason. Symp.,
pp. 1885–1888, 2000.
IFMBE Proc. 2005;9: 114

Medical ultrasound
AN ULTRASONIC METHOD FOR DETECTION OF FLUID PROPERTIES
IN THE PARANASAL SINUSES
T. Jansson*, B. Ask*, P. Walfridsson*, P. Sahlstrand-Johnson**, H. W. Persson*,
N.-G. Holmer***, and M. Jannert**
* Department of Electrical Measurements, Lund University, Lund, Sweden
** Department of Oto-Rhino-Laryngology, Malmö University Hospital, Malmö, Sweden
*** Department of Biomedical Engineering, Lund University Hospital, Lund, Sweden
tomas.jansson@elmat.lth.se
Abstract: We propose a method for detection of the
degree of infection in the paranasal sinuses utilizing
a previously published method whereby the viscosity
in a sealed container may be measured using an
ultrasound Doppler method. As ultrasound
propagates in a liquid medium, due to attenuation,
the resulting pressure gradient will cause the liquid
to move in the propagation direction - the wellknown
effect of acoustic streaming. The streaming velocity
will, for a given acoustic output, be proportional to
the viscosity of the fluid. In this study, we verify that
acoustic streaming can be induced in an
anthropomorphic sinus phantom cast from a human
cranium. The sinus phantom was made from agar
with added graphite providing sound attenuation
prior to the sinus cavity corresponding to an in vivo
situation. A number of water-glycerol solutions with
scattering particles, were prepared to mimic a
clinically interesting range of viscosities (7-47 mPas).
Using a 4.2 MHz continuous wave Doppler probe,
clearly detectable Doppler shifts in the range of 6.5
to 20 Hz were recorded. A linear relationship was
found between the Doppler shifts and 1/viscosity
(R 2 =0.94, corrected for the square-law dependence of
sound speed variation due to varying glycerol
concentration).
Introduction
In the case of a sinus infection, a more or less
viscous fluid may accumulate in the sinus. The
condition is often treated by irrigating the sinus, which
involves penetrating into the sinus cavity with a thick
needle. Needless to say, this is a very uncomfortable
procedure for the patient. Further, if the fluid is serous
(with a low viscosity), the treatment is unnecessary, as
the fluid will resorb spontaneously, whereas only if it is
mucous (and thereby highly viscous), antibiotics is
called for. Today, fluid can be detected using either
ultrasound or X-rays, but nothing can be said about
whether the fluid is serous or mucous. A way of noninvasively
determining the fluid properties would
naturally be of great benefit for both patient and society
in terms of reduced cost.
Dymling et al [1] suggested a method to detect
viscosity changes in sealed containers using ultrasound
†
Doppler. The emitted sound field induces acoustic
streaming in the fluid, and if the fluid contains particles
with other acoustic impedance than the surrounding
fluid, the scattered waves will give rise to a Doppler
shifted signal. The idea proposed here, is that this
method also could be used for detection of the severity
of sinus infection.
This initial study will examine whether acoustic
streaming can be generated in an anthropomorphic sinus
phantom, by modelling the proper shape of the sinus,
and include proper acoustic attenuation prior to entering
the sinus.
Materials and Methods
The maximum flow velocity v max induced by
acoustic streaming is
v max
= Ir2 a
ch Y
where I is the sound intensity, r is the radius of the
sound beam, a is the sound absorption coefficient, h is
the viscosity, c is the sound speed, and Y is a constant,
depending on the geometry of the sound beam and the
vessel that holds the liquid.
Since the Doppler shift and not the velocity is
measured, the combination of the equation above and
the Doppler equation, gives that Df µ 1/(c 2 h). In order
to compensate for the c 2 dependency a constant k was
introduced in the measurements, as the ratio of a
nominal sound speed and the actual sound speed of the
liquid sample in question (see below). As an estimate of
the maximum speed, the mean velocity was used since
the measured spectrum is rectangular.
A phantom of the sinus was manufactured in the
following way. From a human cranium, a cast was made
of the sinus cavity. This “plug” was then used to form
the mould of tissue-equivalent material, in which fluid
could be filled. As tissue-equivalent material, agar was
used with added graphite powder to serve as attenuating
component. The agar-based phantom was manufactured
by adding agar (Bacto-Agar, Difco Laboratories,
Detroit, MI) and sodium benzoate to distilled water at
concentrations of 40 g and 5 g per liter, respectively.
IFMBE Proc. 2005;9: 115

Medical ultrasound
After the solution was heated to 90°C and clarified,
90 g graphite (type 1.04206.2500, Merck, Darmstadt,
Germany) per litre was added and carefully stirred into
the solution. The solution was poured into a form, in
which the sinus cast was fixated, and quickly cooled
down long enough for the agar to solidify. The resulting
velocity of sound in the agar during the measurements
was 1492 m/s. The resulting phantom is seen in Fig 1.
Figure 2: The resulting mean Doppler shift resulting
from water-glycerol mixtures with 1/(k 2 viscosity).
Discussion
Figure 1: The tissue mimicking phantom with the
Doppler probe visible to the right.
Measurements were performed using a custom
made continuous-wave ultrasound Doppler system. The
transducer was a commercially available 4.2 MHz dual
element type from Park Medical Electronics Inc (Aloha,
OR, USA). The transmitter was driven with a voltage of
13 Vpp, producing a spatial-peak-temporal-average
intensity of 79 mW/cm 2 . The intensity level was
measured using a calibrated needle hydrophone
(Precision Acoustics, UK). After placing the ultrasound
probe against the wall of the phantom, the mean
Doppler shift was recorded from a number of water and
glycerol solutions, with varying degree of viscosity.
Here, the viscosity was varied between 7 to 47 mPas, a
range that falls in the clinically interesting range. A
small amount of Sephadex (G10, Pharmacia, Uppsala,
Sweden) was introduced as scattering particles in the
fluid contained in the phantom.
The results show that it is possible to initiate acoustic
streaming in fluids having viscosities in the range of
serous to mucoid in a sinus shaped cavity. The acoustic
energy necessary to initiate the streaming is below the
recommended limit of deposited ultrasound intensity
(100 mW/cm 2 ), even with a realistic acoustic
attenuation prior to the cavity. The bone (fossa canina)
has not been included in the model, and if so, could
naturally affect the measurement. This bone is however
extremely thin, and even if proving to add significant
attenuation at this frequency, that effect should be lower
at a lower carrier frequency.
In order to obtain a Doppler shifted signal from a
fluid within the sinus cavity, the fluid must contain
scattering particles. Whether this is the case is still an
open question.
References
[1] DYMLING SO, PERSSON HW, HERTZ TG,
LINDSTRÖM K, (1991) ‘A new ultrasonic method for
fluid property measurements’, Ultrasound Med Biol;
17:497-500.
Results
The resulting mean Doppler shift from water-glycerol
mixtures with varying viscosities is seen in Figure 2.
There is a linear (R 2 =0.94) relationship between the
recorded Doppler shift and 1/viscosity, in accordance
with theory. The constant k is the sound speed of the
water glycerol mixture in question, divided by the sound
speed for the mixture in the middle of the range. This is
to correct for the square-law dependence of sound speed
due to varying glycerol concentration.
IFMBE Proc. 2005;9: 116

Medical ultrasound
NEW IMPROVED METHOD FOR 2D ARTERIAL WALL MOVEMENT
MEASUREMENTS
Magnus Cinthio 1 , Åsa Rydén Ahlgren 2 , Tomas Jansson 1 , Hans W Persson 1 , Kjell Lindström 1
1 Department of Electrical Measurements, Lund University, Lund, Sweden
2 Department of Clinical Physiology, Lund University, Malmö University Hospital, Sweden
magnus.cinthio@elmat.lth.se
Abstract: We have recently reported that the
inner layers of the arteries, the intima-media
complex, of common carotid artery, move as
much in the longitudinal direction as in the
radial direction, during the cardiac cycle. In
order to study this phenomenon we have
developed a high-resolution ultrasonic method
that can simultaneously record both the
longitudinal and the radial movements of the
arterial wall non-invasively in vivo. However,
in young individuals with large movements
and with thin intima-media complex it
happens sometimes that the echoes from the
adventitia region interfere. To be able to
minimise the size of the region-of-interest in
the radial direction, we suggest that the radial
movement of the arterial vessel is first
measured and that the radial movement is
used as a priori information when the
longitudinal movement is measured. The
mean difference between the two methods is 8
µm and 2 standard deviation is 24 µm.
Introduction
It has been assumed that the longitudinal
movement of the arterial wall during the cardiac
cycle is negligible in comparison with the radial
movement. However, recently, we [1] have
reported that the inner layers of the arteries, the
intima-media complex, of common carotid
artery, move as much in the longitudinal
direction as in the radial direction, during the
cardiac cycle [2]. In order to study this
phenomenon we have developed a highresolution
ultrasonic method that can
simultaneously record both the longitudinal and
the radial movements of the arterial wall noninvasively
in vivo [2, 3].
To allow measurement of the longitudinal
movement in vivo a preselected distinct echo
from an inhomogeneity or irregularity must be
visible in the ultrasound images throughout
several cardiac cycles with no other surrounding
echoes interfering during the measurements.
However, when measuring the longitudinal
measurement of the intima-media complex it
happens sometimes that the echoes from the
adventitia region interfere. This happen because
the size of region-of-interest must be at least
twice the largest expected movement that occur
between two images during the cardiac cycle.
This is especially a problem in young individuals
with large movements and with thin intimamedia
complex. To be able to minimise the size
of the region-of-interest in the radial direction,
we suggest that the radial movement of the
arterial vessel is first measured with a recent
develop one-dimensional vessel wall tracking
method [4]. After that the longitudinal
movement is measured using a smaller regionof-interest
in the radial direction using the radial
movement as a priori information.
Materials and Methods
The longitudinal and radial movements of the
intima-media complex of the far wall of the
common carotid artery in 5 healthy subjects were
measured using B-mode ultrasound. All
investigations were performed by a commercial
ultrasound system (HDI ® 5000, Philips Medical
Systems, ATL Ultrasound, Bothell, WA, USA).
First, the longitudinal movement was
measured with the conventional method for
longitudinal movement without any a priori
information of the radial movement [3]. The size
of the region-of-interest was approximately 0.5 x
0.7 mm in radial and longitudinal direction,
respectively. Thereafter the radial movement of
the far wall was measured in the same 5 subjects
with a recent developed vessel wall tracking
method [4]. The resulting movement trace in
radial direction was used as a priori information
when measuring the longitudinal movement. The
size of the region-of-interest was in this case
approximately 0.25 x 0.7 mm in radial and
longitudinal direction, respectively. Three
distinct movements were measure every
heartbeat and compared.
Secondly, a subject, where the conventional
method without any a priori information of the
radial movement [3] does not work, was tested
with the new modified method that used the
radial movement as a priori information.
IFMBE Proc. 2005;9: 117

Medical ultrasound
Results
Three distinct movements were measure
every heartbeat and compared. The difference
between the methods is presented in a Bland-
Altman diagram. The mean difference for all
three movements between the methods is 0.008
mm and 2 standard deviation is 0.024 mm
(Figure 1). This indicates that the new modified
method works. The test on the subject, where the
conventionally method [3] lost the contact with
pre-selected echo after about approximately 0.7
s, showed that the new algorithm could solve the
problem with the combinations of large radial
movements and thin intima-media thickness.
Conclusions
The 2-D wall movement measurements can
be improved by first measuring the radial
movement of the arterial wall and use the radial
movement as a priori information so the regionin-interest
can be minimised in the radial
direction. This is especially necessary and
effective in young individuals with large vessel
wall movements and thin intima-media complex.
Difference Movement (A Priori - Normal) (mm)
0.08
0.06
0.04
0.02
0
- 0.02
- 0.04
0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9
Average Movement (mm)
Figure 1. Difference vs. mean for the first
longitudinal movement in the direction of the
blood flow (x), the longitudinal movement
opposite the blood flow (o) and the second
longitudinal movement with the blood flow (+)
of the intima-media complex of the far wall in
the common carotid artery in 5 health subjects,
measured with a conventional 2-D arterial wall
movement method [3] and a new modified
method where the radial movement is used as a
priori information. The mean difference between
the methods is 8 µm and 2 standard deviation of
the difference is 24 µm.
Longitudinal Movement (mm)
1
0.8
0.6
0.4
0.2
0
- 0.2
- 0.4
- 0.6
- 0.8
- 1
0 0.5 1 1.5 2 2.5
Time (s)
Figure 2. Results from a subject where the
conventional 2-D arterial wall measurements
method [3] did not work due to the large arterial
wall movements and thin intima-media
thickness. After approximately 0.7 seconds the
conventional method ( … ) lost the direct contact
with the pre-selected echo and soon thereafter
did not follow at all. The new modified method
there the radial movement was used as a priori
information (—) did not loose the contact and it
could thereafter follow the pre-selected echo for
several cardiac cycles
Reference
[1] Persson M., Rydén Ahlgren Å., Eriksson A.,
Jansson T., Persson H. W., and Lindström K.
(2002): "Non-invasive measurement of
arterial longitudinal movement," IEEE
Ultrasonics Symp Proc, pp. 1739-1742
[2] Persson M., Rydén Ahlgren Å., Jansson T.,
Eriksson A., Persson H. W., and Lindström
K. (2003): "A new non-invasive ultrasonic
method for simultaneous measurements of
longitudinal and radial arterial wall
movements: first in vivo trial," Clin Physiol
Funct Imaging, vol. 23, pp. 247-251
[3] Cinthio M., Rydén Ahlgren Å., Jansson T.,
Eriksson A., Persson H. W., and Lindström
K. (In Press.): "Evaluation of an ultrasonic
echo-tracking method for measurements of
arterial wall movements in two dimensions,"
IEEE Trans. Ultrason., Ferroelect., Freq.
Contr.
[4] Cinthio M., Jansson T., Eriksson A., Persson
H. W., and Lindstrom K. (2004) "A robust
and fast algorithm for automatic arterial
lumen diameter measurement with
ultrasound," in Ultrasonic Methods for 2D
Arterial Wall Movement Measurements.
Lund
IFMBE Proc. 2005;9: 118

Medical ultrasound
ULTRASONIC ATTENUATION IMAGING USING COHERENT
PROCESSING IN ULTRASONIC COMPUTER TOMOGRAPHY
Filipík, A., Jiřík, R., Jan, J.
Dept. of Biomedical Engineering, FEEC,
Brno University of Technology,
Kolejní 4, 612 00 Brno, Czech Republic
xfilip10@stud.feec.vutbr.cz
Abstract: This paper presents an improving
modification to a recently published ultrasonic
attenuation imaging method [1]. In this method, the
attenuation coefficients are estimated in a volume,
where directly transmitted, reflected, and scattered
ultrasonic waves can be recorded - an ultrasonic
computer tomography system. All of the recorded
signals are used for the attenuation estimation.
More information is used for the image
reconstruction than in other known approaches;
thus the distribution of local attenuation can be
reconstructed more accurately. However, this
approach has a theoretical limit. It introduces a
simplified model of the imaged volume, where only
a small number of scatterers / reflectors are
assumed. When responses from multiple
reflectors / scatterers combine at the receiving
transducer, the estimation based on such signal is
not correct. In real tissue, obviously, this
simplifying assumption can not be guaranteed.
The, modifying idea consists in incorporating
coherent preprocessing of the recorded signals in a
small neighborhood (subarray) of the receiving
transducers. By using phased subarray signal
processing, the responses from individual
reflectors / scatterers can be distinguished, thus
overcoming the limitation.
direct, reflected and scattered waves. Knowing the
distances and propagation speeds, it is possible to
determine the ultrasonic beam paths along which the
signals are attenuated. Each path corresponds to a short
segment of the recorded radiofrequency signals. In
contrast to other known approaches, where only the
first segment (corresponding to a direct transmission) is
used, here all of the segments are used for attenuation
coefficient estimation. The spatial distribution of local
attenuation can thus be reconstructed more precisely.
Unfortunately, this approach has a limitation. The
method is only valid for a simplified model of the
imaged volume, where only a small number of
reflectors / scatterers is assumed. When two or more
reflector / scatterer responses meet at the same moment
at the receiving transducer, they are added together to
form the recorded signal. When estimating the
attenuation along one of these contributing paths,
attenuations along the other paths are not taken into
account, which yields an incorrect estimate.
Introduction
There has been a considerable interest in estimation
of ultrasound attenuation coefficient for several
decades. Most methods are based on the assumption of
linear dependence of the attenuation coefficient on
frequency [2]. So far, the available methods estimate
the attenuation coefficient only in fairly large tissue
regions, resulting in poor resolution images.
The recently published method [1], based on logspectra
estimation of the ultrasonic attenuation
coefficient, is taking advantage of the possibility to
record directly transmitted, reflected, and scattered
signals in an ultrasonic computer tomography system
(USCT) [3]. In such a system, the scanned volume is
enclosed by several thousands of ultrasonic
transducers. Each transducer can be in the emitting or
receiving mode. While one transducer is emitting, all
of the other transducers are receiving a mix of the
Figure 1: The ultrasonic beam paths in a USCT system.
The receiving transducers record signals from all
directions. Without coherent subarray processing, the
attenuation estimates do not correspond with the real
attenuation values along the intended path.
In this paper, we introduce a modification of the
method. The modification enables that the responses
from individual reflectors or scatterers can be
IFMBE Proc. 2005;9: 119

Medical ultrasound
distinguished, and correct estimates of the ultrasonic
attenuation coefficient along the corresponding paths
can be made.
Methods
The discrimination of responses from individual
reflectors / scatterers is made possible by coherently
processing a small neighboring set of the received
radiofrequency signals – signals from a subarray of the
receiving transducers (Figure 1). The subarray is
treated as a phased array, thus enabling directional
discrimination of signals (also known as beam
steering). Larger size of the subarray allows better
focusing, unfortunately also corresponds to a wider
path along which the attenuation can be estimated,
resulting in a deterioration of spatial resolution. It can
be shown that an optimal subarray size depends on the
distance from the reflector / scatterer. The farther it is,
the larger size of the subarray is necessary for proper
focusing.
After the signals of the subarray are coherently
preprocessed, the corresponding attenuation coefficient
can be estimated (e.g. via the log-spectrum method).
The attenuation image is then reconstructed using
estimated attenuation coefficients along all paths. As
for the reconstruction method, only the unfiltered
backprojection can be obtained by “smearing” the
estimates along the respective ultrasonic beam paths.
Filtered backprojection is not possible, because the
reflected / refracted broken beams don’t form parallel
projections. A better choice is reconstructing the image
via an algebraic reconstruction technique (ART) [4],
enabling an arbitrary geometry of the integration paths.
Results
An ultrasonic phantom was scanned in a USCT
system in Forschungszentrum Karlsruhe, Eggenstein,
Germany. First results of the reconstructed distribution
of the local attenuations are very promising. The
reconstructed images show contrast and resolution
superior to those obtained by the classical approach.
Discussion
So far, a constant sound speed in the imaged
volume was assumed in our experiments. In order to
precisely focus the recorded signals, the exact time-offlights
of the reflected and scattered signals have to be
known.
We suggest using a map of the local sound speeds,
reconstructed by e.g. filtered backprojection of the
direct ultrasonic waves sound speeds. By integrating
the local sound speed values along individual beam
paths (and knowing the travel distance) the precise
time of flight of a reflected / refracted wave can be
obtained. This could further improve the algorithm
performance.
Conclusions
A modification to a recently published work has
been described. The proposed coherent subarray
preprocessing step makes it possible to overcome a
previously limitative condition.
Further research is planned, especially on deciding
when it is necessary to use the coherent subarray
processing (and thus reducing the spatial resolution)
and how large the subarray should optimally be.
We also plan to incorporate a more accurate
technique to estimate the times of flight of the
reflected / scattered waves.
Acknowledgements
This work has been supported by the Ministry of
Education of the Czech Republic research program MS
1850023, and by the joint program of the German
Academic Exchange Service and Czech Academy of
Science (grant. no. D-CZ 22/05-06).
We’re also very grateful to Nicole Ruiter and
Rainer Stotzka (Forschungszentrum Karlsruhe,
Eggenstein, Germany) for providing the indispensable
measurement data.
References
[1] JIRIK R., STOTZKA R, TAXT T. (2005):
‘Ultrasonic Attenuation Tomography Based on
Log-Spectrum Analysis,’ Proceedings of SPIE,
Medical Imaging 2005. Volume: 5750, 2005.
[2] SCHMITT, M.R., et al (1984): ‘Error Reduction in
Through Transmission Tomography Using Large
Receiving Arrays with Phase-Insensitive Signal
Processing’, IEEE Transactions on Sonics and
Ultrasonics, vol. SU-31, no. 4, July 1984.
[3] STOTZKA, R., et al (2002): ‘Medical Imaging by
Ultrasound-Computertomography’, SPIE Medical
Imaging, 2002 / 25.
[4] KAK A.C, SLANEY M.(2002): ‘Principles of
Computerized Tomographic Imaging’, IEEE Press,
1988.
IFMBE Proc. 2005;9: 120

Mean pixel intensity
Fractional moving blood volume
Medical ultrasound
deviation (WD), and linear regression analysis and
Pearson’s correlation coefficient were calculated.
Results
The results of fetal lung PDU evaluation are presented
in figures 1 and 2. There were significantly higher
values of FMBV (p

Medical ultrasound
VERSATILE MICROCHIP UTILISING ULTRASONIC MANIPULATION OF
MICROPARTICLES
M. Nilsson 1 , T. Lilliehorn 2 , L. Johansson 2 , M. Almqvist 1 , U. Simu 2 , S. Johansson 2 , T. Laurell 1 , J.
Nilsson 1
1 Department of Electrical Measurments, Lund University, Lund, Sweden
2 Department of Engineering Sciences, Uppsala University, Uppsala, Sweden
mikael.nilsson@elmat.lth.se
Abstract
This paper presents the concept and initial work on a
microfluidic platform for bead-based analysis of
biological sample. The core technology in this project
is ultrasonic manipulation and trapping of particle in
array configurations by means of acoustic forces. The
platform is ultimately aimed for parallel multistep
bioassays performed on biochemically activated
microbeads (or particles) using submicrolitre sample
volumes. A first prototype with three individually
controlled particle trapping sites has been developed
and evaluated. Standing ultrasonic waves were
generated across a microfluidic channel by integrated
PZT ultrasonic microtransducers. Particles in a fluid
passing a transducer were drawn to pressure minima
in the acoustic field, thereby being trapped and
confined laterally over the transducer. It is
anticipated that acoustic trapping using integrated
transducers can be exploited in miniaturised total
chemical analysis systems (µTAS), where e.g.
microbeads with immobilised antibodies can be
trapped in arrays and subjected to minute amounts of
sample followed by a reaction, detected using
fluorescence. Preliminary results indicate that the
platform is capable of handling live cells as well as
microbeads. A first model bioassay with detection of
fluorescein marked avidin binding to trapped biotin
beads has been evaluated.
Introduction
Microsystem technology (MST) is currently making a
dramatic progress within the field of life science. A
driving factor for this is the tremendous need of
technology advancements in bioanalytical techniques to
meet the expected demands as we are now entering the
post genomic era and face much more complex biological
queries than in the process of decoding the human
genome. One of the most targeted areas are the efforts to
find new technologies to perform protein analysis
commonly called protein chips.
Microbeads are by several groups considered to be a
fundamental base for this development due to its
inherently large surface area and thereby high analytical
sensitivity. To be able to use these particles to
dynamically generate protein arrays a microfluidic
system making it possible to trap and manipulate the
particles is needed. A microfluidic system based on
particles will result in smaller sample volumes, higher
throughput and a more flexible analytical system.
Methods
A device with three acoustic trapping sites was fabricated
and evaluated. The lateral extension of each trapping site
was essentially determined by the corresponding PZT
microtransducer dimensions, 0.8 x 0.8 mm 2 . The flowthrough
volume of the device was approximately 1 µl and
the active trapping site volumes about 100 nl each.
The device was fabricated in a modular fashion with a
microstructured SU-8/glass channel plate being clamped
to a printed circuit board (PCB) carrying the transducer
array as well as the electrical and fluidic connections, see
figure 1.
Figure 1: Two trapping sites, I and II, designed as
acoustic microresonators consisting of glass reflector,
fluid layer and ultrasonic transducer. The close-up shows
a simplified view of beads trapped by acoustic forces due
to pressure gradients in the fluid layer.
Figure 2: Assembled device with an array of three
transducers visible through the examination window.
The fluidic channels were designed to create a pressure
minima in the middle of the channel. With a working
frequency of 10 MHz the channel depth were 71 µm.
IFMBE Proc. 2005;9: 123

Medical ultrasound
To evaluate the performance of the device a model
bioassay using biotin-tagged microparticles and FITCtagged
avidin was performed. The biotin-particles were
ultrasonically trapped and perfused with FITC-avidin and
the fluorescent intensity was monitored over time.
Results
Bead trapping
The channel height was confirmed to be 71 µm as
measured by confocal imaging. Particles trapped in the
channel were mainly positioned in the centre of the
channel, 36 µm above the transducer, see figure 3. A
small fraction of beads were however found on the
surfaces of the transducer and glass reflector.
channels between subsequent arrays. The confocal
measurements showed that the beads mainly were
trapped in the middle of the channel. This is
advantageous since this location has the highest flow and
also minimizes the contact between beads and the
channel walls. Some beads were present at other surfaces
as well, but this can hopefully be corrected by
considering the materials used and the fluidic system.
The results from the bioassay shows that avidin has
bound to the biotin-coated beads, which can be seen as an
increase in fluorescent response in figure 4. The epoxy
used to cover the electrical connections on the PCB
showed a strong unwanted autofluorescence in the same
wavelengths as the FITC. By using other materials the
background intensity can be minimized and a higher
intensity contrast achieved.
Some initial experiments with live cells have also been
done. The indiciations so far is that there is no difference
to manipulate cells or particles regarding the ultrasonic
trapping. Of course the environment in the microfluidic
channels must be monitored and made sure to fit the cell
type used.
Figure 3: Trapped fluorescently labelled polystyrene
particles in a typical laterial distribution.
Model bioassay
The fluorescent read-out from the bioassay is plotted in
figure 4.
The dynamic arraying is believed to be expandable to two
dimensions, thus with a prospect of performing targeted
and highly parallel protein analysis in microfluidics. With
the possibility to handle cells as well, this technique can
be a powerful and versatile tool in the near future.
References
[1] LILLIEHORN, T (2003), ‘Piezoactuators for
Microfluidics - Towards Dynamic Arraying. (Doctoral
Thesis)’, Department of Materials Science. 2003,
Uppsala University, Sweden.
[2] LILLIEHORN T, NILSSON M et al (2005),
‘Trapping of microparticles in the near field of an
ultrasonic transducer’, Ultrasonics, 43, pp 293-303
[3] LILLIEHORN T, NILSSON M et al (2005),
‘Dynamic arryaing of microbead for bioassays in
microfluidic channels’, Sensor and Actuators B, in press
[4 ] GRÖSCHL, M (1998), ‘Ultrasonic separation of
suspended particles, I, Fundamentals’, Acustica - Acta
Acustica, 84, pp 432-447
Figure 4: Loading of FITC-tagged avidin over trapped
biotin-tagged particles. The switching from avidin to
washing flow is indicated in the figure. ∆ indicates the
response of the actual biotin/avidin binding. The plotted
intensities are normalized to the maximum intensity in
the series.
Discussion
To minimize carry over from one bioassay to the next it
is essential to make sure that no beads are left in the
IFMBE Proc. 2005;9: 124

Medical ultrasound
INVESTIGATION OF THE FETAL HEART CIRCULATION IN AN
ANIMAL MODEL USING CONTRAST ENHANCED ULTRASOUND
T. Jansson*, E. Hernandez-Andrade**, G. Lingman**, Peter Malcus**,
David Ley**, and K. Marsál**
* Department of Electrical Measurements, Lund University, Lund, Sweden
** Department of Obstetrics and Gynecology, Lund Universtiy, Lund, Sweden
tomas.jansson@elmat.lth.se
Abstract: Aim: To assess the distribution of blood
from the umbilical vein (UMV), inferior vena cava
(IVC) and superior vena cava (SVC) to either side of
the fetal lamb heart by using ultrasound enhanced
contrast agent imaging.
Material and methods: By injection of ultrasound
contrast agents (UCA) in UMV, IVC or SVC, the
blood from these vessels was tracked, as it then was
highly echogenic. By evaluating the image intensity
within the heart ventricles, the relative
concentrations of blood could be determined. The
study was performed in 19 near term fetal lambs of
mixed breed, with a mean gestational age of 136 days
(range 134-136). Ultrasound contrast agent was
injected at a constant rate of 1 ml/min, to ensure that
a constant level of contrast agent would be obtained
in both sides of the heart.
Results: the median percentages of blood distributed
to the left ventricle when injecting contrast in UMV,
IVC, and SVC, was 68%, 67%, and 21%
respectively.
Discussion: these numbers compare well with
previously published data, except the recorded
percentage distributed from the IVC. This could be a
methodological error as well as a result of the mild
hypoxia, or an actual increased capacity of the left
ventricle at this gestational age.
Introduction
During fetal life, oxygenated and deoxygenated
blood travels together in the fetal vascular system.
Highly oxygenated blood from the placenta reaches the
inferior vena cava (IVC) and the right atrium through
the ductus venosus (DV), from where it is forwarded to
the left cardiac chambers. Deoxygenated blood from the
lower and the upper part of the fetal body reaches the
right atrium through the IVC and the superior vena cava
(SVC), respectively. Consequently, the right atrium
receives and distributes all the fetal venous returning
blood.
Rudolph et al., using a radionuclear microsphere
(RMS) reference sample technique in fetal lambs, have
given numbers on the output blood flow from each
ventricle and from which vessel it originates, but only in
healthy subjects. Furthermore, these values are static,
i.e. the flow values can not be tracked over time. To
easily follow the distribution of blood in real time, as
well under normal, as under hypoxic conditions, we
propose the use of UCA. If injected in UMV, IVC or
SVC, respectively, the blood from these vessels can be
tracked, as it then will be highly echogenic. By
evaluating the image intensity within the heart
ventricles, the relative concentrations of blood can be
determined.
The aim of this study was to evaluate the technique,
and to assess and quantify the distribution of blood from
the DV, IVC and SVC within the fetal heart by using
ultrasound enhanced contrast agents images.
Materials and Methods
The study was performed in 19 near term fetal
lambs of mixed breed. The mean gestational age
calculated from the day of conception was 136 days
(range 134-138). A midline laparotomy was performed
in the pregnant ewe under isofluorane anesthesia.
Catheterization was performed in IVC (n=11 lambs),
SVC (n=3), and UMV (n=7), respectively. In one lamb,
a tibial vein and an umbilical vein catheter were
inserted, and in another lamb a jugular vein and an
umbilical vein catheters were placed. In total, 21
estimations were performed.
Ultrasound recordings were performed using an
ATL HDI-5000 (Philips, Bothell, WA) ultrasound
scanner with a curvilinear 3-5 MHz probe. The US
probe was placed directly on the uterine wall. Settings
were kept identical for all examinations with an output
power of MI=0.9 in fundamental mode.
Infusion of UCA (Levovist 400mg/ml, Shering AG,
Berlin) was performed with a pre-calibrated pump
(Terumo STC-521, Tokyo, Japan) at a rate of 1 ml/min.
The constant infusion was used so that a constant level
of contrast agent in would be obtained in both sides of
the heart. Ultrasound recordings were performed at the
time of the contrast injection, so that a “timeamplitude”-curve
was obtained, and thereafter
trqansferred to a personal computer for off-line analysis
in dedicated software (HDI-Lab 1.81 Philips, Bothell,
WA).
Regions of interest (ROI) were chosen inside each
ventricle. Care was taken so that echoes from the
myocardium were not present in the ROI in any part of
the heart cycle throughout the sequence. The mean
IFMBE Proc. 2005;9: 125

Medical ultrasound
amplitudes over time were exported to a file, as for the
in vitro experiment. The files were then reanalyzed in
MATLAB where the time-amplitude curves were
plotted. To ensure that a constant level of UCA was
present, regions of the plots where the contrast intensity
had reached a plateau were reselected and means for the
two ROI:s were calculated. Finally, the ratios of either
mean amplitude to the sum of the mean amplitudes were
calculated. This gave the percentages of blood directed
to the left or right ventricles. Descriptive statistics of
distribution of UCA in the fetal heart in relation to the
place of injection were used.
Results
Boxplots of the data can be found in Figures 1, 2, and 3,
showing the measured percentages found in each
ventricle when injection was performed in the UMV,
IVC, and SVC, respectively. The median measured
percentages of blood distributed to the left ventricle
when injecting contrast in UMV, IVC, and SVC, was
68%, 67%, and 21% respectively.
Figure 1: Percentage of blood to each ventricle: contrast
injected in UMV, four-chamber view (n=15)
Figure 3: Percentage of blood to each ventricle: contrast
injected in SVC, four-chamber view (n=4)
Discussion
These numbers compare well with previously
published data, except the recorded percentage
distributed from the IVC. This could be a
methodological error as well as a result of the mild
hypoxia or an actual increased capacity of the left
ventricle at this gestational age.
Possible methodological errors include shadowing
of contrast in the anterior ventricle, causing the
posterior ventricle to appear to have a lower value of
contrast concentration. Manual inspection of the tissue
signal posterior to both ventricles before and after
injection of contrast, seem however to indicate that this
effect is small.
Another source of error is possible inclusion of
tissue signal from the ventricular walls in systole. This
would erroneously elevate the signal from within the
ventricle. A possible solution would be to make the
measurement in the outflow tract where movement is
not a large problem.
References
[1] BERMAN W JR, GOODLIN RC, HEYMANN MA,
RUDOLPH AM. (1975) ‘Measurement of umbilical
blood flow in fetal lambs in utero. J Appl Physiol.
1975 Dec;39(6):1056-9
Figure 2: Percentage of blood to each ventricle: contrast
injected in IVC, four-chamber view (n=15)
IFMBE Proc. 2005;9: 126

Biomedical instrumentation
RESONANCE SENSORS FOR BETTER HEALTH CARE
O. Lindahl 1
1 Centre for biomedical engineering and physic, c/o TFE, Umeå University, Umeå, Sweden
olof.lindahl@tfe.umu.se
Introduction
Resonance sensors that are constructed to have a
mechanical resonance frequency or relative phase of
oscillation dependent on the measured parameter are of
considerable practical interest. These kinds of sensors are
developed for a wide range of applications in industry
including sensors for liquid or gas density and viscosity,
liquid level, mass and mechanical force and fluid flow
rates.
Resonance sensors are also used to develop measurement
systems in health care. For example, resonance sensor
systems have been developed for detection of breast
cancer and for measuring muscular elasticity (Omata and
Terunuma 1992) and also for modelling micturation
characteristics based on prostate stiffness (Omata and
Constantinou 1995). In other studies it has been shown
that pitting oedema of the skin (Lindahl and Omata 1995)
and elasticity or spring constant of human skin (Lindahl
et al 1998) could be estimated with a resonance sensor
combined with measurement of force and position.
In resent studies the ability of the resonance sensors to be
related to area of contact, has been used. For example a
new instrument for measuring the intraocular pressure
(IOP) in the human eye has been developed
(Eklund 2002). The measurement principle is based
on Imbert Ficks law that states that IOP=F/A where F is
force (N) and A is area (mm 2 ). The resonance frequency
shift of a piezoelectric element when it attaches the
cornea of the eye gives the area of contact and a force
sensor in contact with the resonance sensor gives the
force of contact. Thus, the IOP can be measured.
At the moment there is also ongoing research concerning
the use of resonance sensors to detect stiffness of human
prostate with the aim to detect prostate cancer (Jalkanen
et al 2003). The background is that cancerous tissue is
generally regarded as "harder" than non-cancerous tissue.
The degree of "hardness" of prostate is estimated with
the use of stiffness parameters as measured with
resonance sensor systems and the research aims at
classifying and visualizing stiffness changes in prostate
tissue. The final goal is to establish a method for early
detection of cancer tissue non-invasively in prostate.
It can be concluded that the use of resonance sensors in
health care applications are of great interest, e.g. when
biomechanical parameters like contact area and/or
stiffness of human organs are to be estimated for clinical
diagnosis. There exist a wide variety of application
possibilities for resonance sensors in health care.
References
Omata, S. and Y. Terunuma, New tactile sensor like the
human hand and its applications. Sens. & Actuators,
1992. 35: p. 9-15.
Omata S and Constantinou, Modeling of Mictiration
Characteristics Based on Prostatic Stiffness Modulation
Induced using Hormones and Adrenergic Antagonists,
IEEE Trans. Biomed. Eng., 1995, 42, 843-848
Lindahl, O.A. and Omata, S. : Impression technique for
the assessment of oedema-Comparison with a new tactile
sensor that measures physical properties of skin.
Medical & Biological Engineering & Computing, 33, 27-
32, 1995.
Lindahl, O., Omata, S., Ängquist, K.-A.: A tactile sensor
for detection of physical properties of human skin in
vivo, J. of Medical Eng. & Technology, 22, 147-153,
1998.
Eklund, A.: Resonator sesnor technique for medical use-
An intraocular pressure measurement system, Umeå
university medical dissertations 801, ISSN 0346-6612,
2002
Jalkanen, V., Eklund, A., Lindahl, O: Force and
frequency shift from a resonance sensor for detection of
prostate cancer, Proceedings of the World Congress on
Medical Physics and Biomedical Engineering, Sydney,
Australien, Augusti, 2003.
Conclusions
IFMBE Proc. 2005;9: 127

Biomedical instrumentation
EYE-PRESSURE MEASUREMENT WITH APPLANATION RESONANCE
TONOMETER
A. Eklund 1, 3 , P. Hallberg 1, 3, 4 , K. Santala 2 , T. Bäcklund 1 , C. Lindén 2
1 Biomedical Engineering and Informatics, Umeå University Hospital, Umeå, Sweden
2 Department of Clinical Science, Ophthalmology, Umeå University, Umeå, Sweden
3 Center for Biomedical Engineering and Informatics, Umeå University, Sweden, ,
4 Department of Applied Physics and Electronics, Umeå University, Sweden, ,
anders.eklund@vll.se
Abstract
Traditionally, IOP is determined by applanation
tonometry, a method where the force needed to flatten a
certain area of the cornea is measured. In a recently
implemented method for IOP-measurement, the contact
area was measured with a resonator sensor device. A
force transducer mounted in the same device as the
resonator sensor measured simultaneously the contact
force. For this study a new sensor design was proposed.
The purpose of the study was to evaluate the new
applanation resonance tonometer system in a clinical
setting. The study included measurement on totally 150
eyes from healthy volunteers and patients with elevated
IOP. The ART mounted in a biomicroscope was
compared with Goldmann applanation tonometry (GAT).
We found that the new sensor design resulted in an
improved precision.
Introduction
Elevated intraocular pressure (IOP) is one of the major
risk factors for glaucoma 1 . Since glaucoma is a leading
cause of blindness, reliable methods for measuring the
IOP are therefore important.
IOP is most commonly determined by applanation
tonometry, methods where the force needed to flatten a
certain area of the cornea is measured. The Imbert-Fick
law states that the contact force is balanced by the IOP
times contact area. With current applanation methods, for
example Goldmann applanation tonometry, different
sources of error such as corneal thickness and corneal
curvature has been shown to influence the
measurements 2 , thus resulting in an error in the
estimation of IOP. This suggests that new methods are
needed. A new applanation sensor for an easy measuring
of the IOP, based on resonance technique has recently
been proposed 3 . The contact area is measured with a
piezoelectric element. The element, oscillating in its
resonance frequency, produces a frequency shift
proportional to the contact area. The applanation
resonance tonometer, ART, estimates the IOP by
continuous sampling of both contact force and contact
area (frequency) and calculates the IOP from the slope
between force and frequency. ART data from in vitro
laboratory bench setting, where sensor was carefully
centralised on to the cornea, was very consistent with
good precision in the determination of IOP 3 . In
evaluation with clinical setting, both in vitro 4 and in
vivo 5 , when the centralisation has been dependent on the
skill of the operator, the unavoidable off-centre
placement of the sensor has resulted in a reduced
precision. This off-centre dependence has also been
confirmed in in-vitro studies with controlled off-centring 3
and edge-effects have been identified as a potential cause
of the deviations.. These studies have suggested that
further development should focus on technical
improvement to over come these difficulties. In
particular, a cornea contact piece with a larger diameter
has been suggested in order to decrease edge-effects. This
development has been performed and the aim of this
clinical study was to evaluate a new ART with a more
symmetric sensor, a larger sensor tip and an improved
sight light.
Methods
To decrease the off-center position dependency a more
symmetric sensor was suggested. The previously 3, 5 used
rod shaped (23 x 5 x 1 mm) piezoelectric element, PZT,
was replaced with a pipe shaped (25x5 mm, 1 mm wall
thickness) PZT-element with a pickup part at the end of
the element (fig. 1). The PZT was moulded in an
aluminium container with silicon rubber.
Figure 1 Applanation resonance sensor probe.
The diameter of the contact piece surface was increased
from a previously of D=4 mm to D=7 mm.
IFMBE Proc. 2005;9: 128

Biomedical instrumentation
An open randomised study included measurement on
totally 150 eyes evenly distributed in three pressure
intervals. Below 16 mm Hg, between 16 and 23 mm Hg
and above 23 mm Hg. GAT was used as reference
method. The ART sensor was mounted on a
biomicroscope in a similar position as GAT. A quick
applanation against the cornea was performed with the
ART. Force and frequency data was recorded with a 1
kHz sampling rate. Linear regression was used to analyse
the continuous force increase against continuous
frequency shift during the applanation. The forcefrequency-slope
in the frequency interval (-30 to -300Hz)
was determined and interpreted as proportional to IOP.
Mean (n=6) ART was compared with mean (n=6) GAT
readings.
1. Sommer A. Am J Ophthalmol 1989;107(2):186-188.
2. Whitacre MM, Stein R. Surv Ophthalmol
1993;38(1):1-30.
3. Eklund A, Hallberg P, Linden C, et al. Invest
Ophthalmol Vis Sci 2003;44(7):3017-3024.
4. Hallberg P, Santala K, Linden C, et al. J Medicin
Engineering and Technology In Press.
5. Hallberg P, Linden C, Lindahl OA, et al. Physiol Meas
2004;25(4):1053-1065.
6. Midelfart A, Wigers A. Br J Ophthalmol
1994;78(12):895-898.
7. Kontiola AI. Acta Ophthalmol Scand 2000;78(2):142-
145.
Acknowledgements
The authors thank Michael Hansson for skillful
measurements.
Results
In a preliminary analysis, the correlation between IOP
according to ART and GAT was R= 0.95 (p

Biomedical instrumentation
DETECTION OF PROSTATE CANCER WITH A RESONANCE SENSOR
V. Jalkanen 1, 4 , B. Andersson 1, 4 , A. Bergh 2 , B. Ljungberg 3 , O. Lindahl 1, 4
1 Department of Applied Physics and Electronics, Umeå University, Umeå, Sweden
2 Department of Medical Biosciences, Pathology, Umeå University, Umeå, Sweden
3 Dept. Surgical and Perioperative Science, Urology and Andrology, Umeå University, Umeå, Sweden
4 Centre of Biomedical Engineering and Physics, Umeå University, Umeå, Sweden
ville.jalkanen@tfe.umu.se
Abstract
Prostate cancer is the most common cancer for men and
tumours are generally regarded as harder tissue than
surrounding normal tissue. In this study we used a
resonance sensor system for measurements on in vitro
prostate tissue stiffness to detect if tumours could be
separated from normal tissue. A morphometric
investigation was performed for comparison with tissue
stiffness data and we proposed a new parameter that
could differentiate between tumour and normal tissue in
three out of seven cases. Further studies are needed to
examine the full value of the resonance sensor.
Introduction
Prostate cancer is the most common cancer for men in
Europe and the US. It is usually diagnosed by a high
blood PSA, rectal palpation and ultrasound examination
of prostate followed by histological examination of
biopsies. Malignant tumours are generally regarded
harder than normal healthy tissue [1].
A non-invasive tactile resonance sensor that measures
physical properties of soft material e.g. human tissue has
been presented [2]. This technique measures the change
in resonance frequency, ∆f, when a vibrating ceramic rod
touches the surface of an object as human tissue. The
ceramic rod is set to vibrate with its resonance frequency
through an electronic feedback circuit and the resonance
frequency change has been shown to describe the
stiffness of an object [1,2].
It has earlier been shown that harder tissue like
prostate stones can be detected with resonance sensors on
fixed human prostate tissue from a patient with benign
prostate hyperplasia with good reproducibility [1]. The
aim of this study was to investigate if a resonance sensor
system could detect cancer in human in vitro prostate
tissue and compare with morphometrical investigations.
Methods
A resonance sensor system, Venustron® (Axiom Co.,
Ltd., Koriyama Fukushima, Japan), was used in the
experiments. It consists of the vibration sensor for
measuring the resonance frequency, a force sensor and a
position sensor; all arranged in a motorised mounting
attached to a stable stand and connected to a computer.
The sensor tip could be lowered towards an object with
the motor. The resonance frequency change (∆f), the
force (F) and the impression depth were sampled with
200 Hz during both the impression and the retraction of
the sensor tip. A maximum preset impression depth was
set to 2 mm and the impression speed was set to 1 mm/s.
Data were saved on a file with Venustron® software and
processed in MATLAB® (Comsol AB, Sweden).
Measurements were performed on ten 1-1.5 cm thick
prostate tissue slices from radicalprostatectomy. Each
tissue slice was pinned onto a foam plastic plate and kept
moist with regular application of saline solution.
Reference markings were marked on the foam plastic and
thereafter a picture was taken with a digital camera. The
position of reference markings and measurement points
on the tissue surface relative the sensor tip were
controlled with a linear XY-precision positioning stage
(Parker Hannifin Corporation - Daedal Division, Irwin,
PA, USA). 12-20 measurements were done on the surface
of each slice, and on 4-6 of these measurements five
repeated measurements were performed. After the
measurements the tissue was fixed in formalin and
thereafter embedded in paraffin.
A morphometric investigation was performed on the
most superior 5 mm cut section from each tissue slice.
Reference markings and the digital photo were used to
locate the measurement points on the 5 mm cuts. The hit
percentage for the sensor tip on a tumour area was
determined and a hit percentage of >75% was considered
as a secure hit on tumour, while a hit percentage of

Biomedical instrumentation
The Wilcoxon rank-sum test shows that in three
prostates there is a significant difference between
measurements on tumour and normal tissue, see Figure 1.
The percentage deviation is positive, thereby indicating
on harder tumour tissue compared to normal tissue.
Similar analysis shows that for four prostates no
differences could be found between tumour and normal at
p

Biomedical instrumentation
AN IMPROVED RESONANCE SENSOR SYSTEM FOR DETECTING
CANCEROUS TISSUE IN THE PROSTATE
P. Lindberg 1, 4 , B. Andersson 1, 4 , A. Bergh 2 , B. Ljungberg 3 , O. Lindahl 1, 4
1 Dept. Applied Physics and Electronics, Umea University, Umea, Sweden
2 Dept. Medical Biosciences, Pathology, Umea University, Umea, Sweden
3 Dept. Surgical and Preoperative Science, Urology and Andrology, Umea University, Umea, Sweden
4 Centre for Biomedical Engineering and Physics, Umea University, Umea, Sweden
peter.lindberg@tfe.umu.se
Abstract
Prostate cancer is the most common cancer form for
men. The purpose of this study was to improve and
evaluate a resonance sensor system for prostate
cancer. System reliability was estimated by
measuring the relative hardness of silicon. The
improved resonance sensor system could measure the
frequency shift and impression depth. The results
showed that frequency shift and impression depth
could describe the relative hardness of silicon (n = 50,
P

Biomedical instrumentation
Results
Silicon
The frequency shift, ∆f, were plotted against the cone
penetration values (fig. 1).
ISO standard 2137 [4]. The improved resonance sensors
frequency shift and impression depth could with
statistical significance detect the hardness of silicon, and
also distinguish between cancer a normal tissue in one, in
vitro, prostate slice. The RST has also been successful
when measuring and quantifying the hardness of lymph
nodes [5] and to detect differences in tissue composition
in non-malignant human prostate tissue [4]. Further
studies on prostate tissue have to be performed before the
full value of the sensor can be determined.
References
Figure 1: Mean value of the frequency shift with standard
deviation for each one of the 5 silicon mixes
The frequency shift, ∆f, and the impression depth, l p ,
could distinguish between the five different silicon mixes
(n=50, P

Biomedical instrumentation
Realtime wireless measurement of mechanical data
for a javelin throw
Jerker Delsing, Jerry Lindblom, Daniel Sjölund and Per Lindgren
EISLAB
Luleå University of TEchnology
Email: jerker.delsing@ltu.se
Abstract— Technology for the real time measurement of mechanical
data from a javelin throw has been developed. The
javelin is instrumented with an ineartial measurement unit
measuring, IMU, acceleration, angle speed and direction to the
earth magnetic field all in three dimensions i.e. in total nine
parameters. The IMU is buildt into the javelin still maintaining
the javelin properties and keeping it within the IAAF specifications.
The instrumentation is build using the EIS architecture
thus incorporating TCP/IP support including an Internet server.
The wireless communication technology choosen is Bluetooth that
connects to Internet through either a Bluetooth enabled mobile
phone or a stationary Bluetooth accesspoint.
I. INTRODUCTION
Athletes like javelin thrower is in a never ending search
for improved throwing technique. For this purpose most often
athletes today use high speed cameras. The equipment is
expensive and difficult to transport. Evaluation of the film
material is done by hand, trying to extract release speed, angle
of attach and other essential parameters. Se figure 1
EIS concept [1] [2]. A unit that can gather necessary values,
store them in a memory and transmit these raw data to for
example a laptop after the throw. To be able to reconstruct position,
orientation and velocity in the movements of the javelin
the unit must measure both acceleration and rotation in three
orthogonal axes. Such a unit is called IMU 1 with 6 degrees of
freedom. In addition a three-axis magnetometer was added that
perhaps can improve accuracy. Two extra accelerometers were
used to measure rotation when the angular velocity exceeds
the gyros maximum range, 300 degrees/second. The digital
components are a Renesas M16c microcontroller, a 512kb
memory, a 12bits A/D-converter and Bluetooth for wireless
communication.
The prototype unit consists of two main parts, an analog part
with the transducers and A/D-converter and a digital part with
microcontroller, memory and bluetooth. All functions have
been tested and work well. Transfer rate between the unit and
a computer with bluetooth is approximately 2.5kb/s. With the
chosen sampling frequency 250Hz approximately 80seconds
of all channels can be stored. The complete design is given in
[3].
III. INITIAL RESUTS
Initial tests with the IMU gives data like in figure 2.
Fig. 2. Data from 1 diemnsion of each of the three sensors, acclerometer,
Magnetometer and gyro for a simple rotation of the javelin.
Fig. 1. Parameters of high interest for a succesfull javelin throw, release
speed v, angle of attachand angle of throw
For the purpose an IMU to fit inside the javalin was
designed. The objectives was to fullfull the IAAF specification
of a javelin and to be able to measure the follwoing parameters:
• Javelin release speed 40 m/s resolution 0.1 m/s
• Angle of attach with resolution better than 0.5 degree
• Throwing angle with a resolution better than 0.5 degree
II. DESIGN
Here we describe the idea and the implementation of a
system based on transducers placed inside the javelin. The
system architecture for the measurement unit is based on the
The obtained data clearly shows that the measurements are
possible with an IMU mounted in the javelin. If the resolution
and accuracy obtained is sufficient for the goal has still to be
determined. The signal processing algorithms needed to obtain
high level measurement data is yet to be developed.
IV. CONCLUSION
Initial tests shows successfully the new javelin EIS instrumentation
capability to measure nine mechanical parameters of
gthe javelin and transfer such data to a reciever using standard
Internet wireless communication hardware and protocols. The
translation of data into simple and easily interpreted imformation
is still to be made. Here the neccesary signal processing
algorithms still is to be developed.
The approach will enable simple feedback to the athlete
and potentially more important also the capability to in real
IFMBE Proc. 2005;9: 134

Biomedical instrumentation
time transfer data directly to TV. Thus generating entierly new
information of possible interest to a TV audiance.
ACKNOWLEDGMENT
The authors would like to thank Nordic Sport AB for
interest and support in this work. In prticualry we like to
thank Kent Johansson of Nordic Sport AB. The project has
been partly funded by EU structural funds.
REFERENCES
[1] J. Delsing and P. Lindgren, “An architecture for mobile internet enabled
sensor networks,” Measurement Science and Technology, 2005, submitted
for publication.
[2] J. Delsing, P. Lindgren, and Å. Östmark, “Mobile internet enabled
sensors using mobile phones as access network,” ITcon, vol. 9, pp.
381–388, 2004, special issue on Mobile Computing in Construction,
http://www.itcon.org/2004/27.
[3] D. Sjölund, “Jimu - javelin inertial measurement unit,” Master’s thesis,
Luleå University of Technology, Luelå, Sweden, 2004.
IFMBE Proc. 2005;9: 135

Biomedical instrumentation
TEMPERATURE INDEPENDENCE OF AN ELECTRO ACOUSTIC
CAPNOGRAPH
M. Folke* and B. Hök**
* Department of Computer Science and Electronics, Mälardalen University, Västerås, Sweden
** Hök Instrument AB, Västerås, Sweden
mia.folke@mdh.se
Abstract: End tidal carbon dioxide measurement
with an electro acoustic sensor has recently been
demonstrated. The sensor consists of an acoustic
resonator coupled to a low cost electro acoustic
element. The aim of this study was to verify if
ambient temperature variation would affect the
measurements. By simultaneous measurements with
a reference sensor, the electro acoustic capnograph
was tested on subjects performing exercise, hypoand
hyperventilation. The output from the
experimental device correlated well with the
reference CO 2 readings with a correlation coefficient
of 0.91 at varied temperature and relative humidity.
Introduction
A new capnograph technique in main stream
application has recently been presented [1].
A linear relationship was obtained within the
accuracy of the measurement, with a correlation
coefficient of 0.976 when measured at constant ambient
temperature and humidity [1].
The sensor used in the capnograph is based on the
measurement of the impedance of an electro acoustic
element coupled to an acoustic resonator [2]. The
impedance characteristic is depending on the sound
velocity within the gas mixture contained in the acoustic
resonator.
It has been shown in an earlier study that there is an
approximately linear relation between the acoustic
impedance and the CO 2 concentration [2]. The sensor
principle has also shown a fast response to increasing
CO 2 concentration in laboratory experiments [3].
At constant temperature and humidity, the sound
velocity is determined by the average molecular weight
of the gas, and is therefore influenced by the CO 2 -
concentration, since CO 2 is considerably heavier than
oxygen and nitrogen, which dominates the average
molecular weight of air.
The aim of this study was to verify if ambient
temperature variation would affect the measurements.
Materials and Methods
The sensor system used for this study is explained in
[1]. The voltage output signal from the sensor system
was correlated to a reference sensor connected in a sidestream
configuration (Microcap ® Plus, Oridion Inc.
Israel, www.oridion.com).
Several repetitions of exercise, hyper- and
hypoventilation were performed, resulting in different
CO 2 -concentrations. In all measurements, approximate
steady state conditions were established before readings
were made.
The temperature in the room was varying from 19 to
26 °C and the relative humidity was varying from 28 %
to 33 % during the tests.
Baseline variation due to temperature variations
were also analysed from 19 to 25 °C at a constant
relative humidity at 33 %.
Results
Figure 1 shows the results of 67 end tidal level
measure ments obtained from the electro acoustic sensor
plotted against the recorded end tidal CO 2 -concentration
measured using the reference capnograph.
Voltage (V)
1,4
1,2
1
0,8
0,6
0,4
0,2
0
2 3 4 5 6 7 8
End tidal CO 2 concentration (kPa)
Figure 1. The relationship between the CO 2 -
concentration of the reference sensor and the output
voltage of the electro acoustic sensor system.
A linear relationship is obtained within the accuracy
of the measurement, with a correlation coefficient of
0.91. If the values above 5 kPa are excluded, because of
less accuracy of the reference in this interval, the
correlation coefficient will be 0.95.
IFMBE Proc. 2005;9: 136

Biomedical instrumentation
The average of measurement errors is estimated to
0.3 kPa for all the measurements.
Voltage (V)
2,5
2
1,5
1
0,5
0
14 19 24 29
Temperature (ºC)
Figure 2. Temperature influence of the baseline.
The ambient temperature affected the base line with
0.6 Volt from 19 to 23ºC, Figure 2. The end tidal
measurements were affected with 0.1 Volt in the same
temperature interval, Figure 3.
Voltage (V)
1,1
1
0,9
0,8
0,7
0,6
0,5
0,4
but not as much as for the baseline. One explanation is
that the sensor may act as a heat exchanger.
The measurements were shown to correlate well
with those obtained with the reference sensor, with an
estimated accuracy of 0.3 kPa as discussed above,
which is believed to be adequate in clinical applications.
The influence of temperature and humidity to the
measurements must be further analysed.
The electronic activation and detection mode used in
this study has several shortcomings, and should be
considered provisional. The impedance characteristic
for each individual sensor is not considered in this set
up.
Conclusions
The results of this study indicate that the use of
elements for effective heat and humidity exchange
reduces the influence of temperature and humidity to
acceptable levels.
References
[1] FOLKE M., HÖK B., EKSTRÖM M., and BÄCKLUND
Y. (2004): ‘End Tidal Carbon Dioxide Measurement
Using an Electro Acoustic Sensor’, Proc. of EMBC
2004 - the 26 th Annual International Conf. of the
IEEE Engineering in Medicine and Biology Society.
San Francisco, USA, 2004. p 362
[2] GRANSTEDT F., FOLKE M., BÄCKLUND Y., and HÖK
B. (2001): ‘Gas sensor with electroacoustically
coupled resonator’, Sensors and Actuators B. 78,
pp.161-165
[3] GRANSTEDT F., HÖK B., BJURMAN U., EKSTRÖM M.,
and BÄCKLUND Y. (2001): ‘New CO 2 sensor with
high resolution and fast response’, Proc. of EMBC
2001 - the 23 rd Annual. International Conf. of the
IEEE Engineering in Medicine and Biology Society.
(EMBC 2001), Istanbul, Turkey, 2001.
0,3
2 3 4 5 6
End tidal CO 2 concentration (kPa)
Temp: 26 ºC RH: 28 % R=0.982
Temp: 24 ºC RH: 33 % R=0.946
Temp: 23 ºC RH: 32 % R= 0.978
Temp: 19 ºC RH: 32 % R=0.995
Linear (Temp: 26 ºC RH: 28 % R=0.982)
Linear (Temp: 24 ºC RH: 33 % R=0.946)
Linear (Temp: 23 ºC RH: 32 % R= 0.978)
Linear (Temp: 19 ºC RH: 32 % R=0.995)
Figure 3. Temperature influence of the output signal
from the electro acoustic sensor system.
Discussion
The increase in output signal compared to the
reference, at 23 ºC compared to that of 19 ºC, shows
that cold inspired air would affect the measurements,
IFMBE Proc. 2005;9: 137

Biomedical instrumentation
Evaluation of a surgeon-centered laparoscopic surgical tool design
M.S. Hallbeck 1 , D. Oleynikov 2
1 Industrial Engineering, University of Nebraska, Lincoln, NE USA
2 Surgery, University of Nebraska Medical Center, Omaha, NE USA
Hallbeck@unl.edu
Abstract
Surgeon-centered design principles
were employed to design an articulating
laparoscopic tool. Evaluation of this tool
by 38 expert laparoscopic surgeons
demonstrated that they believed the new
tool could significantly reduce back,
shoulder, arm, wrist and hand pain and
stiffness. They preferred the new design to
conventional designs for comfort and
general impression. The added articulation
at the grasper tip was deemed a useful
addition by 92% and 89% of the surgeons
would purchase the tool once it was on the
market.
This study demonstrates that good
surgeon-centered design can improve a
standard laparoscopic tool. It further
demonstrates that given a choice between
current tools and ergonomically designed
tools, laparoscopic surgeons will select the
more comfortable, useful tool.
Introduction
Laparoscopic or minimally invasive
surgery employs small incisions for ports into
the body. These ports allow for inflation of the
area and a camera and tools to enter the body
to perform the surgery. This allows faster
healing (1 night in hospital and 1-3 weeks
before back to work) and lower rates of
infection compared to conventional procedures
(5-7 nights in hospital, 6-7 weeks before back
to work). There are approximately 500,000
laparoscopic procedures performed in the US,
with that number rising each year.
Laparoscopic surgery is rising in
popularity since the minimally invasive
procedures allow for reduced hospitalization (1
day vs 5-7 depending on the operation), time
away from work (1 week vs 3-7 weeks),
reduced post-operative pain and lower
infection rates that conventional “open”
surgery. While the benefits to the patients are
considerable, they come with a cost to the
surgeon. The time to perform the operation
can double with laparoscopic surgery as
compared to “open” surgery, in awkward
postures with poorly designed tools. Although
the advantages of minimally invasive surgery
have been clearly established for the patient,
studies have shown that the surgeon is faced
with numerous disadvantages caused by
poorly designed instrument handles, including
the potential of harm to the surgeon due to
awkward postures, high repetition and high
force exertions, and that there is the likelihood
of harm to the patient due to poorly designed
tools. Thus, the design of these instruments is
critical to the result of the surgery.
Current laparoscopic instruments have
been found to be very poorly designed
ergonomically and it is likely that ergonomics
were not considered at all. Berguer et al.
(1998) found 8-12% of practicing laparoscopic
surgeons frequently experience post operation
pain or numbness. This is generally
attributable to pressure points on the
laparoscopic tool handle. Matern et al. (1999)
studied four different handle designs used on
laparoscopic tools (shank, pistol, axial, and
ring handle) and found that all resulted in
either painful pressure spots or caused
extreme ulnar deviation.
To gather surgeon feedback on
laparoscopic tools currently being used during
laparoscopic surgeries, a questionnaire was
administered to 18 expert surgeons at the
University Medical Center after a session
learning a new advanced laparoscopic
technique was to examine the limitations and
problems associated with conventional tools.
The percentage of respondents who indicated
experiencing either slight or substantial
problems in the indicated areas during or after
use of the conventional grasper tools are over
50% for shoulder arm, hand and wrist pain and
stiffness, 60% for instruments awkward to
manipulate and 47% for not able to perform
fine or precise motions (Doné, et al 2004).
Another question asked surgeons to
identify, on a picture of a hand, where they felt
pain during or after laparoscopic surgery and
how painful the area was. Painful areas of the
hand were identified by 61% of the
respondents with an astounding 22% reporting
numbness in the thumb or fingers after
surgery. Based upon these data, ergonomic
evaluation of current tools and surgeoncentered
design principles of ease and
efficiency of use for error minimization,
accommodation of users to lead to subjective
satisfaction, an ergonomic articulating
laparoscopic grasping tool was designed. The
resulting tool contains several important
features including an ergonomic handle with
IFMBE Proc. 2005;9: 138

Biomedical instrumentation
an articulating end effector which is controlled
intuitively. This study is the evaluation of the
prototype developed using surgeon-centered
tool design.
Methods
Subjects: Thirty-eight laparoscopic
surgeons from across the U.S. attending
advanced laparoscopic surgical training at the
University of Nebraska Medical Center
volunteered to evaluate the tool. They were
asked to compare a conventional ring-type tool
with a surgeon-centered tool design prototype
using a questionnaire.
Apparatus: The Intuitool, an
ergonomic articulating laparoscopic grasping
tool was compared to a conventional tool
using a questionnaire that had questions from
the first questionnaire (summarized above)
and some additional questions that directly
compare the prototype to a conventional tool.
Procedure: Each surgeon was asked
to report the pain they felt using the
conventional tool during the surgery session
they had just completed. These questions
were identical to the questions in the predesign
survey asking about the pain. They
were then asked to use the prototype tool and
the standard ring-type tool in a clear plastic
torso to practice some laparoscopic skills.
After the practice with both tools, the
questionnaire was presented to the surgeon
and s/he was asked to complete the survey.
They were then asked if prototype tool would
relieve any of the problems they experienced
with conventional grasping tools (the same list
that was initially presented).
Experimental Design: Ordinal data
were collected throughout this questionnaire;
therefore, a Wilcoxon Signed Rank test was
used to analyze each hypothesis test. The
level of significance for all statistical tests was
0.05. All of the statistical tests were performed
using Minitab 14 (Minitab, Inc.).
Results
Surgeons were then asked to indicate
which, if any, of the problems they believed
would be relieved with use of the prototype
tool after using it in the clear plastic torso. The
results are shown in Figure 1 with stars to
indicate those percentages statistically
different from zero (α=0.05).
There was statistical preference
towards the comfort of the prototype handle
(p

Biomedical instrumentation
SKIN TEMPERATURE EFFECTS ON SKIN BLOOD FLOW AT
AREAS PRONE TO PRESSURE SORE DEVELOPMENT
A. Jonsson*, M. Lindgren**, M. Lindén*
* Department of Computer Science and Electronics, Mälardalen University,
Västerås, Sweden
** Department of Medicine and Care Nursing Science, Faculty of Health Sciences,
Linköping University, Linköping, Sweden
annika.jonsson@mdh.se
Abstract: The microcirculation in tissue areas, in
response to increased skin temperature has been
studied with laser Doppler flowmetry. To represent
areas that are prone to pressure sore development,
the heel and shoulder have been investigated and the
back of the upper arm has been chosen as a control
tissue area. The perfusion at the upper arm
increased in average 27 times and at the shoulder 17
times, when the skin temperature was increased to
38 °C. The perfusion at the heel showed large
fluctuations at all temperatures. The result
strengthens the opinion that the microclimate is a
vital parameter when evaluating antidecubitus
mattresses.
Intr oduction
Antidecubitus mattresses are used as preventing the
development of pressure sores for persons at risk and
also as a part of the pressure sores treatment. A
drawback with foam and latex rubber support surfaces
are that heat is accumulated, leading to an increased
interface temperature and thereby increased skin
temperature [1,2,3]. On the other hand air-loss
mattresses might lower the skin temperature [4], which
is not preferable since that can cause a vasoconstrictor
response and decrease the blood supply to the tissue [5].
The primary factor in developing pressure sore is
pressure on the tissue that reduces or even occludes the
blood flow, so called is chemia. A contributing factor to
pressure sore development is increased tissue
temperature since the tissue then requires a larger
amount of nutrition and oxygen. Normally an increased
blood flow should compensate for an increased skin
temperature. So even if the pressure on the tissue does
not occlude the blood flow, the tissue may suffer from a
relative ischemia due to insufficient skin blood flow
during prolonged heat provocation [6].
In addition one study have shown that spinal cord
injured have a higher skin temperature than able bodied
persons regardless of the type of wheelchair cushions
they were using [2].
The aim of this study was to investigate the
relationship between microcirculation and skin
temperature at areas prone respectively not prone to
pressure sore development.
Materials and Methods
Three heating probes (developed and designed at the
department) was used to heat the skin surface to pre-set
temperatures. The blood flow was recorded with laser
Doppler (Perimed’s PeriFlux laser Doppler Flowmetry
df2).
Thirteen subjects within the age 25 to 39 were
included in the study. Blood pressure, body temperature
and BMI were registered for every subject. The ambient
temperature was held constant. The individuals were to
acclimate and relax for approximately ten minutes
before laid in prone position. The heating probes, 2.5
cm in diameter, were placed at the left shoulder and left
heel to represent areas prone to pressure sore
development. The posterior side of the left upper arm
was used as a control site. Blood flow measurements
were performed after stabilization of the heating for five
minutes at each temperature: normal skin temperature
and with the tissue areas heated to 32 °C, 35 °C and 38
°C.
The Student’s t-test for paired data was used for the
statistical analysis.
Results
The subjects were afebrile and had normal BMI and
a normal blood pressure. During the measurements, the
ambient temperature in the room was kept constant at
23 ± 1 ºC.
Table 1: The mean perfusion and standard deviation at
the upper arm and shoulder.
Perfusion
(V)
Upper
arm
Normal
temp.
0.058
±0.037
Shoulder 0.088
±0.035
32 ºC 35 ºC 38 ºC
0.052
±0.031
0.011
±0.057
0.20
±0.18
0.31
±0.16
1.6
±0.86
1.5
±0.97
IFMBE Proc. 2005;9: 140

Biomedical instrumentation
perfusion than the upper arm at a lower skin
temperature might be that it is not an extremity.
The ambient temperature was held constant during
the study but it may have been to low since several
subjects experienced coldness during the measurement.
Today the main point in evaluation of antidecubitus
mattresses is pressure distribution. But different
antidecubitus mattresses have different possibilities to
maintain a healthy microclimate in the interface
mattress/person. This study shows that evaluation of the
support surfaces in aspect to heat accumulation is
important.
Conclusion
Figure 1: Mean perfusion values in the upper arm and
shoulder at different skin temperatures.
The perfusion in the shoulder, heel and upper arm
increases with increased temperature, table 1. The
perfusion in the upper arm increased in average 27
times and in the shoulder 17 times when the skin
temperature was increased to 38 ºC. The result shows
that the perfusion increases significantly between a skin
temperature of 35 ºC and 38 ºC, for both, the upper arm
(p

Biomedical instrumentation
BLUETOOTH ECG MONITORING SYSTEM
J. C. Tejero-Calado 1 , C. López 2 , A. Bernal 3 , M. A. López 2 , G. Quesada 4 , J. Lorca 5
1 Electronic Department, University of Málaga, Málaga, Spain
2 R&D Department, Andalusian ICT Centre (CITIC), Málaga, Spain
3 R&D Department, IMABIS, Málaga, Spain
4 Critic Care Unit, Carlos Haya Hospital Complex, Málaga, Spain
5 Directorship, revistaesalud.com journal, Málaga, Spain
jctejero@uma.es
Abstract: The aim of this project is the development
and implementation of a high level integration
wireless electrocardiograph, which allows wireless
monitoring of patients. To reach this, the following
sub-objectives have been achieved successfully: sense
and condition of the biosignal; A/D conversion;
digital processing, in order to reduce the noise level;
and transmission functionality. The wireless
capability has been implemented using a
BlueTooth TM module.
Introduction
There are signals which must be monitored
continuously or periodically in patients. The most
important ones are: pulse-oximetry, non-invasive blood
pressure, electrocardiography (ECG)… The
electrocardiogram is the record of the biopotentials,
measured on the body surface, originated by the
electrical activity of the heart.
The ECG BlueTooth favours at-home
hospitalization, reducing the affluence to sanitary
assistance centres to make routine controls. This fact
especially causes a really favourable social impact,
especially for elder people, rural-zone inhabitant,
chronic patients and handicapped people. Furthermore,
it offers new functionalities to physicians and will
reduce the sanitary cost. Among these functionalities,
biomedical signals can be sent to other devices (screen,
PDA, PC…) or processing centres, without restricting
the patients’ mobility.
BlueTooth TM [1-2] technology allows the
electrocardiograph to communicate with GPRS/UMTS
mobile devices. The implementation of wide range
homecare network can be carried out in an easier and
cost effective way. As well, Personal Area Network of
biomedical sensors can be implemented, using
BlueTooth concentration elements.
Methods
This section tries to show the solutions taken in the
design and implementation phases of the different
subsystems. The system has been divided into:
biomedical signal acquisition and conditioning, A/D
conversion and digital signal processing[3], and
transmission facilities (Figure 1).
Sensor
Biosignal
Microcontroller
Conditioning
Subsystem
Figure 1: Block diagram of the ECG.
A/D Conversion
Signal processing
Communication
Wireless communication
subsystem
Biomedical signal acquisition and conditioning:
This block senses the biosignal, by means of four
electrodes placed on the patient’s thorax, with the aim to
register several leads. The record obtained across
different pairs of electrodes results in different
waveform shapes and amplitudes; these different views
are called leads[4]. The four used electrodes generate
the three bipolar limb leads and the three unipolar limb
leads. If the unipolar chest leads want to be measured,
six additional electrodes are required.
The low-level amplitudes of these biopotentials,
range of miliVolts, force to condition the signal. An
analogue circuitry with a precision instrumentation
amplifier, with a gain around 1000, is implemented to
reach this purpose[4].
To protect the patients against over-current, serial
resistors are placed in the electrode inputs and outputs.
Their values have been chosen to limit the maximum
current to 13,6mA when the supply voltage is 4,5V.
These resistors are used, as well, to form a low pass RC
antialiasing filter with 110Hz cut-frequency. The ECG
bandwidth spreads from 0,05 Hz to around 100Hz. This
fact allows the sampling of the signal 250 times per
second.
A/D conversion and digital processing: To digitalize
and process the signal, a Texas Instrument
MSP430F149 has been utilized. This is a low
consumption 16-bit microcontroller, with eight 12-bit
converters though only four of them have been used,
one per lead.
IFMBE Proc. 2005;9: 142

Biomedical instrumentation
Once the signal has been digitalized, the
microcontroller processes it in order to reduce the
diverse kinds of noise[5]. Implementing a Notch filter
centred on 50Hz and two Kaiser (a high pass band and a
low pass band) windowed filters, the 50Hz noise[6], the
breath effect, the motion artifact and most of the EMG
noise are reduced. To minimize the EMG noise in band
with the ECG, two signal states must be identified. One
of them, when the signal carries information about the
electrical activity of the heart, and the other one,
between these electrical activity intervals. In the first
state, an average of three heart beats is made. In the
other one, the signal is filtered by a type I Chebychev
filter.
Transmission: Finally, the communication facility
must be implemented into the electrocardiograph in
order to make possible the connection with other
commercial BT devices.
For this purpose, a BT module (provided by
CETECOM TM ) has been integrated into the design. It
meets the v1.1 specifications, has got a Serial Port
Profile (SPP) and the transmission power is 20dB m
(Class I), allowing communications with other BT
devices within a range of 100m. Data transfer between
the microcontroller and the BT module is made through
an UART interface.
Results
As a result, a small size ECG (38 x 47 mm) with a
high transmission capability (up to 100m range) has
been designed and implemented (Figure 2). The
consumption is low enough (40mA in transmission
mode) to achieve 25 hour battery lifetime using a
1000mAHour battery.
Figure 3: Test scenario.
With the aim of confirming the right operation of the
device, tests have been carried out with different people
and conditions: at rest, walking, working at the office…
being the result satisfactory in all cases. The quality of
the obtained signal is equal than one of a commercial
wired ECG, and better when the patient is in movement.
Discussion
The main emphasis of this paper has been to explain
how an innovative ECG BlueTooth, which will increase
the patient’s wellbeing and the functionalities offer to
physicians, has been designed, implemented and tested.
The principal goals were to achieve size and
consumption as reduced as possible, and a high quality
signal. As it has been exposed above, these objectives
have been successfully reached.
This project has been supported by the Andalusian Health Service
(SAS204/03), Andalusian ICT Centre (CITIC) and The Mediterranean
Institute for Advance in Biotechnology and Sanitary Investigation
(IMABIS).
References
38mm
Figure 2: Top and bottom sides of the wireless ECG.
To confirm the quality of the obtained signal and to
analyse the performances of the ECG, applications
based on Windows XP and Windows Mobile 2003
platforms have been developed. Those let the user
choose the patients to monitor and the leads to display.
The test scenario can be observed in Figure 3. To
establish the wireless communication between the ECG
and the PC, an USB-BT adapter has been used. The PC
application manages the connection as a serial port and
display the signal according to the user settings. The
used PDA model has been a DELL TM Axim TM X30
combo wireless, 624MHz. It has a class 2 BT module.
The PDA application has been developed with
Embedded Visual C++ 3.0 and it handles the BT
connection in the same way as the PC application does.
[1] BRAY J., SENESE B. (2001): ‘Bluetooth
Application Developer´s Guide’, (Syngress
Publishing, Rockland)
[2] BlueTooth TM , Internet Site Address,
www.bluetooth.org.
[3] YUXING Y., DONGYUAN Y, RICHARD F.
(2002): ‘Development of a digital signal processorbased
new 12-lead synchronization
electrocardiogram automatic analysis system’,
Computer Methods and Programs in Biomedicine,
Vol. 69, no. 1, pp.57-63
[4] CARR, J.J., BROWN J. M. (1993): ‘Introduction to
biomedical equipment technology’, (Prentice Hall,
New Jersey)
[5] RAMOS CASTRO J (1997): ‘Detección de
micropotenciales auriculares de alta frecuencia’.
PhD Thesis, Universidad Politécnica de Barcelona.
[6] PROAKIS J. G., MANOLAKIS D. G. (1996):
‘Digital signal processing: principles, algorithms,
and applications’, (Prentice Hall, New Jersey)
IFMBE Proc. 2005;9: 143

Biomedical instrumentation
WIRELESS WEARABLE EMG AND EOG MEASUREMENT SYSTEM FOR
PSYCHOPHYSIOLOGICAL APPLICATIONS
N. Nöjd*, M. Puurtinen*, P. Niemenlehto***, A. Vehkaoja**, J. Verho**, T. Vanhala***,
J. Hyttinen*, M. Juhola***, J. Lekkala**, V. Surakka***
*Ragnar Granit Institute, Tampere University of Technology, P.O. Box 692, FIN-33101, Tampere,
Finland
**Measurement and Information Technology, Tampere University of Technology
***Department of Computer Sciences, University of Tampere
Abstract: Wireless technology enables us to build
unobtrusive, wearable measurement devices. In this
paper a wireless headband measuring bioelectric
field of the eye movements and muscle activation on
the forehead is introduced. The headband includes
embedded textile electrodes, an integrated amplifier
and wireless data transceiver system.
Introduction
niina.nojd@tut.fi
This work is part of a project called Wireless
Technology and Psychophysiological computing. The
aim of the project is to study and develop lightweight
wireless measuring technology for monitoring human
physiological and psychophysiological responses, such
as electrocardiogram (ECG), electro-encephalogram
(EEG), electro-oculogram (EOG), and electromyogram
(EMG).
Wireless measurement systems offer new
possibilities for physiological measurements. They
enable unobtrusive measurement of
psychophysiological responses and increase the
usability of measurement devices, as leads are no
longer needed.
This paper introduces a wearable wireless headband
measurement system with integrated textile electrodes
for measuring eye movements and facial muscle
activity. Textile electrodes improve the
comfortableness and usability of the headband.
Eye movements are often monitored with
appliances which utilize cameras [1, 2]. Those
solutions however are based on image processing made
after and during the measurements. Also EOG (electrooculogram)
has been used [3]. Our wearable device
will enable monitoring of eye movements as well as
muscle tension of the forehead in real time without
wires limiting the comfort or movement of the subject.
Materials and Methods
The measurement device is composed of head band
with its electrodes, amplifier, AD-converter (analog-to
digital converter), and wireless data transceiver.
Figure 1 illustrates the headband with its components.
Figure 1: Wearable wireless headband for EMG and
EOG recording. Amplifier and AD-converter lie on the
upper, and data transceiver on the lower circuit board.
The boards are exposed for demonstration from the
pocket designed to carry them.
There are five electrodes integrated into the
headband. Electrodes are embroidered with polyester
threads covered with silver. The textile electrodes are
embroidered in square shape with size of 20 mm x 20
mm. Material of the headband is flexible elsewhere but
non-flexible on the forehead part. Non-flexibility at
forehead part retain the inter electrode distances and
electrode areas constant. The developed cableelectrode
connection is illustrated in Figure 2.
Figure 2: Electrode-cable-connection. The form of the
copper fiber fan before embroidering is illustrated on
the right, and the completed electrode-cable-connection
in which the fiber fan lies under the embroidering is
illustrated on the left.
IFMBE Proc. 2005;9: 144

Biomedical instrumentation
As Figure 2 illustrates, copper fibers of cables are
fanned out and the conductive silver threads are
embroidered over the cables. Thus the electrode-cable
connection has high conductivity and is mechanically
sound.
Placement of the electrodes on the headband is
illustrated in Figure 3. Electrodes were placed in such a
way that the vertical and horizontal eye movement, and
the activity of facial muscles; corrugator supercilii and
frontalis muscle, could be detected. Each desired
activity is registered from a preset bipolar electrode
pair among these five electrodes.
Figure 5: Wirelessly recorded EOG signal. The signal
shape follows from the to-and-fro movement of eye.
Discussion
Figure 3: Location of the electrodes on the headband.
The amplifier used for signal measurement has six
measurement channels. An AD-converter with 16-bit
resolution and a sample frequency of
1000 Hz is used. The wireless data transfer is arranged
with ZigBee standard compatible radio transceiver
operating at 2,4 GHz frequency band.
For testing the system, eye movements and the
activity of forehead muscles were recorded wirelessly
with the head band. Recordings were implemented
with one channel. High- and low-pass filters were used
for extracting EMG and EOG signals, respectively,
from the original data.
Results
The obtained signals are illustrated in Figure 4
and 5. EOG signal (Fig. 5) follows from moving eye
diagonally to-and-fro. EMG signal (Fig. 4) originates
from furrowing, mainly from frontalis muscle.
Comfortable wireless measurement system for
EOG and EMG monitoring was developed. Connection
between textile electrodes and cables was created and
implementation of textile electrodes using
embroidering was tried and trusted.
Signals recorded with the head band are of good
quality and from those the eye movement and muscle
activity can be detected.
In the future there is possibility to substitute the
existing cables for conducting textile wire. That makes
the cable-electrode-connection easier to implement and
removes nonflexible part from the device. Conductive
textile wires have already been tested in our institute.
Acknowledgement
The project is supported by The Academy of
Finland [202186] and by a grant from The Finnish
Cultural Foundation. The authors gratefully
acknowledge Markku Honkala from Tampere
University of Technology SmartWearLab for helping
with the electrode realization.
References
[1] LAND, M. F., MENNIE, N. and RUSTED, J.
(1999): ‘The Roles of Vision and Eye Movements
in the Control of Activities of Daily Living’,
Perception, 28, pp. 1311-1328
[2] PELZ, J. B., CANOSA, R. L., KUCHARCZYK,
D., BABCOCK, J. S., SILVER, A. and KONNO,
D. (2000): ‘Portable Eyetracking: a Study of
Natural Eye Movements’, Human Vision and
Electronic Imaging V, SPIE Proceedings, 3659.
Figure 4: Wirelessly recorded EMG signal. Three burst
signals follow from furrowing.
[3] JOYCE, C. A., GORODNITSKY, I. F., KING, J.
W. and KUTAS, M. (2002): ‘Tracking Eye
Fixations with Electoocular and
Electroencephlographic
Recordings’,
Psychophysiology, 39, pp. 607-618
IFMBE Proc. 2005;9: 145

Biomedical instrumentation
AN INFLATABLE HIP PROTECTOR FOR PREVENTING INJURIES
FROM FALLS
Toshiyo Tamura*, Takumi Yoshimira**, Masaki Sekine *
* Department of Biomedical Engineering, Chiba University, School of Engineering Chiba, Japan
** Department of Information Technology, Tokyo Metropolitan College of Technology, Tokyo,
Japan
E-Mail tamurat@faculty.chiba-u.jp
Abstract: We developed a hip protector with a
telemetry acceleration monitoring system and an
airbag. The telemetry system evaluates the body
movement of the subject using accelerometry. The
monitor consists of a tri-axial accelerometer and
transmitter. The airbag inflation system consists of
a hip protector, a battery, and a gas cartridge .The
system is small, light, and able to be worn without
discomfort. In addition, the system is designed to
operate with no complex setting or handling. Our
system was attached to young healthy subjects, who
then mimicked falling. When a subject fell, the freefall
point was observed 100~200 ms before the fall
and the hip protector inflated successfully.
I Introduction
Falls are a serious problem for the elderly and
others prone to falling. One-third to one-half of the
population age 65 and over experience falls. Half of
the elderly people who fall do so repeatedly. Falls, a
complex phenomena that suggest the presence of
disease and predict future disability, are caused by
interactions between the environment and dynamic
balance, which is determined by the quality of sensory
input, central processing, and motor responses. Falls
are the leading cause of injury in elderly people and
the leading cause of accidental death in those over age
85. Even falls that do not result in injury can have
serious consequences. Psychological trauma and fearof-falling
produce a downward spiral of self-imposed
activity reduction, which leads to loss of strength,
flexibility, and mobility, thereby increasing the risk of
future falls. Therefore, we developed an inflatable hip
protector for preventing injuries during falls.
operation make it portable. Figure 1 shows an operation
of this device.
Figure 1. Concept of operation
The self-contained protective system is
designed to protect the hips, pelvis, buttocks, and
coccyx areas of the subject. The device can be worn
outside clothing. As it is small and lightweight (250 g),
it is easily put on and removed and does not interfere
with body movements. It contains an inflatable air bag
folded into the protector, a battery, a gas cartridge,
sensors to determine acceleration, a triggering/valve
mechanism to release the gas, and a relief valve. Figure
2 shows a block diagram of the system.
II Materials and Methods
2.1 Apparatus
We developed an inflatable hip protector to absorb the
shock of a fall and reduce the impact on the human
body by automatically inflating an airbag when the
wearer falls. An accelerometer that detects a fall was
attached to the back of the subject’s waist with the hip
protector. Its compact design and battery-powered
Figure 2 Block diagram
When the subject falls, a trigger signal is
activated, gas from the cartridge is released
IFMBE Proc. 2005;9: 146

Biomedical instrumentation
automatically, and the airbag assembly is inflated,
forcing the folded protector to fully cover areas of the
subject’s body. The response time of the airbag is
around 100 ms. After use, the relief valve is opened to
release air from the airbag assembly; the protector is
reinserted into the system; and the device is ready to
reuse once the spent cartridge has been replaced.
The signal that triggers the release valve is
obtained from the acceleration signal. We assumed
that free-fall is reached within a short time during a
fall. The tri-axial acceleration of the subject shows
zero gravity when the standing subject’s acceleration
is 9.8 m/s 2 .
2.2 Experiment set-up
The reliability, stability, and repeatability of the
triggering signal based on acceleration were tested and
the threshold value for releasing the valve and inflating
the airbag was determined.
An experiment was performed on 13 normal
young subjects (25.5±5.7 years old, 62.3±4.0 kg
weight, and 173.3±4.8 cm tall), after obtaining
written informed consent. The device for detecting
acceleration was attached to the subject. Each subject
mimicked three falls. Then, we had ten subjects mimic
falls while wearing the airbag assembly.
III Result
Figure 3 shows a typical example of the waveform
during a fall. After the measurement started, the
impact accelaration was observed at 1.8 seconds.
Before fall, the subject was standing with a vertical
acceleration of 9.8 m/s 2 . During a fall, vertical
acceleration decreased reached 0 m/s 2 just before
falling. This point is the free-fall point. In the 39 trials,
the free-fall point shown in Fig. 2 was attained in the
subjects an average of 156±74 ms before falling.
Some subjects did not show a zero value, because they
did not lump away from the floor, bt remained
supported on their feet. From the results, a threshold
value of less than ±3 m/s 2 was deemed appropriate. If
the acceleration dropped below 3 m/s 2 , the valve was
opened.
In the test of airbag inflation, 9 out of 10 trials
were successful, while there was no trigger signal in one
trial, because the response time was too short to open
the valve.
Figure 3. Typical example of the acceleration waveform
during a fall
IV Discussion
In our test, a triggering signal with a magnitude of less
than ±3 m/s 2 was reliable and had good reproducibility.
However, the response time of 156 ms was too short to
inflate the airbag. The response time of the airbag itself
is 100 ms and some subjects fell faster than the valve
opened.
Further study is needed to 1) determine the
effectiveness of the acceleration signal, 2) identify
another algorithm, such as matching the pattern of
walking for a fast response, and 3) examine angle
monitoring with a gyroscope.
In conclusion, the use of our fall sensor and
airbag-equipped hip protector should save lives and
reduce injuries from falls at construction sites and other
locations.
This work was partly supported by a grant-in-aid of
Longevity Science, and grants from the National
Institute for Geriatric Medicine, Secom Science and
Technology Foundation, and Suzuken Memorial
Foundation, Priority Research of Chiba University
IFMBE Proc. 2005;9: 147

Biomedical instrumentation
ACOUSTIC MONITORING OF LUNG SOUNDS FOR THE DETECTION OF
ONE LUNG INTUBATION
S. Tejman-Yarden 1 , L. Weizman 2 , A. Zlotnik 1 , J. Tabrikian 2 , A. Cohen 2 , G. Gurman 1
1 pediatric devision, Soroka medical center, Beer Sheba, Israel
2 FACULTY OF ENGINEERING SCIENCES, Ben-Gurion University, Beer Sheva, Israel
tegmanya@hotmail.com
Abstract
Analysis of lungs sounds for monitoring and diagnosis
of pulmonary function is well known. One of the
applications of this method is detection of One Lung
Intubation (OLI) during anesthesia or intensive care.
In this work we have applied a new algorithm
which assumes a MIMO (Multiple Input Multiple
Output) system, in which a multi-dimensional AR
(Auto-Regressive) model relates the input (lungs) and
the output (recorded sounds). The unknown AR
parameters are estimated, and a detector based on the
estimated eigenvalues of the source covariance
matrix detects one lung ventilation situations . Testing
the algorithm on real breathing sounds, which were
recorded in a surgery room, shows more than 90%
accuracy in OLI detection.
Introduction
One lung intubation (OLI) is the most common
complication of endotracheal intubation. This incident
has serious complications such as atelectasis, hypoxemia
and pneumothorax and as of today none of the
monitoring techniques used in the operating room has
proved successful in detecting OLI and alarming when it
occurs. In a preliminary study, based on lung sounds
sampling we proved that the acoustic monitor can be of
use for the detecion of unintended OLI. As a result of that
experiment we recognized that each microphone attached
to the chest samples both ipsilateral and contralateral
lungs, each with a different acoustic contribution and an
algorithm which relates each microphone to its ipsilateral
lung can not evaluate properly the ventilation status of
the patient. The aim of this study is to develope a reliable
new algorithm for the determination of the number of
ventilated lungs every respiration for the detection of
OLI.
Methods
24 adult surgical patients schedualed for routine surgical
procedures were included. The patients’ lung sounds
were sampled were sampled by four piezoelectric
microphones attached to the patients’ back. Sound
sampling was done during induction and tube
positioning. To achieve OLI, after induction, the tube was
inserted and advanced down the airway so that no left
breath sounds were heard; the tube was then withdrawn
stepwise until equal breath sounds could be heard and the
final position of the tube was confirmed by fiberoptic
bronchoscopy. The sampled lung sound were processed
by a new algorithm which assumes a MIMO (Multi Input
Multi Output) system; in which a multi-dimensional autoregresive
(AR) model relates the input (lungs) and the
output (recorded sounds). The unknown AR parameters
of each respiration were estimated, and a classifier based
on the estimated second highest eigenvalue of the
covariance matrix was used to indicate the number of
lungs ventilated on each recorded respiation without
retrieving the original breath sounds.
Results
After filtration and processing the second highest
eigenvalue was calculated for every respiration. The
results of the algorithm were consistent over 24
experiments performed on patients. Since the database is
not comprehensive enough to be used for both training
and validation of the proposed method, estimation of the
performance of the system was performed as follows.
Twenty experiments were used to extract the histograms
in order to train the system and define the values of the
second highest eigenvalue and the other four were tested.
This procedure was repeated 6 times.
The analysis of the results is based on the probability of
detection of OLI. There are two types of errors in OLI
detection: P miss, the probability of a true OLI to be
wrongly detected as bilateral ventilation, and false alarm,
P fa , the probability of bilateral ventilation to be detected
as OLI. The Detection Error Tradeoff (DET) curve is a
common mean to display these errors. The DET curve
provides information about the device’s performance,
where each point on the curve shows the P fa and P miss for
a given threshold. The threshold of a real monitoring
system should be calculated according to the requested
sensitivity of the system, while taking into consideration
the allowed P miss of the system.
The Equal Error Rate (EER) point is defined as the point
on the DET curve where P miss = P fa , in this system it is
4.8%. Meaning that by using the offered method an OLI
correct detection probability of 95.2% with P miss of 4.8%
and P false of 4.8% can be achieved.
IFMBE Proc. 2005;9: 148

Biomedical instrumentation
Discussion
We have developed a new device for the detection of
proper ventilation during mechanical ventilation. This
device is based on the acoustic sampling of lung sounds
by piezoelectric non-invasive microphones attached to
the patient’s back. The lung sounds which were
considered in the past as non-reliable are analyzed by an
algorithm, which does not retrieve the original source
from the sampled sound, but rather estimates the
eigenvalues of an auto-regressive model. These values
represent the four major sounds in the chest- right lung,
left lung, right lung noise and left lung noise. By
assuming the four eigenvalues of the estimated matrix
represent the four sounds, we extract the second highest
to be the one who represents the left lung. We show that
by doing so we can detect OLI situations with the
probability of 95.2% with 4.8 of false alarms and
depending on the DET curve we can achieve even 98%
detection but with 9% of false alarms.
Conclusions
This Model is a preliminary prototype for a device to be
used in the operating rooms and intensive care units for
the follow up of proper ventilation of the patients.
IFMBE Proc. 2005;9: 149

Biomedical instrumentation
ELECTROMUSCULAR INCAPACITATING DEVICES
J. G. Webster
University of Wisconsin-Madison/Department of Biomedical Engineering, 1550 Engineering Drive,
Madison WI 53706 USA webster@engr.wisc.edu
Abstract: An electromuscular incapacitating
device (EMD) temporarily incapacitates combative
individuals so they can be apprehended with
minimal harm to themselves, bystanders, or law
enforcement agents. Many newpaper articles [1]
have suggested that EMDs can kill those
apprehended. A peer reviewed study on
anesthetized swine reports that there is a large
safety factor that prevents EMDs from killing [2].
Biomedical engineers, with knowledge of the effects
of electricity on the body can provide information to
clarify this conflict.
Introduction
An electromuscular incapacitating device (EMD)
presents law enforcement with another less-lethal
option for subduing violent individuals who pose a
threat to themselves and others. More commonly
known as stun guns, or as a Taser®, they deliver a high
voltage electric charge that disrupts the nervous
system, causing the suspect to lose control and fall to
the ground. Less lethal weapons are designed to
temporarily incapacitate, confuse, delay, or restrain an
adversary in a variety of situations.
Methods
It is useful to understand the operation of EMDs.
All generate voltages of about 50 kV, currents of about
2 to 15 A, pulse durations of about 10 to 50 µs,
repetition rates of about 20 pulses/s, for about 5 s, and
can ionize an air gap to form an arc about 50 mm long.
Early EMDs, called stun guns, required close contact
with the person, and since the electrodes are about 50
mm apart, only affected a small target area. Later
EMDs used a compressed gas propellant to fire barbed
darts with trailing insulated wires about 7 m long at an
8° angle so the darts would stick on clothing or the skin
a distance of about 50 cm apart to cause muscle
contraction and pain over a larger group of muscles.
Fig. 1 shows a typical circuit for the Taser M26 [3].
The battery drives a 500 Hz oscillator with transformer
step up from 2.5 to 6 V up to 2 kV, which is rectified
and forms a high-voltage power supply to charge up
the capacitor. When the capacitor voltage reaches
about 2 kV, the spark gap breaks down and the 2 kV is
delivered across the primary of the transformer, which
steps it up to about 50 kV, which will ionize an air gap
of 50 mm. If the barbs strike the skin, the 15 A through
the typical body resistance of 300 Ω yields 4500 V. If
the barbs strike clothing, the arc jumps through the air
but the body voltage is still 4500 V. Numerous groups
of skeletal muscles contract, and the person loses his
ability to maintain an erect, balanced posture and falls
to the ground and is temporarily incapacitated.
Fig. 1: When the 0.88 µF capacitor voltage reaches 2
kV, the spark gap breaks down. The 1:25 transformer
creates 50 kV. From [4].
The effective charge delivered to the subject that
might excite the heart and cause ventricular fibrillation
(VF) is contained in the first one-half sinusoidal
waveform and is 9 µs duration times 15 A peak times
0.63 (to convert to average current) = 85 µC.
Fig. 2 shows an improved Taser X26 EMD [3].
Capacitor 1 generates 50 kV at the output to break
down the air resistance. Then capacitor 2 provides a
sustaining current at lower voltage. The result is
lowered battery requirement.
Fig. 2: 0.07 µF capacitor 1 breaks down the air gap in
1.5 µs (arc phase). 0.01 µF capacitor 2 yields a 50 µs
low voltage sustaining current (stim phase). From [4].
Fig. 3 shows a standard waveform of about 2 A for
about 50 µs for a charge of 100 µC. The larger currents
were used to determine cardiac safety factor, which
IFMBE Proc. 2005;9: 150

Biomedical instrumentation
varied from 15× to 42×, as swine weight increased
from 30 to 117 kg.
Fig. 3: The standard waveform from the TASER model
X26 EMD delivers about 2 A for about 50 µs. From [2]
Results
To estimate safety, we can use [5] Fig. 22. For
durations of 1000 µs there is no VF up to 1.4 A (1400
µC). For durations of 100 µs there is no VF up to 7 A
(700 µC). While a durations of 9 or 50 µs are not on
Fig. 22, extrapolation would yield no VF well in excess
of the M26, 85 µC or X26, 100 µC. [5] Fig. 18 shows
that for the low duty cycle of less than 0.1, the
threshold for VF for multiple shocks as used in the
Taser does not change.
Geddes and Baker [6] have developed the strength–
duration curve shown in Fig. 4. They estimate the time
constant τ for cardiac muscle is about 2 ms. For an
EMD duration of 50 µs, current must be increased over
that required for long durations by a factor of 50 to
cause cardiac excitation.
Figure 4: Short duration pulses require higher currents
to cause excitation. From [6]. Charge Q = Id remains
constant for short pulses.
Skin depth for human tissue is (460 m)/(square root
f), where f = frequency [7]. For M26, f = 1/(18 µs) =
55,000 Hz. Skin depth = 1.96 m. Thus there is little
skin depth effect in human tissue at Taser frequency.
Discussion
We are developing a computer model of currents
that flow through the body in response to EMD darts at
a variety of locations. Intuition suggests that a dart is
more likely to induce VF if near the heart rather than
farther away. We will map the contours of cardiac
safety factor for inducing VF. We will verify the model
by tests in which we place a 64-electrode Constellation
catheter within the anesthetized swine heart. We will
also test larger currents from the surface as shown in
Fig. 3 and develop a bench test for any EMD.
If the EMD electric shock induces VF, the blood
pressure would drop to near zero in about 5 s [8]. The
human would lose consciousness within 30 s, which
does not occur in the large majority of deaths following
EMD. We conclude that the electric shock did not
directly cause VF. Alternative hypotheses for death
following EMD shock include positional asphyxia,
skeletal muscle damage causing hyperkalemia and
acidosis, heat, and drugs [9].
References
[1] BERENSON, A. (2004): 'As police use of Tasers
rises, questions over safety increase', The New York
Times, CLIII (52,914), Sunday July 18
[2] MCDANIEL, W. C., STRATBUCKER, R. A.,
NERHEIM, M., BREWER, J. E. (2005): 'Cardiac safety of
neuromuscular incapacitating defensive devices',
PACE Supplement 1, 28, pp. S284-7
[3] Taser M26 and X26 manuals
http://www.taser.com/index.htm
[4] NERHEIM, M. H. (2004): 'Electronic disabling
device', International patent application
WO 2004/073361 A2
[5] IEC (1987): 'Effects of current passing through the
human body IEC 479-2, 2nd ed.', (International
Electrotechnical Commission, Geneva)
[6] GEDDES, L. A., AND BAKER, L. E. (1989):
'Principles of applied biomedical instrumentation' 3rd
ed., (John Wiley & Sons, New York), p. 460
[7] F. OLYSLAGER AND D. DE ZUTTER, Skin effect, in
J. G. WEBSTER (Ed.): Wiley Encyclopedia of Electrical
and Electronic Engineering, (John Wiley & Sons, New
York) Vol. 19, p. 315
[8] ANTONI, H. (1998): 'Electrical properties of the
heart', in REILLY, J. P. (Ed): 'Applied bioelectricity
from electrical stimulation to electropathology',
(Springer, New York), p. 181
[9] LAUR, D. (2004): 'Excited delirium and its
correlation to sudden and unexpected death proximal to
restraint', (Canada: Victoria Police Department)
http://www.taser.com/facts/medical_info.htm
This project was supported by Grant No. 2004-IJ-CX-
K036 awarded by the National Institute of Justice,
Office of Justice Programs, US Department of Justice.
Points of view in this document are those of the author
and do not necessarily represent the official position or
policies of the US Department of Justice.
IFMBE Proc. 2005;9: 151

Biomedical instrumentation
A WIRELESS USB BASED VIEWING BOX FOR
CEPHALOGRAM ANALYSIS
Kuo-Sheng Cheng, Ching-Lin Li, Cheng-Yu Chen, Wen-Hung Ting, and Yen-Ting Chen
Institute of Biomedical Engineering, National Cheng Kung University,
Tainan, Taiwan, ROC.
kscheng@mail.ncku.edu.tw
Abstract: In this paper, a multi-functional viewing
box with wireless USB port is developed for X-ray
film analysis. It is a combination of touch panel and
traditional light box. Based on the look-up table for
calibration, the coordinate of touching point on
attached X-ray film may be precisely located and
transmitted to PC wirelessly. From the experimental
results, its resolution, stability, and repeatability are
demonstrated to be good for cephalogram analysis.
It may be easily extended to other clinical
application. Besides, it may be very helpful in e-
orthodontics.
Introduction
Touch Panel
A/D Converter
Microcontroller
Graphic User Interface
USB
Module
Cephalogram is the basic X-ray film used for the
diagnosis and treatment in orthodontics. Based on it,
orthodontist usually locates the landmarks and makes
the measurements for analysis. Although several
approaches for automated cephalogram analysis have
been proposed in the past[1-5], the process of
landmarking is generally done by manual method. Some
problems still need to be solved. First of all, the data
obtained from the tracing paper are not in digital form
for on-line processing and e-orthodontics. Second, there
is a storage problem in tracing papers. Third, the tracing
papers drawn before and after treatment may not be
easily registered for superimposition analysis.
In order to solve the first problem, a wireless USBbased
multi-function viewing box is designed and
developed in this papre. ssion. Finally, a summary is
made in the conclusion.
Materials and Methods
Fig. 1 shows the block diagram of the cephalogram
analysis system with wireless USB-based multi-function
viewing box..
This system has the great potential in clinical
application due to its two features. Firstly, it is better
than a serial or parallel port for its plug and play,
multiple devices at one time, small power supply, and
fast transfer rate. Secondly, the wireless module is
designed to operate in the worldwide 2.4GHz Industrial,
Scientific, and Medical (ISM) frequency band (2.4GHz-
2.4835GHz). Several devices may be linked
simultaneously for teaching and education purpose
[6][7].
Wireless
Transmitter
Wireless
Receiver
Figure 1: The block diagram of the proposed system
In this study, the whole system is divided into three
major parts: (1) the data communication driver between
the software design and wireless USB port, (2) teh
wireless module circuit and driver, and (3) the data
communication between the microcontroller and the
touch panel.
Here, the chip of DMA-USB FX (CY7C64613) is
used to develop the USB module. It provides the 8051
kernel to develop the firmware program which receives
coordinate data from wireless and transmits coordinate
data to PC. Chip (CYWUSB6935) of CYPRESS
company is applied to realize the wireless module.
The CYWUSB6935 transceiver is a single-chip 2.4-
GHz Direct Sequence Spread Spectrum (DSSS)
Gaussian Frequency Shift Keying (GFSK) baseband
modem radio. It is a receiver which connects directly to
a USB module. It also is a transmitter which connects
directly to a microcontroller (AT89C2051).
The data communication between the
microcontroller and the touch panel is very important. It
is involved in the data format and resolution for the
accuracy of location. IC AD7843 that has 16-bit
resolution is used to convert analog signal into digital
signal. The value of the reference voltage will determine
the input range of the converter. The output coding of
the AD7843 is in binary.
The data format of the coordinate is illustrated in
the following.
IFMBE Proc. 2005;9: 152

Biomedical instrumentation
Check
Byte
X-High
Byte
Data String
X-Low
Byte
Y-High
Byte
Y-Low
Byte
0xFF
1 byte 1byte 1byte 1byte
The coordinate data is transferred to the computer
one at a time through wireless module. The correctness
of the data may be checked by checking byte. Here X-
high byte represents the high byte of the X-coordinate
data. It contains four bits. The data structures for the X-
high byte and Y-high byte are the same. The X-low byte
and Y-low byte include the same eight bits. The x-
coordinate and y-coordinate may then be obtained as
X =XH*256+XL (12bits)
Y=YH*256+XL (12bits)
Results and Discussion
(1)
The performance of the proposed system was tested
in regards to the resolution, stability, and reproducibility.
1. Resolution: Five sequences of points locating on
different parts of the panel are uniformly distributed and
selected. Each sequence has 9 points with different
intervals from 0.04 to 5.00 mm, and every point is
repeatedly located 20 times. A student t-test is then
applied to evaluate the distributions of each pair of
adjacent points. From the experimental results, it is
shown that the resolution of the panel with X-ray film is
about 0.155 mm.
2. Stability: This test evaluates the stability of the output
signal of the control circuit when a position at the panel
is continually pressed. The stability mainly depends on
the performance of the A/D converter and the related
circuits. 10 individual points are randomly sampled on
the panel. Then, a plastic pen held by the PCB
engraving machine is placed on each position with the
constant force for a while, and the output signal of the
control circuit is uniformly sampled 300 times. The
standard deviations for these data of 10 points range
from 0.15 to 0.24 mm on the x-axis and 0.32 to 0.50
mm on the y-axis, respectively.
3. Reproducibility: It is important to test whether the
generated signal retains the same value to represent the
same place for several touches. This reproducibility is
assessed as follows. Each of four randomly selected
positions on the panel is touched ten times. The results
show that the standard deviations of these data range
from 0.07 to 0.16 mm on the x-axis and 0.07 to 0.26
mm on the y-axis, respectively. The panel is
demonstrated to be quite precise in response to each
touches of the same position.
In the user-interface design, the Microsoft VC++ is
employed to implement the cephalogram analysis
system and coordinate data manipulation. An example
of saved tracing file for cephalogram is shown in Figure
2.
Figure 2: An example of the saved tracing file
Conclusions
A novel multi-functions X-ray film viewing box
with wireless USB is developed for cephalometry. The
device can also be easily extended to other applications
such as X-ray film analysis, medical image analysis. In
the future, a see-thru vertical scanner for digitizing X-
film may be incorporated to acquire images. It may be a
stand-alone devices in varied medical applications.
Acknowledgments
This work is supported in part by National Science
Council of ROC under the Grant#NSC93-2213-E006-
044, NSC91-2213-E006-070, NSC90-2213-E006-131,
NSC 89-2213-E006-160.
References
[ 1 ] S. Parthasarathy, S. T. Nugent, P. G. Gregson, and
D. F. Fay, “Automatic landmarking of
cephalograms,” Comput. Biomed. Res., vol.22, pp.
248-269, 1989.
[ 2 ] W. Tong, S. T. Nugent, P. H. Gregson, G. M.
Jensen, and D. F. Fay, “Landmarking of
cephalograms using a microcomputer system,”
Comput. Biomed. Res., vol. 23, pp. 358-379, 1990.
[ 3 ] J. Cardillo and M. A. Sid-Ahmed, “An image
processing system for locating craniofacial
landmarks,” IEEE Trans. Med. Imag., vol. 13, no.
2, pp. 275-289, 1994.
[ 4 ] D. J. Rudolph, P. M. Sinclair, and J. M. Coggins,
“Automatic computerized radiographic
identification of cephalometric landmarks,” Am. J.
Orthod. Dentofac. Orthop., vol. 113, no. 2, pp.
173-179, 1998.
[ 5 ] Y.-T. Chen, K.-S. Cheng, and J.-K. Liu,
“Automated cephalogram landmarking,” Chinese
J. Med. Biol. Eng., vol.16, no.2, pp. 199-213,
1996.
[ 6 ] Jan Axelson, USB Complete: Everything you need
to develop custom USB peripherals, 2 nd edition,
Madison, 2001.
[ 7 ] John Hyde, USB Design by Example: A practical
guide to building I/O devices, New York: John
Wiley & Sons, 1999.
IFMBE Proc. 2005;9: 153

Biomedical instrumentation
TISSUE PERMEABILIZATION STUDIED ON A MATHEMATICAL
MODEL OF A SUBCUTANEOUS TUMOR IN SMALL ANIMALS
N.Pavšelj 1 , Z. Bregar 1 , D. Cukjati 1 , D. Batiuskaite 2 , L.M. Mir 3 , D. Miklavčič 1*
1 University of Ljubljana, Faculty of Electrical Engineering, SI-1000 Ljubljana, Slovenia
2 Vytautas Magnus University, Department of Biology, LT-3000 Kaunas, Lithuania
3 UMR 8121 CNRS, Institute Gustave-Roussy, F-94805 Villejuif Cédex, France
*corresponding author (damijan@svarun.fe.uni-lj.si)
Abstract: Anti-tumor efectiveness of
chemotherapeutic drugs is increased by local
application of short intense electric pulses. This
causes an increase of the cell membrane
permeability and is called electropermeabilization.
In order to study the course of
tissue permeabilization of a subcutaneous tumor in
small animals, a finite elements model was built. The
results of the model were compared to experimental
results from the treatment of subcutaneous tumors
in mice and a good agreement was obtained.
Introduction
The application of electric pulses to cells causes
structural changes in the cell membrane, which becomes
more permeable [1]. If pulse amplitude, duration and
number of pulses are correctly set, the change in the
membrane permeability facilitates the cell uptake of
ions, molecules such as DNA or drugs, which otherwise
can not cross the membrane. The cell membrane is
permeabilized when a threshold transmembrane voltage
is reached, i.e. when the external electric field is above
the reversible threshold value. The goal is to apply an
electric field that will be between the reversible and
irreversible permeabilization threshold which destroys
the cell. [2] The specific conductivity of the tissue
increases when permeabilized [3,4], which could be
used as an indicator of the level of the
electropermeabilization in the tissue in question.
We built a mathematical model of a subcutaneous
tumor in small animals and studied the electroporation
process in tissue during the electrochemotherapy taking
into account the increase in tissue conductivity due to
cell membrane electroporation [5]. We modeled the
dynamics of the electroporation process with a sequence
of static models that describe electric field distribution
at discrete time steps during the process. We tuned the
model to the experimental data on different single
tissues. Finally the model was compared to the
experimental data on subcutaneous tumors with plate
electrodes pressed against the skin.
Materials and Methods
Numerical calculations were made by means of a
commercial program EMAS (Ansoft, Pittsburgh, PA,
USA) based on finite elements method.
This work is based on the data collected during a
study within the EC project Cliniporator (QLK-1999-
00484). We were studying the response of different
tissues to high voltage pulses delivered through plate
electrodes to rat skeletal muscle, skinfold and mouse
tumors. For the electroporation protocol a train of 8
square-wave pulses of 100 µs duration, delivered at a
repetition frequency of 1 Hz and generated by a PS 15
electropulsator (Jouan, St. Herblain, France). All
experiments were carried out in accordance with the
Guide for the care and use of laboratory animals.
Results
First we modeled the in-vivo experimental tissueelectrode
set-ups and the electroporation process of each
tissue separately. The results of the models were
compared to the experimental data (current-voltage
dependencies and 51Cr-EDTA uptake). We fine-tuned
the electroporation parameters of the single tissue
models, such as initial specific conductivities,
electropermeabilization thresholds and the changes in
specific conductivity, until we established good
agreement between the output of the model and the
experimental data. After setting these values for all
tissues, the subcutaneous tumor model was built (Figure
1).
3,12 mm
24 mm
6,43 mm
7mm
24 mm
0,4 mm
0,4 mm
9,2 mm
Fig. 1. Model made in EMAS for subcutaneous tumor.
The model is composed of four different tissues:
skin, underlying connective tissue, tumor and muscle.
IFMBE Proc. 2005;9: 154

Biomedical instrumentation
The electric pulses were applied on the skin with plate
electrodes pressed against the skin (distance: 8 mm).
In Figure 2 six steps of the electroporation process in
the model are shown (voltage: 1000 V). Time intervals
between steps are in general not equal. Different steps
follow a chronological order but do not have an exact
time value associated with them. Step 0 denotes the
electric field distribution as it was before the
electroporation process started, thus when all the tissues
had their initial specific conductivities.
between electrodes show anti-tumor effectiveness of
electrochemotherapy for amplitudes exceeding 720 V.
Anti-tumor effect increased at 1200 V, however this or
higher amplitudes can result in ulceration due to
massive tumor destruction. Nevertheless, anti-tumor
effects (increased tumor growth delay) can be achieved
also at 1040 V, where minimal anti-tumor effect of
electroporation itself is observed and most of the tumor
cells are permeabilized. Our model distributions
together with the reversible and irreversible electric
field thresholds obtained from in vivo measurements
agree well with the effectiveness of the
electrochemotherapy and the necrosis stage of tumors,
depending on the electroporation pulse amplitude.
References
Fig. 8. Six steps of the sequential analysis at 1000 V
between two plate electrodes with distance of 8 mm.
The last step of the sequential analysis, step 5, at
1000 V (Figure 2) the tumor is entirely permeabilized,
in some areas the electric field is also above the
irreversible threshold (800 V/cm). At 500 and 1500 V
(Figure 3, only the last step shown), we can observe that
the tumor is partially permeabilized already at 500 V
(reversible threshold is 400 V/cm) while at 1500 V large
part of the tumor is above the irreversible electric field
threshold. Observing all the steps of the process, we can
see that at first the electric field is the highest in the skin
layer (permeabilized first). After that the electric field
"penetrates" to the deeper layers of the model and
causes the rest of the tissues being permeabilized as
well.
[1] J.C. Weaver and Y.A. Chizmadzhev, “Theory of
electroporation: A review,” Bioelectrochemistry and
Bioenergetics, vol. 41, pp. 135-160, 1996
[2] D. Miklavčič, D. Šemrov, H. Mekid and L. M. Mir,
“A validated model of in vivo electric field
distribution in tissues for electrochemotherapy and
for DNA electrotransfer for gene therapy,” Biochim
Biophys Acta, vol. 1519, pp. 73-83, 2000
[3] U. Pliquett, R. Langer and J.C. Weaver, “Changes in
the passive electrical properties of human stratum
corneum due to electroporation,” BBA, vol. 1239,
pp. 111-121, 1995
[4] M. Pavlin, D. Miklavčič, “Effective conductivity of a
suspension of permeabilized cells: a theoretical
analysis,” Biophys. J., vol. 85, pp. 719-729, 2003.
[5] N. Pavšelj, Z. Bregar, D. Cukjati, D. Batiuskaite,
L.M. Mir, D. Miklavčič, “The course of tissue
permeabilization studied on a mathematical model of
a subcutaneous tumor in small animal,” IEEE Trans.
Biom. Eng. (in press)
[6] G. Serša, M. Čemažar and D. Miklavčič, “Antitumor
Efectiveness of Electrochemotherapy with cis-
Diamminedichloroplatinum(II) in Mice,” Cancer
Res, vol. 55, pp. 3450-3455, 1995
[7] G. Serša, M. Čemažar, D. Miklavčič and D.J.
Chaplin, “Tumor blood flow modifying effect of
electrochemotherapy with bleomycin,” Anticancer
Research, vol. 19, pp. 4017-4022, 1999
Acknowledgment
Fig. 9. The last step of the sequential analysis at 500
and 1500 V, respectively. Plate electrode distance was 8
mm.
This work was supported by the Ministry of Higher
Education, Science and Technology of the Republic of
Slovenia and EC under the FP5 grant Cliniporator
(QLK-1999-00484).
Discussion
We compared the results of the model with some
experimental results of electrochemotherapy [6,7].
Experimental results for treatments with 8 mm distance
IFMBE Proc. 2005;9: 155

Biomedical instrumentation
Production of Nano/Micro Beclomethasone Dipropionate particles
Using Supercritical Carbon Dioxide
M. R. Golriz 1 , E. A. Matida 2 and B. M. Andersson 1
1 Umea University, Dept. of Applied Physics and Electronics, Umeå, Sweden
2 Carleton University, Dept. of Mechanical and Aerospace Engr., Ottawa, Canada
mohammad.golriz@tfe.umu.se
Abstract
The rapid expansion of supercritical solution
(RESS) process and Gas/ Supercritical Anti-
Solvent (GAS/SAS) process were studied for
the recrystallization of beclomethasone dipropionate.
Beclomethasone dipropionate is used
as an inhaled steroid for treatment of asthma.
Methanol was used as solvent or co-solvent.
The results indicated that the process
conditions strongly influenced the particle size
and morphology. In the RESS process, submicrometer
size particles were formed with
the morphology changing from rod-like to
spherical as pressure was decreased. For the
GAS process, rod-like crystals of
beclomethasone dipropionate were formed of
about 4 µm in diameter.
Introduction
Design of small particles, ranging from nanometers
to hundreds of micrometers, have
attracted interest in the pharmaceutical, food,
nutraceutical, chemical, paint/coating, and polymer
industries. The important properties of these
products are narrow particle size distribution,
uniform morphology, and enantiomeric purity.
Asthma ranks among the most common
chronic conditions in the United State, affecting
an estimated 14.9 million persons in 1995 and
causing over 1.5 million emergency department
visits, about 500,000 hospitalizations and 5,500
deaths 1 . The goal of our research has been to
study the GAS and RESS processes for the
formation of beclomethasone dipropionate of
desired particle size and shape for inhaled use for
the treatment of asthma. Particles of 2-5 µm in
range are desirable for this application 2, 3 .
Materials and Methods
A high-pressure view-cell was utilized for both
RESS and GAS experiments, as shown in Fig. 1.
RESS Process
In the RESS process, a solute is dissolved in a
high-pressure supercritical fluid (SCF). This
supercritical solution is then expanded through a
nozzle, where instantaneous nucleation and
crystal growth take place at very high
supersaturation. During the expansion, the
solvent density decreases considerably, causing
the solute to be rejected from the solution due to
low solubility at the gas-like solvent density.
2
1
4
3
V-1
V-2
5
6
7
8
11
PT
TT
V-3
10
RESS
9
V-4
GAS
V-5
V-6
1) drug solution
2) CO 2 tank
3 & 4) high-pressure pump
5 & 6) check valve
7) high-pressure view cell
8) magnetic stirrer bar
9) capillary nozzle
10) rapture disc
11) filter
V1 to V3) valve
V4) 3-way plug valve
V5 & V6) control valve.
Figure 1. Schematic diagram of the experimental set up.
IFMBE Proc. 2005;9: 156

Biomedical instrumentation
GAS/SAS Process
In the GAS/SAS process, a high-pressure gas (as
an anti-solvent) is dissolved into the liquid phase
solution, which lowers the equilibrium solubility.
Precipitation of the dissolved compound then
occurs. After the precipitation, the solvent is in
the gaseous phase so that solvent-free and dry
products can be achieved, hence eliminating
extra wash and drying steps. By varying the
process parameters, the particle size distribution
and morphology can be “tuned” to provide
desired material.
Material
Laboratory reagents methanol and ethanol
(Aldrich) and beclomethasone dipropionate
(Steraloids) were used as received.
Results and Discussions
The scanning electron microscopy
(SEM)
photomicrograph of the original beclomethasone
dipropionate sample, Fig. 2, indicates a wide size
distribution of particle sizes and particle
agglomeration.
Figure 3. SEM of Beclomethasone Dipropionate
processed by RESS.
images to be 3.2 µm and a narrow particle size
distribution as seen in Fig. 4. This result is very
promising since the particle size is very close to
the desired 3 µm diameter for asthma treatment,
Zanen and Lammers 2 .
Figure 4. SEM of Beclomethasone Dipropionate,
processed by GAS.
Figure 2. SEM of unprocessed. Beclomethasone
Dipropionate.
Beclomethasone dipropionate particles pre-
dipropionate after processing by GAS process
cipitated by the RESS process were significantly
smaller than the original material. Upon
processing in scCO 2 with 5 wt % methanol, the
beclomethasone dipropionate crystals, Fig. 3,
become rod-like with mean particle size of 400
nm and the maximum aspect ratio was 3
(T=50°C and P=245 bar). This aspect ratio
increased to 10 when the pressure was increased
to 293 bar. Higher pressures gave smaller
diameter particles with higher aspect ratios.
Figures 4 shows SEM image of beclomethasone
(Batch) and the experiment conditions were
38°C and 89 bar. The crystals were found to be
spherical to rod-like (low aspect ratio). The mean
particle diameter was estimated from SEM
Conclusions
• Particles precipitated from the RESS process
are significantly smaller than those from the
GAS process.
• Process conditions strongly influence the
particle size, size distribution and morphology.
• By increasing the expansion pressure in the
RESS process, the observed precipitate
morphology varied from micron size spherical
to rod shape.
• For beclomethasone dipropionate, superior
material size and morphology was produced
for inhalation therapy.
References
1. National Institutes of Health - National Heart, Lung,
and Blood Institute, Data Fact Sheet, Jan. 1999.
2. Zanen, P., Lammers, J-W., Reducing Adverse Effect of
Inhaled Fenoterol through Optimization of the Aerosol
Formulation, J. of Aerosol Med., 12,4, 241-247, 1999.
3. Golriz, M., Rohani, S., Charpentier, P. Particle
Formation of Phenanthrene and Beclomethasone
Dipropionate Using Supercritical Carbon Dioxide, in
Proc. 15 th Int. Symposium on Industrial Crystalization,
Sorrento, Italy, Sep. 2002, pp. 1047-1052.
IFMBE Proc. 2005;9: 157

Biomedical instrumentation
Aerosol Deposition Simulation in an Idealized Mouth
with a Dry-Powder Inhaler Mouthpiece Inlet
E. A. Matida 1 , W. H. Finlay 2 and M. R. Golriz 3
1 Carleton University, Dept. of Mechanical and Aerospace Engr., Ottawa, Canada
2 University of Alberta, Dept. of Mechanical Engr., Edmonton, Canada
3 Umea University, Dept. of Applied Physics and Electronics, Umea, Sweden
edgar.matida@mae.carleton.ca
Abstract
The deposition of monodisperse particles
(d p = 4.05 µm diameter) in an idealized mouth
geometry with a Dry-Powder Inhaler
mouthpiece inlet has been studied numerically.
The primary flow is solved using a Reynolds
Averaged Navier-Stokes shear stress
transport turbulence model at an inhalation
flow rate of 90.0 L/min. The particulate phase
is simulated using a random-walk /
Lagrangian stochastic eddy-interaction model.
Simulated deposition patterns, which have
similar order of magnitude of total deposition
(TD = 91%) when compared to separate
experiments 1 (TD = 67%), show non-uniform
deposition characteristics due to the nonsymmetric
swirling flow coming out of the
DPI mouthpiece.
Introduction
Nebulizers, pressurized metered dose inhalers
(pMDIs) and dry powder inhalers (DPIs) are
devices used to generate medication in the form
of solid or liquid particles 2 , which are often
inhaled through the oral cavity by patients in the
treatment of lung diseases. Although the lung is
the final target, part of the dose will deposit on
the walls of the extrathoracic region (from the
mouth opening to the end of the trachea), giving
losses and departure from the ideal delivery.
Aiming at deagglomeration of particles, DPIs
normally have very complex outlet flows 3
(including swirling flows, grid turbulence, jets
and impinging jets) and relatively small outlet
diameter (up to 10.0 mm), which undesirably
will increase particle deposition in the mouth
cavity 4 .
In the present work, a highly turbulent aerosol
coming from a complex swirling flow from a
DPI mouthpiece is simulated using Reynolds
Averaged Navier-Stokes (RANS) equations
along with a Lagrangian random-walk eddyinteraction
model (EIM) to track individual
particles in the computational domain.
Materials and Methods
Fluid flow simulations using RANS equations in
an idealized mouth geometry (the Alberta
geometry) with a DPI mouthpiece inlet
(Turbuhaler ® , Astra Pharma Inc., Mississauga,
Ontario) are performed using CFXTascflow
(version 2.12, Ansys, Inc.). This idealized mouth
geometry is the same as the one used by DeHaan
and Finlay 4 in particle deposition experiments
on casts (see Fig. 1). The DPI mouthpiece
consists of a double-helical structure (two
internal guide walls) rotating 300 o over 13.5 mm
of length. A structured grid having
approximately 1,000,000 hexahedric elements,
with biased accumulation of nodes towards the
wall (see Fig. 1) is used. Analysis of grid
convergence in similar geometry (concerning
Figure 1: Grid at the midplane and a cross
section of the 3D geometry.
mean velocities, turbulence kinetic energy and
deposition efficiencies) indicates the size of the
grid to be adequate 5 . A modified linear profile
scheme that gives second order error reduction in
most instances, is used in the discretization of
the equations and the fluid flow is solved using
the shear stress transport model 6 . For the inlet
conditions, a steady mass flow rate of 90.0
L/min, a turbulence intensity of 10% of the mean
velocity and a turbulence length scale of 10% of
IFMBE Proc. 2005;9: 158

Biomedical instrumentation
the inlet diameter are used. A zero gauge
pressure condition was applied at the outlet.
10,000 particles with 4.05 µm diameter and
density ρ p = 953 kg/m 3 are released in the
calculated flow domain and particle deposition
positions are recorded. Calculations are
performed on AMD-Opteron workstations.
particles and a steady inhalation flow rate of
90 L/min (giving a ρ p d p 2 Q of approximately
23,430 g µm 2 s -1 ), show an overprediction of
total deposition efficiency, TD = 91%, when
compared to experimental values obtained in
separate measurements 1 , TD = 67%.
Results and Discussions
The magnitude of mean velocities and
corresponding turbulence kinetic energy, k, for
an inhalation flow rate, Q = 90.0 L/min, are
shown in Figures 2 and 3 respectively, at the
midplane of the idealized mouth and two
different cross sections. Note that the inlet
swirling jet flow has a non-symmetrical
preferential path on one side of the mouth (right
Fig. 4. Particle deposition distribution (buildup
mass flow rate, 0 – 3.1 x 10 -8 kg/s)
Conclusions
Fig. 2. Magnitude of mean velocities (0-86 m/s range)
Our present simulation shows the particle
deposition simulation using a relatively simple
RANS model for the primary flow with EIM for
the particulate phase. Simulated deposition
patterns show non-uniform deposition
characteristics due to the non-symmetric swirling
flow coming out of the DPI mouthpiece. It is
conjectured here that the design of inhalation
devices can be significantly improved when both
numerical simulation and particle deposition
experiments are combined.
References
Fig. 3. Turbulence kinetic energy (0-86 m 2 /s 2 range)
side from a patient’s perspective) with localized
and partial impingement of the inlet swirling jet
on the anterior part of the mouth with relatively
high levels of turbulence kinetic energy
distribution.
Figure 4 shows the particle deposition distribution
(buildup mass flow rate) on the surface of
the idealized geometry. Following the swirling
fluid flow patterns, particles deposited more on
the anterior part of the mouth, particularly on the
right side (again from a patient’s perspective).
The total deposition of particles for 4.1 µm
1. Matida, E.A., Rimkus, M. Grgic, B., Lange, C. F.
and Finlay, W. H., A New Add-On Spacer
Design Concept for Dry-Powder Inhalers. J.
Aerosol Sci., 35, 823-833, 2004.
2. Finlay, W. H., The Mechanics of Inhaled
Pharmaceutical Aerosols: An Introduction.
Academic Press, 2001.
3. Borgström, L., On the use of dry powder inhalers
in situations perceived as constrained. J. Aerosol
Med., 14:281–287, 2001.
4. DeHaan, W. H. and Finlay, W. H., Predicting
extrathoracic deposition from dry powder
inhalers. J. Aerosol Sci., 35, 309-331, 2004.
5. Matida, E. A., DeHaan, W. H., Finlay, W. H. and
Lange, C. F., Simulation of Particle Deposition in
an Idealized Mouth with Different Small
Diameter Inlets. Aerosol Sci. & Technol., 37,
924-932, 2003.
6. Menter, F. R., Two-equation eddy-viscosity
turbulence models for engineering applications.
AIAA J., 32(8):1598–1605, 1994.
IFMBE Proc. 2005;9: 159

Biomedical instrumentation
CHANGE OF ARTERIAL PULSE WAVE IN PATIENTS WITH
HYPERLIPIDAEMIA
Irina Hlimonenko 1 , Kalju Meigas 1 , Rein Vahisalu 2
1 Tallinn University of Technology, Biomedical Engineering Center, Tallinn, Estonia,
2 Tallinn Central Hospital, Institute of Lipids, Tallinn, Estonia
Abstract: In this study was investigated the
relationships between mechanical properties of
arteries and arterial pulse wave. Were studied 26
patients, aged 19 to 77 years. Signals were detected
by piezoelectric sensor. Measurements were
performed at two points: radial artery of wrist and
brachial artery of elbow. For signal analysis was
used time marker τ. It was found that τ (radial) has
good correlation with cholesterol concentration. The
preliminary results indicate that it is possible to use
piezoelectric signal for noninvasive indirect
estimation of elastic properties. Cholesterol
concentration exerts influence on the artery
properties and as a result on the pulse wave shape.
Keywords: pulse wave, pulse wave velocity,
cholesterol concentration, arterial compliance,
arterial stiffness, elastic properties.
Introduction
The determination of arterial mechanical properties
using noninvasive methods is of great interest in
medicine. Pulse wave analysis helps to study largeartery
damage, a major contributor to cardiovascular
disease, which is the common cause of mortality and
morbidity in industrialized countries. This high
incidence emphasizes the importance of early evaluation
of the arterial abnormalities, which constitute the
common lesion of major organ damages due to
cardiovascular risk factors [1,2]. Cholesterol is a key in
the development of atherosclerosis, the accumulation of
fatty deposits on the inner lining of arteries. Mainly, as a
result of this, cholesterol increases the risks of
ischaemic heart disease and other vascular diseases.
Several methods can be used to analyze the structure
and function of the arteries. Among the noninvasive
methods of evaluating arteries, pulse wave analysis can
be used as an index of arterial elasticity and stiffness
[3]. After each heart beat a pulse radiates out to the
peripheral circulation. The pulse can be detected noninvasively
using and piezoelectric sensor. Estimation of
arterial mechanical properties using pulse wave was in
our special interest.
measurements piezoelectric sensor (MLT 1010 pulse
transducer, AD Instruments) was used (Fig. 1, block 2).
The special laboratory instrument for signal amplifying
was designed. Piezoelectric signals were recorded
simultaneously for 20 seconds at a sampling rate 500
Hz. A special National Instruments data acquisition
board (DAQ) PCI-MIO-16E-1 (block 2) to digitize the
signals locally and transmit the digital data to the
personal computer is used in this system.
1
Piezosignal
DAQ
2
control
acquired
3
signal
Personal
computer
Figure 1: Block diagram of the equipment
Signals are directed to the DAQ board where acquisition
and analog-to-digital conversion take place. Once data
is acquired and received it is possible to process and
manipulate it using LabVIEW software packet (block
3).
Signal measurements were obtained from 26 subjects (9
male, 17 female); their mean age was 52 (range 19-77).
Measurements were performed in a laboratory. Each
subject was asked to relax and lie supine on a
measurement coach. An operator then attached
piezoelectric sensor to the radial artery at the wrist. It is
important to have a comfortable arm position in order to
keep the hand relatively motionless for a stable and
repeatable recording. The experimental session involved
two sessions: during the first session piezoelectric
sensor was located at the radial artery, during the second
session at the brachial artery.
Analysis
The waveforms were analyzed offline using LabVIEW
programs. For analysis of pulse wave was used time
marker τ, which is time between pulse wave maximum
peak (inflection point) and minimum endpoint of pulse.
τ
Measurement system
The multi-site measurement and analysis system is
shown in Figure 1. For piezoelectric signal
IFMBE Proc. 2005;9: 160

Biomedical instrumentation
Piezosignal
Figure 2: Pulse wave and time marker τ.
The length of the recorded signal was 20 seconds. The
process of transition from one pulse to another was
manual, which allowed editing of any poorly recognized
pulse landmarks. The pulses were smoothed using a
digital low-pass filter (16 Hz) and high-pass filter (8
Hz) to remove the low frequency baseline. The times of
interest were calculated beat-by-beat. The mean values
of τ defined the baseline levels of the measures.
Results
Figure 3 shows how time marker τ changes with
increase of cholesterol concentration.
τ, ms
300
250
200
150
100
τ
50
3,5 4,5 5,5 6,5 7,5 8,5
Cholesterol, mmol/l
Figure 3: τ (radial) as a function of cholesterol
concentration
Discussion
The main objective of this thesis was to find if time
marker τ changes with cholesterol concentration using
pulse wave analysis.
In fact, it is well established that the pulse wave must be
analyzed as a superposition of two separate waves: the
incident traveling wave from heart to periphery, and the
reflected wave traveling from the periphery and the site
of wave reflection to the heart. The incident wave
depends on the left ventricular ejection and the arterial
stiffness, whereas the reflected wave is related to
arterial stiffness and the potential sites of wave
reflection. The inflection point is dividing the pulse
wave into early and mid-to-late systolic parts. This
inflection point is indicating two different waveform
components: a forward and a backward wave. When the
reflected wave is attenuated and/or delayed the shape of
pulse wave might change. This characteristic change in
the shape of the pulse wave is attributed to an increase
in aortic stiffness and pulse wave velocity, with earlier
return of reflected waves from peripheral sites which
influences change in time marker τ. Cholesterol
concentration exerts influence on the artery properties
and as a result on the pulse wave velocity. It causes the
increasing pulse wave velocity and shifting of peak.
In this study, after analyzing the correlation between
cholesterol concentration and τ, it is suggested to
continue estimating of mechanical properties of arteries
using τ marker.
Conclusions
In this research the time τ showed the dependence on
cholesterol concentration in blood. The average τ of
subjects with high cholesterol concentration (over 8
mmol/l) is more than 20% longer than τ of subjects with
normal cholesterol concentration (less than 5 mmol/l).
In this experiment two points of signal measurement
were used (radial and brachial arteries). Comparing
results for radial signals correlation coefficient was 0.56
and for brachial signals it was 0.33. Probably this
difference is due that signals from brachial arteries was
more difficult to get than signal from radial artery which
might be caused by physiological location of arteries.
Further tests in a clinical environment are necessary to
classify various pathological conditions with the
waveform analysis. We suggest that this type of analysis
can provide a simple inexpensive and noninvasive
means for studying changes in the elastic properties of
the vascular system with diseases.
Acknowledgment- This research has been supported by
Estonian Science Foundation, Project No 5888.
References
[1] Asmar R. (1999): “Arterial Stiffness and Pulse
Wave Velocity. Clinical Applications”, Paris,
1999
[2] Izzo J.L., Shykoff B.E. (2001): “Arterial
Stiffness: Clinical Relevance, Mesurement, and
Treatment”, Cardiovascular Medicine, 2, 2001.
[3] Hlimonenko I., Meigas K., Vahisalu R. (2004):
“Pulse wave transit time in patients with
hyperlipidaemia and hypertension”, Medicon
2004, Italy.
[4] X.F. Teng, Y.T. Zhang, “Continuous and
Noninvasive Estimation of Arterial Blood
Pressure Using a Photoplehtysmographic
Approach”, EMBC 2003, pp. 3153-3156.
IFMBE Proc. 2005;9: 161

Biomedical instrumentation
Data processing: A skin impedance measurement
gives the complex impedance at a number of
frequencies. Because of the complexity of the material
studied it would be naïve to just select a simple
equivalent circuit and to pretend that it would represent
known phenomena. Therefore, it is necessary to use all
impedance/frequency variables in the analysis. This can
be done by data reduction methods using projection,
such as PCA. After exclusion of three obvious outliers
the resulting matrix comprised 57 objects and 51
variables.
Results
In fig 2 results from all measurements are displayed and
it is obvious that the spread in data is large in the case of
no pressure control and no soaking of the skin. No
tendency to separation of the various skin types could
be observed in these cases. However when the skin was
soaked with a saline solution and constant pressure was
employed a good reproducibility was obtained. (fig 2C)
Moreover, in this case a tendency to separate different
skin types was observed in the PCA score plot (fig 3).
Fig 3) 2D score plot for impedance. A tendency of separating different
skin types using impedance measurements. Green is inside of foot
ankle, Red is back of hand, Blue is the shoulder, and Yellow is the
back of the calf.
Discussion and conclusions
Skin impedance measurements have the potential to
become an important tool for diagnoses of different skin
conditions such as Malignant Melanoma. A
combination of this technique, NIR spectroscopy, and
digital photography will probably separate the different
skin types even better than impedance measurements
alone. However, to ensure a good reproducibility in skin
impedance data the skin must be soaked with saline
solution and a constant pressure of the probe is
essential. More tests will be done in order to reveal the
impact of coffee drinking, nicotine use, and hirsuteness
on the skin impedance measurements.
Acknowledgement
The European Union Structure Foundation Objective 1
is recognised for their financial support;
References
FRICKE, H. (1925) The electric resistance and capacity
of blood for frequencies between 800 and 4.5
million cycles. Journal of General Physiology,
9, 137-152.
NOREN, L. (2004) Hudimpedansmätare. Institutionen
för systemteknik, avdelningen för
signalbehandling. Luleå, Luleå tekniska
universitet.
NYSTROM, J., LINDHOLM-SETHSON, B.,
STENBERG, L., OLLMAR, S., ERIKSSON,
J. W. & GELADI, P. (2003) Combined nearinfrared
spectroscopy and multifrequency bioimpedance
investigation of skin alterations in
diabetes patients based on multivariate
analyses. Med Biol Eng Comput, 41, 324-9.
SCHWAN, H. P. (1957) Electrical properties of tissue
and cell suspensions. Adv Biol Med Phys, 5,
147-209.
Fig 2 A) 2D score plots for impedance components 1 and 2. Red is
measurements with soaking and pressure. Blue is measurements with
no pressure but soaking and Green is measurements with neither
pressure nor soaking. B) Enlargement of the square in fig A. C)
Enlargement of the square in B
IFMBE Proc. 2005;9: 163

Biomedical instrumentation
DEVICE FOR CONTINUOUS BLOOD PRESSURE MEASUREMENTS
K. Meigas 1 , J. Lass 1 , D. Karai 1 , R. Kattai 1 , J. Kaik 2
1 Tallinn University of Technology, Biomedical Engineering Centre, Tallinn, Estonia
2 Estonian Institute of Cardiology, Tallinn, Estonia
Email: kalju@bmt.cb.ttu.ee
Abstract: This paper is a part of research which is
focused on the development of the convenient device
for continuous non-invasive monitoring of arterial
blood pressure by non-invasive and nonoscillometric
way. The blood pressure estimation
method is based on a presumption that there is a
singular relationship between the pulse wave velocity
in arterial system and blood pressure. The
measurement of pulse wave velocity involves the
registration of two time markers, one of which is
based on ECG R peak detection and another on the
detection of pulse wave in peripheral arteries. Sixtyone
subjects (healthy and hypertensive) were studied
with the bicycle exercise test. As a result of current
study it is shown that with the correct personal
calibration it is possible to estimate the beat to beat
systolic arterial blood pressure during the exercise
with comparable accuracy to conventional
noninvasive methods.
Materials and Methods
A portable experimental blood pressure
monitoring device consists of two analogue signal
acquisition modules, one for ECG and another for PPG
signal. The digital part of the device is based on a
Motorola signal processor (DSP56F803). The frequency
of discretisation for the signals is 500Hz. Pulse wave
velocity is calculated online, both analogue signals and
the measured values are stored on a flash memory card
and can later be reviewed by PC. The device can be
calibrated in order to make online conversion of time
values to systolic arterial pressure. A regular blood
pressure measurement device can be used as a
calibrator. Calibration means the measurement of
arterial blood
Introduction
Blood pressure and cardiovascular pulsation
are fundamental indicators of cardiovascular disease.
The pulse is considered as one of the four most
fundamental medical parameters. Abnormal shape and
rhythm of arterial pulsation are directly connected to
diverse cardiovascular disorders. Small and weak pulses
can be related to heart failure, shock, or aortic stenosis.
Large and bounding pulses can represent hyperkinetic
states, aortic regurgitation, or abnormal rigidity of
arteries. Arrhythmia can lead to a changing amplitude or
irregularity of pulsation. Abnormal amplitude of
pulsation can also be related to the blood pressure.
Potentially useful and convenient parameter for
continuous monitoring of blood pressure could be pulse
wave velocity between different regions of human body.
It has been demonstrated that systolic blood pressure
estimation from this parameter is possible with
acceptable accuracy by personal calibration of the
method for particular patient [1,2]. However, most of
previous studies are focused on utilizing such
measurement on patients in critical conditions, the data
of experiments with healthy subjects is quite limited.
Current paper is an extension of our previous studies
and is focused on development of continuous blood
pressure monitoring device with a simplified
noninvasive calibration [3].
Figure 1: Inside view of an experimental device
pressure during relaxed condition simultaneously with
the time measurements and manual entering measured
blood pressure values into device. The device thereafter
calculates coefficients for linear conversion formula –
slope and intercept. For correct calibration two
measurements of blood pressure is needed in different
levels. The ECG electrodes are placed on wrists and a
regular pulse oximetry finger sensor is used for PPG
registration.
Special software was adapted to the processor.
The software detects R peaks from ECG and starting
front of PPG pulse, thereafter it calculates time between
ECG R-peak and 50% rising front of PPG pulse. The
IFMBE Proc. 2005;9: 164

Biomedical instrumentation
50% location of PPG pulse was chosen because this is
the point in PPG where the signals’ change is the
sharpest and it can more easily be detected compared to
the beginning of the PPG pulse. Our earlier study
showed that there is no remarkable difference in
correlation of the time measured from the foot of the
PPG pulse compared to 50% rising edge measurement
[3].
Sixty-one subjects were included in this study
with the mean age of 42±15 years. The group contained
38 healthy persons (mean age 36±13) and 23 persons
with hypertension (mean age 50±12) and 10 subjects of
the hypertension subgroup incorporated also CAD
(mean age 48±11). To increase the arterial blood
pressure a bicycle test was used. The test consisted of
cycling sessions of increasing workloads during which
the HR changed from 60 to submaximal heart rate. The
workload was increased in steps of 25W after every
third minute until the submaximal heart rate was
achieved, thereafter 5-10 minute recovery period
followed until the heart rate returned back to normal.
The recording was made throughout the exercise and
recovery.
In parallel to time delay measurement the
reference blood pressure values were registered by PC
(auscultatory and Finapres) and synchronized by time
markers with time delay measurements. The computer
later interpolated the auscultatory blood pressure signal
in order to get individual value for every heart cycle.
Results
In Fig. 2 we can compare the dependence of
pulse wave velocity or pulse wave transit time (PWTT)
and systolic blood pressure measured with blood
pressure monitor Finapres 2300. This is recording of
one patient as example. With the help of bicycle test the
systolic blood pressure was increased from 170 mm/Hg
to 230 mm/Hg. The time delay, measured
simultaneously in the same conditions, decreased from
250 ms to 210 ms correspondingly.
Figure 2: Systolic blood pressure (mm/Hg, upper curve)
and PWTT (ms, lower curve). X-axis represents the
number of heart beats.
Dynamics of both changes are comparable. Sudden rise
of systolic blood pressure between 300 and 400 heart
beats and decreasing of time delay in the same time are
both clearly visible.
Discussion
Assuming linear relationship between the
PWTT and systolic pressure at least two calibration
points have to be taken into account and at least
15mmHg difference of systolic pressure between the
calibration points in order to get reliable accuracy (one
measurement for intercept determination and one for
slope determination). In practical purposes this kind of
calibration is quite complicated because a stable
pressure difference should be created between the two
measurements. In order to simplify the calibration
procedure mean slope parameters could be used for
calibration (age and/or health dependent). That way
only one blood pressure measurement made at rest for
determination of the intercept could be used for
calibration. At the same time it is clear that the
simplification of calibration procedure reduces the
precision of the algorithm, which based on our reference
data, was between 5-10%.
Conclusions
The measurement of beat to beat systolic
arterial blood pressure during the exercise is possible
with comparable accuracy to conventional noninvasive
methods. Individual calibration is still necessary.
Acknowledgment- This research has been supported by
Estonian Science Foundation, Project No 5888.
References
[1] N. Lutter, H.G. Engl, F. Fischer, R.D. Bauer, “Noninvasive
continuous blood pressure control by pulse
wave velocity,” Z Kardiol., vol. 85, Suppl. 3, pp.
124-126. 1996.
[2] Y. Sugo, R. Tanaka, T. Soma, H. Kasuya, T. Sasaki,
T. Sekiguchi, H. Hosaka, R. Ochiai, “Comparison of
the relationship between blood pressure and pulse
wave transit times at different sites,” Proc. of the
First Joint BMES/EMBS Conference Serving
Humanity, Advancing Technology, Atlanta, 1999,
CD-ROM.
[3] J.Lass, K.Meigas, R.Kattai, D.Karai, J.Kaik, and
M.Rosmann, “Optical and Electrical Methods for
Pulse Wave Transit Time Measurement and its
Correlation with Arterial Blood Pressure”,
Proceedings of Estonian Academy of Sciences, Vol.
10, pp. 123-136, 2004.
IFMBE Proc. 2005;9: 165

Biomedical instrumentation
LATENCY OF RANDOM SEARCH SACCADES
N. Ramanauskas 1 , G. Daunys 1 , V. Laurutis 1 , V. Vysniauskas 1
1 Department of Electronics, Siauliai University, Siauliai, Lithuania
Abstract
The latency of saccades was investigated. The latency of
reflexive saccades was compared to that of random
search saccades, which may appear, when there was no
conventional target in subject visual field. We suppose,
that similar situation rises, when subject has retina's
visual field deficits. The latency of random search
saccades was significantly greater than latency of
reflexive saccades. This presents an opportunity for using
such method for examination of patients' visual field and
deficits detection.
n.ramanauskas@tf.su.lt
Introduction
Saccades are fast movements of the eyes made to bring
retina foveal region onto a visual target. One of the most
important saccades parameters is a latency. Earlier was
established that saccades latency is various for different
experimental paradigms[1]. Usually saccades are divided
to two groups: reflexive and voluntary[2]. Reflexive
saccades are made in response to a novel peripheral
stimulus. Voluntary saccades are made under a symbolic
cue or instruction. The one of the kinds of voluntary
saccades are the anti-saccades, which are made in the
opposite direction to the peripheral stimulus.
It was established that reflexive saccades have smaller
latency[1]. The smallest latency is obtained during gap
paradigm[3], when peripheral target appears with delay
after central target fades. In opposite anti-saccades show
longer latency[4]. R. Walker et al[2] found, that symbolic
saccades, when initiating for saccade is some symbol, for
example arrow, have biggest latency.
In our study we didn't give any instruction in case of
absence of the peripheral stimulus. So we imitated
situation, when there is a deficit in periphery of subject’s
visual field. Results of investigation would help to build
system for examination of a visual field of a patient[5].
Methods
Six male subjects aged from 21 to 47 years participated
in the experiments. Stimuli were generated on planar
vertical table with red light emitting LEDs. The LEDs are
placed on the table in horizontal, vertical and two
diagonal lines, as is shown in Figure 1.The distance
between the LEDs could be evaluated as 5 degrees of the
subject’s viewing angle, when the distance from the
subject eye to table is 60 cm.
Figure 1: The table with LEDs for stimuli presentation
The eye movement were recorded with
videooculographical eye movement system, which was
developed in our laboratory[6]. The camera’s frame rate
was set to 100 frames per second. The subject eye was
illuminated by infrared light. Eye movement recordings
were done in four separate blocks of 40 trials. In the
beginning of each trial the central LED was switched on
for 1.2 seconds. In the first three blocks simultaneously
with switch off of central LED the diode in periphery was
switched on. In the last block there were 10 trials, when
the switch on of the LED in periphery was missed.
The eye movement recordings analysis was carried out
off line. For saccade detection we used velocity modulus
method. First we differentiated position signal using
smoothing by Savitzky-Golay filter. Filter window was
15 samples and order 5. After, we had calculated velocity
modulus as square root from velocity components
squares' sum. Later we had detected start time of
saccades, when saccade velocity modulus exceeds
velocity threshold (40 deg/s). We calculated the saccade
latency as a time difference between saccade’s start time
and peripheral target onset time (or the time, when
central LED was switched off).
Results
The statistical parameters of saccades’ latency were
calculated. Before statistical processing some results
were rejected. Usually in the block the first saccade had a
larger latency. It signals about lack of subject’s attention
at the beginning of eye movement recording. Also there
were removed values, which distances from median are
bigger as three standard deviations.
IFMBE Proc. 2005;9: 166

Biomedical instrumentation
The mean and standard deviation were calculated for the
each block. The results are presented in Table 1. Here BN
(N=1,2,3,4) denotes a block of trials. For the forth block
the parameters where separately calculated for trials with
a target in periphery (B4T) and without target (B4RS).
Table 1: Latency mean values and standard deviations
Subject B1 B2 B3 B4T B4RS
DB 194±22 187±28 185±17 196±26 319±88
DD 224±22 226±37 220±36 232±42 315±79
GD 223±27 223±29 239±28 197±32 425±96
KM 193±23 180±16 183±16 183±11 323±67
NR 177±24 178±26 189±22 197±32 420 ±82
VV 226±30 217±40 231±38 249±47 452±106
[4] HALLET, P. E., ADAMS W. D. (1980): The
predictability of saccadic latency in a novel oculomotor
task‘, Vision Research, 18, pp. 1279-1296.
[5] LAURUTIS V., KRISCHUNAS K.(1994): ‘Deficits
in visual fields, accomodation, and ocular aligment
determined by eye-movement responses’, in
‘Contemporary Ocular Motor Control and Vestibular
research’, Thieme Medical Publishers, New York, pp.
130-132.
[6] DAUNYS G., RAMANAUSKAS N.(2004): ‘The
accuracy of eye tracking using image processing’, in
Proceedings of the NordiCHI 2004, ACM Press, Finland,
pp. 377-380.
As one could see from results, there is a difference in
saccades’ latency between subjects. Younger subject
have a shorter latency. Exception is subject DD, which is
25 years old, but saccades’ latency is similar as subjects
GD and VV, which are 46 and 47 years old respectively.
There is no significant difference between 1-3 blocks of
one subject. Saccades without target (random search
saccades) had a bigger latency as reflexive saccades. Also
the standard deviation had increased. For some subjects
latency for reflexive saccades also had increased.
Discussion
The mean values of latency of reflexive saccades good
agree with results of the other studies[1,2,3]. The mean
values of random search saccades are biggest among all
types of saccades. Other hand there is correlation
between latencies of reflexive and random search
saccades. For subjects, who have a smaller latency for
reflexive saccades, random search saccades’ latency also
is smaller. So a problem rise, when we need discriminate
random search saccade from reflexive saccade, when the
mean latency of subject saccades is unknown.
As we expected reflexive saccades direction have a small
deviation from target deviation. Of course random search
saccades in most cases have a very big direction
deviation. This help to detect random search saccades
from reflexive. Using eye movement analysis online there
is possibility to repeat targets with the ambiguous results.
References
[1] CLARK, J. J.(1999): ‘Spatial attention and latencies
of saccadic eye movements‘, Vision Research, 39, pp.
585–602.
[2] WALKER R., WALKER, D. G., HUSAIN M.,
KENNARD C.(2000): ‘Control of voluntary and
reflexive saccades’, Experimental Brain research, 130,
pp. 540-544.
[3] COUBARD O., DAUNYS G, KAPOULA Z.(2004):
‘Gap effects on saccade and vergence latency‘,
Experimental Brain Research, 154, pp. 368–381.
IFMBE Proc. 2005;9: 167

Biomedical instrumentation
WIRELESS SYSTEM FOR REAL-TIME RECORDING OF
HEART RATE VARIABILITY
M. Karlsson*, F. Ragnarsson*, N. Östlund* , **, U. Edström*, T. Bäcklund*,
J.S. Karlsson* , ** and U. Wiklund* , **
* Department of Biomedical Engineering & Informatics, University Hospital, and
**Department of Radiation Sciences, Umeå University, Umeå, Sweden
E-Mail: urban.wiklund@vll.se
Abstract: An inexpensive and wearable wireless
measurement system for recording of different
physiological signals has been developed. Although
there are many potential applications for this and
similar systems, in this study we have focused on its
use for real-time analysis of heart rate variability
signals. electrocardiographic (ECG) data are
recorded from several channels, and the system
incorporates a novel adaptive multichannel filter, in
order to achieve a robust detection of individual
heart beats.
Introduction
Modern wireless communication technologies offer
new possibilities for patient monitoring in hospitals, as
well as at home or in outdoor environments. This paper
presents a modular based digital multi-channel wireless
system for recording of physiological signals, such as
ECG, respiration and electromyography (EMG). The
modular-constructed system can be assembled with
respect to different medical applications and patients'
situations. In this paper, we present an assembly of a
system for real-time analysis of heart rate variability
(HRV) signals. The system also incorporates a novel
adaptive multichannel ECG-filter, which has been
developed in a parallel study, where the aim is to
achieve the robust detection of individual heart beats
that is needed in real-time HRV applications [2].
converters are used together with software based digital
signal conditioning, i.e., filtering. In addition, sampling
rate, resolution, and number of channels can be selected
and optimised in order to minimise the amount of data
to be transmitted.
Data acquisition: In this application, the wireless
multichannel data acquisition unit was configured for
recording of two up to six ECG channels at 500 Hz. The
ECG was measured bipolar and the electrodes were
placed on the chest. One channel was used for recording
of respiration with two nasal thermistors and one
thermistor placed over the mouth (Nihon-Kohden).
Multichannel filter: The correct detection of all heart
beats is crucial for the calculation of HRV, and
algorithms for error detection and correction are
therefore essential for the performance of a system for
real-time HRV analysis. Recently, we have incorporated
an adaptive multichannel filter for detection of heart
beats [2]. A spatial and temporal finite impulse response
filter is used to extract the heartbeat events from the
ECG signal. The filter coefficients are adaptively
updated with an independent component analysis (ICA)
algorithm [3], to maximise the super-gaussianity of the
output signal. At the filter output the heartbeat events
occur as distinct peaks and are detected with a threshold
detector.
Methods
System design: The system consists of a data
acquisition unit, with a digital transfer of data to an
ordinary PC [1]. The modular-constructed acquisition
unit (see Fig. 1) consists of, a main module with a
single-chip microprocessor (8051-core), an applicationspecific
signal conditioner module (including A/D
converters), and a digital wireless module (Bluetooth).
All these modules can be exchanged depending on the
specific measurement situation. Thus, the equipment has
flexible design, and can easily be configured for
different medical applications.
To eliminate the need of highly expensive and space
consuming gain and frequency band adjustable
amplifiers, high resolution and oversampling A/D
Figure 1. Data acquisition unit, configured for eightchannel
recording and wireless transfer (Bluetooth) of
data to a host computer.
IFMBE Proc. 2005;9: 168

Biomedical instrumentation
Software for real-time HRV analysis: The real-time
analysis of HRV is performed in the frequency domain
auto-regressive modelling of the beat-to-beat
fluctuations in heart rate (HR), or by utilising Poincaré
graphs, i.e., plots of each R-R interval against the
subsequent value. At present, we have selected Poincaré
graphs as the preferred visualisation of the HRV during
the recording, since the geometric appearance is less
sensitive to spurious errors in HR data than the
estimated power spectrum. However, detection errors,
artefacts and spurious arrhythmic beats can be removed
when data are further analysed on a later occasion.
Results
Figure 2 shows a part of one HRV recording in a
healthy subject. In this picture, the HRV is illustrated as
Poincaré graphs, but it can also be shown in the
frequency domain (Figure 3).
The performance of the adaptive multichannel ECG
filter is demonstrated in a recording from a healthy male
subject (age 32) shown in Figure 4, where frequent
disturbances are present in all six input channels. The
time instants for the spikes in the output signal indicate
that the algorithm was able to suppress the noise.
Figure 4: Example of the performance of the mutichannel
filter for detection of heart beats. In this data
segment, touching and removing electrodes caused
severe additional noise. Six different ECG channels are
shown, and the bottom trace shows the output from the
filter.
Discussion
In this study we have developed a wireless
multichannel system and software for real-time analysis
of HRV signals. The system has a modular design and
can easily be modified for recording of other
physiological signals. The adaptive multichannel ECG
filter is being further developed to achieve a robust
detection and classification of heart beats also in
patients with arrhythmic beats. At present, we are also
incorporating a memory module in the unit, which will
enable the use of the data acquisition unit as a data
logger.
Figure 2. Data acquisition software for real-time HRV
analysis. In this figure the ECG is shown at the top. The
beat-to-beat fluctuations in R-R intervals are shown as
Poincaré graphs for one-minute segments (middle) and
as a function of time (bottom). Total HRV is quantified
by the area of the Poincaré graphs (third curve).
Figure 3. The middle diagrams show the power
spectrum for the last minute, and successive HRV
power spectra for the complete recording .
Acknowledgements
This study was supported by grants from the
Swedish Research Council (number 2003-4833), and
the European Union Regional Development Fund. The
development of the system is performed within the
Centre for Biomedical Engineering and Physics at
Umeå University, Sweden.
References
1. KARLSSON JS, BÄCKLUND T, EDSTRÖM U (2003):
‘A new wireless multi-channel data system for
acquisition and analysis of physiological signals’,
Proc. 17th International Symposium on
Biotelemetry, September, Brisbane, Australia,
September.
2. RAGNARSSON, F, ÖSTLUND N, WIKLUND U (2005):
'Adaptive multichannel filter heart beat detection’,
Proc. 13 th Nordic Baltic Conference, June 14-17,
Umeå, Sweden.
3. HYVÄRINEN, A: ‘The fast ICA package for Matlab',
http://www.cis.hut.fi/projects/ica/fastica/.
IFMBE Proc. 2005;9: 169

Biomedical instrumentation
WAYS TO DECREASE THE ADHESION OF PSEUDOMONAS AERUGINOSA
BACTERIA TO SURFACES OF ENDOTRACHEAL TUBES
M. Ramstedt 1 , H. Mathieu 1
1 Materials Science Institute, École Polytechnique Fédérale de Lausanne (EPFL), Lausanne, Switzerland
madeleine.ramstedt@epfl.ch
Abstract
The objective of the research presented here is to
develop non-adhesive surface coatings on
endotracheal tubes in order to prevent bacterial
growth. Many hospital-acquired pneumonias start
with colonisation of the intubation devices by
Pseudomonas aeruginosa. This could be prevented by
making the surface of the tubes non-adhesive to
bacteria. Furthermore, such tubes would prevent
infection without the use of antibiotics and, thus,
discourage the formation of antibiotic resistance in
bacteria.
The surface of the endotracheal tube can be modified
using plasma treatment and wet chemical treatment.
The hydrophilicity of the surface is increased in this
way which hinders the adhesion of substances in body
fluids that can allow bacteria to anchor. Also, by
including silver ions, the tube surface can be made
toxic to bacteria (but not to humans), which further
decreases bacterial growth. The research presented
develops and examines different surface modifications
of medical grade poly(vinyl chloride) tubes using
surface-sensitive techniques, including X-ray
Photoelectron Spectroscopy, Atomic Force
Microscopy and contact angle measurements, to
obtain chemical information about the surfaces.
Introduction
Hospital patients undergoing mechanical ventilation are
subject to an increased risk of acquiring infections such
as nosocomial (hospital-acquired) pneumonia. A large
portion of these infections have lethal outcome [1,2], and
their prevention is a major concern for the health system
and society.
Approximately 90% of ventilator associated pneumonia
is preceded by colonization of intubation devices by
Pseudomonas aeruginosa3. A serious complication is
that many pneumonia-causing pathogens are antibioticresistant,
which seriously decreases the possibility for
pharmacological treatment. Thus, an attractive option to
reduce these infections is to alter the surface of the
endotracheal tubes in order to decrease the adhesion and
growth of bacteria on the tube walls. In a study by
Triandafillu et al.4, the adhesion of several strains of
Pseudomonas aeruginosa to endotracheal polyvinyl
chloride (PVC) tubes was investigated. It was found that
strains originally isolated from patients with bacteraemia
possessed the highest capacity to adhere to the PVC
material, which indicates that bacterial adhesion to the
endotracheal tubes is an important step leading to
infection. Further studies by Balasz et al.[2], showed that
a drastic decrease in bacterial adhesion can be obtained
by making the tube surface more hydrophilic. The main
reason for this is suggested to be that hydrophilic surfaces
inhibit adhesion of e.g. proteins that function as
anchorage for bacteria. Balasz et al. also showed that
adhesion of bacteria could be prevented by incorporating
silver in the tube surface although this increases the
adhesion of proteins [5]. The toxic effect of silver was
thought to be a result of the presence of Ag + -ions. These
ions have a high affinity for sulfhydryl functional groups
on the bacterial surface and cause cell death by inhibiting
the main energy transfer system of the bacteria. In
mammalian cells these functional groups are not exposed
on the cell membranes [6], and thus monovalent silver is
not toxic to mammalian cells. Sondi and Sondi [7]
showed that the presence of silver particles completely
inhibited bacterial growth on agar plates by causing
“pits” in the bacterial cell walls. However, on mammalian
osteoblast cells no toxic effect could be observed [8].
Furthermore, clinical studies have indicated that the
presence of a silver hydrogel in endotracheal tubes delays
the onset and severity of bacterial colonisation in the
lungs of dogs [9].
In the studies by Balasz et al, the surfaces of
commercially available endotracheal tubes were modified
using plasma polymerisation and wet chemical treatment
(AgNO 3 and NaOH) [3,4,5]. The modified surfaces were
analysed using X-ray Photoelectron Spectroscopy (XPS),
contact angle measurements, Atomic Force Microscopy
(AFM) and Time of Flight Secondary Ion Mass
Spectrometry (Tof-SIMS) to obtain a chemical
description of the new surfaces and to be able to correlate
the chemistry of the surface to the adhesion of bacterial
and proteins. Furthermore the biological response to the
modified polymer surfaces was tested [3,4,5].
In a continued study by Triandafillu [10] it was found
that although the surfaces modified by Balazs et al.
[3,4,5] displayed an initial decreased bacterial adhesion,
the effect decreased with time. In a continuous flow
cultivation system, biofilm was still formed on the
surfaces with increased hydrophilicity and also on the
IFMBE Proc. 2005;9: 170

Biomedical instrumentation
surfaces containing silver. The presence of silver delayed
the formation of the biofilm but bacteria was able to
adhere after a few hours of continuous flow due to
progressive dissolution and disappearance of the silver in
the modified surfaces [10].
Conclusions
The surface modifications obtained so far of endotracheal
tubes are very promising. By including silver in the tube
surface the adhesion of bacteria can initially be reduced
despite the increase in surface hydrophobicity. However,
the methods for obtaining the surface coatings should be
improved with the aim of creating a surface that can resist
bacterial adhesion for long periods of time. Furthermore,
a method should be found that can be used as part of an
industrial process. Thus, research has been initiated to
find an effective long lasting surface modification
method that is also attractive for the suppliers of
endotracheal tubes.
References
[1] McCrory, R. Jones, D. S. Adair, C. G. Gorman, S.P.
(2003), J. Pharm. Pharmacol., 55, pp 411-428
[2] Kollef, M (1999) The New England Journal of
Medicine, 340(8) pp·627-634
[3] Balazs, D. J. Triandafillu, K. Chevolot, Y. Aronsson,
B.-O. Harms, H. Descouts, P. Mathieu, H. J. (2003) Surf.
Interface Anal., 35, pp 301-309
[4] Triandafillu, K. Balazs, D.J. Aronsson, B.-O.
Descouts, P. Quoc, P. Delen, C. Mathieu, H. J. Harms, H.
(2003) Biomaterials, 24, pp 1507-1518
[5] Balazs, D. Triandafillu, K. Wood, P. Chevolot, Y. van
Delden, C. Harms, H. Hollstein, C. Mathieu, H. J. (2004)
Biomaterials, 25, pp 2139-2151
[6] Davies, R. L. Etris, S. F. (1997) Catalysis Today, 36,
pp 107-114
[7] Sondi, I. Salopek-Sondi, B. (2004) J. Colloid Interf.
Sci. 275, pp177–182
[8] Bosetti, M. Massè, A. Tobin, E. Cannas, M. (2002)
Biomaterials, 23 pp 887-892
[9] Olson, M. E. Harmon, B. G. Kollef, M. H. (2002)
Chest, 121, pp 863-870
[10] Triandafillu, K. (2003) PhD Thesis no 2799, École
Polytechnique Fédérale de Lausanne, Switzerland
IFMBE Proc. 2005;9: 171

Biomedical optics
IN VITRO IMAGING OF HUMAN CARTILAGE – CONTRAST
IMPROVEMENT BY OPTICAL WAVELENGTH SELECTION
A. Johansson * , T. Sundqvist ** , J-H. Kuiper *** and P. Å. Öberg *
* Department of Biomedical Engineering, Linköping University, Linköping, Sweden
** Department of Molecular and Clinical Medicine, Division of Medical Microbiology,
Linköping University, Linköping, Sweden
*** Institute of Science and Technology in Medicine, Keele University Medical School,
Stoke-on-Trent, United Kingdom
andjo@imt.liu.se
Abstract: Cartilage thickness is studied in vitro at a
number of sites on human osteoarthritic condyles.
We evaluate a newly developed technique, utilising
the difference in optical absorption spectra between
cartilage and subchondral bone. A comparison between
obtained cartilage thicknesses and reference
measurements at wavelengths 542, 600 and 940 nm is
presented. Furthermore, Monte Carlo simulations
are performed at these wavelengths for investigating
contrasts in cartilage thickness images. Our results
indicate good prospects for cartilage thickness measurement
using the new technique (r = 0.751 at λ =
940 nm) as well as improved contrast in cartilage
imaging by utilising the chosen wavelengths.
Introduction
Osteoarthritis is a very common disorder in the
joints affecting large population groups and having
great social and economical consequences. Degeneration
of the cartilage layer is the most evident consequence
of this disease. Several studies have shown that
damage of the cartilage layer may be hindered or
slowed down at an early stage of degeneration by
medical or surgical interventions [1]. From this perspective,
it is very important to develop instruments and
methods, the use of which can result in a quantified
view of the degeneration of the cartilage layer. In a
recent paper Öberg et al. [2] give a review of the
methods used so far as well as the introduction of a new
optical principle for cartilage thickness measurement.
We believe that optical methods are well suited for
cartilage quality assessment in clinical investigations
because optical components can be minimised today.
Small optical fibres can carry information to and from
the inside of a joint with minimum trauma to the patient.
There are also interesting possibilities within the field of
arthroscopy.
The present work is a study of human cartilage
thickness using optical reflectance spectroscopy, digital
imaging and Monte Carlo studies of a cartilage–bone
model. The aim of this paper is to evaluate and improve
the methodology for cartilage thickness measurements
by in vitro studies of human condyles, removed from
patients undergoing total knee replacement surgery.
Materials and Methods
Tibial plateaus were obtained from nine osteoarthritic
patients undergoing total knee replacement
surgery. The surgical procedures were performed at the
Robert Jones and Agnes Hunt Orthopaedic and District
Hospital, Oswestry, England. The condyles were stored
in physiological saline in a refrigerator before the measurements
took place. Prior to measurement, a reference
grid pattern consisting of 10x10 mm squares was outlined
by drilling small holes in the cartilage surface of
the condyles.
Optical reflection spectra were recorded by using a
fibre optic spectrometer, equipped with a broad spectrum
tungsten lamp. The light was guided by two optical
glass fibres (NA = 0.35) with the emitting and detecting
fibres separated. Measurements were recorded from
each square centre with one spectrum per square taken.
The diffuse reflectance spectrum taken from a white
surface (BaSO 4 ) was used as a reference.
Cartilage thickness was estimated from each optical
spectrum by studying the quotients between important
absorption regions of cartilage (940 nm) and subchondral
bone (542 nm) and a reference wavelength 600 nm
[2]. The condyles were cut and reference cartilage thicknesses
were measured by means of digital photography.
The derived quotients were matched to the reference
cartilage thicknesses according to an exponential regression
model [2].
A Monte Carlo model was developed for calculation
of light propagation in cartilage covered bone (Figure
1). The tissue model consisted of a semi-infinite layer of
blood perfused bone, covered by a 2 mm thick cartilage
layer. In the cartilage, blocks were removed to create
eight regions of varying cartilage thickness. The optical
properties of the tissues have been found in the literature
and are presented elsewhere [2]. Simulations using
this tissue model were performed for the wavelengths
542, 600 and 940 nm. In each simulation, pathways for
5⋅10 6 photons were determined.
IFMBE Proc. 2005;9: 172

Biomedical optics
5 10
45
1.75
2.00
1.50 1.25
1.00 0.75
Inf
2
35
Cartilage
Subchondral
bone
0.50
0.25
0.00
Figure 1. The tissue model used in the Monte Carlo simulations. Eight regions of varying cartilage thickness were
created (thicknesses are given for each region). Incident photons were uniformly distributed over the model surface.
Results
Cartilage thickness was accurately measured by the
new technique at both investigated wavelengths. For λ =
542 nm intensity was reduced with decreasing cartilage
thickness, while the opposite was seen for λ = 940 nm.
The best correlation coefficient was r = 0.751, seen for
λ = 940 nm.
The results of the Monte Carlo simulations are
presented in Figure 2. Brighter regions represent higher
intensities. No apparent effect of cartilage thickness was
noted for the reference wavelength at λ = 600 nm.
Discussion
Cartilage shows a relatively constant absorption for
visible wavelengths and increased absorption in the near
infrared region due to its high water content. In contrast,
subchondral bone absorbs strongly within the visible
range because of its blood content. By the presented
method it is thus possible to measure cartilage thickness,
either by using wavelengths corresponding to haemoglobin
(542 nm) or cartilage absorption (940 nm).
In the material under study there was a natural variation
in cartilage thickness because of osteoarthritis. An
accumulation of blood in the cartilage could underestimate
the thickness but in pure degradation this is not the
case.
Conclusions
Previous studies have showed good results for cartilage
thickness assessment in a bovine material using the
new method [2]. In the present work on a human osteoarthritis
material we demonstrate even better correlations.
By Monte Carlo simulations we show the prospects
of contrast enhancement in joint surface imaging,
e.g. during arthroscopy. In vivo investigations remain to
be performed.
References
[1] ITAY, S. A., ABRAMOVICI, A., ZERO, Z. (1987):
‘Use of cultured embryonal chick epiphyseal chondrocytes
as grafts for defects in chick articular cartilage’,
Clin. Orthop., 220, pp. 284-303
[2] ÖBERG, P. Å, SUNDQVIST, T., JOHANSSON, A.
(2004): ‘Assessment of cartilage thickness utilising reflectance
spectroscopy’, Med. Biol. Eng. Comput., 42, pp.
3-8
542 nm
600 nm
940 nm
PD (norm)
(n = 5000k)
(n = 5000k)
(n = 5000k)
Mean PD (norm)
1.2
1.1
1
0.9
0.8
0.7
0.6
0 1 2
1.2
1.1
1
0.9
0.8
0.7
0.6
0 1 2
Cartilage thickness (mm)
1.2
1.1
1
0.9
0.8
0.7
0.6
0 1 2
Figure 2. Monte Carlo results for the tissue model in Figure 1. Photon densities (PD) are presented for the three
investigated wavelengths.
IFMBE Proc. 2005;9: 173

Biomedical optics
CLEARANCE VARIATIONS MONITORED BY ON-LINE UV-ABSORBANCE
DURING HAEMODIALYSIS
F. Uhlin 1 , I. Fridolin 2 , M. Magnusson 1 , L. Lindberg 3
1 Dept. of Nephrology, University Hospital of Linköping, Linköping, Sweden
2 Centre of Biomedical Engineering, Tallinn Technical University, Tallinn, Estonia
3 Dept. of Biomedical Engineering, University Hospital of Linköping, Linköping, Sweden
fredrik.uhlin@lio.se
Abstract
On-line monitoring of UV-absorbance during dialysis
treatment shows similar response to clearance
variation as the blood-urea and the ionic dialysance
method. This indicates that on-line UV-absorbance
could be used for monitoring deviations in urea
clearance during dialysis. In addition the high
sampling rate allows quick feedback to the operator
after on-line adjustments.
The Kt/V determined by the two dialysate- based
methods were then compared to blood when Kt/V was
calculated from pre- and post- dialysis blood-urea.
Introduction
Several studies indicate that the main criteria for the
adequacy of dialysis, the urea-based dialysis dose (Kt/V),
correlates with the clinical outcome in dialysis patients’
[1]. Monthly control of dialysis dose has been
recommended using blood samples [1]. Monitoring of
haemodialysis (HD) by an on-line system makes it
possible to provide an adequate dialysis dose consistently
given to the HD-patient.
Our group has earlier shown a good correlation between
ultraviolet (UV)-absorbance and several waste solutes
such as urea in the spent dialysate [2] and present the
possibility to estimate urea-Kt/V by UV-absorbance [3].
The aim of this study was to compare the response to a
manipulated clearance reduction, using three methods for
calculations of dialysis dose: UV-absorbance, ionic
dialysance and blood-urea.
Figure 1. Experimental set-up
Assuming that urea is distributed in a single pool (sp)
volume in the body, spKt/V, can be calculated according
to the second-generation Daugirdas formula [4]. Using
the UV-absorbance slope values the Daugirdas based
monocompartmental equation can be written as [3]:
Methods
5 patients on chronically haemodialysis were included in
the study. The patients were studied during 2 consecutive
dialysis treatments each (10 treatments totally) during the
mid-week and last of week treatment. The urea clearance
during the last of week treatment (n=5) was reduced by
an decrease in the preset blood flow. The patients were
monitored on-line with UV-absorbance at the wavelength
of 297 nm, using a double-beam spectrophotometer
(UVIKON 943, Kontron, Italy) connected to the dialysate
outlet of the dialysis machine and with a commercially
available on-line clearance monitor (OCM, Fresenius
medical care, Germany). Fig 1 shows the experimental
set-up.
where Sa is the slope of the natural logarithmic fitting
curve of UV-absorbance, T is the dialysis session length
in minutes, V is the distribution volume of urea in the
body in mL, UF is the total ultrafiltration in kg and W is
the patient’s dry body weight in kg.
IFMBE Proc. 2005;9: 174

Biomedical optics
clearance during dialysis, which also is demonstrated in
Fig. 2. The difference in mean spKt/V between blood and
UV is probably an effect of measuring other solutes than
urea with somewhat lower elimination rate. Also the
number of samples may affect the determination of Kt/V,
e.g. in case of UV 500 samples and in case of blood only
2 samples. The difference in spKt/V between blood and
OCM is probably a result of an overestimated V for
OCM when calculating spKt/V [5].
Figure 2. shows an example of a troublesome dialysis
treatment during 5 hours where UV-absorbance is
plotted against time. The linear fitting curve of the
natural logarithm of the UV-absorbance values gives
Sa, which is inserted in the equation above.
The spKt/V value for OCM was calculated automatically
by the dialysis machine, which was used for comparison.
Results
Figure 2 shows an example of a troublesome dialysis
treatment due to lack of optimal flow in the artery needle.
This results in reduction in clearance. The blood flow at
the dialysis machine was pre-set to 300 ml/min. During
50 to 145 min the artery pressure was below -200 mmHg
(normally -50 to -160 mmHg). At 8, 125, 145 and 205
min needle corrections were made to improve the blood
flow, here seen as a response in clearance (arrows).
Figure 3 shows the response to the clearance reduction in
spKt/V (mean ± SD) for the three methods when
presented separately for the mid-week and the last of
week session. Somewhat different spKt/V values were
obtained. The mean clearance reduction in absolute value
and percentage was 0.26 (19%) for UV-absorbance, 0.26
(20%) in case of OCM and 0.25 (16%) for blood.
Figure 3. shows the mean spKt/V from the three
methods when given separately for the mid-week
(n=5) and the last of week session (n=5). During last of
week session the clearance was decreased due to
reduction in blood flow.
Conclusions
The UV-method shows sensitiveness in the same
magnitude to manipulated clearance reduction compared
to the other methods. Continuos monitoring of UVabsorbance
with high sampling rate gives the opportunity
to verify variations in clearance e.g. due to poor blood
flow as well as give the nursing staff feedback after online
adjustments.
References
[1] NKF K/DOQI guidelines 2000. Guidelines for
hemodialysis adequacy, Guideline
6.http://www.kidney.org/professionals/kdoqi/guidelines_
updates/doqiuphd_ii.html#6
[2] Fridolin I, Magnusson M, Lindberg L-G. (2002)‘Online
monitoring of solutes in dialysate using absorption of
ultraviolet radiation: technique description’. Int J Artif
Organs.,25, pp.748-761
[3] Uhlin F, Fridolin I, Lindberg L-G, Magnusson M.
(2003): ’Estimation of delivered dialysis dose by on-line
monitoring of the ultra violet absorbance in the spent
dialysate.’ Am J Kidney Dis., 41. pp. 1026-1036
[4] Daugirdas JT (1995): ‘Simplified equations for
monitoring Kt/V, PCRn, eKt/V, and ePCRn.’ AdvRen
Replace Ther.,2, pp. 295-304 Ren Replace Ther.,2, pp.
295-304
[5] Goldau R, Kuhlmann U, Samadi N et al (2002): ‘
Ionic dialysance measurements is urea distribution
volume dependent: A new approach to get better results.,
Int J Artif Organs.,26, pp. 321-332
Acknowledgements
The authors wish to thank all dialysis patients who
participated in the experiments, Per Sveider and Jan
Hedblom technical assistance. Ann- Katrin Persson for
assist during the time of data collection. Supported in part
by the Swedish Competence Centre for Non-invasive
Medical Measurements, NIMED. Fridolin's work was
supported by NATO Reintegration Grant EAP.RIG
981201. Reintegration Grant EAP.RIG 981201.
Discussion
All three methods showed a substantial and similar
response to the reduced clearance when spKt/V from two
dialysis days with a normal and reduced clearance were
compared (Fig. 3). This indicates that on-line UVabsorbance
is capable to follow the deviations in urea
IFMBE Proc. 2005;9: 175

Biomedical optics
MEASUREMENTS OF ALA METHYLESTER DIFFUSIVITY IN
NORMAL SKIN IN VIVO: A PILOT STUDY
N. Yavari 1,2 , H.R. Mobini Far 3 , N. Gustavsson 4 , F. Torabi 3 , C. Anderson 5 ,
B. Danielsson3, P.O. Larsson3, K. Svanberg6 and S. Andersson-Engels1
1Department of Physics, Lund Institute of Technology, Box 118, SE-22100 Lund, Sweden
2Department of Physics and Technology, University of Bergen, N-5007 Bergen, Norway
3 Department of Pure and Applied Biochemistry, Box 124, SE- 22100 Lund, Sweden
4 Department of Plant Biochemistry, Box 124, SE-22100 Lund, Sweden
5Department of Dermatology, SE-58185, Linköping, Sweden
6 Department of Oncology, Lund University Hospital, SE- 22100, Sweden
Nazila.Yavari@fysik.lth.se
Abstract: In photodynamic therapy the patient
is given a tumor-selective drug, e.g. Methyl-
Aminolevulinic Acid, which will convert to the
fluorescent compound Protoporhyrin IX in the
cells. By illuminating Protoporhyrin IX with
635 nm, the excitation energy will be
transferred to oxygen molecules, which forms
very reactive singlet oxygen, leading to tumor
destruction. It is important to know the
different processes of the treatment in detail in
order to obtain the best treatment result. In
our work we have identified the diffusivity of
the drug as an important factor, because it
determines the distribution and concentration
of the drug in tissue, and is important for the
efficacy of photodynamic therapy. These
measurements are the first step towards our
next goal, being the construction of a
dosimetry model for the treatment. As the
topically applied substance (an ALA
derivative) is non-fluorescent, no such
measurements have been successful so far. We
have used microdialysis on normal skin, and
traced the drug with both High Performance
Liquid Chromatoghraphy-Flourimetry, and
Liquid Chromatography Mass Spectrometry.
Introduction
5-aminolevulenic acid (5-ALA) is a precursor
compound in biosynthesis of haem which as an
intermediate product has a photosensitizer called
protoporphyrin IX (PpIX). In Photodynamic
Therapy (PDT), PpIX is excited with 635 nm
light, frequently resulting in that a nearby oxygen
molecule is transferred to its excited singlet state
by exchanging energy with the excited PpIX
molecule. Singlet oxygen is a cytotoxic
compound that can destruct and kill tumor cells.
5-ALA can be administrated topically as a cream
in PDT of basal cell carcinoma. The distribution
of the drug (5-ALA) depends on molecular
diffusion properties and tissue vascularity [1].
Because of the poor penetration of drugs into the
skin [2], drug transport properties can limit the
treatment of thick lesions [1]. In PDT the
concentration and distribution of the drug is an
important dosimetric parameter. Since ALA is a
hydrophilic molecule, its penetration through
cellular membrane and into interstitial space of
tissue is low. Ester derivatives of ALA, such as
ALA methylester (M-ALA), are more lipophilic
than the free acid, can also be used as drugs for
topically administration in PDT [3]. As ALA and
its derivatives are non-fluorescent, and also
because the concentration of these molecules in
the skin after topically administration is very low,
no diffusivity measurements of 5-ALA or its
esters have been reported yet. In our project we
administered M-ALA cream (METVIX 20 %,
OSLO, Norway) to normal human skin, and used
microdialysis for sampling its concentration at a
depth of 0.5 mm inside the skin. Microdialysis is
a technique to monitor the chemistry of the extracellular
space in living tissue. The principle of
microdialysis is based on the passive diffusion of
a compound along its concentration gradient. The
samples, collected by microdialysis, were
analyzed using two different methods, High
Performance Liquid Chromatography-
Flourimetry (HPLC-Flourimetry), and Liquid
Chromatography Mass spectrometry (LC-MS). In
this paper we describe how these methods were
employed for analyzing our samples.
Materials and methods
Metvix was applied on the skin of the forearm
skin of a volunteer. A microdialysis device was
used to sample the subcutaneous fluid at certain
time intervals. A flow rate of 0.3 µl/min and a
permeable membrane with 100 KD cut-off were
used for microdialysis sampling. The permeable
membrane was placed at a depth of 0.5 mm in the
skin. Microdialysis fluid was pumped through the
system, allowing M-ALA and other compounds
with molecular weight less than 100 KD to be
transferred to the liquid by diffusing through the
IFMBE Proc. 2005;9: 176

Biomedical optics
membrane. Samples were collected in 500-µl
Ependorf tubes and were kept in -80 ◦ C until
analysing. For measurement by HPLC-
Fluorimetry, M-ALA in the sample was coupled
with a fluorescent molecule 4-(2-
cyanoisoindonyl) phenylisothiocyanate (CIPIC)
[4] and the derivative was injected into the HPLC
system, using a reverse phase column Synergi
Fusion-RP (C18 with polar embedded groups and
low MS bleed), mobile phase containing 0.1 %
formic acid in water. The sensitivity for detection
of M-ALA using MALDI-TOF MS was assayed
using a preparation of pure M-ALA in buffer. All
MALDI-TOF experiments were performed on a
4700 Proteomics Analyzer (Applied Biosystems,
USA). Di-hydroxy benzoic acid (DHB) was used
as the MALDI matrix. The signal at 147 D,
corresponding to M-ALA, overlapped with a
signal from DHB and could thus not be used by
itself for detection of M-ALA. Instead, the signal
at 147 D was selected for fragment ion analysis
with the instrument operated in MS/MS-mode. In
the M-ALA samples, a characteristic fragment
ion signal at 114 D could be detected and was
used for identification of M-ALA. Samples were
directly injected to LC without any derivitization
[5]. Fractions were collected around the retention
time where M-ALA had shown to elute was
collected and subjected to MALDI-TOF MS
analysis. However, M-ALA could not be detected
in these samples, likely due to coelution of some
unknown substances, interfering with the MALDI
sample preparation.
Results and discussion
HPLC with fluorimetric detection was not
sensitive enough to measure all different
concentrations of M-ALA in the samples. The
detection limit for M-ALA was in the region of
100 ng/ml and 300 ng/ml in buffer and sample,
respectively (Figure 1).
Figure 1: HPLC-Fluorimetric for M-ALA in a
fortified sample.
Using MALDI-TOF MS, the detection limit for
M-ALA was lowered to approximately 5 ng/ml
(Figure 2). Adjusting the mobile phase and
LC/MS procedure to obtain the desired retention
and separation are under way.
Figure 2: MS/MS spectra of M-ALA signal at m/z
146.2
Conclusions
This paper presents a study aiming at measuring
M-ALA concentration as a function of time after
topical application at a specific depth (0.5 mm) of
normal skin in vivo using microdialysis. One of
the aims of this study is to understand the
diffusivity of the M-ALA, which would be
necessary for developing an improved dosimetry
model for PDT. For further studies the presented
project will be performed for tape stripped normal
skin.
References
[1] Svaasand L.O., Tromberg B. J., Wyss P.,
Wyss-Desserich M. T., Tadir Y., and Berns M.
W. (1996): 'Light and drug distribution with
topically administered photosensitizers', Lasers in
Medical Science, 11, pp. 261-265
[2] Moser K., Kriwet K., Naik A., Kalia Y.N.,
and Guy R.H. (2001): 'Passive skin penetration
enhancement and its quantification in vitro',
European Journal of Pharmaceutics and
Biopharmaceutics, 52, pp. 103-112
[3] Juzeniene A., Juzenas P., Iani V., and Moan J.
(2002): 'Topical application of 5-aminolevulinic
acid and its methylester, hexylester and octylester
derivatives: Considerations for dosimetry in
mouse skin model', Photochemistry and
Photobiology, 76, pp. 329-334
[4] Lu J., Lau C., Morizono M., Ohta K., and Kai
M. (2001): 'A chemiluminescence reaction
between hydrogen peroxid and acetonitrile and its
applications', Analytical Chemistry, 73, pp. 5979-
83
[5] Felitsyn N.M., Henderson G.N., James M.O.,
Stacpoole P.W. (2004): 'Liquid chromatographytandem
mass spectrometry method for the
simultaneous determination of delta-ALA,
tyrosine and creatinine in biological fluids'.
Clinica Chimica Acta, 350, pp. 219-30.
IFMBE Proc. 2005;9: 177

Biomedical optics
ADVANCED FIBRE-OPTIC SPECTROMETRY TECHNIQUE FOR SKIN
REFLECTANCE AND FLUORESCENCE DIAGNOSTICS
J. Spigulis 1 , L. Gailite 1 , I. Kuzmina 1 , A. Lihachev 1
1 Institute of Atomic Physics and Spectroscopy, University of Latvia, Riga, Latvia
janispi@latnet.lv
Abstract
A portable fibre-optic spectrometry set for complex
skin in-vivo diagnostics has been assembled and
clinically tested. Healthy and pathologic skin areas
were compared from the points of colour and spectral
features. Skin surface and diffuse reflectance, as well
as the blue and green laser excited auto-fluorescence
spectra in the visible and near-infrared ranges were
studied.
Introduction
Skin cancer and other pathologies are diseases with
tendency to grow in many, especially Nordic, countries.
Dermatologists are looking for fast and non-invasive
methods and portable devices for early skin diagnostics
both in stationary and field conditions. Fibre-optic skin
spectrometry may have good potential in this respect.
Data on skin colour, reflectance and fluorescence can be
obtained this way, with further analysis of the specific
parameters and/or markers of the diseases [1]. This work
was aimed at development of methodology for complex
studies of skin pathologies using these three optical
techniques simultaneously.
Methods
Scheme of the assembled fibre-optic spectrometry set is
presented at Fig. 1. The skin spectra are recorded by the
AVANTES BV mini-spectrometer AvaSpec 2048-2 with
two fibre-optic inputs covering the spectral range
250…1100 nm with resolution of about 2 nm. The data
are displayed, stored and processed by means of a laptop
computer. Four light sources are exploited – a stabilised
halogen lamp AVALIGHT-HAL for broadband reflection
spectroscopy and colorimetry, and three sources for
fluorescence excitation – blue and green lasers and a blue
light-emitting diode AVALIGHT-LED-400. The blue
and green lasers are products of B&WTec, Inc. – models
BWB-405-40-pig and BWN-532-20-pig, respectively.
All light sources are supplied with the SMA-type output
couplers for convenient connection to optical fibre cables.
Three design versions of the skin fibre-optic contact
probes have been developed and tested. Version A is a Y-
shaped bundle with flat common end surface comprising
a central detection fibre (connected to the spectrometer)
surrounded by 6 illumination fibres, all with silica core of
200 micron diameter. This probe is used either in direct
contact with skin for the diffuse reflectance
measurements, or can be mounted in a sloped cylindrical
holder (Fig. 1, A) for distant measurements of skin
colour, surface reflectance and/or fluorescence. The fibre
separation distance for this probe does not exceed 1 mm,
so reflectance only from the superficial layers of skin can
be detected in the contact mode To allow deeper
penetration of the probed radiation, a dual-fibre contact
probe with variable inter-fibre distance (2…6 mm) was
constructed (version B). In order to raise quality of skin
fluorescence spectra, a double-sloped probe comprising a
slot for filter absorbing the scattered light at the
laser/LED wavelengths has been designed (version C).
The above-described fibre spectrometry set is compact
and portable – it fits well in a 46x33x16 cm hand-case.
Battery-powered measurements in the field conditions are
possible, as well.
Fig. 1. Scheme of the fibre-optic spectrometry set.
Results
To illustrate the system’s capacity, some data obtained
from healthy and pathologic skin areas of the same
person are compared on Figures 2 - 4. Figure 5 presents
the healthy skin fluorescence bands excited by blue and
green laser radiation.
Shift of the skin colour coordinates in the case of
hemangioma was observed (Fig. 2); this measurement
was taken by the probe option A. Such quantitative
colorimetric data are clinically important, e.g. for skin
recovery monitoring after therapy/surgery.
The diffuse reflectance recordings (Fig. 3) were taken
using the same probe, but in direct contact of the bundle
IFMBE Proc. 2005;9: 178

Biomedical optics
end with the skin surface. If normal skin and melanoma
areas are compared, some spectral differences appear. To
emphasize such differences, several spectra processing
approaches were tested. One possibility is to deal with
ratios of the normalised spectra, e.g. dividing the
pathology spectrum by that of the healthy skin. As
presented on Fig. 4, different pathologies show different
spectral dependences of such intensity ratios; eventually
this might be exploited as a diagnostic marker in future.
Regarding the laser fluorescence studies, only clean
healthy skin has been investigated so far – see Fig. 5.
Some decay of integral fluorescence intensity from the
skin area irradiated by the green laser has been observed
within few minutes. Further studies of fluorescence from
pathological areas of skin are needed to find out if such
decay could be used as a criterion for diagnostic
purposes.
Fig. 4. Intensity ratios of normalized diffuse reflectance
spectra for various skin pathologies (2-fibre probe,
lesion/healthy interface vs healthy skin).
Fig. 2. Colour coordinates of healthy skin and
hemangioma (a – red/green, b – yellow/blue).
Fig. 5. Healthy skin auto-fluorescence bands at 405 nm
and 532 nm laser excitation.
Discussion
Portable multi-functional fibre optic spectrometry
technique seems to have good prospects for fast and noninvasive
early diagnostics of skin pathologies. Further
studies are necessary to optimise diffuse reflectance
processing schemes and to specify the diagnostic
potential of laser-excited skin fluorescence.
Fig. 3. Diffuse reflectance of cancerous and healthy skin
at the red-infrared spectral region.
References
[1] TORICELLI A., IFFERI A., TARONI P.,
GIAMBATISTELLI E., CUBEDDU R. (2001): ' In vivo
optical characterization of human tissues from 610 to
1010 nm by time-resolved reflectance spectroscopy',
Phys. Med. Biol., 46, pp. 2227-2237.
Acknowledgements
Part of the described equipment set was purchased using
the European Regional Development Funds; this
financial support is highly appreciated.
IFMBE Proc. 2005;9: 179

Biomedical optics
DEVELOPMENT OF OPTICALLY FUNCTIONAL FIBER-REINFORCED
COMPOSITE FOR MEDICINE AND DENTISTRY
J. Lehtinen 1 , T. Laurila 1 , T. Reivonen 2 , E. Levänen 2 , T. Mäntylä 2 , L. Lassila 3 , S. Tuusa 3 , P.
Kienanen 3 , P. Vallittu 3 , R. Hernberg 1
1 Institute of Physics, Optics laboratory, Tampere University of Technology, Tampere, Finland
2 Insitute of Materials Science, Tampere University of Technology, Tampere, Finland
3 Institute of Dentistry, University of Turku, Turku, Finland
janne.lehtinen@tut.fi
Abstract
In this work optically functional non-resorbable
composites have been developed and studied for
medicine and dentistry. The objective of this study
was to enhance the photo initiated polymerisation of
dimethacrylate monomer systems by embedding
optical fibers into the resin-fiber structure. First,
fundamental optical properties, i.e., absorption
coefficient and refractive index, of
BISGMA/TEGDMA resin as a function of the
polymerisation time were experimentally determined.
The acquired experimental data was then used as an
input for the optical modelling of the composite
system. Experimental verification of the selected
designs is reported. The application of modern blue
semiconductor lasers to composite curing is also
discussed.
Introduction
Polymer based composites are becoming more and more
important as biomaterials in dentistry and medicine.
Load-bearing capacity for non-resorbable composites can
be obtained by continuous reinforcing fibers. As the size
of the prosthesis or device increases, problems with the
polymerisation of the composites emerge in certain
applications. The photo initiated polymerisation can be
performed ex vivo or in vivo.
The light from a conventional quartz-tungsten-halogen
light-curing unit cannot penetrate deeply into the
composite [1]. In order to ensure proper degree on
monomer conversion, embedding of optical light guides
into the composite has been proposed. By embedding an
optical fiber with proper light guiding and scattering
properties into the composite, it is possible to transfer
light deeper into the composite. Also, by controlling the
light-to-fiber – coupling and by modifying the properties
of the fiber, a more uniform intensity distribution for the
polymerisation can be obtained. At first, this method is
applied to the in situ polymerisation of fiber reinforced
root canal posts for anchoring crowns and reinforcing
remaining tooth structure [2].
Methods
Optical fiber based techniques have many advantageous
properties in medicine: non-contact operation is possible,
optical light guides can be made biologically inert and
immune to several chemical agents. Despite passive
optical applications, also active optical sensing
techniques can be applied to diagnostics and structural
monitoring.
Results
In this work the fundamental optical properties, i.e., the
index of refraction and the absorption coefficient, of
several dental polymer admixtures were first
experimentally determined at the curing wavelength of
470nm. To authors’ knowledge, no such series of
measurements of the fundamental optical properties of
biomedical polymers have been conducted before. The
index of refraction was measured with an ABBE –
refractometer using four different admixtures of
BISGMA/TEGDMA polymer. Each admixture was
measured after various curing times, from 0 to 1800
seconds. The degree of monomer conversion and
Vickers-hardness of the samples were also determined.
Absorption coefficient was determined also as a function
of the curing time. Measurements were made by using a
conventional halogen light source. A monochromator was
used to select the 470 nm wavelength for the
measurements. The absorption coefficient of the polymer
admixtures was determined by illuminating the samples
of different lengths and measuring the transmittance of
each sample.
Discussion
FRED ray tracing software has been used for optical
modelling of the composite system. Parameters for the
optimal light intensity distribution for a 20 mm long
composite system have been studied. Applying the
acquired data from the refraction and absorption
measurements to the optical model, novel fiber-enhanced
curing techniques have been simulated. Potential curing
schemes have been experimentally tested. Preliminary
modelling of the light scattering in optical fiber structures
is also discussed.
IFMBE Proc. 2005;9: 180

Biomedical optics
As far as the light-curing units are concerned, novel
semiconductor light sources, i.e., light emitting diodes
and semiconductor lasers, operating in the blue spectral
range offer many advantages over the conventional
incandescent light sources. Semiconductor devices are
compact and have high efficiency in converting electric
energy into light. The output spectral range of the blue
semiconductor devices match well with the absorption
range of camphorquinone photo initiator broadly used in
dental polymers. In this work the application of novel
high-brightness blue semiconductor lasers to the polymer
curing is also studied.
References
[1] Ranjit D. Pradhan, Noureddine Melikechi, Frederick
Eichmiller, Dental Materials
18
(2002) p. 221-226
[2] L. Lassila, J. Tanner, A.-M. Le Bell, K. Narva, PK.
Vallittu, Dental Materials 20
(2004) p. 29-36
Acknowledgements
This research was financially supported by the National
Technology Agency of Finland (TEKES).
IFMBE Proc. 2005;9: 181

Biomedical optics
LIPID DIFFUSION IN COCHLEAR MEMBRANES BY FRAP
J. Boutet de Monvel*, W.E. Brownell**, and M. Ulfendahl*
* Center for Hearing and Communication Research, Karolinska Institutet, Stockholm, Sweden.
** The Bobby R. Alford Department of Otorhinolaryngology and Communicative Sciences,
Baylor College of Medicine, Houston, USA.
j.boutet.de.monvel@cfh.ki.se
Abstract: We used fluorescence recovery after
photobleaching (FRAP) to investigate the membrane
diffusion properties of several cellular components
from the cochlea of the inner ear. In particular, we
analyzed the two-dimensional pattern of membrane
diffusion on isolated sensory outer hair cells (OHCs),
and found this pattern to be orthotropic, as
suggested by the specialized structure of the OHC
lateral wall. We also performed a series of in situ
FRAP experiments on various cellular membranes
within in vitro temporal bone preparations of the
hearing organ, and found that lipid mobility is
significantly higher in the sensory cells of the organ
of Corti than on its supporting structures.
Introduction
Cochlear outer hair cells (OHCs) display rapid length
changes in response to changes in membrane potential,
a property termed electromotility [2]. This unique
ability is believed to mediate an active feedback process
in the cochlea, responsible for the extreme sensitivity
and frequency selectivity of the mammalian ear [6].
Using fluorescence recovery after photobleaching,
Oghalai et al [4] showed that lipid membrane mobility
in OHCs is also voltage dependent. The OHC lateral
membrane is built upon an orthotropic cytoskeleton,
which gives the cell different mechanical properties in
the longitudinal and the circumferential directions. Here
we used two-dimensional FRAP on isolated OHCs to
see if this orthotropy affects lipid membrane diffusion in
these cells. We also performed a series of FRAP
experiments directly on membranes of the intact hearing
organ, within excised temporal bone preparations in the
guinea pig. On the basis of these experiments, we
present a comparative account of membrane mobility in
some of the main sensory and supporting structures of
the cochlea.
Materials and Methods
Isolation and staining of outer hair cells Outer hair
cells (OHCs) were isolated from the hearing organs of
guinea pigs, and stained with the lipid dye Di-8-
ANEPPS (Molecular Probes). In isolated OHCs, Di-8-
ANEPPS stains specifically the plasma membrane and
do not internalize significantly [3]. All animal
procedures were approved by the Swedish ethics
committee.
Temporal bone preparations Young pigmented guinea
pigs were decapitated, their temporal bones excised and
fixed in a custom chamber containing tissue culture
medium. The middle ear cavity was exposed, and a
small opening was made in the apical turn of the
cochlea for optical access. Another opening in the basal
turn allowed perfusion and application of fluorescent
dyes, namely the stiryl dye RH795, the lipid dye di-8-
ANEPPS, or the membrane marker FM1-43 (all from
Molecular Probes), depending on the experiment.
Confocal microscopy The preparations were examined
with a Zeiss LSM 510 confocal microscope (Carl Zeiss,
Germany), using a water-immersion objective lens
(40X, NA0.8; Zeiss). The dyes were excited with laser
light at 488 nm for di-8-ANEPPS or FM1-43, and at
543 nm for RH795. Detection was performed with a
low-pass filter (505 nm or 560 nm).
FRAP experiments Experiments on isolated OHCs
FRAP [1] was applied by bleaching a small region of
the cell membrane, and monitoring the recovery process
in the focal plane. Two kinds of bleaching
configurations were used. In profile bleach
configuration (Figure 1A) the focal plane was adjusted
vertically to the middle of the cell body. In surface
bleach configuration (Figure 1B), the focal plane was
positioned to coincide approximately with the upper
surface of the cell membrane. In this configuration the
diffusion process is monitored in two dimensions.
Figure 1. Experimental setup for bleaching experiments on
isolated OHCs. A: Profile configuration. B: Surface
configuration. C-F: Different phases of the recovery process
in the surface configuration.
IFMBE Proc. 2005;9: 182

Biomedical optics
In-situ experiments For experiments in temporal bone
preparations, the laser was focused on different
structures of interest, including the sensory hair cells,
the pillar cells, the Hensen cells, and the Reissner
membrane. Most experiments were performed with a
profile configuration, as good contrast was diffucult to
obtain in the surface configuration.
FRAP analysis The recovery process was modelled
with the diffusion equation used by Siggia et al [6],
modified in order to account of a possible anisotropy in
the diffusion. For a given experiment, this equation was
used to simulate the recovery in a selected region of the
images. A standard least-squares fit to the data was then
performed (using the Levenberg-Marquard algorithm)
to estimate the principal diffusion axes and the
corresponding diffusion rates. Processing was done in
Matlab (The MathWorks), using the Expokit package
(Sidje, 1998, http://www.maths.uq.edu.au/expokit) with
custom codes to solve the diffusion equation.
Results
Experiments on isolated outer hair cells Figure 2
illustrates the results obtained for typical bleaching
experiments performed on isolated OHCs stained with
di-8-ANEPPS. The recovery process exhibited a distinct
orthotropic pattern, with diffusion rates 2-4 times higher
along the cell's axial direction than along its
circumferrential direction. Most OHCs stained with di-
8-ANEPPS showed a bright staining localized to the
membrane. The observed mobility was therefore
assumed to be little affected by cytoplasmic diffusion.
For most OHCs the principal diffusion axes were well
matched with the longitudinal and circumferential axes
of the cell. Performing an average over 30 OHCs whose
lengths ranged between 25-82 mm, we found D l = 0.36
± 0.17 µm 2 /s and D c = 0.17 ± 0.10 µm 2 /s for the
longitudinal and circumferential diffusion rates,
respectively. We also measured the axial diffusion
angle, defined as the angle between the axis of fastest
diffusion and the longitudinal axis of the OHC, for 20
cells. This angle was not significantly different from
zero on average (4.4 ± 12°), but its distribution showed
a significant spread, larger than expected from the
estimated error (6.1°) for single measurements.
Figure 2. Analysis of surface bleach experiment on an isolated OHC.
A: The cell just after the bleach, with the diffusion axes superimposed.
Scalebar 10µm. B: Intensity curves averaged over two small regions
around the bleach spot. Superimposed (dashed line) are the results of
the simulation.
Experiments with excised temporal bone preparations
The diffusion rates estimated with all dyes used (di-8-
ANEPPS, RH-795, and FM1-43) showed typical values
in the range 0.1-1 µm 2 /s, comparable to the ones
measured on isolated OHCs [4]. These values are also in
the range found for other cells and generally expected
for biological membranes. For the two dyes RH-795 and
FM1-43 the diffusion rates showed relatively large
dispersions, with no clear differences for different
structures of the organ of Corti. In experiments made
with the di-8-ANEPPS, for which intracellular staining
was less significant, a clear difference was observed
between sensory and supporting cells (Figure 3), lipid
mobility in IHCs (0.46±0.23, n=8) and OHCs
(0.30±0.23, n=9) being 3-20 times faster than in the
Hensen cells (0.023±0.01, n=5) and the pillar cells
(0.09±0.004, n=2) and 3-5 times faster than in the
Reissner membrane (0.11±0.06, n=2). The percentages
of fluorescence recovery measured with di-8-ANNEPS
were also higher in the sensory cells (66±18% and
75±10% for IHCs and OHCs, respectively) than in the
supporting cells (13±6% and 53±1% in the Hensen cells
and in the pillar cells, respectively).
Figure 3. Examples of in situ FRAP experiments with the lipid dye
di-8 ANEPPS. A: Bleach of sensory inner hair cells. B: Bleach of
supporting Hensen cells. The graphs show the recorded intensity
curves in each cases. Note the faster recovery in case A.
Discussion
Our results demonstrate that the orthotropy of the OHC
lateral wall affects the lipid mobility in this membrane.
This could reflect interactions with the dense array of
particles that populate the plasma membrane, which
appear under electron microscopy to follow closely the
actin filaments of the cytoskeleton. The finding that in
situ sensory hair cells have much more fluid membranes
than the supporting cells is also significant. This
underscores the relevance of membrane diffusion as a
pertinent parameter in the study of living cells.
References
[1] Axelrod D., Koppel D.E., Schlessinger J., Elson E., Webb
W.W. 1976. Mobility measurement by analysis of
fluorescence photobleaching recovery kinetics. Biophys J.
16:1055-69.
[2] Brownell, W. E., C.R. Bader, D. Bertrand and Y. de
Ribaupierre. 1985. Evoked mechanical responses of isolated
cochlear outer hair cells. Science, 227: 194-196.
[3] Oghalai J.S., Patel A.A., Nakagawa T., Brownell W.E.
1998. Fluorescence-imaged microdeformation of the outer hair
cell lateral wall. J Neurosci. 18:48-58.
[4] Oghalai J.S., Zhao H.B., Kutz J.W., Brownell W.E. 2000.
Voltage- and tension-dependent lipid mobility in the outer hair
cell plasma membrane. Science 287:658-61.
[5] Santos-Sacchi J. 2003. New tunes from Corti's organ: the
outer hair cell boogie rules. Curr Opin Neurobiol. 13:459-68.
[6] Siggia E.D., Lippincott-Schwartz J., and Bekiranov S.
2000. Diffusion in Inhomogeneous Media: Theory and
Simulations Applied to Whole Cell Photobleach Recovery.
Biophys. J. 79:1761-1770.
IFMBE Proc. 2005;9: 183

Biomedical optics
DIFFUSE REFLECTANCE SPECTROSCOPY OF SKIN PATHOLOGIES
I. Kuzmina 1 , L. Gailite 1 , A. Lihachev 1 , R. Karls 2 , J. Spigulis 1
1 Institute of Atomic Physics and Spectroscopy, University of Latvia, Riga, Latvia
2 Department of Dermatology, Riga Stradinsh University, Riga, Latvia
Abstract
Diffuse reflectance of skin has been studied.
Intensity ratio of lesion to healthy skin spectra were
obtained in order to investigate skin lesion diffuse
reflectance characteristics. Derivation and linear
approximation of intensity ratio spectra is proposed
as a possible processing method. The processed
spectra of malignant and non-malignant pathologies
are compared.
kuzminailona@inbox.lv
Introduction
Most of the cancer diagnostic methods are invasive
and unpleasant for patients; furthermore they need long –
up to few days – processing time. One of the promising
non-invasive diagnostic methods is the diffuse reflectance
spectrometry [1].
Methods
The measurements were taken using a minispectrometer
(Avantes), connected to a halogen light
source via optical fibres. Two kinds of skin contact
probes for measurements on the pathology and on the
border (pathology/healthy skin) were used. The
spectrometer output was connected to a laptop computer.
The “Avantes AvaSoft” software allowed controlling the
measurement parameters and recording the optical
spectra. The recordings of the experimental set-up are
presented in [2].
Skin diffuse reflectance spectra from healthy skin,
different kind of nevus, pigmentation disorders and
melanoma were recorded. The following algorithm of
data processing was applied:
· The noise from each spectrum was subtracted.
· The normalized spectra of pathology or border
(pathology/healthy skin) were divided by the normalized
spectrum of healthy skin.
· The derivative and approximation of acquired
results were computed.
Figure 1: Intensity ratio spectra for Clark and dermal
nevi.
Results
Intensity ratio (pathology/healthy skin) curves of
Clark and dermal nevi border’s spectra (Fig.1) have
different slopes in the 700-930 nm region that depend on
the pigmentation degree.
Figure 2: The role of lesion pigmentation.
The first derivatives of the spectra intensity ratio
(Fig.2) show remarkable differences between Clark and
IFMBE Proc. 2005;9: 184

Biomedical optics
dermal nevi after linear approximation. The 1 st derivative
increases with wavelength increment in the case of Clark
nevi, while it decreases for the dermal nevi. If linear
approximation of the intensity ratio curve (Fig.1) and
then the first derivative are calculated, slope parameters
of these curves are obtained. Table 1 presents slope
parameters for different kinds of nevi with different
pigmentation degrees. The slope parameter correlates
with the value of pigmentation degree: it diminishes with
decrease of pigmentation degree. The highest slope
parameter values are for the hyperpigmented nevi
(Table1).
Table 2: Slope parameter for melanoma and
hyperpigmented nevus.
Table 1: Slope parameter for the border (lesion/healthy
skin) of different kind of nevi.
Discussion
Although the number of inspected patients is not high
enough to draw convincing conclusions, it seems that the
proposed processing method of spectra enables to
distinguish between Clark and dermal nevi and shows
correlation between slope parameter of intensity ratio and
pigmentation degree of disease.
Slope parameter of the hyperpigmented Clark nevus
is very high compared to the other nevi and almost
reaches the slope parameter of melanoma. That can be
explained, since the Clark nevus is a precursor of the
melanoma. It means that Clark nevus can transform to
melanoma with very high probability.
The same correlation is seen in Figure 2, where first
derivative curves are placed higher for nevus with greater
pigmentation degree. It is observed for both Clark and
dermal nevi.
Melanoma, possible melanoma and hyperpigmented
nevus are compared in Table 2. In this case, melanoma
has a higher slope parameter than the hyperpigmented
nevus. Possibly melanoma – diagnose that is not
confirmed yet – has a higher slope parameter than
melanoma. In accordance with our results, this diagnose
should be true.
References
[1] CORDO CHINEA M., SENDRA SENDRA J.R.,
LOPEZ SILVA S.M. and VIERA RAMIREZ A. (2003):
‘Pigmented Skin Lesions by VIS-NIR Diffuse
Reflectance Spectroscopy’, Proc. of SPIE 2003 –
Bioengineered and Bioinspired systems. Maspalomas,
Gran Canaria, Spain, vol. 5119, pp. 157-167.
[2] SPIGULIS J., GAILITE L., KUZMINA I. and
LIHACHEV A. (2005): ‘Advanced Fibre-optic
Spectrometry Technique for Skin Reflectance and
Fluorescence Diagnostics’ (present volume).
Acknowledgements
This work was mainly supported by European Social
Fund.
IFMBE Proc. 2005;9: 185

Biomedical optics
STUDY ON OPTICAL QUALITY OF INTRAOCULAR LENSES
A.F. Shkapa 1 , S.M. Kulikov 1 , L.V. L’vov 1 ,
A.N. Manachinsky 1 , S.A. Sukharev 1 , L.I. Zykov 1
1
Russian Federation Nuclear Centre –VNIIEF,
Institute of Laser Physics Research
607190, Russia, Nizhni Novgorod Region, Sarov
shkapa@otd13.vniief.ru
Abstract:
To date it is adaptivnot unusual in
ophtholmosurgery to perform operations for
exchanging defective eye crystalline lens for a manmade
lens that came to be called an intraocular lens
(IOL). An essential feature of the IOL is its angular
resolution or acuity of vision. In this report, two
techniques are presented to measure the acuity of
vision for the IOL. One method consists in
determination of the vision acuity based on
measurements of IOL aberrations made by a
wavefront detector. In the second technique, the
IOL was placed either inside the cell with planeparallel
windows or in the human eye model. The
acuity of vision was measured by an alphabet and a
dash test board.
of deflection PWR of the nearest approximation sphere
were found by the restored wavefront surfaces. The
angular resolution θ for the IOL equals the root mean
square value for the angles of inclination θ i and
platforms that constitute the wavefront surface. The
acuity of vision V is equal to the inverse value of θ
expressed in terms of angular minutes.
In the second technique the acuity of vision was
estimated by analysis of images of alphabet and dot test
boards produced by the IOLs. In the subsequent
investigations, an IOL was placed in the eye optical
system model. Eye model includes the cornea model
fabricated from polymethylmethacrylate, intraocular
lens, and the vitreous humour (water). CCD camera
captured the test board image which was entered the
computer for analysis that followed.
Materials and Methods
Results
The optical quality of IOLs fabricated from
polymethylmethacrylate was studied. In the first
technique [1] by Hartmann method IOL wavefront
distortion was measured. A detector comprising a
kinoform raster and CCD camera detected the
wavefront. On entering a probe radiation on the
wavefront detector in the absence of IOL of interest at
the acceptance camera an ordered system of focal spots
is formed. On placing the IOL in the model the focal
spots are displaced from the regular grid knots. The
wavefront surface is restored by spot displacement (see
Fig. 1).
The quality of IOLs with 4.5 mm aperture and optical
power in the range from –25 D to +40 D was studied.
In the first technique, the wavefront of the high quality
probe radiation with RMS=0.02 µm on passing through
the IOL being in air was distorted to RMS=0.2-0.3 µm.
The acuity of vision for tested IOLs was determined
based on the results of measurement of the wavefront
surfaces which was V=0.6-1.0 D. Parameters of the
wavefront distortions brought in by the IOL and values
for the vision acuity measured are summarized in
Table 1.
Table 1: Parameters of the wavefront distortions
brought in by the IOL and values for the
vision acuity
IOL
optical -25 -20 -10 0 +10 +20 +40
power, D
RMS, µm 0.27 0.17 0.32 0.02 0.15 0.19 0.21
PWR, µm 0.30 0.21 0.29 0.01 0.14 0.22 0.24
V 1 0.7 0.9 0.6 – 1.0 0.9 0.8
Figure 1: Wavefront plane surface after IOL distortion
The wavefront surface consists of a set of platforms
with different orientation of normal vectors. The root
mean square inclination from the plane RMS and arrow
In the other technique the acuity of vision of the IOLs
under test was in the range of V=0.8-1.2 D. Fig. 2 is an
example of the dash board image formed by the IOL.
IFMBE Proc. 2005;9: 186

Biomedical optics
Figure 2. Example of the dash board image formed
by the IOL
Comparison of the results showed that values for the
acuity of vision by two techniques coincided in the
range of about 20%. Image patterns produced by the
IOLs placed in distilled water were detected. The
values for the acuity of vision found for IOLs placed in
water were 1.5-2 times higher than for IOLs in air.
Experimental measurements of the acuity of vision and
contrast of images for the eye model as a whole were
performed. The investigations showed that the quality
of the eye optical system model was suffice to test
IOLs with the acuity of vision up to V=1.5.
Discussion
The optical quality of IOLs with optical power in the
range from –25 D to +40 D was studied. Two methods
for estimating the IOL acuity of vision: by measuring
the wavefront and visual assessment of the acuity of
vision on the test boards. The measurement of the
acuity of vision estimated via both techniques give
coinciding results with a spread about 20%. Values for
the acuity of vision for the IOLs placed in distilled
water simulating the aqueous humour of the eye
anterior chamber and the vitreous humour are
V in water ≥ 1.5, that by 1.5-2 times higher than that for
the IOLs placed in air - V in air ≤ 1.0.
Conclusion
Analytical techniques for estimating the optical quality
of the IOL based on the measurements of the wavefront
and the estimate of the acuity of vision by the test
boards with the use of the eye experimental model may
be employed in the developing and fabricating novel
IOLs.
The work was financially supported by ISTC within
the framework of Project # 1159.
References
1. M.A.Vorontsov, A.V. Koryabin, V.I. Shmal’gauzen
– Controlled optical systems.- M., “Nauka”, 1988
(Russia).
IFMBE Proc. 2005;9: 187

Biomedical optics
Experimental eye model for intraocular lens tests and demonstrations
L.I.Zykov, S.M. Kulikov, L.V. L’vov, A.N. Manachinski, S.A. Sukharev, A.F. Shkapa,
S.N. Bagrov *
Russian Federation Nuclear Centre - VNIIEF, Institute of Laser Physics Research
607190, Russia, Sarov, Nizhni Novgorod Region, e-mail:zykov@otd13.vniief.ru
*Intersectoral Scientific and Technical Complex “Eye Microsurgery” Research
Experimental Production Co. Ltd. 127550, Russia, Moscow, e-mail:nepmg@mail.ru
Abstract: A full-size experimental model of the
human eye optical system is proposed. An image
from the retina model the function of which is
performed by the CCD camera matrix is sent to the
computer. The image quality was assessed on
placing a monofocal and bifocal intraocular lens
into the model.
Introduction
Nowadays surgeries for exchanging cataract impaired
human eye crystalline lens for a man-made intraocular
lens (IOL) are applied widely [1]. Novel IOL models
[2], for example multifocal ones that compensate
somehow for the loss of the accommodation of eye with
an implanted IOL have been designed. It is important
for the ophthalmologist and of course the patient to see
visually the image pattern and to assess its quality in
the course of deciding on IOL before surgery. In this
work, an experimental eye model and its application for
the assessment of the quality of images formed by
monofocal and bifocal IOLs is described.
Materials and methods
An experimental eye model [3] that almost entirely
simulates the full-size human eye optical system was
employed to test IOLs under conditions close to real
ones. It consists of a model (Fig. 1) of the cornea made
of polymethylmethacrylate, an intraocular lens (IOL) of
interest and a vitreous humour model (water). An image
is captured on the retina model the function of which is
performed by a matrix of the CCD camera. The size of
this image at the retina site is 1.1 × 0.8 mm, the size of
one receptor is 1.5 µm. The quality of the system
optical elements enables to detect an image picture with
angular resolution not worse than 1 angular minute.
The visual field of the experimental model varies in the
range of 1 to 5 degrees on testing an IOL with optical
power of minus 25 to plus 40 diopters. A computeraided
analysis of image allows one to obtain a contrast
transfer function of the pattern of the dot board placed
at the distance of 63 cm away from the eye model.
Results
Comparison of the alphabet patterns obtained on trial
monofocal and bifocal IOLs on long and short
focusing, inclinations and transverse displacement was
performed on the experimental eye model. It was
shown that transverse displacement of a monofocal IOL
+20 D by 2 mm resulted in the degradation of pattern
perception at the acuity of vision equals 1 to 0.8 and
the inclination by 10 degrees slightly affected the acuity
of vision. The IOL with the acuity of vision 1 focused
on distant objects gives low quality image when
passing to viewing at close range with the acuity of
vision not higher than 0.4. A bifocal IOL of acryl with
optical powers of +22 and +25 D is liable to give by
turn image perception with the acuity of vision not
higher than 0.7 when viewing both distant objects and
near ones (see Fig. 2).
1 2 3
4 5
6
0.55
4.5
20.5 – 57.6 mm
Fig. 1. Model of eye optical system. 1 - cornea, 2 - IOL, 3 – distilled water, 4 – exit window, 5 – image location surface,
6 – CCD camera.
IFMBE Proc. 2005;9: 188

Biomedical optics
+22 D +25 D
2.0 mm
6.0 mm
10.5 mm
a) b) c)
Fig. 2. Configuration of lens regions (a) and photos of images at focusing away from the circle (b) and near the
central part (c) of the lens regions
Discussion
Comparison of these photographs shows that the
bifocal IOL enables to view objects at two specified
distances – away from and nearby with satisfactory
quality while the monofocal lens allows to view images
at a fixed distance only. One has to pay for such
simulation of the focal distance in the bifocal IOL with
some loss in image quality.
Conclusions
The designed experimental model is proposed to be
employed on trials novel IOL models (multifocal,
diffraction, etc.), simulations of variation in vision at
possible IOL migration followed implantation and
demonstration patterns to patients when they choose
IOLs for implantation that lies ahead.
The work was financially supported by ISTC within the
framework of Project # 1159.
References
[1] Fyodorov S. N. (1977): ‘Implantation of man-made
crystalline lens’. Moscow, Medicine
[2] Azar D. T. (2001): ‘Intraocular Lenses in Cataract
and Refractive Surgery (Saunders, Philadelphia, Pa.)
[3] Zykov L.I., Kulikov S.M., L’vov L.V. et al.:
‘Human eye optical system model for assessment of
intraocular lens quality’, XXII Congress of the ESCRS,
Paris, September 2004. Book Abstracts, p. 170.
IFMBE Proc. 2005;9: 189

Biomedical optics
OPTICAL COHERENCE TOMOGRAPHY IN HIGH RESOLUTION IMAGING
B. Kudimov 1 , G. Dobre 2 , A. Podoleanu 2
1 Applied Physics, University of Tartu, Tartu, Estonia
2 School of Physical Sciences, University of Kent, Canterbury, UK
Abstract
B.Kudimov@kent.ac.uk
An optical coherence tomography (OCT) method was
used to obtain enhanced in vivo information on biological
tissue. As a non-invasive method of high resolution
imaging, the technique could be important and useful in
biological and clinical applications and diagnostics,
particularly in the investigations of features on a scale of
1-2 mm in depth. The low-coherence apparatus is based
on a fibre-optic Michelson interferometer, with a
superluminescent diode (SLD) used as a broadband light
source.
Introduction
Optical coherence tomography (OCT) is a promising
non-invasive method for in vivo high resolution imaging
in real time. Providing images of biological tissue up to 1
mm in depth with a spatial resolution up to 1 µm, OCT
may have immediate applications in biomedical research,
and especially in the clinic, where the possibility of in
vivo measurements in real time is particularly attractive
for practitioners. The imaging technique uses light
obtained from a superluminescent diode. Near infrared
(IR) non-ionizing radiation of the source with the power
of less than 1 mW makes the method completely
harmless in terms of safety standards for irradiation of
tissues. Depending on type of a scanning system, the
process consists in the recording of en face or crosssectional
images, which allows further processing of the
data and visualisation in different planes.
Methods
In our system the continuous infrared light emitted from a
superluminescent diode (SLD, Superlum Diodes Ltd.) is
coupled into a fiber-optic Michelson interferometer by
means of two 2x2 fiber couplers, where light is split into
a reference beam and a sample beam (Fig. 1). The SLD
has a FWHM bandwidth ∆λ of 34.8 nm centered at
1316.7 nm (λ) (inset, Fig.1).
Figure 1. OCT system scheme
For Gaussian beam profile that corresponds to a
coherence length l C given by [1]:
The above parameters result in a FWHM value l C = 20
µm. The output power of the SLD is set at less than 1
mW. The beam in the sample arm of the interferometer is
focused onto a finger tip. The reference beam propagates
a scanning system mirrors and recombines then with the
backscattered sample beam. The beams interfere only
when the optical path difference is less than or equal to
the coherence length of the light source. The interference
signal is therefore presented with different path lengths
caused by different optical properties of the scanning
tissue.
Results
We have imaged birch leaf as a biological tissue and
shown a selection of different en-face [2] cross-sections
of the volume data obtained from this type of tissu. Two
of the images with different magnification are preseted in
Fig. 2.
IFMBE Proc. 2005;9: 190

Biomedical optics
Figure 2: OCT image of birch leaf with the transversal
resolution of 1 µm
Discussion
In vivo imaging of biological tissue in real time is very
useful in biomedical, biological and especially in clinical
applications. The optical coherence tomography method
is quite a new technique that has high potential for
clinical monitoring of tissues. First, OCT got clinical
usage in ophtalmology [3], but recently it has several
applications in different branches of material science for
non-invasive assessment such as the testing of optical
components [4], non-destructive testing and evaluation of
ceramic and other materials [5]. In combination with
Doppler flowmetry [6, 7] the method is able to provide an
estimation of peripheral blood microcirculation with the
high resolution in terms of both spatial geometry and
velocity. Furthermore, due to the relatively simple
hardware requirements, OCT may be essentially useful
for research applications such as imaging or monitoring.
circulation with color Doppler optical coherence
tomography”, Opt. Lett. 25, No. 19, pp. 1448-1450
[4] Swanson, E. A., Huang, D., Hee, M. R., Fujimoto, J.
G., Lin, C. P., and Puliafito, C. A. (1992): “High-speed
optical coherence domain reflectometry”, Opt. Lett. 17,
No. 2, pp. 151-153
[5] Duncan, M. D., Bashkansky, M., Reintjes, J. (1998):
“Subsurface defect detection in materials using optical
coherence tomography”, Opt. Exp. 2, No. 13, pp. 540-
545
[6] Salerud, E. G., and Öberg, P. Å., (1987): “Single fiber
laser Doppler flowmetry: A method for deep tissue
perfusion measurements”, Med. Biol. Eng. Comput., 25,
pp. 329-334
[7] Kudimov, B., Dobre, G., Podoleanu, A. (2004):
“Fluid-flow velocity measurement with Doppler optical
coherence tomography”, SPIE Proc of AOMD-4 - 4th Int.
Conf. on Adv. Opt. Materials and Devices, Estonia, 2004,
5946
Acknowledgements
The first author acknowledges the financial support from
Archimedes Foundation under contract O.10-04/09 at the
research site in the School of Physical Sciences,
University of Kent, UK.
Conclusions
We have shown how optical coherence tomography can
be used as a tool for non-invasive imaging in an optically
turbid medium such as birch leaf. Scanning the
superficial layers of the biological tissue have shown
high resolution images. To perform skin blood flow
velocity mapping with the OCT system in combination
with Doppler flowmetry further study is necessary.
Among potential applications one could count fingertip
velocity imaging at user-specified depths of the skin,
which could present a very useful tool to help in the study
of blood microcirculation; this is subject of future
investigations.
References
[1] Fercher, A. F. (1996): “Optical coherence
tomography”, J. Biomed. Opt. 1, No. 2, pp. 157-173
[2] Podoleanu, A. G., Dobre, G. M. and Jackson, D. A.
(1998): “En-face Coherence Imaging Using
Galvanometer Scanner Modulation”, Opt. Lett. 23, pp.
147–149
[3] Yazdanfar,S., Rollins, A. M., and Izatt, J. A. (2000):
“Imaging and velocimetry of the human retinal
IFMBE Proc. 2005;9: 191

Biomedical imaging techniques
CENTER FOR MEDICAL IMAGE SCIENCE AND VISUALIZATION (CMIV) -
A UNIQUE CROSS-DISCIPLINARY ENVIRONMENT FOR MEDICAL IMAGE
PROCESSING RESEARCH.
M. Borga 1
1 Department of Biomedical Engineering, Linköping University, Linköping, Sweden
Abstract
Center for Medical Image Science and Visualization
(CMIV) is a multidisciplinary research center in
collaboration between Linköping University, the
County Council of Östergötland and industrial
partners.
CMIV conducts focused front-line research within
multidisciplinary projects providing solutions to
tomorrow’s clinical issues. The mission is to develop
future methods and tools for image analysis and
vizualization for applications within health care and
medical research.CMIV provides unique research
facilities and infra-structure for research in several
medical problem areas. The organisation of CMIV
comprises researchers from medical and technical
faculties. Around 70 researchers and 20 PhD students
are associated to the center.
IFMBE Proc. 2005;9: 192

Biomedical imaging techniques
CORRELATION CONTROLLED BILATERAL FILTERING OF FMRI
DATA
J. Rydell*, H. Knutsson* and M. Borga*
* Dept. of Biomedical Engineering, Linköping University, Linköping, Sweden
Center for Medical Image Science and Visualization
{joary,knutte,magnus}@imt.liu.se
Abstract: In analysis of fMRI data, it is common to
average neighboring voxels in order to obtain robust
estimates of the correlations between voxel timeseries
and the model of the signal expected to be
present in activated regions. This paper presents a
novel method for analysis of fMRI data, which
extends this approach by averaging only neighboring
voxels whose time-series have similar correlation
coefficients. A comparison between the new method
and two other filtering strategies is also presented,
and the novel method is shown to have superior
ability to discriminate between active and inactive
voxels.
Introduction
In fMRI data analysis, highly sensitive detection of
activated voxels is needed. The high noise levels in
typical fMRI data makes this impossible to achieve
without averaging data from several voxels before the
signal detection. The most common approach, applied
in for instance SPM [1], is to employ spatial low-pass
filtering (with a fixed filter kernel) of the data prior to
the detection of active voxels. This is done under the
assumption that in most small neighborhoods, either
almost all or almost no voxels are part of an activated
region. Naturally, this causes blurring of the edges of
activated areas, and depending on the chosen threshold
either causes regions of detected activation to shrink or
grow. Another method, proposed by Friman et al [2],
uses canonical correlation analysis (CCA) to adaptively
find a spatial low-pass filter that maximizes the
similarity between the filtered data and the model of the
blood oxygen level dependent (BOLD) signal. It is,
however, also important to correctly classify inactive
voxels, and if the similarity in each neighborhood is
maximized, this may cause voxels close to the boundary
of an active region to be falsely declared as active.
To alleviate these problems, we propose an edgepreserving
method for adaptive filtering of fMRI data.
The proposed method is similar to bilateral filtering [3],
but instead of image intensity, a correlation measure,
related to mutual information, between individual time
series and the BOLD model is used as distance measure
for the range filters.
Materials and Methods
When ordinary low-pass filtering is used for noise
reduction, voxels that are spatially close to each other
are treated as samples from one distribution, and a
weighted average of the voxels in a neighborhood is
used as an estimate of the true signal value in the center
of that region. The weights are predetermined and based
on the distance from the center of the neighborhood.
Close to edges in an image, the voxel values are actually
samples from two or more distributions, and using
predetermined weights for averaging causes blurring of
the edges. Bilateral filtering extends low-pass filtering
by also considering the distance between the value of a
certain voxel and that of the center voxel, thereby
creating different filter kernels in each neighborhood.
This approach causes voxels from the “other side” of an
edge to be treated as outliers, and thus their effect on the
estimate of the true signal value is reduced or
eliminated. An example of using low-pass filtering and
bilateral filtering, respectively, of a noisy onedimensional
signal is shown in figure 1. The signal is a
step function with additive gaussian noise, and it is
obvious that low-pass filtering causes blurring of the
edge while bilateral filtering preserves it.
b
Figure 1: a) noisy data, b) result of low-pass filtering,
c) result of bilateral filtering
In bilateral filtering, the filter kernel in each neighborhood
can be expressed as a product of two filter kernels:
the domain filter F d and the range filter F r . The domain
filter is based on spatial distance while the range filter is
based on the difference in image intensity. That is,
given an image I(x, y), the bilateral filter kernel F(i, j) at
image coordinates (x, y) can be written
F( i,
j)
= F
d
( i,
j)
⋅ F r ( i,
j)
, where F
d
( i,
j)
is an
a
c
IFMBE Proc. 2005;9: 193

Biomedical imaging techniques
ordinary spatial filter kernel g( i,
j)
and the range filter
is defined as Fr
( i,
j)
= h(
I ( x + i,
y + j)
− I ( x,
y))
.
A common choice of the filter kernels g and h is
gaussian functions.
Godtliebsen et al [4] have proposed using bilateral
filtering of the raw fMRI data, with a time dimension in
addition to the spatial and range dimensions described
above. What we propose here is similar to bilateral
filtering, but instead of creating the range filter from
differences in image intensity, we use the difference in
correlation between the individual voxel timeseries and
the BOLD model. This means that voxels with similar
correlations will be averaged together. Furthermore,
instead of using the correlation coefficient directly, we
use a mapping of the correlation. Under certain
conditions this measure is equivalent to mutual
information. Hence,
Fr
( i,
j)
= h(
M ( x + i,
y + j)
− M ( x,
y))
,where
2
M ( x,
y)
= log(1 /(1 − ρ ( x,
y)
)) and ÿ(x, y) is the
correlation coefficient at coordinates (x, y).
Assuming that the initial correlation estimate is good
enough, in each neighborhood this yields a filter F(i, j)
that averages over voxels that are spatially close and
have similar levels of activation. These filters are then
used to filter the raw data in each timepoint, after which
each voxel in the resulting data is analyzed separately to
detect activation. It is important to notice that this is
different from calculating the correlation in each voxel
and then performing bilateral filtering of the correlation
map.
The BOLD model we suggest is a linear subspace
model, based on principal component analysis of several
plausible BOLD responses generated using Buxton’s
balloon model [5].
Results
The proposed method has been evaluated on both real
and synthetic data. Figures 2c-e show correlation maps
generated from simulated data using fixed low-pass
filtering, adaptive filtering using CCA and adaptive
filtering using the proposed method, respectively. The
areas where BOLD-like signals were embedded in the
noise are shown in figure 2b. In figure 2a, receiver
operating characteristic (ROC) curves, showing the
sensitivity (ability to correctly classify active voxels)
versus the specificity (ability to correctly classify
inactive voxels) of the different methods, are shown.
The signal to noise ratio of the simulated data is
approximately 5 %. Figure 2f shows activation detected
in real data from a finger tapping task, overlaid on an
anatomical image of the brain. The activation in the
motor area is consistent with the task, but since the
ground truth is unknown, it is difficult to use real data to
evaluate a detection method.
Discussion
It is evident from the ROC curves that the presented
method has superior ability to discriminate between
c
e
Figure 2: a) ROC curves, b) locations of embedded activation,
c) activation detected using low-pass filtering, d) activation
detected using CCA-based adaptive filtering, e) activation
detected using the proposed method, f) activation detected in
real data using the proposed method
active and inactive voxels in the simulated data. This is
also supported by the correlation map in figure 2e,
which shows sharper edges between active and inactive
regions than the correlation maps generated by the CCA
method and the method based on a fixed filter. The
ability to preserve edges in the activation map is clearly
an advantage of the proposed method.
Although the method is here presented for twodimensional
filtering, a generalization to three
dimensions should be trivial and is expected to further
improve the results.
References
1. K. J. WORSLEY, K. J. FRISTON (1995): ‘Analysis of fMRI
time-series revisited – again’, NeuroImage, 2(3):173-181
2. O. FRIMAN, M. BORGA, P. LUNDBERG, H. KNUTSSON
(2003): ’Adaptive analysis of fMRI data’, NeuroImage, 19(3):837-845
3. C. TOMASI, R. MANDUCHI (1998): ‘Bilateral filtering for gray
and color images’, IEEE International Conference on Computer
Vision 98, 839-846
4. F. GODTLIEBSEN, C.-K. CHU, S. H. SØRBYE, G. TORHEIM
(2001): ‘An estimator for functional data with application to MRI’,
IEEE Transactions on Medical Imaging, 20(1):36-44
5. R. BUXTON, E. WONG, L. FRANK (1998): ‘Dynamics of blood
flow and oxygenation changes during brain activation: the Balloon
model’, Magnetic Resonance in Medicine, 39(6):855-864
b
d
f
IFMBE Proc. 2005;9: 194

Biomedical imaging techniques
SEGMENTATION OF VELOCITY ENCODED CARDIAC MAGNETIC
RESONANCE IMAGES
E. Bergvall* , **, H. Arheden** and G. Sparr*
* Centre for Mathematical Sciences, Lund Institute of Technology, Lund, Sweden
** Department of Clinical Physiology, Lund University Hospital, Lund, Sweden
erik.bergvall@med.lu.se
Abstract: This paper presents an extension of the
Active Appearance Model framework to segment
velocity encoded magnetic resonance images of the
left ventricle in the human heart. Such images can be
used to calculate myocardial strain. It is shown that
partitioning the model into smaller sub-models for
endo- and epicardium gives better results than a
single model, and that velocity maps should be used
with care in the model building.
Introduction
Segmentation of cardiac images is an important and
challenging task where robustness and reproducibility is
of high concern.
Phase contrast magnetic resonance imaging
(PCMRI) does not only capture time resolved
anatomical information but also the velocity field within
the human body. Such velocity maps may be used to
e.g. calculate myocardial strain [1]. The ability to
automatically segment and analyze such images will
increase the potential for clinical use of PCMRI.
Materials and Methods
We propose an extension to the Active Appearance
Model (AAM) [2] framework to segment the left
ventricle in PCMRI images. AAM image segmentation
is based on constructing a statistical model for shape
and image texture which is then matched to an unknown
input image. Given a training set of N manually
I
1
delineated heart images { } N i i= , where I i is a m× n
matrix describing gray level intensity, we let I ~
i be the
row vectors obtained by row stacking I i . The shape of
the left ventricle is defined as a coordinate vector
x = x , Κ , x , y , Κ , y ) , where x , y ) is the jth
i
( i1 ik i1
ik
( ij ij
coordinate along a sampled parameterized curve
outlining the left ventricle. The images are aligned to
the mean shape using a piecewise affine warp and a
rigid body motion. A statistical model by principal
component analysis (PCA) of the matrix Q where the
ith row is the vector ( xi I ~ N
1
i ) − m , and = ∑( i I ~
m x i ).
N 1
PCA allows all members of the training set to be written
i=
N −
as a linear combination of variation modes { } 1
ϕ by
j j=
1
N
~
( ) ∑ − i Ii
= m +
j=
x c ϕ ,
where cij
are the model parameters.
Matching of the model to an input image is made by
finding the rigid body motion and model parameters that
minimize the difference between the model and the
input image. This optimization problem can be
efficiently solved by estimating a Jacobian matrix from
the training set [2]. Contour overlapping can be
prevented by constraining the model parameters.
The use of PCA will favour global changes in
deformation and intensity over local changes which may
limit the fine tuning of the model matching. It also
ignores the spatial connectivity of the data. To
overcome this drawback we will decompose our model
into smaller ones to increase the ability to adapt the
model locally to an input image.
Decomposition of the model is accomplished by
partitioning the warped images and shapes by a set of
functions { r ) , where r denotes the image
s } M k (
k=
1
M
coordinates, such that ∑ =
s (r)
= 1, for all r . Every
k
member of the training set will contribute to M different
matrices { Q so that the ith row in the kth matrix is
} M k k=
1
component wise multiplicated by ( ( x ) s ( r)
)
s k i k ,
giving us a set of MN basis vectors and model
parameters. Changing a model parameter will now only
affect the model locally. In this application we will use
this general framework to partition the model into two
sub-models consisting of epi- and endocardium.
We will also investigate the use of velocity maps in
addition to anatomy images. Building a model directly
on the velocity maps is not feasible due to the normal
physiological variation which would dominate the
statistical model. A velocity field discontinuity measure
(VDM) is constructed using an edge detection
algorithm. Constructing a model based on edge structure
has been shown to produce better results than ordinary
AAM [3]. Let L (r)
denote the velocity gradient matrix
i
for velocity map i. We construct a scalar image J (r)
by the formula J i ( r)
= det( Li
( r))
. The ith row in the
~
matrix Q will be augmented to Q
i
= ( x
i
I
i
J
i
) − m and
model building will proceed as described above.
1
k
1
1
ij
j
i
IFMBE Proc. 2005;9: 195

Biomedical imaging techniques
A total of 24 PCMRI images were acquired on a
Gyroscan Intera 1.5 T scanner (Philips Medical
Systems) and the shape of the left ventricle in the first
frame in each acquisition was manually delineated.
The performance of 4 different AAM is investigated:
with and without a partitioned model combined with
and without VDM. Performance is evaluated using a
leave-one-out approach. The difference between manual
segmentation and AAM segmentation is measured by
point-to-curve error (PTCE). PTCE is defined as the
mean distance between the landmarks given by the
segmentation and the manually delineated contour.
Several starting guesses were used in each case. The
model’s ability to generalize is also investigated,
measured by the reconstruction error of an input image,
given a fixed number of variation modes.
Results
Table 1 summarizes the segmentation error for the
four different AAM versions. The error distribution is
shown in Figure 1. The partitioned AAM performs
better than the ordinary AAM in terms of segmentation
error. The use of VDM has worsened the performance
of both ordinary and partitioned AAM. Figure 2 shows
the models ability to generalize. The partitioned AAM
outperforms the ordinary AAM ability to reconstruct
unknown input. Finally, Figure 3 shows an example of a
segmented left ventricle.
Discussion
Figure 1: Histogram of segmentation error for ordinary
AAM (solid), ordinary AAM with VDM (dash-dotted),
partitioned AAM (dashed) and partitioned AAM with
VDM (dotted).
Figure 2: Reconstruction error as a function of number
of modes used, for shapes (left) and textures (right) for
ordinary AAM (solid) and partitioned AAM (dashed).
It has been shown that the partitioned AAM
performs better than an ordinary AAM and is also better
when it comes to generalization ability which is
important for segmentation of pathological ventricles.
The result that velocity discontinuity measure worsens
the segmentation accuracy is surprising and suggests
that care must be taken when using such additional
information. Other measures can perhaps be used with
greater success.
Conclusions
This paper presents an extension of the AAM
segmentation framework to segment the left ventricle in
PCMRI images. It is shown that a partitioned AAM
performs better that an ordinary AAM as local variation
is emphasised, and that the use additional information in
form of a velocity discontinuity measure may worsen
the segmentation accuracy.
Table 1: Segmentation error in pixels for the four
different AAM models.
Model type
Mean Std Median
PTCE PTCE PTCE
AAM 2.68 2.47 1.79
Partitioned AAM 2.09 1.82 1.42
AAM with VDM 3.83 3.29 2.24
Partitioned AAM
with VDM
3.84 3.14 2.22
Figure 3: An example of a final segmentation.
References
[1] E. BERGVALL, P. CAIN, G. SPARR, H. ARHEDEN.
(2004): ‘Very Fast and Highly Automated Method
for Myocardial Motion Analysis with Phase Contrast
Magnetic Resonance Imaging’, Proc. SCMR,
Barcelona, 2004, p. 394
[2] COOTES T.F., EDWARDS G.J., TAYLOR C.J. (2001):
‘Active Appearance Models’, IEEE Trans. PAMI,
23, pp. 681-685
[3] I.M. SCOTT, T.F. COOTES, C.J. TAYLOR. (2003):
‘Improving Appearance Model Matching Using
Local Image Structure’, Proc. Information
Processing in Medical Imaging, 2003, pp. 258-269 b
IFMBE Proc. 2005;9: 196

Biomedical imaging techniques
MORPHONS AND BRAINS - NON-RIGID REGISTRATION FOR ATLAS-BASED
SEGMENTATION OF MRI VOLUMES
A. Wrangsjö * and H. Knutsson *
* Dept. Biomedical Engineering and
Center for Medical Image Analysis and Visualization (CMIV), Linköping, Sweden
Abstract: A method for non-rigid registration of one
MRI volume to another is described. The method is
presented as a way to perform atlas-based segmentation
by deformation of a known prototype volume
to an unknown MRI volume.
Introduction
The art of outlining parts of the brain is a field of ever
increasing interest. There are many applications where
such techniques are required. One example is studies of
the hippocampus and how it alters shape and size as a
result of various neurological conditions such as
Alzheimer’s, chronic stress or physical trauma. Today,
such outlining is performed using manual or semiautomatic
segmentation tools. As far as the authors are
aware, no fully automatic method has yet been taken in
use. The aim of this paper is to present a non-rigid
registration method which we believe will make way for
such a method.
Atlas based segmentation
Atlas-based segmentation is a fairly new but increaseingly
popular class of segmentation methods. Here, a
well known atlas object is deformed to fit the observed
image or volume using some registration scheme. Once
a properly deformed atlas has been fitted to the input
data, we can use all the available information on the
atlas object to draw conclusions about the observed
object.
Non-rigid Registration
Registration is the art of fitting one image or volume to
another. This is required e.g. when several images of the
same object have been acquired at different time
instances or using different imaging modalities. The
same techniques can, however, also be applied when
trying to fit images of different objects to each others.
A registration method can be defined by properties:
• Image features
• Similarity metrics
• Deformation model
The first property, image features, refers to what kind of
image structures we are using to find the suitable
deformation. Typical such features are image intensity
level, gradient magnitude and orientation or some other
disparity measure.
andwr@imt.liu.se, knutte@imt.liu.se
The similarity metrics defines how we compare images
based on the features considered. The simplest measure
is using standard correlations, but several more
elaborate schemes based on e.g. Shannon's mutual
information have been presented.
Finally, a model is typically used to define allowed
deformation types. This limits unrealistic deformations
and can also be a convenient and effective tool to
incorporate prior knowledge on what the interesting
object usually looks like and what it can at all look like.
Typical deformation models are rigid or affine ones
(which have very global properties) or FEM, splines or
freeform models (which are more local and allows for
non-rigid, local deformations).
For more on existing registration methods, please refer
to e.g. the references [1-4]. Put together they give a
fairly good overview of this research area.
Our method - the Morphon
The method we present contributes new methods in all
three areas above. Concerning what features to examine,
we have chosen to use quadrature phase differences as a
measure of disparity. This is a method well researched
within our group and others [3, 4]. Among its benefits is
insensitivity to intensity levels and weak gradients. For
MRI, where the intensity levels differ from patient to
patient and even within a single scan, this is a most
desirable property.
The metrics used to measure similarity between images
is based on the fact that the phase difference between
two image features is locally proportional to the spatial
displacement between the images. Our method estimates
locally the optimal deformation that minimises
the phase difference.
The deformation model used here is based on local
regularisation of the estimated deformation fields. By
simply averaging the estimated deformations in the
local region, a more robust measure is obtained. At the
same time, the object is prohibited from performing too
large local deformations. No regularisation at all would
most likely render an object twisted and deformed
beyond recognition, whereas a suitable regularisation
will yield smoothly deformed prototypes.
An estimate certainty measure has also been incorporated
to further increase the robustness of the disparity
measures. Such certainty measures are easy to find
when using quadrature phase. The proposed measure is
IFMBE Proc. 2005;9: 197

Biomedical imaging techniques
based on the amplitude of the quadrature filters. The
larger the amplitude, the more trustworthy the measure
is considered to be.
The deformation process is iterative and gradually
morphs the prototype into looking more and more like
the observed (target) image. Since quadrature phase (as
well as practically all disparity measures) is by nature
local, a scale-space scheme is used. The deformation
accumulation starts at a very coarse scale and gradually
moves towards a finer scale. Also the accumulation
makes use of the estimate certainty measure to increase
the robustness of the method.
Experimental Results
The morphon method was applied to a number of pairs
of MRI brain volumes. Results from these experiments
are shown in the figures below.
Original prototype image before deformation.
Target image.
Deformed prototype.
Overlay of target image and prototype before
deformation.
Overlay of target image and prototype after
deformation.
Squared difference between target and non-deformed
prototype. (ideally black)
Squared difference between target and deformed
prototype. (ideally black)
Discussion
The method presented here is in no way the only
existing non-rigid registration method applied to MRI
images. Due to the use of quadrature phase in estimating
the deformation field, the Morphon does, however,
achieve a lot of robustness comparable to the elaborate
statistical models in [3] and [4]. It also opens for
explicit incorporation of even quite heuristic clinical
rules on how to deform the prototype. A thorough
comparison is desirable. This is, however, beyond the
scope of a two page abstract.
Conclusions and Future
A method to automatically perform non-rigid registration
between different MRI volumes was presented
and experimental results were shown. The results are
somewhat preliminary but nevertheless very pleasing.
The method was presented as the required registration
step in an atlas based segmentation method. It could,
however, also be used in the creation of atlases. By
summation of a number of deformed prototypes, an
atlas image can be computed. By studying the deformation
field statistics, an application tailored deformation
model can be obtained.
Acknowledgements
Many thanks to our research partners lead by Helge
Malmgren at Göteborgs universitet. This project was
funded by the Vetenskapsrådet.
References
[1] PENNEY G. P., WEESE J., LITTLE J. A., DESMEDT P.,
HILL D. L. G., AND HAWKES D. J. (1998): ‘A
comparison of similarity measures for use in 2-d-3-
d medical image registration’, IEEE Transactions
on Medical Imaging, 17:586-595.
[2] PLUIM J. P. W., MAINTZ J. B. A., AND VIERGEVER
M. A. (2003): ‘Mutual-information-based
registration of medical images: A survey’, IEEE
Transactions on Medical Imaging, 22(8):986-1004.
[3] PITIOT A., DELINGETTE H., THOMSON P., M.,
AYACHE N. (2004): ‘Expert knowledge-guided
segmentation system for brain MRI’, NeuroImage,
23: S85-S96.
[4] FISCHL B., SALAT D. H., VAN DER KPUWE A. J. W.,
MAKRIS N., SÉGONNE F., QUINN B. T. (2004):
‘Sequence-independent segmentation of magnetic
resonance images’, NeuroImage, 23: S69-S84.
[5] HEMMENDORFF M, (2001): ‘Motion Estimation and
Compensation in Medical Imaging’ PhD thesis,
Linköping University, Sweden, SE-581 85
Linköping, Sweden, 2001. Dissertation No 703,
ISBN 91-7373-060-2.
[6] FLEET D. J. AND JEPSON A. D., (1990)
‘Computation of Component Image Velocity from
Local Phase Information’, Int. Journal of Computer
Vision, 5(1):77-104.
IFMBE Proc. 2005;9: 198

Biomedical imaging techniques
GENERATION OF PATIENT SPECIFIC BONE MODELS FROM VOLUME
DATA USING MORPHONS
Johanna Pettersson*, Hans Knutsson* and Magnus Borga*
* Department of Biomedical Engineering, Linköping University, Linköping, Sweden
johpe@imt.liu.se
Abstract: The use of simulator systems for surgical
planning and training is growing as the systems
become more advanced. One important feature of
these systems is the possibility to work on real
patient data. This paper presents a method for
generating patient-specific models of the femoral
bone and the pelvis to be used in a hip surgery
simulator. The bones are segmented from
volumetric CT data using the Morphon method [3],
where a prototype pattern is iteratively morphed to
fit the corresponding structure in the input data.
Introduction
Surgical simulator systems are gradually becoming
a more common tool in the training and planning
process in the clinical environment. These simulators
can be used by the surgeons to get a better look at a
specific patient's anatomy and plan surgeries
beforehand. These systems are furthermore good tools
for educating medical students and train new surgeons.
This work presents a technique for automatic
segmentation of hip bones in order to generate patient
specific models for a hip surgery simulator system.
Usually bone tissue is easy to separate from
surrounding soft tissue in CT data. However, due to
the fact that the patients suffer from osteoporosis the
bone density is very low, which implies that basic
segmentation methods based on e.g. thresholding and
morphological operations are insufficient for the
segmentation task. Besides this, the femoral bone and
the pelvic bone are not clearly differentiable, which
makes it very hard to find a segmentation process that
can automatically separate these bones into different
objects. Methods based on e.g. active contours are not
capable of dealing with these topological changes.
Instead the segmentation process is performed using
Morphons. A compact description of how this method
works is found below. For a complete description we
refer the reader to [3].
Materials and Methods
The Morphon method is a non-rigid registration
technique where a prototype pattern is iteratively
morphed to match the target data. Segmentation with
the Morphon method is done in the following steps:
• Design of a prototype suited for the
specific application.
• Iterative deformation of the prototype onto
the target structure in the data. This step
includes estimation of displacement fields,
accumulation of these field and finally,
deformation of the prototype according to
the accumulated field.
• Segmentation of the target structure from
the deformed Morphon model.
The method employs an iterative, multiscale
approach where the comparison between the images is
first done on a very coarse resolution scale. When the
images are aligned at this scale the algorithm moves on
to a finer scale. This continues until the finest resolution
scale and, hence, the most detailed displacements are
reached.
Designing the prototype To be able to segment a
complex structure in a dataset something that gives
information about which pixels/voxels that belong to the
structure is desired. For the Morphon method this
implies that we want to have a prototype that contains a
easily segmented object that describes the general shape
of the structure. By morphing this general object onto
the corresponding object in the input data we obtain a
deformed prototype with the same shape and size of the
target object as in the input (patient-specific) data. Thus,
the target object can be easily segmented from the
morphed prototype data.
Morphing process In the morphing process the
prototype is iteratively deformed to minimize the
difference between the prototype data and the input
data. This process can be separated into displacement
estimation, deformation field accumulation, and
deformation.
For the displacement estimation a technique based
on quadrature phase is used [1, 2]. By applying a set of
quadrature filters, each one associated to a certain
direction, the local phase in the data can be found and
used as a measure of how the prototype data must be
deformed in order to fit the input data. The problem is
formulated as a least square problem according to the
following equation.
v
T
∑ [ w
i
( n
i
v −v
i
)]
min ,
i
2
IFMBE Proc. 2005;9: 199

Biomedical imaging techniques
where v is the sought displacement field estimate, n i is
the direction of filter i, v i is the displacement field
associated to filter direction i, and w i is a certainty
measure, derived from the amplitude of the phase
difference.
The displacement estimation gives us a temporary
displacement estimate that tells us how the deformed
prototype must be moved at this iteration in order to
better fit the input data. To obtain an accumulated
deformation field that tells us how the original
prototype should be transformed to fit the input data,
the updating procedure in the following equation is
applied.
Figure 1: Axial slices.
d
'
a
=
d
a
c
a
+ ( d
c
a
a
+ c
+ d
k
k
) c
k
Figure 2: Coronal slices.
That is, to obtain the accumulated field, d ’ a, the
accumulated displacement field from the previous
iterations, d a , and the temporary displacement field
from the current iteration, d k , are combined. The
updating scheme also includes certainty measures for
the accumulated field, c a , and the temporary field, c k .
Furthermore, a regularisation of the displacement field
is necessary to avoid tearing the prototype apart with a
too divergent field. In this implementation normalised
averaging is applied as a local regularisation scheme
[1].
Finally, the deformation of the prototype is
performed according to the accumulated deformation
field. To accomplish this conventional bi/tri-linear
interpolation is used.
Segmentation Once the prototype data has been
deformed to fit the input data the segmentation can
easily be done from the morphed prototype, as
described in Designing the prototype.
Results
In the hip data case we have used a prototype that
contains a simple description of a femoral bone and
part of the pelvis. This has been generated from a hand
segmentation of an arbitrary CT dataset. In the
prototype the femoral bone and the pelvic bone are
easily separated into two objects, and the bone is easily
separated from the background.
2D slices of the result are shown in the figures
below. Figure 1 shows an axial slice of the prototype
on top of the corresponding slice in the CT data. The
left part of the figure shows the position of the
prototype before the morphing process has started. The
right part of the figure shows the resulting position of
the prototype when it has been deformed to match the
structure in the CT data. Figure 2 shows the same
result for coronal slices. Figure 3 shows isosurface
plots of the input CT data (left), the deformed
prototype (middle) and the original prototype (right).
Figure 3: Isosurface plot of input data (left), deformed
protoype (middle) and original protoype (right).
Discussion
The results indicate that the Morphon algorithm is a
promising method for automatic segmentation of bones
from CT data. The prototype can hold prior information
about the data necessary for handling the segmentation
problems associated to the low bone density and the
lack of a distinct joint space.
This paper has only shown results for bones without
fractures. The intent is, as mentioned in the introduction,
to generate models of bones containing fractures.
In these cases the prototype must incorporate information
such that in knows how to behave in the part of the
bone where the fracture is located.
References
[1] GRANLUND G. H., KNUTSSON H. (1995): ‘Signal
Processing for Computer Vision’, (Kluwer Academic
Publishers, ISBN 0-7923-9530-1).
[2] FLEET D. J., JEPSON A. D. (1990): ‘Computation of
Component Image Velocity from Local Phase
Information’, Int. Journal of Computer Vision, 5(1):77-
104.
[3] KNUTSSON H., ANDERSSON M. (2005): ‘Morphons:
Paint on Priors and Elastic Canvas for Segmentation and
Registration’, In: Proc. of the Scandinavian Conference
on Image Analysis, Joensuu, June 2005.
IFMBE Proc. 2005;9: 200

Biomedical imaging techniques
MEASURING THE MOTION PATTERNS OF THE HEARING ORGAN
USING A THREE-DIMENSIONAL OPTICAL FLOW METHOD
M. von Tiedemann 1 , A. Fridberger 1 , M. Ulfendahl 1 and J. Boutet de Monvel 1
1 Center for Hearing and Communication Research, Karolinska Institutet, Stockholm, Sweden.
miriam.von.tiedemann@cfh.ki.se
Abstract
We used sequences of confocal image z-stacks
acquired inside the cochlea of the inner ear while a
gradually varying pressure gradient was applied
across the cochlear partition, in order to analyse the
organ's three-dimensional motion patterns at the
quasi static level. A 3D optical flow technique was
used, allowing estimation of the apparent
displacement per frame of the structures in the field
of view, at each pixel of high enough contrast. Here,
we briefly present the technique together with
examples of confocal 3D optical flow measurements,
which can be analysed to obtain detailed information
on the three-dimensional cellular movements
occurring inside the cochlea.
Introduction
The detection of sound depends on the vibrations of a
complex array of sensory hair cells within the cochlea
(see Fig.1). The way this array is set into motion
depends on the mechanical couplings of the sensory
cells with the surrounding fluids, membranes, and
supporting cells of the hearing organ. The intricate
three-dimensional organisation of these components is
thought to have important consequences in the way they
interact together.
use the ability of confocal microscopy to acquire threedimensional
stacks of images within intact tissue, in
order to study the 3D-deformations of the hearing organ
in response to quasi-static pressure changes applied
across the cochlear partition. Our analysis uses an
extended optical flow algorithm for estimating the
three-dimensional displacements vectors between
successive stacks. We analyse the displacements of the
sensory hair cell bodies (see Fig. 2) and estimate their
relative components in the longitudinal and radial
directions of the cochlea. We also analyse the deflection
of the stereocilia produced by the pressure changes.
Such measurements will be useful to characterise the
3D-shearing motion that affects the hearing organ
during low-frequency sound stimulation.
Figure 1. Schematic drawing of the cochlear perfusion set-up and a
cross section of the hearing organ, where the sensory hair cells are
situated; one row of inner hair cells (IHC) and three rows of outer hair
cells (OHC). BM, Basilar membrane; DC, Deiter cell; OP, outer pillar
cell; TC, tunnel of Corti; IP, inner pillar cell; TM, tectorial membrane;
HC, Hensen cell; RM, Reissner's membrane.
A lot of unanswered questions still remain concerning
the actual motion patterns of the hearing organ in three
dimensions. Approaches combining conventional or
confocal microscopy with optical flow techniques were
used recently to obtain a more comprehensive picture of
the organ's motion in the plane of the images. Here we
Figure 2. A 3D reconstruction based on a confocal stack acquired
inside in-vitro preparation of the hearing organ. Showing three rows
of sensory hair cells (OHC, outer hair cell) inside the hearing organ,
with their stereocilia sticking out and a glimpse of an inner hair cell
(IHC). Axes orientation; x-axis in the longitudinal direction, pointing
towards the cochlea’s base, y-axis along the radial direction and z-axis
perpendicular to the reticular lamina (pointing upward);
Materials and Methods
Animal preparation and visualisation The preparation
is described in [1] (see Fig.1). Excised temporal bones
of guinea pigs were attached to a holder in a chamber
containing tissue culture medium. The middle ear cavity
was exposed, and a small opening was made in the
apical turn of the cochlea for optical access. Via an
opening in the basal turn of the cochlea the scala
tympani was perfused (from perfusion reservoir) with
oxygenated medium and fluorescent dyes (RH795 and
calcein AM) to label inner ear structures.
The preparation was visualised with a Zeiss LSM 510
confocal microscope, using a 40X NA0.75 objective
lens (Zeiss), equipped with 15nW Krypton/Argon
(488nm) laser and a Helium/Neon (543nm) laser.
IFMBE Proc. 2005;9: 201

Biomedical imaging techniques
Pressure changes Each experiment consisted of
applying a sequence of pressure changes inside the scala
tympani, acquiring for each pressure level a z-stack of
images within the selected volume. The pressures were
altered by moving the reservoir above and below the
fluid level of the perfusion chamber according to the
sequence 0, +10, 0, -10, 0 cm, to form a cycle of
pressure levels. Each position was maintained long
enough for the organ to reach a new equilibrium
position and to acquire a z-stack. Pressure changes in
the cochlea are linearly related to changes in height of
the reservoir within the conditions of the experiments.
Though large, they remain within physiological range,
and may be assumed within a fraction of Pa or less [2].
Optical flow computations A 3D-optical flow algorithm
was applied to the confocal z-stacks to measure the
movement patterns of the organ. This approach is based
on applying a differential constancy constraint equation
[3] to the images, namely:
where I denotes image intensity (∂ t I and denote its
temporal derivative and spatial gradient, respectively)
and v = (v x , v y , v z ) is the unknown displacement vector,
referring to a particular pixel x and a time t. A direct
extension of the multiscale algorithm described in [4]
was used to tackle the aperture problem and recover the
three components of v(x,t). In short, the z-stacks are
filtered with a bank of 3D-wavelet filters of different
sizes and polarisations, thereby producing a set of
filtered stacks, to which the above constraint equation is
applied. In this way an over-determined system of linear
equations for v(x,t) is obtained, which is solved by
least-squares inversion. Prior to the estimation, wavelet
denoising was applied to the images as described in [5],
to reduce noise levels inevitably present in confocal
microscope images.
Results
Fig. 3 shows the motion response of three rows of outer
hair cells (one of the two types of sensory hair cells in
the organ) to the applied pressure changes. The
trajectories of points at the apex and the base of the
cells, together with their difference are plotted. The
axes’ orientation is as depicted in Figure 2.
The difference trajectories between the apexes and the
bases of the outer hair cells represent the bending
trajectories of cells’ bodies. These bending trajectories
were similar in amplitude and shape, but had different
orientations for different rows. The bending of the 1st
row outer hair cell body was nearly radial (contained in
the (y, z)-plane), while an increasing longitudinal (-x)
component was observed in rows 2 and 3.
Similar analysis (not shown) was made for the inner
hair cells. Despite uncertainties in the measurements, a
significant radial shearing of the inner hair cell body
was seen from base to apex, confirming previous inplane
optical flow measurements [1, 4]. A significant
longitudinal component was also observed in the inner
hair cell trajectories.
Figure 3. The upper two graphs show the basal and apical 3Dtrajectories
of the sensory outer hair cells first row. Below are the
bending trajectories (difference between apex and base trajectories) of
the 3 rows of outer hair cell bodies (labelled 1,2 and 3). The different
orientations of these trajectories might be of significance for models
of cochlear mechanics. x, y and z refer to the same axes as in Fig.2.
Discussion
The optical flow method, combined with confocal
microscopy, is a promising tool for getting direct
information on the three-dimensional motion patterns of
the hearing organ. Needed for the future are more
stability in the preparation, together with more accurate
optical flow estimation and more systematic finite
element analysis. In addition, the upgrade to 3D-optic
flow measurements of sound-induced vibration, though
challenging, is feasible. Some interesting observations
have been made using the present set-up. The finding
that outer hair cells from different rows display different
patterns of bending is significant, as the possibility of
differential vibration modes of the outer hair cell bodies
is usually not considered in cochlear models.
References
[1] FRIDBERGER A., BOUTET de MONVEL J. and
ULFENDAHL M. (2002): 'Internal shearing within the
hearing organ evoked by basilar membrane motion', J.
Neurosci., 22, pp. 9850-9857
[2] FRIDBERGER A., van MAARSEVEEN JT., SCARFONE
E., ULFENDAHL M., FLOCK B. and FLOCK Å. (1997):
'Pressure-induced basilar membrane position shifts and the
stimulus-evoked potentials in the low-frequency region of the
guinea pig cochlea', Acta Physiol Scand., 161(2), pp.239-52
[3] BARRON JL., FLEET DL. and BEAUCHEMIN SS.
(1994): 'Performance of optical flow techniques', Int J Comput
Vision, 12, pp. 43-77
[4] FRIDBERGER A., WIDENGREN J. and BOUTET de
MONVEL J. (2004): 'Measuring hearing organ vibration
patterns with confocal microscopy and optical flow', Biophys
J., 86, pp. 535-543
[5] B0UTET de MONVEL J., Le CALVEZ S.
and ULFENDAHL M. (2001): 'Image restoration for confocal
microscopy: improving the limits of deconvolution, with
application to the visualization of the mammalian hearing
organ', Biophys J., 80, pp. 2445-2470
IFMBE Proc. 2005;9: 202

Biomedical imaging techniques
MICROWAVE PROBING OF COMPLEX DIELECTRIC BODIES
P. Norin 1 , T. Gunnarsson 1 , D. Åberg 1 , P. Risman 2
1 Dpt. of Computer Science and Electronics, Mälardalen University, Västerås, Sweden
2 Microtrans AB, Landvetter, Sweden
peder.norin@mdh.se
Abstract
This study on using a finite time domain simulation tool
for the development of microwave imaging enables
the generation of a starting point image for the
reconstruction algorihtms. The simulation result agrees
well with real measurements and shows the dynamic
signal response of a tumor in a breast phantom.
Introduction
The concept of biomedical imaging using microwaves
dates from the 70’s. Microwave imaging is to be seen as
a compliment to other biomedical imaging techniques as
X-RAY, ultrasound and MR imaging. Researchers have
proof of principle [1, 2] and even temperature imaging is
possible [3], however the technique is far from mature.
Theoretically the method should map the three
dimensional difference in dielectric constant to scattered
microwaves. Microwaves have a high attenuation in
tissues with a high content of water such as muscles, liver
and tumors. Tissues with less water content (bones and
lungs) has a lower attenuation, where ultrasound has
imaging problems. Magnitude and phase will differ due
to the shape and materia of the object which will give
arise of scattered fields. The goal of microwave imaging
is to reproduce the scattered fields to reconstruct the
shape and the dielectric properties of the object.
Reconstruction algorithms have been developed [4],
however, there is a lack of reliable data. The input
parameters for these algorithms is the magnitude and
phase of the signal that is scattered from the investigated
object. To get reliable data it is needed to improve the
hardware of the microwave imaging setup. To optimize
the hardware a reliable simulation environment is needed.
This study investigates the finite difference time domain
simulation tool Quickwave 3D, where a breast with a
tumor in different positions is made. The simulations are
compared with measurementson on ambient water to
verify the simulation tool.
Methods and Materials
An octagonal metal container with a radius and height of
75mm and 210mm, respectively, defines the
measurement domain. The waveguided antennas, referred
to as 1, 2 and 3, are positioned as shown in Figure 1.
Measurements and simulations were made on a breast
phantom, defined as a cylinder with a radius of 50mm.
The phantom had a skin with thickness of 2mm (ε=36,
σ=4 S/m) and contained fat tissue (ε=9, σ=0.4 S/m). A
cylinder with a radius of 5mm (ε=50, σ=4 S/m) was
inserted as tumor phantom, off axis, inside the breast
phantom. Filling medium in the container was water,
which has a reasonable matching dielectric constant,
however, with the drawback of high attenuation. This
scenario was made to investigate signal strength
dynamics at a two-dimensional mammographic scan.
Measurements were made on water ambient at room
temperature (295 K), at 0.3 to 3 GHz, as well as on the
phantom with ambient water. Broadband simulations on
the setup was made with the frequency changed in 21
steps. These simulations are compared to measurements
and shows good correlation for the ambient, as shown in
Figure 2.
Figure 1: Setup
Figure 2: Simulation vs. Measurements on ambient water
Results and Discussion
Figure 2 shows broadband simulations on ambient water
compared to measurements, which correlate very well.
The phantom simulations shows differences when
placing the tumor in different positions in the breast.
Figure 3-5 shows the dynamic signal response for
IFMBE Proc. 2005;9: 203

Biomedical imaging techniques
different locations. Every number in Figure 3-5 is the
diffrence between the presents and absents of a tumor at
that location in the breast. The position of antenna 3 lies
on the border of forward- and reflected wave domains.
This makes it extra sensitive to disturbances, where a
disturbance may block the forward wave and/or give rise
to a reflected wave. Minor displacements of the
disturbing body may result in large wave differences. As
seen in Figure 3-5 there is symmetry for the signal path
for different tumor positions. Due to the large differences
in phase and magnitude it is possible to detect tumors in
the breast. In magnitude for S 21 it is hard to distinguish a
tumor due to a small difference, but the phase of S 21
shows that the phase differs in straight path between
antenna 1 and 2. In the outskirt there are small
differences in the phase.
Figure 3: Phase difference in degrees for S 21 @2.5 GHz
Conclusions
The comparison between simulations and measured
experiments showed good accordance for the forward
domain. The antenna lying on the border between
forward and reflected domain showed good accordance,
however very sensitive to surface aberrations.
Nevertheless, the simulation environment has been
proved able to be used for this pursuit. Figure 5 shows
remarkable amplitude differences due to the tumor
presence, up to 18 dB in one point. In some other position
the tumor presence may only be indicated by the phase
information according to Figure 3 and 4. Simulation
results of this kind may be used as a starting point in the
microwave imaging to improve the performance of the
reconstruction algorithm[4].
References
[1] J.Ch. Bolomey. (1994): ’New concepts for microwave
sensing’, Advanced Microwave and millimeter-Wave
Detectors’, San Diego, California, volume 2275,p. 2 -10
[2] N. Joachimowicz, C. Pichot and J.-P. Hugonin.
(1991): ‘Inverse scattering: an iterative numerical method
for electromagnetic imaging’, IEEE Transaction on
Antennas and Propagation, 39(12):pp.1742-1752.
[3] J.M Ruis, C. Pichot, L. Jofre, J.C. Bolomey, A.
Broquetas and M. Ferrando. (1992) ‘Planar and
Cylindrical Active Microwave Temperature
Imaging:Numerical Solution’, IEEE Transactions on
Medical Imaging,11,pp. 457- 469,
[4] J.J Mallorqui, N. Joachimowicz, A. Broquetas and
J.Ch. Bolomey. (1996): ‘Quantitative images of large
biological bodies in microwave tomography by using
numerical and real data’. IEE Electronic letter, 32,pp.
2138 - 2140,1996.
Acknowledgements
The author acknowledges Ian Ray, Saab Metech and
Rohde&Schwarz for providing with help and equipment
for measurements. This study was financed by the
Swedish Knowledge Foundation in cooperation with
Arbexa Industrier.
Figure 4: Phase difference in degrees for S 31 @2.5GHz
Figure 5: Magnitude difference in dB for S 31 @2.5GHz
IFMBE Proc. 2005;9: 204

Biomedical imaging techniques
BRAIN VENOGRAPHY WITH NMR BOLD CONTRAST AT 3T
D. Zacà 2 , S. Casciaro 1,2 , R. Bianco 2 , S. Neglia 2 , E. Casciaro 1,2 , T. Scarabino 3 , A. Distante 1,2
1 Consiglio Nazionale delle Ricerche/Istituto di Fisiologia Clinica, Sezione di Lecce, Lecce, Italy
2 Istituto Scientifico Biomedico Euromediterraneo/Divisione di Bioingegneria, Brindisi, Italy
3 IRCCS Casa sollievo della sofferenza, S. Giovanni Rotondo, Italy
zaca@isbem.it
Abstract: In this work a potentially usable clinical
tool was developed for obtaining high resolution
brain venographies in a very short time. The analysis
was accomplished by Blood Oxygen Level Dependent
(BOLD) MRI with a voxel size of about 1 mm 3 . The
raw data were analysed through signal processing
techniques to obtain signal phase and intensity maps.
After performing phase unwrapping on the phase
map, the informations contained in the intensity and
phase datasets were merged into a new one in which
contrast between veins and surroundings is
enhanced. On this 3D dataset a minimum Intensity
Projection (mIP) algorithm was applied obtaining a
2D projection of the brain venous structures. A
comparison with the results gained with a direct
application of mIP to raw data showed very slight
differences, indicating that the resolution obtained
with the 3 T scan was high enough to render the
post-processing almost superfluous. Focusing
attention on the clinical application of this means, a
great advantage for patients can be pointed out
considering that the time required to obtain their
venography is reduced practically to the sole time
needed for the MR acquisition, as the mIP can be
accomplished in just a few minutes.
Introduction
A good knowledge of the geometrical parameters
that characterize the cerebral vascular and microvascular
system is a fundamental base for the study of
the functionality of the brain and to detect the tissues
and the morphology of the brain itself. This is extremely
important for clinical purposes, as it can be a very
helpful tool for the characterization of pathologies and
malformations of the veins in patients affected by
hypertension, who could suffer from cerebral ischemia,
thrombosis and so on (e.g. [1]). The vascular system can
be traced using imaging techniques, such as Magnetic
Resonance. In this field, there are different methods
which allow to obtain venographic images [2], such as
Time Of Flight (TOF), Phase Contrast Angiography
(PCA) and Blood Oxygen Level Dependent (BOLD)
contrast. In this paper the latter was chosen as it does
not create the image through measures of blood flow but
it exploits instead the different magnetic properties of
venous blood and arterious blood. In this way even the
small vessels, in which the blood flow is slow, can be
detected [3].
Materials and methods
The experiments were carried out on a 25 years old
healthy volunteer. The scanner used was a GE Signa 3T
owned by the Istituto di Ricovero e Cura a Carattere
Scientifico “Casa Sollievo della Sofferenza” in S.
Giovanni Rotondo (FG, Italy).
A 3D scan was taken with gradient-echo sequence
and 60 slices were obtained along the axial direction,
each 1 mm thick and zero distance between two
consecutive slices. Each slice had a Field Of View
(FOV) of 24 x 24 cm 2 and was imaged through a matrix
of 512x256 voxels. The image resolution was of about 1
mm along the sagittal direction and of about 0.5 mm
along the coronal direction, obtaining a voxel size of 0,5
x 1 x 1 mm 3 .
The image was acquired with a NEX equal to 1, Flip
Angle of 25°, TR of 30 ms, TE of 25 ms and Band
Width equal to 31.2 Hz. We obtained 60 anatomic
images of the brain, in terms of both the modulus of the
magnetization vector and its phase.
Two softwares were used for the image processing
phase, AFNI © [4] and MATLAB (MathWorks, MA).
At first, in the modulus images, the values of all the
voxels that did not belong to the brain were set equal to
zero (Fig. 1) as to eliminate the part of the image
representing the braincase. On the phase images a phase
unwrapping [5] procedure was implemented in order to
linearize the phase signal. On the images obtained after
this step a program written with MATLAB in our labs
was applied. This assigned a value equal to zero to the
phase voxels with a positive value, representing the
arteries. In this way, the information left represented
only veins and parenchyma.
On the images obtained, a phase mask filter was
applied to enhance the difference between the venous
and surrounding structure. To do this, the graylevels of
the voxels are altered accordingly to a function that
linearizes the pixel negative values assigning values
from 0 to 1 and sets to a value of 1 the voxels with
starting positive values [2].
IFMBE Proc. 2005;9: 205

Biomedical imaging techniques
phase unwrapping procedure. It also shows some loss of
information on the air-tissue interface.
Discussion
Fig. 1: Images before (left) and after (right) the brain
skull removal.
The phase values were then multiplied by the
respective modulus values obtaining a map in which
phase and modulus informations were merged (Fig. 2
left), allowing a better enhancement of the veins.
Finally, a minimum Intensity Projection (mIP) onto the
central slice of the chosen brain slab was performed,
obtaining a 2D projection of the brain venous system of
17 slices 1-mm thick (Fig. 2, left).
It is immediate that the differences between the two
venographies compared in Fig. 3 are so small that the
images obtained without processing can be used with no
limitations compared to the processed ones with a large
saving in time. Other studies have compared scans at
1.5 T with ones at 3 T, obtaining a supralinear
dependence of the Contrast to Noise Ratio from the
field strength, with a 25% shorter echo time of the 3 T
scan [6], thus confirming the higher efficiency in terms
of time and resolution of the latter.
The time factor can be further underlined
considering that an estimation of the total duration of
the scan and the post-processing could result in about
one hour. This means that the patient can profit by a
great reduction of the time necessary for his/her test.
The importance of this result is directly linked to the
importance of time saving in clinical procedures,
typically very high.
Conclusions
Fig. 2: Merged map (left) and application of mIP (right)
Results
In Fig. 3 a comparison is shown between a
venography obtained by applying a mIP on the modulus
images and a venography obtained with all the
processing operations above described.
Fig. 3: Comparison between non processed (left) and
processed (right) venographies.
The filtered image results in a slightly more clearly
delineated venography. This is due to the operations on
the two groups of images (modulus and phase) and to
their fusion, that allows to enhance the contrast between
veins and surrounding tissues. On the other hand, the
image on the right shows some areas were the vein
structures are not well defined, due to a non-optimal
The NMR based on the BOLD technique is today
the only imaging method that allows not only to gain a
very high resolution, which allows to see vessels with
diameters of even 1 mm (voxel dimension), but is also a
non-invasive technique, as even the contrast agent used
is endogen, being the blood susceptibility. Applying the
mIP directly to the raw data, highly defined
venographies are obtained and virtually requiring no
time at all for the processing procedures. This result
should represent a further point in favour of the
exploitation of MRI for vascular clinical purposes.
References
[1] CONNOR S.E.J., JAROSZ J.M., (2002), ‘MRI of
Cerebral Venous Sinus Thrombosis’, Clinical
Radiology, 57, pp 449-461;
[2] REICHENBACH J., ESSIG M., HAACKE E., LEE B.,
PRZETAK C., KAISER W., SCHAD L. (1998), ‘High
Resolution venography of the brain using magnetic
resonance imaging’, MAGMA, 6, pp 62-69;
[3] OGAWA S., LEE T.-M. (1990), ‘Magnetic Resonance
Imaging of Blood Vessel at High Fields: In Vivo and
in Vitro Measurements and Image Simulation’. Mag.
Res. in Medicine, 16, pp 9-18;
[4] AFNI, Internet site address http://afni.nimh.nih.gov/
[5] CASCIARO S., ‘Image Processing and Functional
MRI Methods: Pulse Sequences, High Resolution
Venography, A.S.L. and B.O.L.D.’, PhD Thesis AA
1999-2002, Pisa University.
[6] OKADA T., YAMADA H., ITO H., YONEKURA Y.,
SADATO N., (2005), ‘Magnetic Field Strength
Increase Yields significantly greater Contrast-to-
Noise Ratio increase: measured using BOLD contrast
in the primary visual area’, Acad. Radiol., 12, pp 142-
147.
IFMBE Proc. 2005;9: 206

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THE INFLUENCE OF CORNEA TO FORMATION OF THE 2D IRIS IMAGE
E. Paliulis 1 , G. Daunys 1
1 Department of Electronics, Siauliai University, Siauliai, Lithuania
Abstract
The computer simulation of light rays propagation
through human eye cornea and anterior chamber was
carried out. The simulation revealed that because of
influence of cornea optics iris image has nonlinear
transformation in direction of eye rotation. The model
with spherical iris surface was proposed. The results of
computer simulation were found to agree with
experimental data of iris sections length changes versus
eye rotation angle.
egidijus.p@tf.su.lt
Introduction
In the recent years videoculography[1,2] became
prevailing eye tracking method. Because the systems
exploiting this method are contactless, they are widely
used not only in laboratory experiments and clinical
diagnosis, but also in domain of Human Computer
Interaction. The advantage of videoculography is
possibility to evaluate three dimensional eye rotations
[3,4,5].
The correct mathematical model is essential part of every
eye tracking system. 2D eye trackers are using pupil
centre coordinates. 3D eye trackers are using the natural
landmarks of the iris for extracting the angular position of
the eye. For both trackers the influence of eye cornea to
the formation of eye image on video sensor is important.
Usually influence of cornea is evaluated as linear
transformation[6].
The purpose of this research is to estimate the influence
of cornea to the formation of the image of iris and to
propose computationally effective model for taking it into
account.
Methods
Computer simulation was used to investigate the
propagation of light rays through cornea and anterior
chamber. The statistical human eye data were used in
mathematical model. The cornea was described as
transparent environment (refraction index n 2 =1.376)
occluded by two spherical surfaces. The curvature radius
of external surface is R1=7.8 mm, internal surface –
R2=6.5 mm. The curvature centres have displacement.
This causes different cornea thickness in centre and in
periphery. The refraction index of anterior chamber very
close to one of cornea (n3=1.336). The iris is planar, disc
shaped. We selected its diameter D=11 mm. The used
disposition of eye surfaces is presented at Figure 1.
Figure 1: Cross-section of frontal eye surfaces
The light ray refraction at boundary between to
environments is described by Snell’s law of classical
optics: n 1 sinα=n 2 sinβ, where α– angle of incidence, β –
angle of refraction, n 1 – refraction index of air. From the
simulation results length of sections from left pupil edge
to limbus (L L ) and analogical distance on the right side
(L R , see Figure 2) were evaluated versus eye rotation
angle in direction of eye rotation.
Figure 2: Evaluation of length of sections from pupil to
limbus
The same distances were manually evaluated from
experimental images. During experiment the subject sited
before screen with red light LEDs. The LEDs were
switched on by random order in horizontal direction.
Results
The propagation of parallel light rays through eye cornea
and anterior chamber was simulated. The results are
shown at Figure 3. The horizontal axis represents the
distance from eye centre. The vertical axis represents
IFMBE Proc. 2005;9: 207

Biomedical imaging techniques
distances between ray images on iris, when the initial
distance was 0.1 mm. One could see that when
eccentricity increases, the distances between images also
increase. Because of principle of reversibility light will
follow exactly the same path if its direction of travel is
reversed. So equally distributed points on iris will be find
closer to one other in eccentric positions than in centre.
The results suggested us to evaluate cornea optics
influence by changing planar iris disc to spherical surface
with curvature radius bigger than cornea curvature radius.
Figure 3: Distance between light images on iris versus
eccentricity of incident rays
Figure 4: Simulated changes of left section‘s length
versus the eye rotation angle (spherical iris – solid line,
discoid iris– dashed line)
Figure 5: Measured changes of sections' lengths L L and
L R versus the eye rotation angle (points). Solid lines
represent computer simulation results for spherical iris
with curvature radius 26.2 mm, dashed line - discoid iris
The computer simulation of section from pupil centre to
limbus boundary length changes versus eye rotation angle
was carried out with new model. The results are shown at
Figure 4 by solid lines. By dashed line is plotted cosinus
function, which describes length changes in discoid iris
model[7]. Simulation results were tested experimentally.
8 healthy subjects (5-male and 3-female) participated in
experiments. The experimental results from all 8 subjects
are plotted in Figure 5 together with the simulated lines.
The results exhibit good correspondence with computer
simulation.
Discussion
Computer simulation revealed that eye cornea causes
nonlinear distortion of the iris image after eye rotation.
The simulation results suggest using the simplified model
for the iris image formation, where influence of cornea is
taken into account by using iris of the spherical form.
Experimental investigation of the iris section length
changes in rotation direction confirmed the assumption.
The results of the study could help in the advancement of
the videooculographical eye movement recording
systems. The analysis of pupil centre displacement would
reduce systematic error of its centre coordinates
detection. Iris image distortion would enable to correct
the eye torsion measurements during eye rotation.
References
[1] DUCHOWSKI, A. T. (2003): ‘Eye Tracking
Methodology: Theory and Practice’. Springer-Verlag,
London. 252p.
[2] DELL’OSSO F., DAROFF R. (1987): ‘Eye
movement characteristics and recording techniques’,
Clinical Ophthalmology, 2, pp. 1 – 17.
[3] SUNG K., RESCHKE, M. F. ( 1997): ‘A Model-
Based Approach for the Measurement of Eye Movements
Using Image Processing’, Visiting Scientist, University
Space Research Association., pp.– 44.
[4] Wildes, R. P. (1997): ‘Iris Recognition: An Emerging
Biometric Technology’, Proceedings of The IEEE, 85(9),
pp. 1348-1363.
[5] HATAMIAN M., ANDERSON, D.J. (1983): Design
Considerations for a Real-Time Ocular Counterroll
Instrument, IEEE Trans. on Biomedical Engineering,
30(5), pp. 278-288.
[6] JAMES P. IVINS, PORRILL J. (1998): ‘ A
deformable model of the human iris for measuring small
three-dimensional eye movements’, Machine Vision and
Application, 11, pp. 42-51.
[7] Paliulis E., Daunys G., Vysniauskas V. (2004): ‘A
mathematical simulation of iris planar image’,
Electronics and Electrical Engineering (Kaunas:
Technologija), 51, pp. 57 – 61.
Acknowledgements
This research was carried out during project
ECOGASYD of the Eureka program with the cofinancing
of the Lithunian State Science and Studies
Foundation.
IFMBE Proc. 2005;9: 208

Biomedical imaging techniques
HIGH RESOLUTION VENOGRAPHY AS A VASCULAR MASK FOR
ACTIVATION SITES OF AUDITORY CORTEX AT 3T
S.Casciaro 1,2 , D.Zacà 2 , G. Palma 2 , R. Bianco 2 , S.Neglia 2 , E. Casciaro 1,2 , A. Distante 1,2
1 Consiglio Nazionale delle Ricerche/Istituto di Fisiologia Clinica, Sezione di Lecce, Lecce, Italy
2 Istituto Scientifico BioMedico EuroMediterraneo/Divisione di Bioingegneria, Brindisi, Italy
casciaro@ifc.cnr.it
Abstract: A high resolution vascular investigation
and a functional study were performed by means of
Blood Oxygen Level Dependent (BOLD) contrast
MRI in the auditory cortex of a healthy volunteer in
a 3T MR scanner. Activation maps of auditory
cortex were drawn by fMRI cross correlation
analysis and superimposed to the corrispondent
venous map depicted exploiting T2* effects showing
vessels having a diameter smaller than 1 mm. After
obtaining the activation regions, the local venous
architecture was used as anatomical images, in
order to correlate functional and anatomical
informations. A clear spatial correpondence
between activated regions and vein vessels was
found. These results are discussed referring to
hemodynamic models available in the literature:
future perspectives and applications of the method
are introduced.
resolution venograms were obtained simply through a
minimum Intensity Projection (mIP) across 10 slices
and a grayscale inversion to highlight the small dark
venous structures (Figure 1). Finally, functional maps
were combined with venographic images (Figure 2).
After a preliminary visual inspection, we carried out a
pixel by pixel signal analysis comparing the signal time
course between vascular and non vascular activated
voxels (Figure 3).
Introduction
The aim of this study was to demonstrate how
BOLD MR imaging at 3 Tesla, combining the results
of high resolution venography [ 1 ] and functional
analysis, can open new windows on underlaid
relationships between brain hemodynamic and BOLD
activation. At 3T the T2* images can be easily
processed to obtain high resolution vein maps, as
described in the next paragraphs. A method for
suppressing functional artefacts due to venous blood
contribution to the signal is suggested.
Materials and Methods
A fMRI event-related experiment was performed by
providing a series of five acoustic stimuli of different
durations to the subject. AFNI © suite softwares were
used to analyse both functional and anatomical data. A
sets of 10 T2*-weighted, 3 mm thick, EPI axial slices
covering auditory cortex were acquired, with TR =
1000 ms, TE = 30 ms, FA = 90°. The matrix
acquisition size was 64 x 64 and the field of view 20
cm. A cross correlation analysis was carried out to
find activation maps (r>0.6131) taking the Cohen
model as ideal reference function [ 2 ]. For the
venogram maps 60 T2*-weighted anatomical slices 3D
were also acquired (0.5 x 0.5 x 1 mm voxel size). High
Figure 1: Grayscale inverted mIP of 10 axial slices.
Results
The functional maps obtained in the auditory cortex
showed selective activation in the Heschl’s gyrus and
was consistent with those observed in other similar
experiments [3]. Figure 1 shows the high resolution
venographic image used as anatomical base for the
combined functional study. In fact Figure 2 (image on
the right) represents the fusion between functional
maps and venography. An interpolation of the
functional low resolution and the anatomical high
resolution dataset was necessary. In the activated
regions, a pixel by pixel analysis of the signal time
course and intensity deviation from baseline was done,
dividing them into two groups: vascular and non
vascular voxels. Ten main categories of deviation were
found as shown in Figure 3.
IFMBE Proc. 2005;9: 209

Biomedical imaging techniques
According to the hemodynamic models available in the
literature [4], the signal deviation from baseline (%
change of the maximum of the signal relative to the
baseline) is proportional to the fractional blood volume
in the voxel.
Figure 2: Venogram and functional map of the auditory
cortex images. On the left, zoomed image of the veins
only and on the right, vein image combined with the
regions of activation.
Figure 3: Histograms of signal deviation from baseline
for vascular and non vascular voxels in the activated
(r>0.6131) regions.
Discussion
The magnetic susceptibility difference between oxyand
deoxygenated hemoglobin increases with the
magnetic static field [4], allowing at 3 T to depict
voxels containing very small veins with sufficient
contrast respect to parenchyma and arteries in T2*
images. The image with superposition of functional
map and venography suggests a causal relationship
between the metabolic demand of neuronal local
activity and the density of small veins network in both
emispheres showing activations, as already
demonstrated for capillary beds in auditory cortex [5].
As shown in Figure 3, the highest deviation from
baseline belongs to the vascular voxels, as expected for
vascular signal change at 3T [6]. These results prove
how the obtained venography and functional map
based on BOLD effect are coherent to each other and
that a combination of informations is possible.
Conclusions
The discussed results suggest a method to suppress
vascular artifacts with a high resolution venography
based mask: it can be helpful to better localize the
neuronal activation site. Integration of informations
about function and anatomy of a brain region can
improve the understanding of the relationships between
functional measured signals and brain pathology. This
has potential clinical applications, providing important
additional information for preoperative neurosurgical
planning [ 7 ]. The results obtained in our first
experiment encourage to extend the method to a multisubject
analysis and to investigate other brain
activation sites.
References
[1] REICHENBACH J.R., VENKATESAN R., SCHILLINGER
D.J., KIDO D.K., HACCKE E.M. (1997):’Small
Vessels in the Human Brain: MR Venography with
Deoxyhemoglobin as Intrinsic Contrast Agent’,
Radiology, 204, pp 272-277
[2] COHEN M.S. (1997): ‘Parametric analysis of fMRI
data using linear systems methods’, Neuroimage, 6,
pp. 93-103
[3] ROBSON M.D., DOROSZ J.L., GORE J.C. (1998):
‘Measurement of the Temporal fMRI Response of
the Human Auditory Cortex to Trains of Tones’,
NeuroImage, 7, pp 185-198
[4] HACCKE E.M., BROWN R.W., THOMPSON M.R.,
VENKATESAN R. (1999): ‘Magnetic Properties of
Tissues: Theory and Measurement’ in ‘Magnetic
Resonance Imaging-Physical Principle and Sequence
Design’, (J. Wiley and Sons, New York), pp. 741-
781
[5] HARRISON R.V., HAREL N., PANESAR J., MOUNT
R.J. (2002): ‘Blood Capillary Distribution Correlates
with Hemodynamic-based Functional Imaging in
Cerebral Cortex’, Cerebral Cortex, 12, pp. 225-233
[6] OGAWA S., MENON R. S., TANK D.W., KIM S.G.,
MERKLE H., ELLERMAN J.M and UGURBIL K. (1993):
‘Functional Brain Mapping by Blood Oxygenation
Level-Dependent Contrast Magnetic Resonance
Imaging: A Comparison of Signal Characteristics
with a Biophysical Model’, Biophys. J.,64, pp. 803-
812
[7] MAJOS A., TYBOR K., STEFANCZYCK L., GORAY B.
(2004): ‘Cortical Mapping by Functional Magentic
Resonance Imaging in Patients with Brain Tumors’,
Eur Radiol, In Press
IFMBE Proc. 2005;9: 210

Biomedical imaging techniques
BPA QUANTIFICATION AND DETECTION IN PHANTOMS USING
THREE DIMENSIONAL 1 H MAGNETIC RESONANCE SPECTROSCOPIC
IMAGING
M. Timonen* , **, S. Savolainen** , *** and S. Heikkinen**
* Dept. of Physical Sciences, Univ. of Helsinki, POB 64, FIN-00014, Helsinki, Finland
** HUS Helsinki Medical Imaging Center, Univ. of Helsinki, POB 340 FIN-00029 HUS, Helsinki,
Finland
*** Boneca Corp., POB 700, FIN-00029 HUS, Finland
sauli.savolainen@hus.fi
Abstract
Applicability of three dimensional 1 H magnetic resonance spectroscopic imaging (3D 1 H MRSI) to detect and
quantitate boron neutron capture therapy (BNCT) boron carrier, boronophenylalanine (BPA), was studied with
phantoms. BPA-fructose/creatine phantom was used to study the quantitativity. Quantification was performed
with both single voxel 1 H MRS and 3D 1 H MRSI using creatine as an internal standard. The results obtained in
this study suggest that 3D 1 H MRSI could be very useful for BPA detection and quantification in vivo.
Introduction
Boron neutron capture therapy (BNCT) is an experimental radiotherapy that has typically been used to treat malignant
gliomas [1]. Effective BNCT necessitates selective boron accumulation in tumour. Currently, there is not a direct
method to monitor boron accumulation in patient brain during or after boron carrier compound infusion. L-pboronophenylalanine
fructose complex (BPA-F) is widely used as a boron carrier compound in BNCT. It has been
reported that aromatic protons of BPA can be detected in vivo using proton magnetic resonance spectroscopy ( 1 H MRS)
[2-4].
In this work the applicability of 3D 1 H MRSI to detect and quantitate BPA is studied. In principle, the traditional
single voxel 1 H MRS would be an optimal MRS method for in vivo BPA detection and quantification. In practise,
however, positioning of the MRS voxel successfully can be problematic in in vivo. Patients may have undergone a
surgery prior to BNCT and thus the resulting resection cavity could complicate the positioning of the MRS voxel. Also,
possible necrotic areas in unoperated tumours will hamper the voxel positioning. Furthermore, gliomas are typically
very heterogenous leading also to difficulties in voxel positioning. Quite frequently, several potential tumour voxel
locations can be found, but due to time restrictions only few single voxel spectra can be recorded. These problems
encountered in single voxel MRS can be circumvented by using 2D, or more efficiently, 3D 1 H MRSI.
Methods
BPA-F aqueous solutions with BPA concentrations of 1.75 and 3.0 mM were prepared. The 3.0 mM BPA-F phantom
contained also 10.1 mM creatine for quantification purposes. Solutions were in 50 ml round bottomed flasks. For MRS
measurements the flask was positioned in a spherical plastic phantom filled with 0.4 % NaCl-solution doped with
MnCl 2 (0.1 mM).
Measurements were performed at 1.5 T using clinical MRI scanner (Siemens Magnetom Sonata). Hamming
apodization with 400 ms FWHM (Full Width Half Maximum) was applied prior to spectral Fourier transformation in
both single voxel and 3D 1 H MRS-studies. In all measurements TE of 30 ms and TR of 1500 ms were used. 3D 1 H
MRSI (PRESS sequence) was measured from both phantoms using a standard clinical protocol with duration of ~17
min. The VOI size was 10 x 10 x 10 cm 3 and the nominal voxel size was 10(RL) x 10(AP) x 7.5(SI) mm 3 (one-fold zero
filling in SI-direction). Single voxel 1 H MRS spectrum with duration of ~17 min (PRESS sequence, 680 scans) and
voxel size of 15x15x15 mm 3 was measured from 3.0 mM BPA-F phantom. Creatine methyl proton T 1 was determined
from spectra recorded using TR-values of 1310, 1800, 2500, and 4000 ms (TE=30 ms, 128 scans, 20 x 20 x 20 mm 3
voxel size).
Quantification of BPA was done for 3.0 mM BPA-F phantom using both single voxel 1 H MRS and 3D 1 H MRSI.
Creatine signal was used as an internal standard. Gaussian lineshape fitting was used to determine intensities of
aromatic proton signals of BPA and creatine methyl signal. Intensities were corrected for T 1 -effects using T 1 -values of
935 ms for BPA aromatic protons [3] and 1545 ms for creatine methyl protons. BPA signal-to-noise ratio (SNR) was
determined for both single voxel and 3D MR spectra. Noise was determined as 2SD from the region 8-10 ppm. Signal
was determined as amplitude of BPA peak located at 7.3 ppm. The FWHM of BPA signal located at 7.3 ppm was
approximately the same for both single voxel and 3D MR spectrum.
IFMBE Proc. 2005;9: 211

Biomedical imaging techniques
Results
Single voxel and 3D MRSI spectra used in BPA quantification are presented in Figure 1. BPA quantification results and
SNR values are presented in table 1.
Figure 1. 1 H MRSI (A) and single voxel 1 H MR (B) spectra recorded from aqueous BPA-F/creatine phantom (3.0 mM
BPA, 10.1 mM creatine). BPA aromatic signals are indicated by white arrow. Inserted MR-images show the voxel
positioning.
Table 1. BPA concentrations and signal-to-noise ratios in single voxel 1 H MRS and 3D 1 H MRSI phantom
measurements.
Study
True BPA
concentration
[mM]
Measured BPA
concentration
[mM]
SNR
3D 1 H MRSI 3.0 3.01 21.3
Single voxel 1 H MRS 3.0 2.81 4.7
3D 1 H MRSI 1.75 11.3
Discussions
The SNR in 3D 1 H MRSI spectrum is over fourfold compared to SNR of the single voxel spectrum recorded using the
same measurement time as for 3D measurement (Table 1, 3.0 mM phantom). As FWHMs of BPA signals are
approximately the same in both spectra, indicating that the high SNR in 3D MRSI is due to large actual voxel volume.
In fact, the SNR of 3D spectrum measured from 1.75 mM BPA-F phantom exceeds the SNR of single voxel spectrum
recorded from 3.0 mM phantom (Table 1). High SNR of 3D MRSI increase probability of detection and identification
BPA signals in vivo. However, it should be noted that 3D MRSI suffers from problems due to non-uniform excitation
within VOI. This results from non-ideal slice selection profiles of RF-pulses and chemical shift originated VOI
displacement [5]. These problems are not of significant concern in single voxel 1 H MRS.
Conclusions
This phantom study suggests that 3D 1 H MRSI is a feasible method for BPA detection and quantification in vivo. Single
voxel 1 H MRS measurements will now be replaced by 3D 1 H MRSI in the Finnish BNCT project.
References
[1] DIAZ A. Z., CODERRE J. A., CHANANA A. D., and MA R. (2000): ’Boron neutron capture therapy for malignant
gliomas’, Ann. Med., 32, pp. 81-85
[2] ZUO C. S., PRASAD P. V., BUSSE P., TANG L., and ZAMENHOF R. G. (1999): ’Proton nuclear magnetic
resonance measurement of p-boronophenylalanine, (BPA): A therapeutic agent for boron neutron capture’, Med. Phys.,
26, pp. 1230-1236
[3] HEIKKINEN S., KANGASMÄKI A., TIMONEN M., KANKAANRANTA L., HÄKKINEN AM., LUNDBOM N.,
VÄHÄTALO J., and SAVOLAINEN S. (2003): ’ 1 H MRS of a boron neutron capture therapy 10 B-carrier, BPA-F:
phantom studies at 1.5 and 3.0 T’, Phys. Med. Biol., 48, pp. 1027-1039
[4] TIMONEN M., KANGASMÄKI A., SAVOLAINEN S., and HEIKKINEN S. (2004): ‘ 1 H MRS phantom studies of
BNCT 10 B-carrier, BPA-F using STEAM and PRESS MRS sequences: Detection limit and quantification’, Spectrosc-
Int. J., 18, pp. 133-142
[5] NELSON S. J., VIGNERON W. D. B., STAR-LACK J. and KURHANEWICZ J. (1997): ’High Spatial Resolution
and Speed in MRSI’, NMR in Biomed., 10, pp. 411-422
IFMBE Proc. 2005;9: 212

Biosensors
An Introduction to Biosensors
Britta Lindholm-Sethson
Dep of Chemistry, Biophysical Chemistry, Umeå University, Umeå, Sweden
Centre for Biomedical Engineering and Physics, Umeå University, Umeå, Sweden
Britta.sethson@chem.umu.se
Abstract: The basic features of biosensors are
presented. It is also discussed how biosensors
generally are classified according to certain
common characteristics based on signal
transduction and the molecular recognition event.
Introduction
A biosensor is a device which transduces a biological
signal using optical [1], electrochemical, [2, 3] or mass
sensitive techniques. [4] High selectivity can be
obtained since a biochemical mechanism is exploited
as recognition event. This is of high importance
especially in analytical chemistry applications in the
case of complex matrixes. Moreover, conclusions can
be drawn from investigations of specific biochemical
effects on the impact on living biological tissue or
specific biological functions.
Analyte in solution or
gas phase
Bioreceptor
Molecular
Recognition
Transducer
Measurement
Recording and
Display of Data
Fig 1: Principle of a Biosensor
The bioreceptor is responsible for the molecular
recognition of the analyte, i.e.: the selective binding of
the analyte to the sensor.
The molecular recognition event can be divided into six
major categories based on the type of the exploited
interaction.
a) antibody/antigen
b) nucleic acid/DNA
c) enzymatic
d) cellular structures/ cells
e) tissue/whole (higher) organism based
biosensors
f) biomimetic materials such as synthetic
bioreceptors.
The signal transduction of the molecular recognition
event is categorised in three major classes: i.e.: optical,
electrochemical or mass sensitive techniques, which are
described below.
Optical Sensors
When an optical sensor is chosen for signal transduction
in a biosensor, a large variety of subclasses are at hand.
These are based on for instance: absorption, Raman,
fluorescence, phosphorescence, SERS, refraction,
dispersion spectrometry etc. Different spectrochemical
properties are recorded such as amplitude, energy and
polarisation. Especially for analytical purposes the
amplitude is of high importance, since it is often related
to the concentration of the analyte. Affinity based
sensors with optical detection play an important role in
many important areas such as proteomics, clinical
diagnostics and drug discovery. [5]
Mass sensitive Sensors
In this case the analysis relies on a shift in the resonance
frequency of a piezoelectric crystal. The frequency shift
is directly related to addition or reduction of mass from
the surface of the crystal as a consequence of the
molecular recognition event. The technique is sensitive
to very small changes in the mass and detection limits in
the picogram range are reported. [6]
Electrochemical Sensors
Electrochemical detection is an attractive choice for
transduction of a biological related event since it allows
for mass production of disposable and cheap sensors.
For instance, a variety of commercial sensors are now
IFMBE Proc. 2005;9: 213

Biosensors
available for glucose analysis in whole blood based on
either amperometric [7] or coulometric detection [8].
Other types of electrochemical sensors are based on
potentiometric or chronoamperometric assays, various
puls or voltammetric techniques and in some cases
impedance measurements are employed.
Biochip
The word biochip is often associated with a solid
support, glass or polymer, with more than 100 000
microscopic spots containing a specific
oligonucleotide or cDNA. [9] This rapidly developing
field allows massive and simultaneous determination
of various binding events mostly detected by
measurements of fluorescence intensities. It is used for
instance in the screening of cancer genes. [10]
Discussion and Future outlook
The growing interest in the field of biosensors and
biochip technology has resulted in a large and
increasing amount of scientific papers. Not only are
several types of commercial and reliable analytical
biosensors developing. Detailed and important insight
is also gained from biosensor measurements
concerning the biology of living organisms such as the
kinetics of protein binding to oligomers or antibodies.
The large challenge in the future lies in the further
development of the biochip technology especially in
the case of membrane based biosensor.
References:
[1] HANEL C., GAUGLITZ G. (2002): ‘Comparison of
reflectometric interference spectroscopy with
other instruments for label-free optical detection’
Anal Bioanal Chem 372; 91
[2] MUNTEANU FD, OKAMOTO Y, GORTON L (2003):
‘Electrochemical and catalytic investigation of
carbon paste modified with Toluidine Blue O
covalently immobilised on silica gel’ Anal Chim
Acta 476; 43
[3] GOODING JJ, PUGLIANO L, HIBBERT DB, EROKHIN
P (2000): ‘Amperometric biosensor with enzyme
amplification fabricated using self-assembled
monolayers of alkanethiols: the influence of the
spatial distribution of the enzymes’ Electrochem
Comm 2; 217
[4] CUNNINGHAM B, WEINBERG M, PEPPER J, CLAPP C,
BOUSQUET R, HUGH B, KANT R, DALY C, HAUSER
E (2001): ‘Design, fabrication and vapor
characterization of a microfabricated flexural plate
resonator sensor and application to integrated
sensor arrays’ Sens Actuator B-Chem 73: 112
[5] BAIRD C.L. MYSZKA D.G. (2001): ‘Current and
emerging commercial optical biosensors’ J. Mol
Rec. 14; 261
[6] FREUDENBERG J, SCHELLE S. BECK K VON
SCHICKFUS M HUNKLINGER S (1999): ‘A
contactless surface acoustic wave biosensor’
Biosens Bioelectron 14: 423
[7] FELDMAN B, MCGARRAUGH G, HELLER A,
BOHANNON N, SKYLER J, DELEEUW E, CLARKE D
(2000): ‘FreeStyle: A small (300 nL) Volume
Electrochemical Sensor for Home Blood Glucose
Testing’ Diabetes Technol. Therapeutics 2; 221
[8] CHEN T, FRIEDMAN KA, LEI I, HELLER A (2000):
‘In situ assembled mass-transport controlling
micromembranes and their application in implanted
amperometric glucose sensors’ Anal Chem 72;
3757
[9] ZHU H,SNYDER M (2003): ‘Protein chip
technology’ Curr Opin Chem Biol 7; 55
[10] HACIA J, WOSKI S, FIDANZA J, EDGEMON K,
MCGALL G, FODOR S, COLLINS F (1998):
‘Enhanced high density oligonucleotide array-based
sequence analysis using modified nucleoside
triphosphates’ Nucleic Acids Res 26; 4975
IFMBE Proc. 2005;9: 214

Biosensors
CUSTOMIZING THE COMPUTER SCREEN PHOTO-ASSISTED
TECHNIQUE FOR EVALUATING QUICK DIAGNOSTIC TESTS
D. Filippini*, I. Lundström
Division of Applied Physics, Department of Physics and Measurement Technology, Linköping University, S-
581 83 Linköping, Sweden. E-mail: danfi@ifm.liu.se
Abstract: A computer screen used as a controlled
light source and a web camera as an image
detector is used for the spectral fingerprinting of
color indicators from a commercial test for eight
urine parameters. The measuring and
classification strategy of the method is presented
and general conclusions about the choice of color
indicators for custom designed assays are
discussed.
Introduction
The computer screen photo-assisted technique (CSPT
[1,2]) utilizes computer screens as programmable
light sources and regular web cameras as image
detectors for evaluating assays based on color or
fluorescent substances. The broad availability and
versatility of such as measuring setup, essentially just
a computer set with a web camera, makes CSPT an
attractive method for determinations carried out at
homes, doctor's offices or primary health care
centers, where other alternatives are either less
advanced or more expensive.
Since CSPT is an imaging method it can easily be
adapted to diverse assay layouts providing a unique
instrument for multiple tests. The method has been
demonstrated compatible with a number of sensing
principles (e.g. whole cell assays [2], cell viability
tests [1], DNA detection using polytiophene
derivatives [3], ELISA tests [4], gas sensing devices
[5], etc.).
Here the CSPT concept is introduced and its
application to the analysis of opaque samples such as
present in commercially available quick tests is
discussed. The limitations and possibilities of the
method are illustrated by considering the evaluation
of a eight parameters urine test.
Materials and Methods
In a typical CSPT experiment arrays of color
substances are characterized by illuminating them
with different light colors provided by the computer
screen, whereas the web camera captures the image
of the assay in synchronism with the illumination.
From the resulting video stream the spectral
fingerprints from selected regions of interest (ROI) in
the recorded images are extracted.
Since neither computer screens nor web cameras are
intentionally designed for analytical purposes,
measuring strategies must be developed in order to
overcome these weaknesses and to exploit the
inherent computing power.
A CSPT platform composed by a laptop computer
(Dell Latitude D800, with Pentium M processor of
1.6 GHz) with a WXGA screen (operating at a
resolution 1280 x 800 pixels with a color resolution
of 32 bits, and at a refresh frequency of 60 Hz) and a
web camera (Philips PCVC740K ToU Cam Pro, with
a CCD detector operating at a resolution of 320 x 240
pixels) is used for the determination displayed in Fig.
1.
neg.
1+
2+
3+
4+
2nd(9.49)Nitrites
2nd(4.18)Hemoglobin
2nd(7.08)Glucose
7
6
5
4
3
0.2
0.1
0
-0.1
-0.2
-1
-2
-3
-4
-5
-6
hemoglobin
blood
protein
nitrite
leucocytes
ketones
glucose
pH
3+ 3+ 3+ neg. 3+ 3+ 3+ 8
1+
2+
3+
4+
-4 -2 0 2 4
1st PC (91%)
neg.
pos.
-0.3 -0.2 -0.1 0 0.1 0.2 0.3 0.4
1st PC (31%)
1+
normal 2+
3+
-5 0 5
1st PC (88.2%)
4+
2nd(12.5)Blood
2nd(1.08)Leukocytes
2nd(8.21)pH
1
0.5
0
-0.5
1.6
1.4
1.2
1
0.8
0.6
0
-1
-2
-3
-4
-5
3+
1+ 2+
neg.
4+
5
6
7
8
9
-2 -1.5 -1 -0.5 0 0.5 1 1.5
neg. 1+
1st PC (75.9%)
2+
-1 -0.5 0 0.5 1
1st PC (92%)
5
6
7 8
-4 -2 0 2 4
1st PC (88%)
3+
9
2nd(2.44)Protein
2nd(8.66)Ketones
-3
-4
-5
-6
-1
-2
-3
-4
samples
evaluation
chart
strip
neg.
1+
2+
-3 -2 -1 0 1 2 3
neg.
1+
1st PC (94.3%)
2+
-3 -2 -1 0 1 2 3
1st PC (88.7%)
Fig. 1. Evaluation chart layout and one possible strip
outcome with its corresponding CSPT classification.
In the score plots the red circles correspond to the
classification of the references of each parameter and
the blue cruces to the projection of the corresponding
test fingerprints
3+
3+
IFMBE Proc. 2005;9: 215

Biosensors
Fig. 2 collects the classification of samples with color
gradients that are used to estimate the substance
colors associated with the best CSPT performance.
2nd principal component
12
10
8
6
4
2
0
-2
-4
-6
-8
samples
Y01 B01 G01 C01 R01
Y04 B04 G04 C04 R04
classification
YG04
YG03 CG04
G02
YG02
G01
YG01
YR01
Y02
Y03
Y01
YR02
CG03
G03
Y04
YR04 MR04
YR03
R03
R01 MR03
R02
MR01
MR02
R04
G04 C04
B04
CG02
C03 C02
C01
B03 CB02
B02
-10 -5 0 5 10
1st principal component
Fig. 2. Samples with color gradients (Y, B, G, C, R
and M for yellow, blue, green cyan red and magenta
respectively) and the considered regions of interest
(circles). Classification of CSPT fingerprints from the
regions of interest.
Results and Discussion
Quick tests e.g. such as those used for the
determinations of hemoglobin in urine can be
composed by test strips with a sensing patch that
changes color upon exposure to the target analyte.
Individual parameters can be evaluated by visually
comparing with a calibrated color scale. However,
this simple and inexpensive procedure does not
produce permanent records and depends on the visual
condition of the user. Additionally, multiple
parameter tests become cumbersome to evaluate and
dedicated readers (specific to certain brands and test
layouts) are demanded instead.
In the present case CSPT aims at producing this
evaluation at a comparable cost as the visual
inspection, specially assuming that potential users
possess computer sets. Fig. 1 illustrates the case of a
eight parameters urine test (hemoglobin, blood,
proteins, nitrites, leucocytes, ketones, glucose and
pH) and the CSPT evaluation of a particular sample.
B01
CSPT fingerprints of the different assay categories
(ROIs on the evaluation chart) are classified by
principal components analysis, and the fingerprints of
the corresponding test area projected into this space
and evaluate by proximity to the reference clusters
(the clusters are due to the variability between the
fingerprints of all the pixels composing each ROI).
In this way CSPT provides a recordable evaluation of
a multi-parameter test using a controlled illuminating
source independently of the visual capability of the
user. Computer screens and web cameras are not
originally conceived for analytical purposes and
therefore aspects such as stability must not be
assumed. Conversely, in every CSPT measurement
the test is simultaneously evaluated with the
evaluation chart, saving from additional calibration
measurements and providing a fresh reference for
each determination.
In order to assess the limitations of the CSPT
classification of diverse color substances, gradients of
yellow, blue, green, cyan and red samples are
classified (Fig. 2). Samples with combinations of
these colors are also projected into the principal
components space in order to complete the mapping
of color substances. Thus, is possible to observe
larger areas corresponding to orange-yellow-green
substances that indicate a better classification
resolution in these cases.
Conclusions
The strategy of simultaneously measuring samples
and references minimizes the influence of the
platform instabilities on the CSPT classification
capabilities, and eliminates additional calibrations,
which makes the system simpler for end users.
The analysis of the classification capabilities of color
gradients allows identifying preferential color
indicators that help to boost the CSPT performance.
CSPT, a technique able to adapt to any assay layout,
operating on a familiar and highly available hardware,
suggests attractive possibilities for primary care or
home testing applications at a cost comparable with
visual inspection methods.
References
1 D. Filippini, S. P. S. Svensson and I. Lundström,
Chem. Commun. 2 (2003), 240-241.
2 D. Filippini, T. P.M. Andersson, S. P.S. Svensson
and I. Lundström, Biosens. Bioelectron. 19 (2003)
35-41.
3 D. Filippini, P. Åsberg, P. Nilsson, O. Inganäs, I.
Lundström, Sensors and Actuators B (2005), in
press.
4 D. Filippini, K. Tejle, I. Lundström, Biosensors and
Bioelectronics (2005), in press.
5 D. Filippini, I. Lundström, Appl. Phys. Lett. 82
(2003), 3791-3793.
IFMBE Proc. 2005;9: 216

Biosensors
EXTRACTION AND ACTIVITY MEASUREMENT OF HUMAN 11ß-HSD2 FOR
USE IN A SALIVA CORTISOL BIOSENSOR
M. Sandström 1 , T. Shen 1
1 North, National Institute for Working Life, Umeå, Sweden
mattias.sandstrom@arbetslivsinstitutet.se
Abstract
The interest in work related stress is still of utmost
importance and the demand for easy-to-use methods for
measuring biological responses related to stress is strong.
In this paper we describe the improvement of the
enzymatic properties of the enzyme 11ß-hydroxysteroide
deydrogenase, type 2 (11ß-HSD2) that can be used in a
biosensor for measuring cortisol, a hormone related to
stress. The enzyme will be used as the biological
component in an amperometric biosensor for measuring
cortisol in saliva, which is under development. In this
work we have increased the enzymatic activity
considerably, between 10 and 100 times. The
measurements of 11ß-HSD2 activity are performed using
spectrophotometry and liquid spectrometry connected to
a mass spectrometer.
Introduction
The interest in work related stress is still of utmost
importance and the demand for easy-to-use methods for
measuring biological responses related to stress is strong.
It is important to complement other methods of
investigating stress related conditions. Cortisol is one
hormone which levels have been found to be affected in
patients with symptoms related to stress. Most methods
used for measuring cortisol are based on separate
sampling and analysis.[1; 2] However, these methods
delay the time between sampling and response. To reduce
this response delay we are developing a one-shot
biosensor that will be able to measure cortisol. To reduce
a potential problem of further stress due to sampling of
cortisol in patients, we also aim to develop the biosensor
for use on saliva, which is not invasive to the patients.
This paper describes the improvement of the enzymatic
properties required for use in such a device.
The overall goal is the development of an amperometric
biosensor that uses a recombinant form of the human
enzyme 11ß-hydroxysteroide deydrogenase, type 2 (11ß-
HSD2). 11ß-HSD2 is a membrane bound human enzyme.
The role of the enzyme in the human body is to regulate
the levels of cortisol by catalysing the oxidation of
cortisol to cortisone. The reaction requires the cofactor
NAD + , which will be used in the biosensor to determine
the cortisol level by measuring the electron transfer at the
electrode surface.
Methods
The enzyme was expressed in E-coli by transformation of
a HIS-tagged 11ß-HSD2 gene inserted in a pET25
plasmid vector. To ensure that the enzyme could be used
in a biosensor application the following were
investigated: methods of growing bacteria, extracting and
purifying the enzyme to increase the production of the
total amount of enzyme and to increase the activity of the
11ß-HSD2 enzyme.
Parameters investigated in order to increase the
expression of the enzyme were, for instance, the
temperature for growing the E-coli bacteria, the time for
expression of the enzyme and the concentration of the
chemical promoting the expression of the enzyme
(Isopropylß-D-thiogalactoside).
In order to increase the extraction of enzyme from the E-
coli bacteria we investigated ultrasonic treatment,
addition of a nondenaturing zwitterionic detergent
(CHAPS) and addition of Triton to the extraction buffer.
A spectrophotometric method was set up to measure the
production of NADH in the enzymatic reaction, as the
activity of 11ß-HSD2. Although similar methods have
been used when measuring activity for other enzymes, a
spectrophotometric method for measuring 11ß-HSD2
activity has not been found in the literature. To verify the
spectrophotometric method, a method based on liquid
chromatography connected to a mass spectrometer (LC-
MS/MS) was also used. It was based on the measurement
of cortisone, which is produced in the enzymatic reaction.
Results
The best enzyme production was achieved by expressing
the protein at 27°C and at low concentrations of IPTG.
The plasmid DNA electrophoresis showed that E-coli
would only keep the transformed plasmid containing the
HIS-tagged 11ß-HSD2 gene in the cell during the first
generations. The bacteria did grow for several
generations. However, the plasmid DNA, which could be
extracted, was very low. Hence, it is of utmost
importance to use the first generations after
transformation.
The most effective extraction of the 11ß-HSD2 enzyme
was achieved by a combination of ultrasonic treatment
and the addition of CHAPS and Triton to the extraction
buffer.
IFMBE Proc. 2005;9: 217

Biosensors
The spectrophotometric method indicated a 10 to 100
folds increase in enzyme activity, which was also verified
with the LC-MS/MS method.
Discussion
No attempts have previously been made to use the 11ß-
HSD2 enzyme in a biosensor application. It was reported
that the activity of the enzyme is very low (about 100
pmol/hour/mg protein).[3] Thus, increasing the activity
and the production of the enzyme is necessary for the
biosensors ability to detect the low concentrations of
cortisol in saliva.
The biosensor application for measuring cortisol has the
possibility of becoming a quick and easy way of
measuring stress related hormones in patients and in
research projects. However, the need for a high activity
and high stability enzyme for this purpose is apparent and
we find the activity increase, shown in this paper,
important for future work on the cortisol biosensor.
References
[1] Morineau G, Boudi A, Barka A, Gourmelen M,
Degeilh F, Hardy N, Al-Hanak A, Soliman H, Gosling
JP, Julien R, Brerault J, Boudou P, Aubert P, Villette J,
Pruna A, Galons H, Fiet J. (1997): Radioimmunoassay of
cortisone in serum, urine, and saliva to assess the status
of the cortisol-cortisone shuttle. Clinical Chemistry , 43,
pp. 1397-1407.
[2] Okumura T, Nakajima Y, Takamatsu T, Matsuoka M.
(1995): Column-switching high-performance liquid
chromatographic system with a laser-induced flurometric
detector for direct, automated assay of salivary cortisol.
Journal of Chromatography B: Biomedical Applications.
670, pp. 11-20.
[3] Nunez BS, Mune T, White PC. (1999): Expression of
human kidney 11ß-hydroxysteroid dehydrogenase (11-
HSD2) in bacteria. Biochemical and Biophysical
Research Communications. 255, pp. 652-656.
Acknowledgements
Swedish council for working life and social research
(FAS)
Prof. Perrin C. White, The University of Texas
Southwestern Medical School, USA
PhD Lena Sundberg, National Institute for Working Life
– North, Umeå, Sweden
PhD Anna-Lena Sunesson, National Institute for
Working Life – North, Umeå, Sweden
IFMBE Proc. 2005;9: 218

Biosensors
BIOSENSORS BASED ON MEMBRANE ORGANISATION
Paul Geladi*, Andrew Nelson**, Kathryn Bradley** and Britta Lindholm-Sethson***
* The Unit of Biomass Technology and Chemistry, SLU Röbäcksdalen,
P.O. Box 4097, SE - 904 03 UMEÅ, Sweden
** Self-Organising Molecular Systems, Dep of Chemistry, University of Leeds, LS2 9JT, UK
***Department of Chemistry, Biophysical Chemistry Umeå University, SE-901 87 Umeå Sweden
Paul.Geladi@btk.slu.se
Abstract: Electrochemical Impedance Technique
was employed to investigate magainin interaction
with a lipid monolayer on a mercury drop
electrode. It is also demonstrated how multivariate
techniques can be used as a complement to classical
analysis of the data.
Introduction
During the last decade, several research groups have
focused on the creation of artificial membranes on
solid supports. One proposed application area is the
development of biosensing devices, which also gives
the prospect for fundamental studies of membrane
protein function. Much attention has been paid to the
building of supported lipid bilayers, surrounded with
aqueous phases. The rationale is to make room for
incorporated bulky membrane proteins with retained
activity. As suitable methods Langmuir-Blodgetttechniques,
liposome fusion and/or self-assembly have
been proposed. [1]
The purpose of our research is to develop membrane
based biosensors that rely on the identification of
specific interactions between certain ions or molecules
in solution and the membrane surface itself. Thus,
there is no need for construction of supported lipid
bilayers and the focus is therefore on reproducible
creation of only half the lipid bilayer. A fluid
monolayer on a mercury drop is employed in this
particular work. The technique was originally
presented by Miller [2] and has been developed and
refined mainly by Nelson and co-workers. [3]
Electrochemical Impedance Spectroscopy provides an
opportunity to probe relaxation processes with
different time constants in one single experiment. It is
therefore acknowledged as a powerful method for the
characterisation of supported lipid membranes and is
the method of choice in this work.
Classically the impedance data are analysed by fitting
them to an equivalent circuit of resistors, capacitors
and other elements as an approximated model of the
interface [4]. This makes it possible to extract
information on for instance diffusion coefficients or
heterogeneous rate constants provided the model is
correct. Obviously, information from relaxation
processes that are not included in the equivalent circuit
is lost. It is also well known that the more components
that are added to the equivalent circuit the better is the
fit, although it could turn out to be completely nonsense.
We have earlier shown that the relevant information can
be extracted from the data by using methods from
chemometrics, see for instance: [5] This is further
demonstrated in this paper from impedance
measurements on magainin interaction with a lipid
monolayer on a mercury drop.
Materials and Methods
Monolayers of dioleoyl phosphatidylcholine, DOPC
were prepared on a fresh mercury drop as described
earlier. [3,6] 0.1 mol/dm 3 NaCl was used as supporting
electrolyte and a blanket of argon gas was maintained
above the fully deaerated electrolyte during all
experiments.
Impedance measurements were performed on the
electrode system at a potential regime where the
monolayer is expected to be defect free and at 50
frequencies logarithmically distributed between 65 000
and 0.1 Hz, 0.005 V rms. Experiments were performed
either in pure supporting electrolyte or with magainin
added to a total concentration of 10 nM and 30 mM
respectively. Magainin is a toad venom anti-microbial
peptide.
A data matrix was built containing 17 objects for 100
variables, 50 real impedances and 50 imaginary ones.
Principal components were calculated after meancentering
and scaling
Results:
The impedance data is transformed to the complex
capacitance plane through the formula C= 1/jωZ = ReC
+jImC, where j 2 = -1 and ω is the angular frequency.
This transformation emphasis relaxation processes in
the low-frequency range where the real part of the
capacitance is related to the dielectric ‘constant’
although ReC is frequency independent only at very low
frequencies. The imaginary part is related to the
dielectric loss and is the frequency dependent part of the
conductivity. In Fig 1, typical complex capacitance
spectra are displayed from measurements in the three
different solutions. Around 1000 Hz a sharp decrease is
IFMBE Proc. 2005;9: 219

Biosensors
observed in ReC, which can be interpreted as a fast
relaxation process at the lipid/solution interface, i.e.:
charging of the interface. The measurements in pure
0.1 M NaCl and 0.1 M NaCl + 10 nM magainin are
identical whereas at low frequencies a small increase is
observed in the case of 0.1 M NaCl + 30 nM magainin.
Similar observations are made in the ImC component
t2
10
5
10 nM Magainin
30 nM Magainin
pure dopc
3
0
Re(C/µF*cm -2 )
-Im( C/µF*cm -2 )
2
pure DOPC
pure DOPC
10 nM Mag
10 nM Mag
30 nM Mag
30 nM Mag
-5
1
-10
-10 -5 0 5 10
t1
Figure 2. PCA Score Plot from all measurements. t1
and t2, explain 41.4% and 30.2% of the sum of squares.
0
-2 -1 0 1 2 3 4 5
log(f/Hz)
Figure 1. Results from impedance measurements after
transformation to the complex capacitance plane.
Filled symbols refer to ReC and unfilled to ImC.
The score plot from PCA analysis of the seventeen
impedance measurements on the DOPC monolayer is
displayed in Fig 2. It is clear that the scores from
measurements in pure supporting electrolyte form a
cluster with the measurements in 10mM magainin.
Measurements in 30 mM magainin form a separate
cluster
Discussion
A close inspection of the impedance data after
transformation to the complex capacitance plane
reveals a dramatic change in the low-frequency part of
imaginary capacitance when the concentration of
magainin is increased from 10 nM to 30 nM. The low
time constant of the interaction indicates channel
formation due to partitioning of the magainin into the
monolayer that gives rise to ionic motion within the
film. The slight increase in the real capacitance
verifies that the magainin actually has entered the film.
The score plot in figure 2 underlines the observations
above. An addition of 10 nM magainin does not affect
the DOPC monolayer. With 30 nM magainin in
solution the result from principle component analyses
reveals a significant interaction with the monolayer.
Conclusions
The present system is well chosen for studying peptide
interactions with a lipid layer. Electrochemical
impedance spectroscopy is an ideal technique to
investigate these systems, especially in combination
with multivariate analyses.
References
[1] Sackmann E (1996): 'Supported Membranes:
Scientific and practical applications’ Science 271,
pp. 43-48
[2] Miller IR, Rishpon J, Tenenbaum A (1976):
‘Electrochemical determination of structure and
interactions in spread lipid monolayers’
Bioelectrochem Bioenerg 3, pp. 528-542
[3] Nelson A, Benton A (1986): 'Phospholipid
Monolayers At the Mercury Water Interface’ J .
Electroanal. Chem. 202, pp. 253-270
[4] Mac Donald, JR (1987): ‘Impedance Spectroscopy’
John Wiley & Sons; New York.
[5] Lindholm-Sethson, B.; Nystrom, J.; Geladi, P.;
Nelson, A. (2003): 'Gramicidin A Interaction at a
Mercury Supported Dioleoyl Phosphatidylcholine
Monolayer’ Anal. Bioanal. Chem. 375, 350-355.
[6] Leermakers, F. A. M.; Nelson, A. (1990): Substrate
–induced structural changes in electrode absorbed
lipid layers – A self consistent field theory J.
Electroanal. Chem. 1990, 278, 53-72.
IFMBE Proc. 2005;9: 220

Biosensors
ORIENTED IMMOBILISATION OF ANTIBODIES FOR IMMUNOSENSING
I. Vikholm-Lundin and W.M. Albers
Technical Research Centre of Finland, Information Technology, Finland
inger.vikholm@vtt.fi
Abstract: In order to obtain a site-specific orientation
of antibodies for immunosensing and prevent nonspecific
binding, antibody Fab´-fragments have been
immobilised directly onto gold and the remaining free
space in between the antibodies have been coated by a
non-ionic hydrophilic polymer of N-
[tris(hydroxymethyl)methyl]acrylamide. The polymer
possesses low NSB and can be covalently attached
onto the gold surface by disulfide anchors. The novel
immobilisation method has been demonstrated for
antibody fragments specific for human IgG and C
reactive protein both in phosphate buffer and in
serum matrices.
Introduction
The use of immunoassay technology in clinical, food
safety, and environmental analysis will continue too
grow. In immunoassays the antibodies are, however,
mostly adsorbed on the sensor surface or covalently
coupled via functional groups that are not site-specific.
When immobilizing antibodies via amino and carboxyl
groups, the orientation of the protein molecule will be
random. The lack of control over the orientation of the
antibodies limits the proportion of available binding sites,
whereas site-specific immobilisation leads to higher
activity. We have previously demonstrated the covalent
coupling of antibody Fab´-fragments to N-(εmaleimidocaproyl)dipalmitoylphosphatidylethanolamine
embedded in a host monolayer matrix of phosphatidylcholine
to achieve site-directed immobilisation of
antibodies with high antigen binding efficiency. 1-3
Antibody Fab´-fragments can, however, also be directly
coupled onto gold and the space in between the
fragments can be filled up with a non-ionic hydrophilic
polymer. 4-7 In this presentation surface plasmon
resonance (SPR) has been used to investigate the
coupling of various antibody fragments and the polymer
N-[tris(hydroxymethyl)methyl]-acrylamide, pTHMMAA
onto gold. The immobilisation method has been
optimised for Fab´-fragments of polyclonal anti-human
IgG (hIgG) and monoclonal anti-C-reactive-protein
(CRP). The influence of immobilisa-tion order, polymer
length, Fab´-fragment and polymer concentration on the
sensitivity of the layer both in phosphate buffer and
serum has been determined. Preliminary measurements
of patient samples have also been performed.
Materials and Methods
Human IgG and polyclonal goat anti-human IgG
F(ab´) 2 (Jackson ImmunoResearch), as well as CRP and
monoclonal anti-CRP F(ab´) 2 (Medix Biochemica)
were used as model systems. The F(ab') 2 fragments
were split into Fab´-fragments with dithiotreitol (DTT,
Merck) in a 150 mM NaCl, 10 mM HEPES, 5 mM
EDTA buffer pH 6.0, typically over night in a
microdialysis tube. 5
Glass slides coated with a thin film of gold were
cleaned in a hot solution of H 2 O 2 :NH 4 OH:H 2 O (1:1:5)
and rinsed with water. The slides were assembled into
the SPR device (either the SPRDevi from VTT,
Tampere, Finland, or the Biacore 3000 from Biacore,
Uppsala, Sweden) and Fab´-fragments and the
disulphide bearing polymer, pTHMMAA were allowed
to interact with the surface for 5 minutes. 7 The surface
was blocked with 0.5 g/l bovine serum albumin (BSA),
rinsed with buffer and the interaction with antigen or
patient samples was determined in phosphate buffer or
serum/PBS (1/100).
Results and Discussion
Anti-hIgG Fab´-fragments directly coupled to gold
gave a threefold response to human IgG compared to
that of F(ab) 2 -fragments. The layer of F(ab) 2 -fragments
could not be regenerated because the antibody F(ab) 2 -
fragments were adsorbed on the surface in a randomly
oriented fashion. A much higher antigen binding
capacity and an ability to regenerate the antibody layer,
on the other hand, was observed for the layers formed
from Fab´-fragments, indicating site-directed
immobilisation on the gold surface.
When the antibody Fab´-fragments were applied to
the gold surface and the remaining free space in
between the antibodies are filled up with pTHMMAA,
non-specific binding (NSB) was considerably reduced
and a large part of the fragments seemed to be attached
in a site-directed fashion through the free thiol bonds
(Figure 1). Coupling of the antibody Fab´-fragments and
the polymer, and thus both the amount of NSB and
antigen binding, but also the ability to regenerate the
layer is dependent on the immobilisation order. When
Fab´-fragments and pTHMMAA were immobilised on
IFMBE Proc. 2005;9: 221

Biosensors
the surface from the same solution up to 80% of the
antigen could be removed, indicating a high degree of
site-directed immobilisation of the antibody fragments.
About 60% of the antigen could be removed, when the
fragments were coupled directly onto a clean Au surface
before the polymer or if low concentrations of polymer
were attached onto gold before the Fab´-fragments. The
highest response to hIgG was, however, obtained when
the antibody fragments were attached onto the surface
before the polymer. When the polymer was attached to
the surface before the antibody the response was half of
that when the antibodies and the polymer were attached
from the same solution.
S S S S S S S S S S S S S S S S
Antigen
Fab´-fragment/
polymer
Au
Glass
Figure 1: Schematic view of site-directed antibody Fab´fragments
and polymers with low non-specific covalently
attached onto gold and interaction of the layer with
antigen.
In PBS buffer the highest response to CRP was
obtained at a Fab´-fragment concentration of 60 µg/ml.
CRP could be detected in a concentration range of 1
ng/ml – 50 µg/ml from a standard solution in phosphate
buffer. The response was 5-fold to that of a F(ab) 2 -
fragment/polymer layer. The non-specific binding of
BSA was only 1% of the CRP binding and a detection
limit of 1.5 ng/ml could be obtained. If the NaCl
concentration in the buffer was high, the Fab´/polymer
layer seemed to be randomly oriented with a similar
response to CRP as that of a F(ab) 2 -fragment/polymer
layer. Antibody fragments embedded in a pTHMMAA
polymer prepared with 1.1 mol% lipoate initiator was
very stable and did not show any decrease in response
after a few days of storage. Fab´-fragments embedded in
a pTHMMAA polymer prepared with 2 mol% lipoate
initiator, on the other hand, were unstable and the
response to CRP decreased abruptly during the first days
of storage. Thus the pTHMMAA polymer with a longer
chain length (1.1%) seems to preserve the activity of
antibodies much better. CRP was successfully detected in
patient samples with good reproducibility. By
modifications of the polymer such as length and repellent
properties, the immunological response of the layer could
most probably be further improved. The layer would be
sensitive enough to analyse the CRP concentration in
human serum for predicting cardiovascular disease.
Conclusions
Antibody Fab´-fragments can be directly coupled onto
gold and the space in between the fragments can be
filled up with a non-ionic hydrophilic polymer. The
antibody fragments should be embedded in the protein
repellent host matrix in such a way that only the antigen
binding part of the antibody sticks out from the
monolayer surface. NSB to the host matrix and to the
antibody fragments could be substantially reduced and a
high antigen binding capacity could be obtained.
The immobilisation method is generic and can be
used for coupling any antibody in an oriented manner to
the sensor surface. There are several advantages with
the method:
1. It is simple and easy to perform in only a few steps.
2. The site-directed orientation of the antibodies ensures
a high specific binding of antigen with a minimum
amount of antibody needed for immobilisation.
3. The non-specific binding is extremely low due to the
repellent polymer.
4. A membrane-like environment is provided by the
polymeric host matrix that protects the antibodies
from unfolding.
5. The layer is reasonable stable, and can be regenerated
for repeated use.
References
[1] VIKHOLM, I., ALBERS, W.M. (1998): ‘Oriented
Immobilisation of Antibodies for Immunosensing’,
Langmuir 14, pp. 3865-3872.
[2] VIKHOLM, I., ALBERS, W.M., VÄLIMÄKI, H.
HELLE, H. (1998): ‘In situ Quartz Crystal
Microbalance Monitoring of Fab´-fragment Binding
to Linker Lipids in a Phosphatidylcholine
Monolayer Matrix. Application to Immunosensors’,
Thin solid Films 327-329, pp. 643-646.
[3] VIKHOLM, I., VIITALA, T., ALBERS, W.M.,
PELTONEN, J. (1999), ‘Highly Efficient
Immobilisation of Antibody Fragments to
Functionalised Lipid Monolayers’, Biochim. et
Biophys. Acta 1421, pp. 39-52.
[4] VIKHOLM, I., SADOWSKI, J. (2002), ‘Method
and Biosensor for Analysis’, Patent Application.
(2001) No FI20011877, US2003059954.
[5] VIKHOLM, I. (2005): ‘Self-assembly of Antibody
Fragments and Polymers onto Gold for
Immunosensing’, Sensors & Actuators B 106, 311-
316.
[6] VIKHOLM-LUNDIN, I. (2005) ‘Immunosensing
Based on Site-Directed Immobilization of Antibody
Fragments and Polymers that Reduce Non-Specific
Binding’, Langmuir. In press.
[7] VIKHOLM-LUNDIN, I., ALBERS, W. M. (2005):
‘Site-Directed Immobilisation of Antibody
Fragments for Detection of C-Reactive Protein’,
Biosensors & Bioelectronics. In press.
IFMBE Proc. 2005;9: 222

Biosensors
BIOCOMPATIBLE PACKAGING OF
A THREE-AXIS MICRO ACCELEROMETER
K. Imenes*, K. Aasmundtveit*, L. Hoff*, and O.J. Elle**
* Vestfold University College, Faculty of Science and Engineering, Horten, Norway
** Rikshospitalet University Hospital, Interventional Centre, Oslo, Norway
kristin.imenes@hive.no
Abstract: One challenge in developing new
microsystems and -products is the availability of
suitable packaging. When MEMS-chips are
miniaturized, the packaging and connection become
the driving parts regarding the size of the final
product. This paper will present the early progress
in developing packaging methods for an implantable
micro sensor for heart motion monitoring and in
short terms describe the upcoming activities. The
goal set for our project is to develop a miniaturized
biocompatible sensor with approximately 2mm
diameter and ~5mm length. The sensor is electrically
and mechanically connected to a cable for signal
output and for pulling the sensor out of the body
after the postoperative monitoring.
Introduction
Vestfold University College and Rikshospitalet
University Hospital are collaborating in developing a
micro sensor for monitoring heart motion. The idea,
originated from the Interventional Centre at
Rikshospitalet University Hospital, is to use a three-axis
micro acceleration sensor for monitoring patients during
coronary by-pass surgery and the first postoperative
days [1]. The sensor is to be attached to the heart
surface and connected to an analysis system through a
cable for power supply and signal output. This
connection must have the mechanical strength to be
pulled out of the body after the postoperative
monitoring. The physical size of the final miniaturized
heart sensor is estimated to be about 2mm in diameter
and 5mm in length. Finding methods for packaging with
reference to this size is a challenge.
Materials and Methods
The first experiments on monitoring the heart
motion was performed on anesthetized pigs with a
prototype sensor based on commercially available dualaxis
accelerometer sensors. The results reproduce the
heart motion in great detail and reveal patterns that may
be an indication of heart circulation failure [2, 3].
This first prototype sensor was made of two
commercial dual-axis sensors (ADXL310, Analog
Devices Inc, Norwood, MA, USA), separately mounted
on a printed circuit board and assembled perpendicular
to each other to get acceleration data from three axes.
The sensors needed some passive electronics and
electrical connection to 6 conductors (see figure 1). We
had available a 4-lead medical cable from Plastics One
Inc. (Roanoke, VA, USA), and therefore two cables
were necessary. The assembling was done by hand
soldering techniques and we ended up with a system
approximately 15 mm long and 6 mm wide, excluding
encapsulation. Stress relief was achieved by tying a
sewing thread around the cables and through holes in
one of the PCB’s.
These sensors were encased in two ways, one using
heat-shrinkable tubing and the other moulded in
polyester and grinded. They are glued to a plastic plate
with small holes along the edge, see figure 1. This was
done to make the encapsulated sensor easy to be sutured
to the wall of the pig’s heart, and to get a fixed
orientation of the sensor relative to the heart’s axes.
Figure 1: First Prototype with and without
encapsulation.
None of these encapsulated sensors are qualified for
use in humans, nor are they small enough for pulling out
of the body. A second prototype sensor is currently
being developed and qualified for human trials with the
aim to collect data for work on signal interpretation.
This shall be qualified for use on patients during surgery
for a limited period of time - up to one hour.
Our next generation prototype is smaller and based
on a commercial three-axis accelerometer with
dimensions 5x5x1.8mm (KXM52-1050, Kionix Inc,
Itacha, NY, USA). It is a Dual Flat No lead package
(DFN), with 14 terminals underneath and a pitch of
0.5mm. The sensor is assembled on a thickfilm substrate
together with passive components and the cable
connection [4].
IFMBE Proc. 2005;9: 223

Biosensors
Medical cables for use in vivo are usually custom
made, and we have developed a cable for our
application in cooperation with New England Wire
Technologies (Lisbon, NH, USA). The cable has 10
silver plated copper conductors with diameter 0.23mm.
The outer jacket is silicone rubber and has a diameter of
2.06mm. It has twisted pairs to suppress noise and is
very flexible. The last is essential to ensure that the
cable and the sensor do not influence the heart motion.
This sensor must be encapsulated in a biocompatible
material which is electrically insulating and prevents
body fluid from penetrating. We have moulded the
sensor in a biocompatible silicone material. Silicone is a
well established material for use in invasive devices and
is suitable for prototyping. Several different types of
silicone approved for invasive use are available [5]. We
have experimented with some types and evaluated their
suitability for our sensor and application. The
encapsulation must also stand a cleaning and sterilizing
process. Steam autoclaving is the preferred sterilization
method, but EtO gas and gamma irradiation are
alternatives [6, 7]. Another task is to test the
permeability to ensure that body fluid does not penetrate
during the time frame set.
The encapsulation design is important for several
reasons:
− Good fixation to the heart is essential to ensure
that the sensor follows the heart motion as
closely as possible.
− A round shaped sensor will minimize the
damaging of heart tissue and the surrounding
body tissue.
− Smooth surface for proper cleaning.
The prototype is to be sutured on the heart wall through
holes in each corner of the encapsulation. It shall not
have any motion of its own, as this may interfere with
the results and cause misinterpretations. Figure 2 shows
a preliminary encapsulation design.
Discussion
The size of the second prototype is limited by the
fact that the sensor chip has a given size. We had to
minimize the dimensions in all directions so that the
total size of encapsulated sensor became as small as
possible. Soldering of 5-10 leads onto a substrate
together with a good mechanical fastening of the cable
to the substrate challenges the total packaging size. A
flat cable offers easier mounting and also reduces the
pitch, but it does not have the wanted flexibility and
neither the same immunity against noise. The prototype
is assembled partially by hand which gives practical
limits to how small the dimensions could be. The
packaging of the prototype has also been important in
order to gain more experience in working and moulding
with silicone.
References
[1] ELLE, O. J., HALVORSEN, S., GULBRANDSEN, M. G.,
AURDAL, L., BAKKEN, A., SAMSET, E., DUGSTAD,
H., FOSSE, E. (2005): ‘Early recognition of regional
cardiac ischemia using a 3-axis accelerometer
sensor’, Physiol. Meas., 26, in press
[2] GRIMNES, M., HOFF, L., HALVORSEN, S., ELLE, O.
J., FOSSE, E. (2004): ‘Velocity and Position
Approximations from Left Ventricular 3D
Accelerometer Data’, Proc. of the IEEE 30th Ann.
Northeast Bioeng. Conf., Springfield, MA, USA
2004, pp. 25-26
[3] HOFF, L., ELLE, O. J., HALVORSEN, S., ALKER, H. J.,
FOSSE, E. (2004): ‘Measurements of Heart Motion
using Accelerometers’, Proc. of 26th Ann. Internat.
Conf., IEEE Eng. in Med. and Biol. Society, San
Fransisco, USA, 2004, pp. 2049-2051
[4] TUMMALA, R. R. (2001): ‘Fundamentals of
Microsystems Packaging’, (McGraw-Hill, New
York), pp. 717-724
[5] HEIDE C. (1999): ‘Silicone Rubber for Medical
Device Applications’, Medical Device & Diagnostic
Industry, Nov, p. 48
Figure 2: Preliminary encapsulation design.
[6] BLACK, J. (1999): ‘Biological Performance of
Materials: Fundamentals of Biocompatibility’,
(Marcel Dekker Inc., New York), pp. 24-25
[7] RATNER, B. D. (1996): ‘Biomaterials Science, An
Introduction to Materials in Medicine’, (Elsevier
Science, San Diego), pp. 415-419
Acknowledgement
Thanks to Øyvind Kaasa at Vestfold University
College for assembling the first prototype.
IFMBE Proc. 2005;9: 224

Sports and rehabilitation biomechanics
POSSIBILITIES AND CHALLENGES OF MULTI-CHANNEL
SURFACE EMG IN SPORTS AND REHABILITATION
K. Roeleveld*, J.S. Karlsson**, C. Grönlund**, A. Holtermann*, N. Östlund **
*Human Movement Sciences Program, Norwegian University of Science and Technology,
Trondheim, Norway.
**Department of Biomedical Engineering & Informatics, University Hospital, Umeå, Sweden
and Centre for Biomedical Engineering and Physics, Umeå University, Umeå, Sweden
E-mail: karin.roeleveld@svt.ntnu.no
Abstract: In sports and rehabilitation, bipolar surface
electromyography (SEMG) is often used to get
information about muscle activation. Applying many
addition channels improves the estimation of
traditional SEMG variables and allows estimation of
variables that traditionally could not be studied
without penetrating the skin. However, to allow broad
application, several improvements still have to be
made.
Introduction
Surface electromyography (SEMG) is the wellestablished
procedure for the study of muscle function
through electrical signals the muscle emanates, recorded
on the skin surface above this muscle. In the fields of
sports and medicine, this technique offers a great potential
source of information for researchers, trainers and
clinicians. More specific, SEMG is often used to detect 1)
the timing of muscle activation, 2) changes in muscle
activation level due to changes in task, - technique, -
intensity and - duration and 3) localized muscle fatigue.
SEMG also contains information about the force
produced by the muscle, characteristics of active motor
units (MU), and MU recruitment and firing patterns.
Although of high value, due to methodological
difficulties, the latter are much less often used. In
contrast, MU characteristics are often studied in the field
of clinical neurophysiology using invasive needle or wire
electromyography to detect MU abnormalities. Mostly
due to its invasiveness this technique is hardly used in
sports and rehabilitation.
Traditionally, SEMG of one muscle has been
investigated by the application of two electrodes; bipolar
SEMG. Due to the developments in computer science,
nowadays single muscles can be investigated with over 30
and up to 200 electrodes simultaneously. Such multichannel
SEMG (MCSEMG) systems give new
possibilities and different technical problems have to be
solved.
In this contribution we will give an overview of
problems, technical challenges and solutions of using
MCSEMG with the focus on applications in sports and
rehabilitation.
SEMG detection
The quality of any prediction based on SEMG will
ultimately depend on the quality of the SEMG signal
itself. Three main problems have to be dealt with using
SEMG detection; high electrode-skin impedance, powerline
noise and movement artefacts [4]. These sources of
noise can be reduced by adequate skin preparation, the
use of high quality electrodes and cables, and appropriate
recording equipment [1]. Subsequently, further noise
reduction can be obtained by using a variety of filtering
techniques.
With modern, high-input impedance SEMG amplifiers
the electrode-skin impedance becomes less critical, but
one still has to be aware of potential power-line noise and
movement artefacts.
While traditional bipolar SEMG uses electrodes with a
diameter or length of 5 to 10 mm, inter-electrode-distance
(IED) of 1 to 2 cm and often have an electrolytic gel as a
chemical interface between the skin and the metallic part
of the electrode (floating electrodes), MCSEMG
electrodes are usually not larger than 1 mm, with an IED
as small as 3 mm and typically use dry electrodes in direct
contact with the skin. This can cause problematic
electrode-skin impedance, especially since MCSEMG
often aims at quantitative use on MU level.
Besides good electrode-skin contact, it is important
that this contact is fixed. Traditional bipolar SEMG
electrodes are glued to the skin, but electrodes with fixed
IED are hardly commercially available. In contrast, since
MCSEMG electrodes are placed in an electrode grid, they
have fixed IEDs, but they are usually pressed on the skin
by hand or straps around an electrode grid holder.
Recently, an MCSEMG system with floating
electrodes glued on the skin was presented [8]. Although
the quality of SEMG recordings is improved, the long
preparation time limits its applicability.
Another aspect important for signal quality is the
electrode configuration. As stated above, bipolar electrode
arrangement is traditionally used in SEMG and involves a
differential amplifier which suppresses common signals,
i.e., subtracts the two potentials and then amplifies the
difference. Correlated signals common to both sites, such
as distant AC signals from power cords and electrical
devices and SEMG signals from more distant muscles.
Most of the recently published MCSEMG systems use a
monopolar configuration for data collection; recordings
IFMBE Proc. 2005;9: 225

Sports and rehabilitation biomechanics
with a single electrode placed on the skin above the
muscle with respect to a reference electrode. This allows
flexible digital spatial filtering after data collection [9],
but increases the chance of power line interference.
The increased number of electrodes using MCSEMG
doesn’t only increase the demands on the electrodes, but
also on the cables used. Especially using MCSEMG in
dynamic contractions or during whole body movements
(like gait), cables might interfere with the task performed
or movement artefacts can be introduced.
Signal processing:
To get information about muscle force production or
muscle activation level, it is important to get the best
possible estimate of the SEMG amplitude. To quantify the
amplitude, average rectified value and root-mean-square
are most often used. The application of many electrodes
in stead of two, improves the force estimate quality from
SEMG by about 30%, mostly because of increased
electrode surface [10].
Frequency analysis of the SEMG signal during
sustained muscle contractions has been used to indicate
local muscle fatigue, force production, SEMG signal
conduction velocity, muscle fibre type proportion, and
diagnostic classification. The first step towards the
computation of the spectral variables is the estimation of
the power spectral density function of the signal. The
most widely used method for spectral estimation of the
surface SEMG signal is the Fourier transform. However,
even during a sustained voluntary contraction the SEMG
signal is non-stationary although during relatively low
level (20-30% of MVC) and short contractions (20-40
seconds), the SEMG signal may be considered as widesense
stationary, i.e., quasi-stationary. Therefore, to
handle dynamic contractions, time-frequency methods
which do not require stationarity of the signal have been
introduced in SEMG analysis [7]. The additional value of
MCSEMG on the estimation of frequency content has not
been evaluated yet, but similar improvements as to
amplitude estimations can be expected.
The last couple of years, several studies have been
published that show how MCSEMG can be used to
investigate MU characteristics and muscle activation
patterns. The estimate of muscle fibre conduction velocity
is probably the most applied one. The advantage of
MCSEMG over a few bipolar recordings is that muscle
fibre orientation and muscle fibre conduction velocity can
be estimated simultaneously, allowing rapid electrode
placement [3]. Investigation of activation patterns are
mostly based on decomposition of the interference SEMG
signal [2], but also the amplitude distribution over the
skin can be used for this purpose [5]. Although both
method types are applicable, they are far from a stage that
can be considered as finished.
Despite high publication activity on the development
of new MCSEMG analyses techniques, only a few studies
are published that applied these new techniques to answer
questions relevant to sports and rehabilitation and just a
few more related to clinical neurophysiology, at least not
by other authors than the ones that developed the
methods. However, that MCSEMG can positively
contribute to sports and rehabilitation science is supported
by a recent study that could quantify small changes in
activation level due to training [6], while SEMG was
traditionally thought to be too insensitive to do so.
Conclusions
Applying addition channels improves the estimation
of traditional SEMG variables and allows estimation of
variables that traditionally could not be studied without
penetrating the skin. However, to allow this to be applied
in standard studies in sports and rehabilitation; several
improvements still have to be made.
References
[1] CLANCY EA., MORIN EL. AND MERLETTI R. (2002)
Sampling, noise-reduction and amplitude estimation
issues in surface electromyography. J Electromyogr
Kinesiol. 12, 1-16.
[2] FARINA D, MERLETTI R, ENOKA RM. (2004) The
extraction of neural strategies from the surface EMG.
J Appl Physiol.; 96(4):1486-95.
[3] GRÖNLUND C, ÖSTLUND N, ROELEVELD K,
KARLSSON JS. (2005). Simultaneous estimation of
muscle fibre conduction velocity and muscle fibre
orientation using 2D multichannel surface
electromyogram. Med Biol Eng Comput.; 43(1):63-
70.
[4] GRÖNLUND C, ROELEVELD K, HOLTERMANN A,
KARLSSON JS. On-line signal quality estimation of
multichannel surface electromyograms. Med Biol
Eng Comput, in press
[5] HOLTERMANN A, ROELEVELD K, KARLSSON JS.
(2005) Inhomogeneities in muscle activation reveal
motor unit recruitment. J Electromyogr Kinesiol;
15(2): 131-7.
[6] HOLTERMANN A, ROELEVELD K, VEREIJKEN B,
ETTEMA G. Changes in agonist activation level
cannot explain early strength improvement. Eur J
Appl Physiol, in press
[7] KARLSSON S, YU J, AKAY M (2000). Time-frequency
analysis of myoelectric signals during dynamic
contractions: a comparative study. IEEE Trans
Biomed Eng.; 47(2):228-38.
[8] LAPATKI BG, VAN DIJK JP, JONAS IE, ZWARTS MJ,
STEGEMAN DF. (2004) A thin, flexible multielectrode
grid for high-density surface EMG. J Appl Physiol.;
96(1):327-36.
[9] ÖSTLUND N, YU J, ROELEVELD K, KARLSSON JS.
(2004) Adaptive spatial filtering of multichannel
surface electromyogram signals. Med Biol Eng
Comput.; 42(6):825-31.
[10] STAUDENMANN D, KINGMA I, STEGEMAN DF, VAN
DIEEN JH. (2005) Towards optimal multi-channel
EMG electrode configurations in muscle force
estimation: a high density EMG study. J
Electromyogr Kinesiol.; 15(1):1-11.
IFMBE Proc. 2005;9: 226

Sports and rehabilitation biomechanics
MUSCLE ARCHITECTURE AND FIBRE TYPES USING SPATIOTEMPORAL
INFORMATION OF PROPAGATING MOTOR UNIT ACTION POTENTIALS
RECORDED BY 2-D MULTICHANNEL SURFACE EMG
C. Grönlund 1, 2 , K. Roeleveld 3 , S. Karlsson 1, 2
1 Department of Biomedical Engineering and Informatics, R&D, University Hospital, Umeå, Sweden
2 Centre for Biomedical Engineering and Physics, Umeå University, Umeå, Sweden
3 Human Movement Sciences Program, Norwegian University of Science and Technology, Trondheim,
Norway
stefan.karlsson@vll.se
Abstract
The muscle fibre conduction velocity (MFCV) of a motor
unit (MU) is related to its fibre-type and together with
architecture they are the main determinators of muscle
function. In this paper we propose a method to determine
muscle function using 2-D multichannel surface EMG
recordings. A previously developed method detects
propagating MU action potentials (MUAPs), and
estimates their corresponding MFCV, muscle fibre
orientation (MFO), and spatial onset-position of
propagation. In an attempt to separate the estimates of
simultaneously active MUAPs, we propose to examine 2-
D probability distributions of MFCV and MFO estimates,
respectively, against onset-position estimates. The
method was tested on recordings from biceps and
trapezius.
Introduction
Muscle function is dominantly determined by the muscle
fibre-type distribution and architecture. The basic
functional unit of a muscle is the motor unit (MU), which
is defined as a motorneuron and all the muscle fibres that
it innervates. When a muscle fibre is depolarised, a MU
action potential (MUAP) propagates from the endplate to
the tendons with a certain velocity, the muscle fibre
conduction velocity (MFCV). Estimations of MFCV
provides direct physiological information of the muscle
and can be used to investigate fibre types, MU
recruitment, and muscle fatigue. In addition, several
neuromuscular diseases are characterised by a change in
MFCV.
Multichannel surface electromyography (MCSEMG)
recordings using a 2-D electrode grid contain, in
principle, information about muscle architecture as well
as MFCV within the detection volume, i.e., the spatial
amplitude distributions recorded by the electrodes. An
advantage of 2-D MCSEMG, as compared with linear
array MCSEMG, is that it gives information on active
MUs in a large spatial area of a muscle.
In this paper we propose a method to determine muscle
function, in terms of architecture and fibre types, using 2-
D MCSEMG recordings. A new method previously
developed, [1], based on 130-channel surface
electromyography recordings, detects propagating
MUAPs, and estimates their corresponding MFCV,
muscle fibre orientation (MFO), and onset-position of
propagation in the detection volume. However, the higher
the contraction level, the higher the variance of the
estimates. In an attempt to separate the estimates of
simultaneously active MUAPs, we therefore propose to
examine 2-D probability distributions of MFCV and
MFO estimates, respectively, against estimated onsetposition
of propagation.
Methods
Data acquisition and experimental protocol: MCSEMG
data was obtained using a 13 by 10 active electrode-grid
device (a modified ActiveOne, BioSemi, Amsterdam,
Netherlands) with 1.5 mm electrode diameter, 5 mm inter
electrode distance and 2048 Hz sampling frequency. One
subject performed isometric shoulder elevation and
isometric elbow flexion. The electrode-grid device was
placed on the skin above the distal half of the biceps and,
in the middle of the trapezius on the line between the C7
and acromion, respectively.
Data processing: Data analysis was performed off-line
using MATLAB® 6.5 (The MathWorks, Inc., Natick,
USA). Monopolar signals were high-pass filtered using
an eight-order Butterworth filter with a cut-off frequency
of 10 Hz, before data was bipolarly spatial filtered.
The experimental data was processed using the algorithm
proposed by Grönlund et al [1]. The method detects
individual propagating MUAPs as moving components in
the detection volume. The detection results in trajectories
which describes the electrode positions (row and column)
closest to the main peak for the individual MUAPs
throughout the propagation (over time). Next, MFCV,
muscle fibre orientation (MFO) and onset-position of
propagation (x,y) is simultaneously estimated by
matching a surface template to the spatial amplitude
distributions of the 3 x 3 closest electrodes of the MUAP
trajectories. The coordinate system’s x and y axes where
aligned with the electrode-grid’s 10 and 13 electrode
directions, respectively (figure 1).
Since the amplitude distribution at the skin’s surface is
related to the sum of all active MUAPs, the spatial
IFMBE Proc. 2005;9: 227

Sports and rehabilitation biomechanics
amplitude distribution of individual propagating MUAPs
are contaminated with other active MUAPs. In order
separate the estimates of simultaneously active MUAPs,
we propose to examine 2-D probability distributions of
the MFCV and MFO, respectively, against the
corresponding x-position estimates. This is concomitant
to separate the estimates of simultaneously active
MUAPs using their estimated positions in the detection
volume.
The 2-D probability distributions were calculated by
smoothing 2-D histograms. The 2-D histograms were
calculated using 100 by 100 bins in non-overlapping
time-windows of 15 s from the estimates of MFCV and
MFO, respectively, against position. Time window length
was set on empirical basis to obtain averaged
distributions and such that no fatigue manifestation
should occur. The smoothing procedure was based on a
non-parametric (no a priori distribution is assumed) curve
estimation [2] were the histograms are convoluted with a
Gaussian kernel. By normalisation to unit volume the
smoothed distributions are equal to probability
distributions.
Results
Figure 1 presents results from a biceps (I, first column of
the figure) and a trapezius measurement (II, second
column).
time-window of 15 s. The peaks of the 2-D distributions
are the estimates of probable MU populations.
First, the electrode positions and the corresponding
detection system’s co-ordinates for both measurements
are presented. Then the estimated trajectories of the
propagating MUAPs are plotted as lines on a map of the
root-mean-square values of each channel. The last two
figures show the 2-D distributions of MFCV and MFO,
respectively against onset-position. The peaks of the
distributions correspond to high probabilities for
estimates of different MU populations.
Discussion
In this paper we proposed a method to determine muscle
function, in terms of muscle fibre-types and architecture,
using 2-D MCSEMG recordings. Different populations of
MUs were revealed using 2-D distributions of estimates
of MFCV and MFO, respectively, against spatial onsetposition
of propagating MUAPs.
For simplicity and possible visual examination, 2-D
distributions of estimates were used, however,
multidimensional probability distributions could also be
used. In this way, possibly the estimates of the active
MUs can be separated even more.
A limitation of the technique is that the MCSEMG data
needs to be spatially filtered, in which activity of distant
MUs are suppressed, and superficial MUAPs are
enhanced. However, using decomposition, possibly a
larger set of MUs can be detected. The estimation
procedure could therefore be performed on averaged
monopolar MUAPs.
References
[1] Grönlund C, Östlund N, Roeleveld K, Karlsson JS
(2005) : 'Simultaneous estimation of muscle fibre
conduction velocity and muscle fibre orientation using 2-
D multichannel surface electromyogram', Med Biol Eng
Comp, 43, pp. 63-70.
[2] Fan J and Marron JS (1994): ‘Fast implementations of
non-parametric curve estimators’, J Comp Graph Stat, 3,
pp. 35-56
Acknowledgements
This study was supported by the European Union
Regional Development Fund.
Figure 1: Results from a biceps (I) and a trapezius
measurement (II) at 25 % MVC. The estimated
trajectories of the propagating MUAPs are plotted as
lines on a map of the root-mean-square values of each
channel. The results are based on detected MUAPs in a
IFMBE Proc. 2005;9: 228

Sports and rehabilitation biomechanics
GLOBAL MUSCLE ACTIVATION IN SUSTAINED CONTRACTIONS
Andreas Holtermann, Karin Roeleveld
Human Movement Sciences Program, Norwegian University of Science and Technology,
Trondheim, Norway.
E-mail: andreaho@svt.ntnu.no
Abstract: This study examines and compares the
global activation changes in a sustained
contraction with a ramp contraction in the upper
trapezius muscle. The global activation pattern
was equivalent to increased excitatory drive in
both contractions. The findings confirm
fundamental principles of motor unit activation
based on limited motor unit samples.
Introduction
Although recruitment of motor units in a
size-ranked order has been demonstrated in submaximal
sustained contractions (Adam and De Luca
2003), divergent observations from the orderly
recruitment scheme have also been reported (Westad
et al. 2003).
Even though the recruited motor units
throughout a sustained contraction are thought to be
due to an increased central drive, the discharge rate
is observed to remain stable or even decrease in
sustained contractions (Bigland-Ritchie et al. 1986).
The above mentioned principles have
primarily been investigated with needle or wire
electrodes or non-invasive multi-channel
decomposition techniques. However, in both
techniques, the recorded information is limited to a
local area of the muscle comprising a restricted
sample of detected motor units.
To be able to obtain global information of
the activation pattern of a muscle, the changes in
spatial amplitude distribution in a sustained
contraction were quantified and compared with
changes in a ramp contraction. The main aim of this
study was to examine the activation changes in the
upper trapezius muscle in a sustained contraction.
Method
Data from 14 subjects (10 women and 4
men) was used in the study. A Biodex dynamometer
was used to standardize position and record force
generation. The force was generated with the upper
trapezius muscle by isometric shoulder elevation.
The subjects received direct feedback of the
generated torque from a monitor.
Surface EMG data was acquired with a
multi-channel system consisting of a 130-channel
(13 x10) grid with an inter-electrode distance of
5mm. The center of the MCSEMG grid was placed
on the right trapezius muscle in the middle of the
line between the seventh cervical vertebra and the
acromion.
The experiment consisted of two separate
contractions. In the first contraction, the subject
generated three isometric ramp contractions from 0
to 90 % of maximal voluntary contraction (MVC).
Each contraction lasted 10 seconds. Subsequently,
the subjects performed a sustained 3 minutes lasting
isometric contraction at 25 % of MVC.
The EMG channels with low signal quality
were automatically identified (Grønlund et al. in
press), and omitted from further processing. Root
mean square (RMS) and median frequency (MF) of
bipolar MCSEMG leadings was calculated in epochs
of 500 ms. In order to quantify RMS distribution
changes, correlation coefficients were calculated
between RMS values of all electrode pairs at one
time epoch, with the RMS values of the same
electrode pairs at another epoch. Correlation
coefficients were obtained for all possible
combinations within the recording period of the
ramp contraction and the sustained contraction
respectively (Holtermann et al. 2005). In addition, to
compare the change in muscle activation during the
ramp and sustained contractions, correlation
coefficients were calculated between the RMS
values of each epoch of the sustained contraction
with each epoch of the ramp contraction.
Subsequently, throughout each epoch of the
sustained contraction, the epoch with peak
correlation from the ramp contraction was
determined.
Results
In the isometric ramp contraction, the
subjects generated a force from 0 to 87 % of MVC
with a similar change in RMS and RMS correlation.
Five subjects did not increase RMS throughout the
sustained contraction (p=0.35), and neither
significantly changed MF (p=0.08), RMS correlation
(p=0.4), the epoch with peak RMS correlation
(p=0.5), or peak RMS correlation (p=0.06). Nine
subjects increased RMS (p

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Sports and rehabilitation biomechanics
force levels (p

Sports and rehabilitation biomechanics
MYOFEEDBACK ISSUES IN REHABILITATION OF WORK-
RELATED MUSCULAR PAIN
Leif Sandsjö
National Institute for Working Life, Göteborg, Sweden
leif.sandsjo@arbetslivsinstitutet.se
Abstract: Myofeedback training during regular
work at the workplace has potential of becoming a
powerful tool in prevention and rehabilitation of
work-related muscle pain. A system for practical
use must be fully self-administered, leaving
complete control of where and when to use the
system to the user. It must also provide feedback
signals easy to understand and only in situations
relevant to the development of muscle pain. This
presentation briefly describes a myofeedback
system for training at the workplace and discusses
issues regarding when and how an alarm should
be presented to the user.
Introduction
Biofeedback has been defined by Birk [1] as “the
use of monitoring instruments (usually electrical) to
detect and amplify internal physiological processes
within the body, in order to make this ordinarily
unavailable internal information available to the
individual and literally feed it back to him in some
form”. Biofeedback based on muscle activity, often
referred to as myofeedback, has been successfully
applied at workplaces for individual training of work
technique and to increase awareness about working
situations involving potentially harmful muscle
activation, e.g. [2]. This has been accomplished by
means of presenting an alarm to the worker as soon as
the muscle activity exceeds a threshold level
indicating a muscle activation of concern.
However, this approach may be too crude in
many situations as the alarm criterion not primarily
includes the duration and/or variation of the muscle
activation, and, striving for lowered muscle activity in
general may be contra productive as it may lead to a
decrease also in dynamic activity beneficial to the
muscle.
An alternative feedback approach has been
presented based on the Cinderella hypothesis [3,4].
This hypothesis states that a stereotyped recruitment
order of the motor units of the muscle leads to the
same motor units always being active when the
muscle is activated [5]. The consequence is that the
muscle must be completely relaxed in order to relieve
the motor units first recruited. The alternative
feedback approach allows short periods of high
muscle activation as long as the user manages to relax
the muscle in-between [4]. This presentation describes
the technique and discusses issues of concern with
respect to time aspects and threshold level settings for
alarm presentation using this method.
Methods
The system is designed to be used during regular
work at the workplace by subjects suffering from workrelated
neck/shoulder pain. It is fully self-administered,
i.e. easy to don and doff, and requires no calibration
procedure. A harness assists the user to position the
built-in dry electrodes over the same part of the
trapezius muscle each time the system is used. A
control unit records the myoelectric signal and produces
a vibration signal when the chosen alarm criterion is
met [6].
These properties not only make large-scale training
at the workplace during working situations that may
have led to the work-related musculoskeletal problems
possible, but the on-site training is also motivated by
the fact that learned behaviour is often coupled to the
learning environment and/or the situation. Further, this
type of myofeedback training at the worksite can be
done with little or no productivity losses to the
employer.
Apart from providing alarms to the user in
situations of insufficient muscle rest the system also
records the muscle activity patterns. These patterns can
then be used in discussions with an ergonomist or
physiotherapist to identify specific work tasks of
concern and possible alternative ways to perform the
work. This type of dialogue further increases the benefit
of the system as it helps the user to discern between
good and bad working practice.
Discussion
The new type of biofeedback system has been
applied in a first pilot study. The study indicated that
myofeedback based on absence of sufficient muscle rest
can lead to improved muscle activation patterns and
thereby also influence the experience of pain [4].
A general issue regarding biofeedback methods
applied at the workplace is that it operates on the
individual level with problems that may have its origins
at the organisational level. Before applying the method
at a workplace or specific workstation it is crucial that
IFMBE Proc. 2005;9: 231

Sports and rehabilitation biomechanics
the existing situation allows for changes that may be
introduced by the biofeedback method. This includes
not only time and space constraints at each
workstation but also the organisational level as
necessary changes in work organisation may be
prompted by the biofeedback intervention.
More specific issues concerns the way the alarm
is presented to the user to promote less demanding
ways of performing the work. As the neuromuscular
system of each individual has been finely tuned
during its formative years, the feedback signal needs
to be experienced on an intuitive rather than
intellectual level in order to bring change that lasts
beyond the actual use of the system, i.e. changed
behaviour. Apart from the quality of the signal itself
(visual, aural or tactile) the intuitiveness of the
feedback signal of the new method is determined by
the alarm evaluation period and the presentation
period. The evaluation period is the time frame for
evaluation of muscle rest, i.e. if the amount of muscle
rest has been insufficient within the evaluation
period, an alarm situation is at hand and an alarm is
presented. The current system allows the evaluation
period to be set according to the work being
performed. A long evaluation period makes it more
difficult for the user to connect an alarm to the
activity of concern, i.e. less intuitive, while a too
short evaluation period (~seconds) will result in
feedback given from a muscle activity perspective
rather than the amount of muscle rest. (An evaluation
period of zero seconds makes the muscle rest time
evaluation identical to the amplitude level criteria
traditionally used.) In a study comparing 5-, 10- and
20-second evaluation intervals, the 10-second interval
was found to result in the highest level of muscular
rest time [7].
The presentation period, i.e. the time the alarm is
active, is likely even more informative than the onset
of the alarm as the action taken by the user to make
the alarm to go off will be associated with a
successful behaviour to reduce harmful muscle
activations. The most straightforward, and maybe
also the most intuitive, is to simply end the alarm as
soon as the muscle activity is below the threshold
level chosen to discriminate between activity and
relaxation [3].
Finally, the activity threshold level used to
discriminate between the muscle at rest and the active
muscle, can be discussed. From a user-oriented point
of view, an absolute or fixed threshold level is
preferred as it does not require the user to do anything
more than to put the system on and start using it.
However, from a methodological standpoint this is
not optimal as the recording conditions will differ
slightly each time the system is used due to the small
variations electrode positions and the electrode-toskin
contact obtained each time the system is applied.
Under bad recording situations these problems may
lead to alarms being issued due to a noisy recording
rather than muscle activation patterns of concern. An
alternative approach is to use a threshold level
determined from the activity recorded when the subject
is trying to relax. Muscle rest time evaluation based on
this method has shown significant differences between
workers reporting pain compared to their pain-free
colleagues [8]. Although this method ascertains a
threshold level that is above the level recorded during
attempted relaxation and thus not subject to the abovementioned
problem, it is at the expense of the user to
perform calibration activities each time the system is
applied.
Conclusion
Myofeedback training at the workplace has large
potential in prevention and rehabilitation of workrelated
musculoskeletal disorders. New methods based
on the amount of muscle rest during work look
promising. Further research involving the behavioural,
physiological as well as technical disciplines is needed
to present feedback signals to the user that are
unambiguous and effective in learning new, less
harmful, work patterns.
References
[1] BIRK L. (1973): ‘Biofeedback: Behavioural
Medicine’, (Grune & Stratton, New York)
[2] NORD S., ETTARE D., DREW D., HODGE, S. (2001):
‘Muscle learning therapy – Efficacy of a
biofeedback based protocol in treating workrelated
upper extremity disorders’, J o Occup
Rehab., 11, pp. 23-31
[3] HÄGG GM. (1997): ‘New sEMG feed-back
principle for gap training’, Proc. of the second
general SENIAM workshop, Stockholm, Sweden,
June 1997, pp. 88-91
[4] HERMENS HJ., HUTTEN MMR. (2002): ‘Muscle
activation in chronic pain: Its treatment using a
new approach of myofeedback’, Int J Ind Erg., 30,
pp. 325-336
[5] HÄGG GM. (1991): ‘Static work and myalgia – A
new explanation model’, Electromyographical
kinesiology, Amsterdam (Elsevier Science), pp.
141-44
[6] Twente Medical Systems International, Internet
site address: http://www.tmsi.com
[7] VOERMAN GE., SANDSJÖ L., VOELLENBROEK-
HUTTEN MMR., GROOTHUIS-OUDSHOORN CGM.,
HERMENS HJ. (2004): ‘The influence of different
intermittent myofeedback training schedules on
learning relaxation of the trapezius muscle while
performing a gross-motor task’, Eur J Appl
Physiol., 93, pp. 57-64
[8] SANDSJÖ L., MELIN B., RISSÉN D., DOHNS I.,
LUNDBERG U. (2000): ‘Trapezius muscle activity,
neck and shoulder pain, and subjective experiences
during monotonous work in women’, Eur J Appl
Physiol., 83(2-3), pp. 235-238
IFMBE Proc. 2005;9: 232

Sports and rehabilitation biomechanics
EFFECT OF INTERELECTRODE DISTANCE ON MEASURING
SENSITIVITY FOR FACIAL EMG
M. Puurtinen, J. Hyttinen and J. Malmivuo
Ragnar Granit Institute, Tampere University of Technology, Tampere, Finland
merja.puurtinen@tut.fi
Abstract: New miniaturized portable biopotential
measuring devices may require reduced electrode
size and distance. This paper introduces a study
where the effect of reducing interelectrode distance
(IED) to the measuring sensitivity for facial EMG
was modeled. The modeling study was based on 2D
finite difference method (FDM) model that
represented the human forehead. The results
indicate that that IED does not significantly affect
the measuring depth, but it has a greater influence
on measuring width.
Introduction
In recent years, the advances in technology have
brought about portable and wireless systems for
measuring biopotential signals, and this has been
particularly true in ECG and EMG measurements. This
study was conducted in a project called “Wireless
technology and psychophysiological computing”, where
wireless systems for psychophysiological signal
monitoring are being developed. The objective of the
project is to build small and unobtrusive measuring
systems for monitoring human physiological signals.
The aim of the project is to minimize the size of the
measuring unit. For this reason, the objective of this
paper is to model, how reducing the interelectrode
distance (IED) and electrode size affects bioelectric
signal strength obtained from the human forehead
muscles.
It has been reported in literature that decreasing the
interelectrode distance decreases the measuring depth
and decreasing the electrode dimensions increases the
noise from electrodes [1, 2]. A few studies on the effect
of these factors exist, but they are mainly concentrated
on EMG applications on larger muscles [3-6]. The
modeling study presented in this paper was based on 2D
finite difference method (FDM) model that represented
the human forehead.
Materials and Methods
Modeling study was based on Finite Difference
Method (FDM), in which the model composes of cubic
elements and connective nodes forming a resistor
network [7]. The model used in this study was a
forehead model including 6 tissue layers (Figure 1.).
This model is an approximation of the anatomy of the
human forehead, and it was created for illustrating the
behavior of electric fields in the facial muscle layers.
The size of the model is 50x9 mm (500x90 pixels) and
it is composed of two slices and 135000 nodes. The
model includes 6 layers: 0.1 mm of “dry” highly
resistive skin layer, 0.9 mm of lower skin layer, 2 mm
of muscle layer, 1 mm of cortical bone layer, 4 mm of
cancellous bone layer, and 1 mm of cortical bone layer.
The layer thicknesses were determined as an average
according to anatomical atlases and several MRI and CT
images, as no certain values for forehead tissue layer
thicknesses have been published, and the variation is
quite high among individuals. The resistivity values for
the forehead models were adopted from the research
conducted by Gabriel et al. [8], and from the torso
model by Kauppinen [9].
In the modeling study, the calculation of the
sensitivity of the electrodes to measure muscles electric
activity was based on the reciprocity theorem: a
reciprocal current was applied on a pair of electrodes
located on the model, and the resulting lead field in the
muscle was calculated. This field in the muscle
represents the leads i.e. the measuring electrodes’
sensitivity to measure the electric source of the muscle.
The measuring sensitivity was modeled with different
IEDs in order to see how the IED affects the measuring
sensitivity. [10]
Figure 1: Forehead model composed of 6 tissue layers.
Results
Figures 2. and 3. illustrate the behavior of the lead
field in the 6 layered forehead model. From these
figures, it can be seen that the inhomogenities of the
model affect the behavior of the lead field. As the
muscle layer is less resistive than the skin layer, most
current passes through the muscle. Moreover, the
current chooses this path even if the IED is reduced. As
a consequence, decreasing the IED has only a minor
effect on the measuring depth when detecting the
activity of facial muscles. However, the width of the
detected area is directly proportional to the IED.
The results are further illustrated in Figure 4., which
shows how the magnitude of the lead field changes in
the muscle layer as a function of width. The point 25 in
IFMBE Proc. 2005;9: 233

Sports and rehabilitation biomechanics
width corresponds to the location in the middle of the
two electrodes. It can be seen that in the middle of the
electrodes the measuring sensitivity stays practically the
same regardless of the IED. Yet, the width of the
measured area increases as the IED is increased.
muscle. Nonetheless, the IED has a greater influence on
the measuring width.
The results indicate that if the electrodes can be
placed over the desired muscle even very closely
separated electrodes can be used. This enables to design
extremely small devices for facial EMG.
References
Figure 2. Resiprocal current or lead field in the forehead
model with an IED of 12 mm. The size of the 2D model
is 20 mm x 9 mm.
Figure 3. Resiprocal current or lead field in the forehead
model with an IED of 6 mm. The size of the 2D model
is 20 mm x 9 mm.
Figure 4. Current i.e. the lead field in the muscle layer
of the forehead model as a function of width with
different IED's.
Conclusions
In this study, the effect of changing the IED on
measuring sensitivity of facial EMG was studied with
FDM modeling. The results indicate that IED does not
significantly affect the measuring depth and that in the
middle of the electrodes the measuring sensitivity stays
practically same when changing the IED. This is
because irrespective of the IED, the measuring
sensitivity concentrates on the less resistive path, i.e. the
[1] GRIMNES S. and MARTINSEN O. G. (2000):
'Bioimpedance and Bioelectricity Basics',
(Academic Press, New York)
[2] HUIGEN E., PEPER A. and GRIMBERGEN C. A.
(2002): 'Investigation into the origin of the noise of
surface electrodes', Medical & Biological
Engineering & Computing, 40, pp. 332-338
[3] ELFVING B., LILJEQUIST D., MATTSSON E.
and NEMETH G. (2002): 'Influence of
interelectrode distance and force level on the
spectral parameters of surface electromyographic
recordings from the lumbar muscles', Journal of
Electromyography and Kinesiology, 12, pp. 295-304
[4] FARINA D., CESCON C. and MERLETTI R.
(2002): 'Influence of anatomical, physical, and
detection-system parameters on surface EMG',
Biological Cybernetics, 86, pp. 445-456
[5] ROSENBURG R. and SEIDEL H. (1989):
'Electromyography of lumbar erector spinae
muscles--influence of posture, interelectrode
distance, strength, and fatigue', European Journal of
Applied Physiology and Occupational Physiology,
59, pp. 104-114
[6] ZEDKA M., KUMAR S. and NARAYAN Y.
(1997): 'Comparison of surface EMG signals
between electrode types, interelectrode distances and
electrode orientations in isometric exercise of the
erector spinae muscle', Electromyography and
Clinical Neurophysiology, 37, pp. 439-447
[7] JOHNSON J. (1995): 'Numerical methods for
bioelectric field problems', in J. Bronzino, Ed. 'The
Biomedical Engineering Handbook', (CRC Press,
Boca Rato (FL))
[8] Compilation of the Dielectric Properties of Body
Tissues at RF and Microwave Frequencies,
Internet site address:
[9] KAUPPINEN P., HYTTINEN J., HEINONEN T.
and MALMIVUO J. (1998): 'Detailed model of the
thorax as a volume conductor based on the visible
human man data', Journal of Medical Engineering &
Technology, 22, pp. 126-133
[10] MALMIVUO J. and PLONSEY R. (1995):
‘Bioelectromagnetism: Principles and Applications
of Bioelectric and Biomagnetic Fields’, (Oxford
University Press, New York)
IFMBE Proc. 2005;9: 234

Sports and rehabilitation biomechanics
NEAR INFRARED SPECTROSCOPY FOR MEASURING MUSCLE
OXYGENATION
Albert G. Crenshaw*, Bente R. Jensen**
*Centre for Musculoskeletal Research, University of Gävle, Umeå, Sweden
**Department of Human Physiology, Institute for Exercise and Sport Sciences, August Krogh
Institute, University of Copenhagen, Denmark
E-mail: albert.crenshaw@hig.se
Abstract: The advent of near infrared
spectroscopy (NIRS) as a tool for monitoring
muscle oxygenation has allowed for important
physiological data in sports training and
rehabilitation. Commercial methods are
generally user-friendly and the technique is noninvasive.
By projecting a light beam into the
muscle concentrations of
haemoglobin/myoglobin (with and without
oxygen) in the vascular bed consisting of small
arterioles, capillaries and venules can be
determined. Despite its appeal methodological
improvements to account for varying muscle
depths, and to distinguish between arteriolar
and venular contributions separately are
desired.
Background
Since the pioneering work of Jobsis (1977)
nearly 30 years ago, near infrared spectroscopy
(NIRS) has emerged as an important tool for a wide
variety of medical and research applications. This is
evidenced by the vast number of papers published
over the last two decades. The application of NIRS
for exercise training and sport was exemplified in a
recent review article (Neary 2004). The fact that
NIRS is non-invasive and it provides information
on the hemodynamic states of muscles fuels its
appeal.
Principle of operation
In principle, an optical probe consisting of two
optodes – an emitter and a detector, is placed on the
skin over the muscle of interest. The emitter
projects a visible light beam in the near-infrared
region (i.e. 600-900 nm) that penetrates into the
deeper tissues. The detector receives the light from
the muscle. The distance between the emitter and
detector determines the depth of penetration – a
general rule of thumb is that the light will penetrate
to a maximum depth equivalent to 90% of the
distance between optodes, and that the average
measurement depth is equivalent to half the
distance between optodes. As the light traverses
through the muscle tissue it is either scattered or
absorbed. In the near infrared range the strongest
absorbers are oxy-haemoglobin (HbO 2 ) and
deoxyhaemoglobin (Hb); myoglobin also absorbs
but only to a minor degree (less than 10%). The
following algorithm encorporating the modified
Beer-Lambert law allows analysis of these
substances:
A = αcdB + G
where A is the measured light attenuation, α is the
specific extinction coefficient of the absorbing
compound, c is the concentration of the absorbing
compound, d is the distance between the optodes, B
is the differential pathlength factor, and G is a
constant reflecting the scattering of light in the
tissue.
Thus, the detected light provides information about
the dynamics of HbO 2 and Hb during, for example,
an experimental maneuver. For many applications
the outcome measurement is expressed as the
percent saturation (i.e. % of haemoglobin that
contains oxygen). The equation for this is:
[HbO 2 /(Hb + HbO 2 )] x 100.
In addition, the total haemoglobin content (Hb +
HbO 2 ) provides useful information about blood
volume changes.
NIRS measurements express the dynamic
balance between oxygen supply and consumption in
the volume of tissue beneath the probe. This
volume is in part determined by the distance
between the emitting and detecting optodes of the
probe used as explained above. The signals arise
mainly from small vessels, i.e. arterioles. capillaries
and venules deep within the muscle because in
larger vessels NIR light is fully absorbed by the
high haemoglobin concentration. Currently, a
distinction between arteriolar and venular
contributions to the measured signal cannot be
determined, thus warranting further development of
the method for this purpose. NIRS measurements
IFMBE Proc. 2005;9: 235

Sports and rehabilitation biomechanics
determined, thus warranting further development of
the method for this purpose. NIRS measurements
can be influenced by subcutaneous fat and tissue
temperature, amongst others. Therefore these
factors should be accounted for when recording.
Commercial devices
There are a number of different types of devices
for NIRS measurements. These are briefly
described here but a more detailed description is
given in the paper by Myers and co-workers (2005).
Continuous wave spectrometers measure changes in
the attenuation of a few or several wavelengths of
light due to changing concentrations of substances
(see Beer-Lambert law above). Time resolved
spectrometers use pulse lasers and resolve the
amount of time that launched protons remain in the
tissue. Phase resolved spectrometers modulate the
intensity of emitted light in determining the
distance photons travel within the tissue. Spatially
resolve spectroscopy measures an attenuated light
signal at multiple probe spacing distances. The
current system of choice for the contributing
authors of the present paper is a continuous wave
device that employs four wavelengths – 680, 720,
760 and 800 nm, and incorporates a scaled second
derivative absorbance spectrum (INSPECTRA
Tissue Spectrometer – model 325, Hutchinson
Technology Inc., The Netherlands) in determining
absolute tissue percent saturation values.
Previous NIRS experiences of present authors
The authors of the present paper independently
have recent experiences with NIRS measurements.
Jensen and co-workers (1999) combined NIRS
measurements with electromyography (EMG) and
intramuscular pressure measurements for assessing
muscle load and haemodynamics of the
paravetebral muscles. The study showed that a 20%
maximum voluntary contraction corresponded to
intramuscular pressure of 30-40 mmHg, and that
these were associated with a significant drop in
tissue oxygenation (measured with Runman; NIM,
Philadelphia). Heiden and co-workers (2005) used
the Inspectra tissue spectrometer to measure
oxygenation of the forearm extensor muscles in
comparing two different modes of computer mouse
use, i.e. with and without time pressure and
precision imposed. A significant drop in
oxygenation was found when imposing such
restraints that were not observed when not
imposing them. Furthermore males exhibited higher
oxygenation values than females both before and
throughout the work. These finding have
implication in investigating mechanisms of
musculoskeletal disorders in computer users.
distribution of oxygenation, in combination with
EMG, of the vastus lateralis muscle during various
knee extension activities. For this study the
Inspectra tissue spectrometer is employed and two
probes lengths, 25-mm versus 35-mm, are used that
are placed alternately along the muscle length.
Owing to that for a given isometric contraction the
architectural pennation angle of the distal muscle is
greater than that for the proximal, and that the 35-
mm probe measures deeper in the muscle than the
25-mm, it can be speculated that muscle
oxygenation would be more affected at the distal
and deeper areas due to the generation of higher
respective intramuscular pressures. This hypothesis
is being tested. While the current study intends to
provide insight into muscular oxygenation for the
vastus lateralis as a function of depth, more studies
to see whether this might be muscle dependent are
suggested. Additionally, further development of the
method to distinguish between arteriolar and
venular contributions separately to NIRS signals are
desired.
References
Neary JP (2004). Application of Near Infrared
Spectroscopy to Exercise Sports Science. Can
J Appl Physiol 29: 488-503.
Jobsis FF (1977). Noninvasive, infrared monitoring
of cebral and myocardial oxygen sufficiency
and circulatory parameters. Science 198:1264-
1267.
Myers DE, Cooper CE, Beilman GJ, Mowlem, JD,
Anderson LD, Seifert RP, Ortner JP. (2005) A
non-invasive method for measuring local
haemoglobin oxygen saturation in tissue using
wide gap second derivative near-infrared
spectroscopy. J Biomed Opt (In press)
Jensen BR, Jörgensen K, Hargens AR, Nielsen PK,
Nicolaisen T. (1999). Physiological response to
submaximal isometric contractions of the
paravertebral muscles. Spine 24: 2332-2338.
Heiden M, Lyskov E, Djupsjöbacka M, Hellström
F, Crenshaw AG. (2005) Effects of time
pressure and precision demands during
computer mouse work on muscle oxygenation
and position sense. Eur J Appl Physio, (In
press)
Ongoing project
Currently the present authors are working
together on a project to determine spatial
IFMBE Proc. 2005;9: 236

Sports and rehabilitation biomechanics
APPLICATION OF RECIPROCAL ORTHOTIC SYSTEMS WITH
ELECTRICAL STIMULATION OF MUSCLES IN CHILDREN WITH SPINAL
DISEASES
E. Dukendjiev 1 , V. Mihnovich 1
1 Division of Mechanics of Prosthetics, Riga Technical University, Riga, Latvia
apl@stradini.lv
Abstract
There is discussed rehabilitation of a 10 year old patient
with progressive Duchenne dystrophy of muscles by
means of reciprocative orthesic system and standard two
channel electric stimulation of muscles in reciprocative
regime.
Introduction
Authors have elaborated a reciprocative orthesic system,
allowing rehabilitation of children with spinal diseases at
home, with the aid of parents and standard electric
stimulator ELPHA 2000 [3]. Artificial excitation of
muscles-antagonists of a lower limb in corresponding
phases together with a reciprocative orthesic system
(ROS) facilitates recovery, approaches the norm of
stereotype walking and recovery of locomotive ability.
[1].
Methods
Kinematical system of Dukendjiev’s reciprocative system
[2] in this case is simplified because of decreased number
of connections of mobility degrees caused by dynamic
correspondence with the patient. ROS integrates remains
of muscular activity for accomplishment of movement
and due to multilink structure of locomotive system uses
the effect of “non-straight” work of muscles, allowing
redistribution of kinematical moments between adjacent
parts of body.
Electric stimulation of muscles was carried out with
standard equipment which had two aims – rehabilitation
of neuron networks of control starting from the lower
towards the upper hierarchical structures and to obtain
additional muscle strength. The novelty of the approach
is connected with creation of specular pairs of electrodes
and synthesis of programmes based on reciprocative
principle. The programme determines phases of
stimulation of muscles during a step in correspondence
with the location of ROS segments (Fig.1.). Thus two
channels work in opposite phase and stimulate four
muscles. Synchronisation of electric stimulation and
operation of ROS is carried out by means of location
sensors of one of the joints of legs for each limb.
Results
ROS was created for the patient EN (Fig. 2) aged 10,
suffering from progressive Duchenne dystrophy of
muscles. First signs appeared at the age of 4 with atrophy
of pelvic and femoral muscles, there was observed
movement clumsiness, instability, frequent flounder and
falls during the walk, marked locomotive passiveness,
unwillingness to walk caused by the fear of falling and
rapid exhausting. Gradually atrophied muscles of
shoulders and arms, walk was disturbed, later on there
appeared difficulty of movements. The patient has
preserved satisfactory physical condition up to 8 years
and only after acute pneumonia has lost the ability of
unaided movements. By 9 years of age the patient
developed muscular contractures of pelvic, knee and
talocrural joints. While the process is progressing,
equinovarus deformation of feet is developing.
Figure 1: The Sheme of reciprocal ESM
The patient has retained muscular reaction to electric
signal. With 4-6 [mA] there appears sensitivity, with 9-12
[mA] isometric tension of muscles is observed and with
increasing of stimulating signal up to 15-17 [mA] visible
changes in position of joints appear.
For this patient there is used two channel stimulation of
muscles unbending and bending hips (unbending – big
IFMBE Proc. 2005;9: 237

Sports and rehabilitation biomechanics
sciatic muscle, bending – straight muscle of the hip). The
latter participates also in unbending of knee joint.
However walking in ROS is done with closed lock,
preventing moments in the knee. There remains only
function of bending the hip. Knee and reciprocative locks
are opened for sitting.
Figure 3: Sheme of the innervation on segmental level of
CNS
Figure 2: Patient in the Reciprocal Orthotic System with
electrical stimulation
ROS is put on the patient, all caps of cases are fixed,
electrodes of a standard electric stimulator are applied to
muscles to be stimulated so that various channels
correspond to the muscles antagonists of the limb (Fig.3.)
and parameters of the programme (duration of
stimulation – 1.0-1.5 [s], duration of relaxation 2.0-2.5
[s], offset of the phase between the channels 1.5-2.0 [s])
are set. Then the support is lifted to the angle of 45-60
degrees and the patient is put on his legs. Walking is
carried out with the help of an assistant, which helps the
patient to keep his balance and bend the body in the
direction opposite to the forward moving leg. The patient
moving his hands opposite his legs transfers to the latter
strain by means of an integrator of muscle strains.
Discussion
Along with the motor zone of the shell, initiating the
pyramidal system, other shell areas, parts of CNS, joined
in an extrapyramidal system also has a big role in
formation of wilful movements [4]. Thus the initial
locomotive commands contain only the main
characteristics of the coming movement, supplemented
by the necessary details as the activity of other nervous
centres is added. Thus on a segmental level in the
programme of movement structure of the lower limb
there is added the principle of reciprocative slowing of
muscles antagonists of the same limb and the muscles of
the corresponding group of different limbs (Fig.3.).
Conclusions
It is shown that in children with spinal diseases, ROS
with functional electric stimulation have several positive
changes. Functional characteristics of their muscles
increase – their strength, maximal electric activity is
increased and blood circulation is improved;
biomechanical and inertial structure of walking in
improved – they approach to norm, as well as the image
of movements in main joints of legs.
There is observed increasing of stability and big
endurance during walking, decreased use of support on
crutches and a stick.
The meaning of the described method in theory and
practice of prosthetics should be stressed in particular.
Joint use of reciprocative orthesic systems and functional
electric stimulation allows achieving the effect on
absolutely different level. Electric stimulation by means
of ROS provides additional energetic source of
movement, i.e., increased efficiency factor of muscles
and creates circumstances for normal work of nervousmuscular
system of children with spinal diseases.
References
[1] Dukendjiev E, Mihnovich V. New Approach to the
Synthesis of Orthesis Systems for Spinal Patients. The
11 th World Congress of the International Society for
Prosthetics & OrthoticsAugust 1-6, 2004, Hong Kong.
[2] Dukendjiev E. Orthesis pulling system for patients
with paralysis, paresis, defects, amputations. Proceedings
of the 12th Nordic-Baltic Conference on Biomedical
Engineering. Reykjavik, 2002. pp.218-219.
[3] http://www.danmeter.dk.
[4] Aivars Yu, Dukendjiev E., Mihnovich V. Joint active
interdependent participation of elements of system
“extremity - prosthesis/orthosis" during realization of the
movements.Proc.2 th Baltic-Bulgarian Conference on
Bionics, Biomechanics and Mechanics, Varna, Bulgaria,
4-6.06.2001, 13-16 pp.
IFMBE Proc. 2005;9: 238

Sports and rehabilitation biomechanics
COMPARING DYNAMICS OF SKI JUMPING AND DRY IMITATION
JUMPS BY COMPUTER SIMULATION
G.J. Ettema* and S. Bråten**
* Human Movement Sciences Programme, NTNU, Trondheim, Norway
**Olympiatoppen, Trondheim, Norway
gertjan.ettema@svt.ntnu.no
Abstract: Dry imitation jumps are an essential part
of training in ski jumping. Little is understood about
the effects of the different dynamics in the jumping
hill and in the training hall. By computer simulation
we compared the two conditions while the jumps
were optimised for angular momentum at toe-off.
Dynamics, kinematics, and muscle activation
differed substantially between the two conditions.
The differences are likely due to complex
interactions between initial conditions that differ in
the two conditions, the optimisation criterion and the
resulting kinematics, that in its turn affect the
dynamics by muscle actions.
Introduction
Ski jumping has a challenge with regard to
technique training. On average a ski jumper can perform
four jump trials per hour in the actual jumping hill. In
practice, ski jumpers therefore use different dry
activities that imitate (aspects of) actual ski jumping.
However, in all cases, the conditions of the dry training
forms are far from identical to those in a jumping hill. In
dry training, air resistance is usually minimal and
ground surface friction substantial, whereas in real
jumping the opposite is the case. We developed a
computer simulation model to quantify and understand
the differences that occur in these training forms.
The aim of this particular study was to compare a
dry jump (with full ground friction and without air
resistance) with real jumping conditions. We simulated
a hill jump and dry jump that were optimised for the
same criterion, in this case angular momentum of the
body at toe-off as a decisive measure in practice (in
reality, the athlete tries to obtain a small, but yet not
quantified, amount of forward momentum).
Materials and Methods
A forward dynamic model of a ski-jumper was
obtained by adapting a model previously described for
vertical jumping [1]. The model contained four rigid
segments foot+boot+ski, leg, thigh and upper body
(trunk, head+helmet and upper extremities). Segment
parameters were obtained for an average ski-jumper
(mass 60kg, height 1.75m) by adjusting data from [1].
The initial joint angles at onset of push-off were taken
from experimental data.
The athlete started his push-off 2m before the end of
the 105m radius curvature of the hill at a speed of
25ms -1 . Effects of air resistance were modelled per
segment and comprised two components, drag and lift.
A coefficient of friction between skis and snow was
assumed at 0.08. Otherwise, the dry training constraint
was a horizontal full friction surface and zero start
velocity.
Six muscle-tendon complexes of the lower extremity
(see Fig. 2) were embedded in the model and
represented by Hill-type models [1]. In the present
study, the stimulation (ranging between 0 for rest and 1
for fully activated) required to maintain the static squat
position was calculated. The activation could be
enhanced from that initial level at any time to 1
according to the bang-bang principle [1]. The timing of
the onset of activation enhancement for the six muscles
was optimised with regard to optimisation criterion.
The model’s set of equations of motion was solved
and integrated over time using the ode45 integrator in
Matlab (Mathworks).
Results
The zero angular momentum at the end of push off
in the hill is achieved by a moderate backward trunk
rotation accompanied by forward thigh rotation and
backward leg rotation (moderate hip extension and full
knee extension). The dry movement shows less trunk
rotation (Fig 1E). The normal force shows comparable
patterns (Fig. 1A), be it that the force in the hill is
slightly higher. The differences are explained by muscle
behaviour: muscle activation levels need to be higher at
the onset (centripetal effect), which affects the entire
force build up during push-off even though the
centripetal effect itself has disappeared.
The external moments by air resistance and snow
friction generate a forward momentum (Fig. 1B) that is
balanced by a backward moment of the normal force
(which is a reflection of muscle activity). The friction
force (tangent component of ground reaction force) in
the dry condition is irregular, indicating the direction of
the ground reaction force is changing considerably
during push-off. The moment of the entire ground
reaction force shows a rather similar pattern for both
conditions (Fig 1C). The momentum is not directly
nullified (or kept zero all the time during the dry jump),
but fluctuates considerably in both conditions before
approaching the zero value at toe-off (Fig. 1D).
IFMBE Proc. 2005;9: 239

Sports and rehabilitation biomechanics
Normal force (N)
Torque (Nm)
Torque (Nm)
Angular momentum (Nms)
1800
1600
A
1400
1200
1000
800
600
400
200
0
0 0.1 0.2 0.3 0.4
200
150
100
50
0
0 -50
0.1 0.2 0.3 0.4
-100
-150
-200
150
100
50
-100
-150
-200
B
C
D
onset
Hill
Dry
Surface friction
Air resistance
Normal force
GRF torque
0
0 0.1 0.2 0.3 0.4
-50
6
4
2
0
-2
0 0.1 0.2 0.3 0.4
-4
-6
Time (s)
-8
-10
-12
E
0.2s toe-off
Figure 1. Time traces of the simulations: normal force
(A), torques by normal force, air resistance, surface
friction (B), and by the total ground reaction force (C)
and angular momentum of the body (D). E: stick
diagrams at three moments during push-off.
The muscle activation pattern (Fig. 2) indicates that in
the hill the muscles have their onsets of activation closer
together in time. Furthermore, the onset order is not the
same for both conditions.
Discussion
The results of this study indicates that by mimicking
angular momentum in a dry jump as is found in the hill,
clear differences occur in the dynamics and muscle
activation pattern. In practice, athletes tend to imitate
the kinematics of a hill jump and this should give the
same angular momentum. Our simulation did not result
in the same kinematics. This may indicate that, given
the different initial conditions, it is not easy to obtain
similar kinematics in the hill and dry conditions. This is
also reflected by the pattern of the angular momentum
(Fig. 1D), which fluctuates considerably in both
conditions, instead of zooming in to the zero value (an
intuitive and attractive hypothesis). It seems that given
the constraints of the musculoskeletal system, it is
difficult to control angular momentum in a jumping
movement.
Hill
Dry
-0.35 -0.3 -0.25 -0.2 -0.15 -0.1 -0.05 0
Onset of muscle activation before take-off (s)
GLUTEUS
HAMSTRINGS
RECTUS
FEMORIS
VASTII
GASTROCNEMIUS
SOLEUS
Figure 2. Onset of activation of the six muscles during
the push off.
A striking finding is that not the normal torque
but that of the total ground reaction force shows similar
(but not identical) time traces. Apparently, to control
momentum in the dry (full friction) condition, the model
athlete needs to take full advantage of the friction
forces, resulting in a stereotype torque by the ground
reaction force. The differences in dynamics cannot be
considered as a mere down regulation of offsets in
angular momentum and moment by air resistance and
snow friction.
The current results need to be taken with
caution for various reasons. First, the optimisation
criterion use here is likely too simplistic as an athlete
does not only control momentum during a push-off.
Furthermore, various optima may exist in the sixdimensional
space of the present model. This has not yet
been investigated extensively. Further development of
the model may prove useful in assessing different
training techniques.
References
[1] VAN SOEST, A.J., SCHWAB, A.L., BOBBERT, M.F.
AND VAN INGEN SCHENAU, G.J. (1993): ‘The
influence of biarticularity of the gastrocnemius
muscle on vertical-jumping achievement’, J.
Biomech., 26, pp. 1-8
IFMBE Proc. 2005;9: 240

Sports and rehabilitation biomechanics
A 3 DIMENSIONAL MODELING OF HUMAN BODY UNDER VERTICAL AND
HORIZONTAL VIBRATION
M. Behzad 1 , P. Hejazi Dinan 2 , M. Rasolzadeh 1 , F. Farahmand 1
1 Mechanical Engineering, Sharif University Of Technology, Tehran, Iran
2 Physical Education, Alzahra University, Tehran, Iran
m_behzad@sharif.edu
Abstract
A 3D model of seated occupant was developed in order to
study the effects of whole body vibration on different
parts of human body. The model included 12 parts, i.e.,
the head, the neck, the upper torso, the pelvis, the thighs,
the legs, the feet and the seat backrest and cushion.
Excitation was exposed in vertical and horizontal
direction. The effects of level of vibration, its direction,
the way feet were positioned (hanging or fixed), the use
of headrest and the stiffness of seat springs were
explored. Results indicated that in both directions the
maximum acceleration ratio was seen in head. The least
resonance frequency was seen in pelvis. With increase of
vibration magnitude in both directions the acceleration
ratios decreased. Making feet fixed showed an increase in
acceleration ratio for upper parts of body and decrease of
it for lower parts. The secondary resonances disappeared
in pelvis and upper torso by fixing feet. In both directions
removing headrest increased acceleration ratio of upper
parts of the body and secondary resonances were
disappeared. With stiffening seat springs resonance
frequency increased for all parts, acceleration ratio
decreased in upper parts and increased in lower parts of
the body. Apparent mass increased and its resonance
decreased. Behavior of upper body parts were seen the
same
Introduction
Our exposure to mechanically induced vibration occurs
under a wide variety of living and working conditions.
Vibrations that arouse human health concerns are
classified into two main categories: (1) hand-arm
vibrations and (2) whole-body vibrations. Hand-arm
vibration (HAV) is transmitted through the hand-arm
system from a power or impact hand tool and affects the
upper extremities of the body. Whole-body vibration
(WBV) affects the entire body and is transmitted from a
vibrating seat, bed or floor to a person who is in a sitting,
lying or standing position. A large number of
epidemiological studies indicate an elevated risk of
disorders of the lumbar spine and of the connecting
nervous system due to long-term exposure to whole-body
vibration. An increased amount of back pain has been
associated with higher WBV intensity and longer
duration of exposure. Understanding of the mechanical
responses of the body is essential to assist in reducing the
undesirable influences on the health, the activities and the
feelings of occupants caused by vibration.
ISO 2631-1 defines the means to evaluate periodic,
random and transient vibration with respect to human
responses: health, comfort, perception and motion
sickness [‎1]. This standard specifies direction
and location of measurements, equipments to be used,
duration of measurements and frequency weighting, as
well as methods of assessment of measurements and
evaluation of weighted root-mean-square acceleration.
The biodynamic response characteristics of seated
occupants have been shown to be influenced by several
factors, among which body posture, body weigh,
vibration excitation type and amplitude probably
represent the most influential parameters [‎2].
The biodynamic response characteristics of seated human
subjects have been extensively reported in terms of
apparent mass (APMS), driving-point mechanical
impedance (DPMI) and seat-to-head vibration
transmissibility (STHT) [‎3].
It seems that the main focuse in previous works is the
behavior of seat while the effect of vibration on different
parts of body is also an important debate and should be
considered. Most of the papers concentrate on dynamic
properties of seat not behavior of the body under
vibration. The purpose of the present study is to develop
a detailed 3-dimensional model of human and seat in
which anthropometric data can be defined optionally and
also posture is a definable parameter.
Methods
Modeling
This paper models a seat–occupant system undergoing
base vibration. A need for simplified vibration models of
seat–occupant systems always has its own necessity.
The model described in this paper is a three dimensional
model. Modeling consists of following steps: Human
Body Modeling, Seat Modeling and Vibration Excitation
Modeling.
‎The model has 10 parts including: Head, Neck,
Upper Torso, Pelvis, Upper Legs, Lower Legs and Feet.
Arms and Hands were not considered avoiding time
IFMBE Proc. 2005;9: 241

Sports and rehabilitation biomechanics
consuming calculations. Each part was modeled as a
sphere for head and a cylinder for others. Parameters to
be defined were geometrical parameters (radius and
length), mass, moments of inertia and location of joints.
Excitation was applied to the seat base as a function of
seat position through a translational actuator. Excitation
could be in either vertical or fore and aft (horizontal)
direction. Initial phase of excitation was considered to be
zero.
Mathematical Formulation
Method used to model seated dummy in this paper was
based on multi body dynamics method. Parts were bodies
as rigid bodies, constraints, spring dampers and actuators.
Discussion
Great deal of parameters affecting WBV clears the need
of analytical modeling of seated occupants in order to
reduce number of experimental tests and decrease cost of
experimental studies. Although still experimental tests
are needed to validate results of mentioned models.
The measured frequency responses of car-seat occupants
are usually relatively simple with few peaks [‎2],
and some researchers have used as little as two-degreeof-freedom
models to fit to these frequency response
functions. Others have used simple mass–spring–damper
models [‎2].
Both the static and dynamic response characteristics of
occupied seats are important. However, in this paper only
the dynamic response was focused on.
Method used for modeling the Multi body modeling has
limitations and assumptions as bodies and joints are
considered to be rigid, joints are considered to be ideal
joints with no slackness, no impact between bodies is
considered, constraints are holonomic (function of time
and position) and the problem is forward dynamics. All
the springs and dampers were assumed to be linear but
had a range of motion.
The simplified three dimensional model in this paper
contains the basic components of the seat and the
occupant and models the interaction (foam and soft
tissue) between the two with several springs and
dampers. There are geometric non-linearities in the
system but the springs and dampers are assumed to be
linear. The non-linearities in the system model come
from the mannequin and seat geometry, and from the
piecewise linearity in torsional spring dampers of joints
and translational spring dampers of the seat cushion.
Validation of results was based on the fact that according
to simplified assumptions made here to develop model it
was unexpected to have exactly the same results in
magnitude as it is reported in experimental tests. This
model was only expected to have results which follow the
main rules of response of human body to WBV:
The main resonance frequency of the whole body and
body parts was seen in range of 4.5-5.5hz as is reported
in most of previous works [‎1]. The normalized
apparent mass in vertical vibrations started at 1 and
reached to 1.5 times of static mass, but in horizontal
direction it started near to zero. Behavior of upper body
parts were seen the same. Behavior of lower body parts
were seen the same. Secondary resonances in pelvis
seemed to be a cause of legs and lower parts of the body
so by fixing feet they disappeared. Increase of
perturbations like level of excitation or hardening seat
springs decreased the acceleration ratio in upper parts of
the body.
References
1. International Standard Iso 2631-1:1997, Mechanical
Vibration And Shock - Evaluation Of Human Exposure
To Whole Body Vibration - Part 1: General
Requirements.
2. S.K. Kim, S.W. White, A.K. Bajaj, P. Davies Journal
Of Sound And Vibration 264 (2003) 49–90 - Simplified
Models Of The Vibration Of Mannequins In Car Seats
3. M. Demicd And J. Lukicd Journal Of Sound And
Vibration (2002) 253(1), 109-129 -Some Aspects Of The
Investigation Of Random Vibration Influence On Ride
Comfort
IFMBE Proc. 2005;9: 242

Sports and rehabilitation biomechanics
STUDY ON THE NET JOINT MOMENTS OF THE LOWER LIMB IN NORMAL
AND ABOVE KNEE AMPUTED SUBJECTS
P. Hejazi Dinan 1 , T. Rezaeian 2 , M. Behzad 2
1 Physical Education, Alzahra University, Tehran, Iran
2 Mechanical Engineering, Sharif University Of Technology, Tehran, Iran
m_behzad@sharif.edu
Abstract
This study intended to measure the net moments of the
ankle, knee and hip joints of the intact and amputated
limbs of above knee amputees during walking, and to
compare the results with those of normal subjects. Gait
analysis was performed on five transfemoral amputee,
and 5 normal subjects. The motion and force data were
recorded using a digital video camera and a force
platform, respectively, and the anthropometric data were
derived using the regression relationships in the
literature. The kinematics and dynamics of the body
segments of each subject during walking were estimated
using a two-dimensional link segment model. Results
indicated that the sound limbs of the amputated subjects
had a larger average and higher extensional peak (2.08 to
1.68 Nm/Kg) for the hip joint moment, and longer
occurrence time for the above-average hip and ankle joint
moments in comparison with the normal subjects. The
amputated limbs of the amputee subjects had a larger
average for the knee joint moment and longer occurrence
time for the above-average hip joint moment, when
compared to the intact limbs, and a lower flexional peak
for the knee joint moment (0.2 to 1.14 Nm/Kg) and
extensional peak for the ankle joint moment (0.96 to 1.67
Nm/Kg), when compared to the normal subjects. The
longer stance phase duration and higher hip and ankle
joint moments observed in intact limb of amputee
subjects may indicate that the risk of articular cartilage
damage and OA incident is higher in this group in
comparison with normal subjects.
Introduction
The human gait involves the harmonized motion of the
lower extremities (foots, legs and thighs) via the
coordinated function of the muscles crossing the related
joints (ankles, knees and hips). In normal gait, not only
the energy consumption is optimized, but also the
resulting loads are restricted, so that they are tolerated by
the joints without any destructive change in the
articulating cartilages. Following amputation, however,
the normal gait pattern is altered, due to the inevitable
structural and functional changes occurred.
Previous clinical studies have indicated that there is a
higher incidence of osteoarthritis at the knee joint of the
intact limb of transtibial and transfemoral amputee
subjects [1, 2]. Recent studies, however, showed that
osteopenia/osteoporosis occurred significantly more on
the amputated side than the intact side of the above knee
amputee subjects [3, 4].
The purpose of the present study was to quantitatively
analyze the gait cycle of above knee amputee and normal
subjects, to calculate the net joint moments of ankle, knee
and hip joints using the inverse dynamic method, and to
compare the spatiotemporal variables and joint loads of
amputated and intact limbs of amputees with those of
normal subjects.
Methods
The tests were conducted on two groups of subjects. The
amputee subjects included four men with transfemoral
amputation and one man with knee disarticulation. They
had an average age of 37.8±4.48 years, height of
1.74±0.09m, mass of 72.4±11.97kg and body mass index
(BMI) of 24.63±2.99. Before beginning the test, all
subjects were assessed by a prosthetist to ensure that
none of the subjects had residual limb problems (such as
pain, swelling, pressure sore, painful motion, crepitus
with motion, ligamentous instability, limitation of
motion). The normal subjects included five men with
similar BMI and without any lower limb pathology, who
were chosen among volunteer students.
Each subject walked at a free cadence along a 6.5m
walkway. A black screen was placed at the background
for better marker detection and placing calibration
markers. A Kistler force plate (9286A with external 8
channel amplifier) was located approximately at the
midpoint of walkway for force data collection.
Kinematical data were recorded on a PC using a digital
Sony camera (DCR-TRV 330 E EIS), mounted
perpendicular and 10 m far from the walkway at knee
height, and a digital video I link (IEEE 1394 complaint)
fire wire. During walking, the camera recorded the
motion at 25 fps and the force plate collected the force
data at 50 Hz. Synchronization of the camera and force
plate data were conduced using a LED which flashed as
the force plate started to collect the data.
IFMBE Proc. 2005;9: 243

Sports and rehabilitation biomechanics
Results
This study suffers from limitations such as the small
number of subjects and the low speed of videography.
Also, the calibration method and probable motion of the
markers on the cloth of subjects can be assumed as the
main sources of error for the present study. However,
repeatability tests for kinematical and force plate data of
normal subjects indicated an acceptable level of accuracy
for the results. For example, a very good correlation
observed
observed for the ground reaction force and knee angle
variation of one of normal subjects in figures 7 and 8,
respectively. Moreover, the results of the present study
correspond well with those obtained by others, in the
mean patterns of the comparable quantities including the
ankle, knee and hip joints angles, the ground reaction
force, the net joint moments of ankle, knee and hip, the
stance duration, etc. for normal subjects and the net joint
moments of ankle, knee and hip for amputated limb
subjects.
Acknowledgements
The authors wish to thank all amputees who kindly
participated in the tests.
Discussion
In comparison of the net joint moments of intact limb of
amputee subjects and normal subjects, all the hip, knee,
and ankle joints of nonamputated limb of amputee
subjects experienced consistently larger moments.
Particularly, significantly higher results were found for
the second peak of hip joint moment and application time
of the above-average moment in the hip and ankle joints
of this group.
The longer stance phase duration and higher hip and
ankle joint moments observed in intact limb of amputee
subjects may indicate that the risk of articular cartilage
damage and OA incident is higher in this group in
comparison with normal subjects. However, it seems that
the amputee subjects reduce this risk by slowing down
their walking speed and decreasing the inertia force and
moments. Our overall results and conclusion are in good
agreement with the remarks given by other researchers..
References
1. Burke M.J., Roman V., Wright V., Bone and joint
changes in lower limb amputation, Annuals of
Rheumatology Disease, vol. 37, pp. 252-4. 1975.
2. Hungerford D.S., Cockin J., Fate of retained lower
limb joints in Second World War amputees, Journal of
Bone and Joint Surgery, vol. 57-B, pp. 111-7, 1975.
3. Kulkani J., Thomas E., Silman A., Association
between amputation, arthritis and osteopenia in british
male war viterans with major lower limb amputations,
Clinical Rehabilitation, vol. 12, pp. 348-53, 1998.
4. Rush P.J., Wong J.S.W., Kirsh J., Devlin M.,
Osteopenia in patients with above knee amputation,
Archives of Physical Medicine and Rehabilitation, vol.
75, pp.112-5, 1994.
IFMBE Proc. 2005;9: 244

Sports and rehabilitation biomechanics
PLANAR GEOMETRY OF FEET IMPRINTS
AS THE REFLECTION OF THE STRUCTURE
AND CONDITION OF LOCOMOTIVE SYSTEM
Е. Dukendjiev 1 and Т.Оgurtsova 1
1 Division of Mechanics of Prosthetics, Riga Technical University, Riga, Latvia
apl@stradini.lv
Abstract: The norm and the pathology of feet
biomechanics is analysed, taking planar imprints
into account by means of a graphical-calculated
method by E. Dukendjiev.
Introduction
In orthopaedics there is used planar computerised
podometry based on the matrix grid of tensometric
sensors, forming digital images of load in contacting
and surface areas. There is no geometrical analysis of
the imprint, as well as its connection with
biomechanical structure of a foot.
Methods
Author's method unambiguously connects bonejoint
system with the imprint geometry on a flat glass
support. Figure 1 shows geometrical structures in
norm. A patient is walking along the glass path under
which there is a carriage with three video cameras,
filming the feet in three planes [2]. On the planar
imprints obtained in various phases of the walk there
are placed geometrical matrices, areas of all
contacting spots are fixed and colour temperature is
measured (colour scale is differentiated) as the
equivalent to the load. By means of a programme
obtained data are compared with the data in norm and
clinical diagnostics is carried out with the following
synthesis of construction of bionical inner soles [1].
Results
For the first time there is carried out complex
orthopaedic diagnostics of feet condition in walking
process according the parameters of planar geometry,
revealing not only anatomical pathologies but also
biomechanical ones, connected with the condition of
the locomotive system.
Discussion
Objective geometrical parameters – area,
symmetry, power rays and loads – significantly
decrease the influence of the subjective factor
(competence level, form of technical possibilities,
work experience etc. of the orthopaedist). From the
colour zones by means of colour scale the load is
determined, but from their geometrical structure
conclusions are made about form and degree of
pathology. Complex of these data optimise the
topography and architectonics of the bionical inner
soles.
Conclusion
Analysis of more than 420 patients of different
genders and age groups has allowed synthesising
geometric matrices of planar imprints in norm and
formulating the signs of pathology and criteria for its
evaluation. In clinical conditions there has been
achieved minimal time of diagnostics – about 25-40
minutes per patient. At the same time the precision of
determining the form and degree of pathology not
only of feet but the whole locomotive system is
almost 96%. The method may be also applied in
evaluation of the results of prosthetics of the lower
extremities.
References
[1] E. DUKENDJIEV, T. OGURCOVA. (2004):
Examination of Feet During Walking and Synthesis
of Bionical Insoles. Proc. of The 11 th World Congress
of the International Society for Prosthetics &
Orthotics August 1-6, 2004 Hong Kong, p.349
[2] E. DUKENDJIEV, T. OGURCOVA. (2001)
Methods end Examination’s aparatus of feetprints on
the surface. Proc. International Conference for Young
Scientists on Bionics, Biomechanics and Mechanics,
Proceeding – Riga – Varna, 2001 p – 23-25.
Acknowledgement
On The 11 th World Congress of the International
Society for Prosthetics & Orthotics August 1-6, 2004
Hong Kong abstrakt “Examination of Feet During
Walking and Synthesis of Bionical Insoles” has
received the award for prezentation one set of Crystal
Globe
IFMBE Proc. 2005;9: 245

Sports and rehabilitation biomechanics
Figure 1. Geometrical structures in norm
IFMBE Proc. 2005;9: 246

Sports and rehabilitation biomechanics
PROPRIOCEPTIVE ABILITY WITHIN PATIENTS WITH WHIPLASH
ASSOCIATED DISORDERS
H. Grip 1, 3 , G. Sundelin 2 , S. Karlsson 1, 3
1 Department of Biomedical Engineering and Informatics, Umeå University Hospital, Umeå, Sweden
2 Department of Community Medicine and Rehabilitation, Umeå University, Umeå, Sweden
3 Centre for biomedical engineering and physics, Umeå University, Umeå, Sweden
helena.grip@vll.se
Abstract
Objective neck movement analysis can be used as an
assessment tool in the examination and rehabilitation of
patients suffering from chronic whiplash associated
disorders (WAD). In two earlier studies we have shown
significant differences in head movement pattern between
a group of WAD patients and a control group, e.g. head
movement range and rotation angle velocity [7,8].
In this pilot study we evaluated the proprioceptive ability
and neck muscle activity in patients with WAD and a
group of controls.
Introduction
Whiplash Associated Disorders (WAD) are a common
diagnosis after neck trauma, caused by sudden
acceleration and deceleration forces acting on the head
and neck, often related to rear-end car accidents [1]. In
clinical practice, WAD patients report two main areas of
complaints: increased muscle tension and pain during
repetitive arm and shoulder movements and increased
pain and stiffness during repetitive neck movements. The
clinical examination generally includes palpation of
muscles for pain and visual inspection of range-ofmovement
in neck and shoulder joints. It is difficult to
identify subgroups and thus establish a more precise
diagnosis, and imaging techniques such as X-ray and
magnetic resonance therapy seldom reveal any
pathological changes.
Several studies have described pathological neck
movement patterns related to WAD [2-6]. There are also
indications on that muscle activation in patients with
chronic musculosceletal pain differs from normal
patterns, e.g. in trapezius and sternocleidomastoideus
(SCM) muscles [9,10]. Mathiassen and Winkel [11]
developed a method called exposure variation analysis
(EVA), to quantify the electromyographic (EMG) muscle
activity in both amplitude and duration, which allow
comparison of individual differences when performing a
work task.
An objective decision support system based on
movement analysis and EVA might enhance the
development of more precise diagnostic procedures for
WAD patients. In two earlier studies we used movement
analysis based on reflective skin markers to study
repetitive head rotations and showed significant
differences between WAD patients and controls, e.g. in
movement range and rotation angle velocity [7,8].
The aim of this pilot study was to investigate the
proprioceptive ability and the muscle activity during
slow, repeated head rotations in subjects suffering from
WAD compared to subjects with a group of controls
without neck problems.patients suffering from WAD.
Methods
Six subjects with chronic WAD symptoms lasting longer
than three months (40±17 years), and six controls without
neck pain (42±19 years) were selected.
Movement task: The subject started the test session sitting
relaxed in a chair, in front of a board where five positions
were marked (rest position, flexion-extension 25 degree,
right-left rotation 30 degree). An IR-lamp was attached to
a helmet, with the light ray pointing at the board to guide
the subject. Twenty head movements were performed in
random order. After rotating the head to the prescribed
position, the subject returned to rest position with closed
eyes. A button (generating an analog signal) was pressed
down by the subject, to mark estimated rest position. The
eyes were opened and the subject re-positioned the head,
guided by the IR-light.
Figure 1. Marker placements
Movement analysis: Markers were placed as shown in
Figure 1. The 3D motion data was collected at 120 Hz
and the head rotation angle relative to the upper body was
calculated [7] and used to find the difference between
estimated and original rest position.
Muscle activity: Surface EMG signals were collected
from upper trapezius and SCM muscles at 2040 Hz. The
skin was dry shaved and cleaned and electrodes were
attached to the skin. The reference electrode covered the
IFMBE Proc. 2005;9: 247

Sports and rehabilitation biomechanics
seventh cervical vertebra (C7). Reference voluntary
contractions (RVC) estimated to 15% and 65% of
maximal voluntary contraction (MVC) were performed
before the test for trapezius and SCM respectively. The
subject was asked to hold both arm straight at 90 degrees,
in the sagittal plane, for 7 seconds. Then the subject lay
down, and lifted the head slightly during 7 seconds. Data
analysis was performed off-line using MATLAB ® . For
each subject and muscle, the reference voluntary
electrical activity (RVE) was determined as the RMS
value for a selected 2 s period of RVC. The EMG signals
collected during the test were RMS filtered using a
moving average window of 0.1 s and normalised with
RVE before an EVA was performed.
Results
The work task took 6±2 minutes in mean for WAD
patients compared to 4±1minutes for controls. Examples
on EVA are shown in figure 2-3.
Figure 2. EVA from a patient with WAD. Trapezius
show an activity up to 30%MVC, while the SCM
muscle show an activity on up to 7%MVC.
Figure 3. EVA from a control subject. Trapezius show an
activity on

Sports and rehabilitation biomechanics
A PILOT STUDY TO ESTIMATE THE LACTATE THRESHOLD USING AN
ELECTRO ACOUSTIC SENSOR
M. Folke*, L. Gullstrand** and B. Hök***
* Department of Computer Science and Electronics, Mälardalen University, Västerås, Sweden
** Swedish National Sports Complex, Lidingö, Sweden
*** Hök Instrument AB, Västerås, Sweden
mia.folke@mdh.se
Abstract: End tidal carbon dioxide measurement
with an electro acoustic sensor has recently been
demonstrated. The sensor consists of an acoustic
resonator coupled to a low cost electro acoustic
element. The aim of this study was to verify if the
electro acoustic sensor would be able to estimate the
lactate threshold during a steady state ergo meter
bike test. Simultaneous measurements with a
reference carbon dioxide sensor were made during
the whole test. Heart rate and blood lactate
measurements were made in the end of each
workload.
The electro acoustic sensor has its break point at
the same workload as the end tidal carbon dioxide
concentration and shows to be useful to estimate the
lactate threshold in this pilot study.
Introduction
A new carbon dioxide (CO 2 ) sensor for
measurements in expired air in a mainstream application
has recently been presented [1].
The sensor is based on the measurement of the
impedance of an electro acoustic element coupled to an
acoustic resonator [2]. The impedance characteristic is
depending on the sound velocity within the gas mixture
contained in the acoustic resonator.
It has been shown in an earlier study that there is an
approximately linear relation between the acoustic
impedance and the CO 2 concentration [2]. The sensor
principle has also shown a fast response to increasing
CO 2 concentration in laboratory experiments [3].
At constant temperature and humidity, the sound
velocity is determined by the average molecular weight
of the gas, and is therefore influenced by the CO 2
concentration, since CO 2 is considerably heavier than
oxygen and nitrogen, which dominates the average
molecular weight of air, Equation 1.
The lactate threshold can be verified by analyses of
the expiratory gases [4, 5]. A decrease of end tidal CO 2
tension during hard exercise has been seen [5] and end
tidal oxygen tension have shown to increase at the hard
exercise [4].
The aim of this study was to verify if the electro
acoustic sensor would be able to estimate the lactate
threshold using the break point where end tidal CO 2
concentration starts to decrease, eventhough the oxygen
tension in expired air will increase during rapid
respiratory rate because a smaller amonth of gaschange
get time to take place.
where:
RT γ
c = Equation 1
M
c is the sound velocity
R is the general gas constant
T is the absolute temperature
γ is the ratio between the heat capacities at constant
pressure and volume
M is the average molecule mass of the gas mix
Materials and Methods
A steady state ergo meter bike test was performed,
starting at 50 Watt with an increase of 25 Watt every 4-
minute until raised blood lactate levels at 175 watt. The
test was performed on a Monark 839E ergo meter bike
(Monark Exercise AB, Sweden).
The electro acoustic sensor was placed between two
filters in a tube. The purpose of the filters was to reduce
variations in humidity and temperature. The tube was
then placed outside the flow sensor of the reference
equipment during the test.
In the end of each workload heart rate, blood lactate,
end tidal CO 2 level and the output signal from the
electro acoustic sensor have been analysed. Lactate in
micro samples of blood was analysed using a Biosen C-
Line, (EKF Diagnostic, Germany), the end tidal CO 2
level with a Jaeger Oxycon Pro, (Viasys Healthcare
GmbH, Germany) and the heart rate with a heart rate
monitor from Polar (PE3000, Polar Electro Oy,
Finland). Both the Oxycon Pro and the Biosen C-Line
were calibrated before the test.
The output signals from the Oxycon Pro and the
electro acoustic sensor were sampled every 10 ms
during the test and analysed afterwards. The average
values from ten breaths during the last minute on each
workload were calculated.
IFMBE Proc. 2005;9: 249

Sports and rehabilitation biomechanics
Results
Figure 1 shows the plot of workload versus the
output signal from the Oxycon Pro representing the end
tidal CO 2 concentration and the output signal from the
electro acoustic sensor representing the molecular
weight of the respiratory air, respectively. The break
points of the two systems occur at a workload of
150 Watt. The decrease of the signal representing the
end tidal CO 2 concentration was steeper than the one
representing the end tidal molecular weight.
Molecular
weight
Carbon dioxide
concentration
end tidal molecular weight does not have as steep
decrease after the break point as the end tidal CO 2
content. The end tidal molecular weight is also higher
than the end tidal CO 2 content at the workload of
50 Watt. Looking at the oxygen content in the expired
air it will be higher in the beginning and in the end of
the test.
The break point of the end tidal molecular weight
and the end tidal CO 2 content in the respiratory air does
deviate from the workload at the defined lactate
threshold with 20 Watt and the workload at 4 mmol/l
with 10 Watt. The relationship between blood lactate
and the end tidal CO 2 content in the respiratory air will
be further analysed in a parallel study.
During the test, the test subject found the respiratory
resistance irritating, caused by the filters, above
125 Watt. That feeling was verified by increased
oxygen uptake as well as increased ventilation when
taking the filters away the measurements at the final
workload.
Conclusions
0 50 75 100 125 150 175 200
Workload (Watt)
Figure 1. Workload versus the output signals from the
Oxycon Pro and the electro acoustic sensor,
respectively.
Figure 2 shows the plot of workload versus heart
rate and blood lactate, respectively. The blood lactate of
4 mmol/l occurs at a workload of 140 Watt. By defining
the threshold of blood lactate as the intersection of a
horizontal line through the lowest lactate level and the
steepest slope of the lactate curves latest part gives a
break point at a workload of approximately 130 Watt.
The electro acoustic sensor can, although it is not
selective to carbon dioxide, verify the break point of end
tidal carbon dioxide concentration in expired air during
exersice and shows to be useful to estimate the lactate
threshold in this pilot study. However, the influence of
humidity and temperature variations must be reduced
before the method can be further evaluated.
References
[1] FOLKE M., HÖK B., EKSTRÖM M., and BÄCKLUND
Y. (2004): ‘End Tidal Carbon Dioxide Measurement
Using an Electro Acoustic Sensor’, Proc. of EMBC
2004 - the 26 th Annual International Conf. of the
IEEE Engineering in Medicine and Biology Society.
San Francisco, USA, 2004. p 362
Lactate (mmol/l)
8
7
6
5
4
3
2
1
0
100
0 50 75 100 125 150 175 200
Workload (Watt)
Lactate
Heart rate
170
160
150
140
130
120
110
Heart rate (Beats per
minute)
Figure 2. The workloads versus blood lactate and heart
rate, respectively.
Discussion
The end tidal molecular weight has the same break
point as the end tidal content in the respiratory air. The
[2] GRANSTEDT F., FOLKE M., BÄCKLUND Y., and HÖK
B. (2001): ‘Gas sensor with electro acoustically
coupled resonator’, Sensors and Actuators B. 78,
pp.161-165.
[3] GRANSTEDT F., HÖK B., BJURMAN U., EKSTRÖM M.,
and BÄCKLUND Y. (2001): ‘New CO 2 sensor with
high resolution and fast response’, Proc. of EMBC
2001 - the 23 rd Annual. International Conf. of the
IEEE Engineering in Medicine and Biology Society.
(EMBC 2001), Istanbul, Turkey, 2001.
[4] WASSERMAN K., WHIPP BJ., KOYAL SN., and
BEAVER WL. (1973) ’Anaerobic threshold and
respiratory gas exchange during exercise’, J of Appl
Phys. 35 (2), pp.236-243.
[5] WASSERMAN K. and WHIPP BJ. (1975): ’Exercise
Physiology in Health and Disease’. American
Review of Respiratory Disease. 112: pp.219-249.
IFMBE Proc. 2005;9: 250

Sports and rehabilitation biomechanics
MONITORING HEALTH AND ACTIVITY BY SMARTWEAR
L. Berglin 1 . M. Ekström 2 . M. Lindén 2
1
The Swedish School of Textiles, University College of Borås, Sweden. Department of Computing
Science and Technology, Chalmers University of Technology, Göteborg, Sweden
2
Computer Science and Electronics, Mälardalen University, Västerås, Sweden
lena.berglin@hb.se
Abstract
This paper describes how health monitoring and
communication could be integrated in a textile
garment. Initial experiments show how communication
could be integrated in a textile structure
using conductive fibres. Further experiments
show that textile knitted electrodes could be used
for ECG-registration. The project demonstrates
the possibilities of integrating health monitoring
and communication in a garment. The next step is
to integrate the textile electrodes with a wireless
communication system.
Introduction
The development of smart, wearable systems
has become possible because of better performance
and smaller size in electronics, biomedical
sensors and communication technologies. At the
same time material science has developed new
textile materials, so-called smart textiles [1].
These materials react to stimuli from their environment
and thereafter, in different levels, adapt
their behaviour to the circumstances. Combining
electronic devices and new textile materials gives
an opportunity to develop a new type of clothing
with a new type of behaviour and use. An example
of this is the implementation of electronics in
clothing, so-called Smartwear. Several research
projects have been presented in recent years in
different areas of applications in smart textiles
and clothing. There are prototype systems aiming
at multi-function monitoring, as the LifeShirtTM
[2]. A survey of “intelligent biomedical clothes”
have been performed [3]. However, most projects
have been limited to combining existing technology
with textile and do thus not include a total integration
between technology and textile. One of
the first effort to integrate textile and computing
was performed at Georgia Institute of Technology,
designing and developing the Georgia Tech
Wearable Motherboard [4].
The aim of this project is to develop a wearable
system integrated in clothing. The system is focused
on communication and health monitoring.
Material and methods
This project aims at a total integration between
technology and textile. So far, however, existing
technology has been combined with textile and
one standard component at a time has been replaced.
The initial trials include one communication
part and one monitoring part.
Integrated communication
In order to integrate communication in textiles,
four conductive knitted surfaces isolated from
each other were made. In the ends, a microphone
(m), a headphone (h) and a connection to the mobile
phone was attached, see figure 1. The surfaces
were made in different knitting and weaving
techniques. Two types of yarns were tested,
100% stainless steel and 50/50 % stainless steel/
polyester. The electronics were connected to the
conductive surfaces by traditional electronic conjunctions,
snap buttons or a conductive Velcro
fastening. The signal transmission was tested on
frequencies up to 20000 Hz.
Figure 1. Principal scheme and picture of communication
system integrated with textile.
IFMBE Proc. 2005;9: 251

Sports and rehabilitation biomechanics
Health monitoring system, integrated electrodes
To accomplish a textile integrated health monitoring
system, electrodes were made from conductive
yarn, knitted in pieces that can be applied
under an undershirt, see figure 2. Two types of
yarn was used, 100 % stainless steel and 50/50 %
stainless steel/polyester.
textile. In the future, we aim at substituting the microphone
using piezoelectric material in a textile
structure. The initial test of the health monitoring
system includes textile electrodes. Next step will
be to integrate them with a wireless communication
system, e.g. using Bluetooth, from which we
have experience in another project [5].
Conclusion
Figure 2. Textile electrode.
To investigate the performance of these electrodes,
they were used for an ECG-registration
with a standard ECG-machine.
Result
The transmission of communication signals works
in all the tested yarns. Samples in 100% stainless
steel had the best performance. The choice
of conjunctions gives a very little difference between
traditional conjunctions and snap buttons
while the Velcro fastening performs worse and is
also more clumsy and hard to handle.
Both types of conducting yarn could be used
for ECG-registration. In figure 3, an example is
shown.
Figure 3. ECG-registration using textile electrodes.
Discussion
Finally this project aims at a total integration between
technology and textile. To accomplish this,
pilot experiments replacing one standard part at
time, has been performed. So far, communication
has been performed by integrating a standard
microphone, headphone and a mobile phone into
The experiments show that health monitoring
from and communication through a textile garment
is possible. A system where wireless surveillance
can offer both health monitoring and
communication could be applied in many areas,
as health care, sports medicine and surveillance
of persons with extreme working conditions as
firefighters.
References
[1] van Langenhove L., Hertleer C. Smart textiles-an
overview, Proceedings of Autex Conference,
Gdansk, Poland, 2003, pp 15-20.4.
[2] Grossman P. The lifeshirt: a multi-function
ambulatory system that monitors health, disease,
and medical intervention in the real world, New
Generation of Wearable Systems for e-health,
International Workshop December 11-14, 2003.
Lucca, Tuscany, Italy.
[3] Lymeris A., Olsson S. Intelligent biomedical
clothing for personal health and disease management:
State of the art and future vision, Telemedicine
journal and e-health, Volume 9, Number 4,
379-86, 2003.
[4] Park S., Gopalsamy C., Rajamanickam R., and
Jayaraman S. The wearable motherboard: An information
infrastructure or sensate liner for medical
applications, in studies in Health Technology
and Informatics. Amsterdam, The Netherlands:
IOS Press, 1999, vol. 62, pp. 252-258.
[5] Lönnblad J., Castano J.G., Ekström M.,
Lindén M., Bäcklund Y. Optimization of Wireless
BluetoothTM Sensor Systems, , 26th Annual International
Conference of the IEEE Engineering
in Medicine and Biology Society (EMBC 2004),
1-5 Sept 2004, San Francisco, USA.
IFMBE Proc. 2005;9: 252

Sports and rehabilitation biomechanics
THE ASSESSMENT OF THREE DIMENSIONAL ANKLE JOINT
FORCES DURING THE POSTURAL BALANCE CONTROL
MOVEMENT
M.J. Seo and H. Choi
School of Mechanical Engineering, SungKyunKwan University, Suwon, Korea
Abstract: The purpose of the study was to assess
three dimensional reaction forces and
bone-on-bone forces within ankle joint during
postural balance control movement. With
experiments and MATLAB simulation we could
calculate ankle joint kinematic and kinetic data.
The results presented in this study will be useful
data for understanding the injury mechanism of
ankle joint during postural balance control.
Introduction
Postural balance control is a complex process to
adjust posture by various joints, muscles and bones.
Muscle is an important mechanical factor to control
and to excite these movements. Also, muscle is
known for a contributor to contact force of joint.
Until recently, many studies about joint reaction
forces have been developed during various posture
like gait, stair-climbing, balance control, sit to stand
[1-3]. But, most of the biomechanical researchers
analyzed joint movement in two-dimensional plane
to reduce the complexity of analysis. And they did
not properly consider muscle force by active
contraction in dynamic condition.
In this study, we assessed three dimensional
contact force of ankle joint during waist pulling
considering muscle force for dynamic condition.
Materials and Methods
In the experimental process, nine healthy male
adults (age range: 21-27; height range: 169-181cm;
weight range: 58-86kg) participated. We used
6-camera motion analysis system (VICON,
sampling frequency 120 Hz), force platform (AMTI,
1080 Hz) and waist pulling system. The waist
pulling system pulls and pushes the rope which is
connected to subject's waist. Subjects stood with
bare feet on a force plate. They loosely tethered at
chk@me.skku.ac.kr
the waist to the perturbation system. During the
perturbation, we measured body motion and ground
reaction forces. The period of time from the onset of
pulling to after the balance recovery of subject was
normalized to 1.
Ankle joint model is assumed three-degree of
freedom ball and socket joint considering three
rotational movements. And fourteen reflective
markers with a diameter of 15mm were attached on
pelvis, knee, both malleoluses and foot to determine
three dimensional position of lower extremity.
BOBF (bone on bone force) is the contact force
of a joint considering MCF (muscle contraction
force). In this study, we calculated 11 muscle forces
which contribute ankle joint reaction force and
assessed effect of these muscle forces on BOBF.
Results
With the three-dimensional ankle joint model,
we calculated the three-dimensional angular
displacements, moments, JRF (joint reaction force)
and BOBF using double linear optimization. In
Figure 1, we divided the graph into two to present
the results of nine subjects.
In the figure, we presented normalized BOBF to
the subject's weight. BOBF in Y-axis is the superior
direction at joint, X-axis is antero-posterior shear
direction and Z-axis is medio-lateral shear direction
each. The largest BOBF was calculated in Y-axis
due to subject's weight. The maximum range of
BOBF in X-axis was approximately -1.8 ~ -0.3
times of subject's body weight, in Y-axis 2.2~ 7.9
times of subject's body weight, and in Z-axis 0.2~
0.9 times of subject's body weight each [Figure 1].
The time at which maximum bone on bone force of
X-axis appears is 0.2~ 0.4 except subject SH.
Balance was gradually recovered by the action of
muscles.
IFMBE Proc. 2005;9: 253

Sports and rehabilitation biomechanics
study can be applied various disability of diabetic
foot patient, flatfoot, etc and we considered that
biomechanical analysis presented in this study can
Figure 1: Vertical(Y) and shear(X, Z) components of ankle bone-on-bone force for 9 young male
adults during waist pulling, Horizontal-axis: balance recovery cycle (%), Vertical-axis: Force
(N)/Weight (N).
Discussion and Conclusions
To obtain BOBF of ankle joint, we performed
the following procedure. Firstly, we calculated joint
moment by using external force and inertia force of
subject’s weight during waist pulling. Then, we
implemented a optimization procedure using the
anatomical data to calculate muscle forces. Finally,
we calculated BOBF of ankle joint with 11 muscle
forces.
Maximum BOBF considering muscle force is
gradually increased and decreased during the
balance recovery cycle, but the maximum JRF
which is not considering muscle force did not show
the smooth behavior which is reflected by the real
natural movement. Therefore, we could conjecture
that the effects of muscle contraction must be
considered during the analysis of joint kinetics.
We analyzed mechanical factors that could
affect joint injure mechanism during the postural
balance control in a very simple way only using
sway experiment and optimization procedure.
Therefore, the results and the methodology of this
study can be usefully applied to further clinical and
biomechanical studies to understand more about
ankle joint injury mechanism during the postural
balance control movements.
References
[1] CALLAGHEN J. P., PATLA A, E., AND MCGILL, S.
M. (1999): "Low Back Three-Dimensional Joint
Forces, Kinematics, and Kinetics During
Walking", Clinical Biomechanics, 14, pp.
203-216
[2] WYSS U. P., COSTIGAN P. A., LI J., OLNEY S.
J., ZEE B. C., AND COOKE T. D. V. (1994):
"Bone-on-bone Forces at the Knee Joint During
Walking and Stair Climbing", J. Biomechanics,
27, pp. 819
[3] MAK M. K. Y., LEVIN O., MIZRAHI J., CHRISTINA
W. Y., AND HUI-CHAN (2003): "Joint Torques
During Sit-to-Stand in Healthy Subjects and
People with Parkinsons Disease",
Clinical Biomechanics, 18, pp. 197-2
IFMBE Proc. 2005;9: 254

Sports and rehabilitation biomechanics
THE EFFECTS OF EXTERNAL LOAD , TRUNK AND KNEE
POSITION ON THE TRUNK MUSCLES ACTIVITY
S. Kahrizi* , M.Parnianpour**, S.M. Firoozabadi*** and A. Kazemnejad****
*Assistant Professor,Dept. of Physiotherapy, Tarbiat Modares University
** Associate professor, Faculty of Mechanic Engineering, Sharif Industrial University
*** Associate Professor ,Dept.of,Medical Physics, Tarbiat Modares University
****Associate Professor,Dept. of Biostatistics, Tarbiat Modares University
kahrizi20@yahoo.com
Abstract: Eighteen static tasks while holding
three levels of load, two levels of knee
position and three levels of trunk position
were simulated for 10 healthy male subjects.
With highest external load, the electrical
activity of flexor and extensor muscles
increased (p

Sports and rehabilitation biomechanics
70
60
50
40
30
20
10
0
*
*
Muscles
K45
K180
Figure 1: The knee position effect on the eight
trunk muscles in 10 male subjects
Discussion
From 0 kg. to 20 kg., activity of all 8 trunk
muscles were increased. This results were
agreement with some studies [3-5]. To maintain
the realistic conditions, muscle activation must
be recruited to stiffen and stabilize the spine. In
this study the relation between trunk flexion
angle (15 and 30) and trunk muscles activity
weren’t significant. In the other hand, there
were no different between activity of flexor and
extensor muscles as the trunk flexion angle
increased. Some other studies similar to this
study couldn’t find the higher muscle activity
when trunk flexed [7,8]. Tan have reported a
significant reduce in the EMG activity of flexor
muscles due to the mechanical disadvantage of
these muscles as they were shorted, when the
flexion angle of trunk increased [6]. Marrass
have reported a significant decrement in the
EMG activity of the erector spine muscles
between upright posture and forward flexed
posture [7]. Although, results of some studies
aren’t similar to this study [3,4]. In our study
the maximum isometric contraction were
measured in standing and upright position,
while some authors have concluded there are a
linearity relation between increasing trunk
flexion and %MVC [8].
In this study also, the relation between electrical
activity of the extensor muscles (LD and ES)
and knee bending from 0 to full flexion were
significant (p

Sports and rehabilitation biomechanics
AN INVERSE KINEMATIC MODEL OF THE OSTEOPOROTIC SPINE
Z. Yang 1 , J. Griffith 2 , P. Leung 3 , M. Pope 4 , R. Lee 1
1 Rehabilitation Sciences, Hong Kong Polytechnic University, Hunghom, Hong Kong
2 Radiology, Chinese University of Hong Kong, Shatin, Hong Kong
3 Orthopaedics, Chinese University of Hong Kong, Shatin, Hong Kong
4 Medical Physics and Engineering, University of Aberdeen, Aberdeen, UK
rsrlee@polyu.edu.hk
Abstract
An inverse kinematic model which was employed to
determine an optimal intervertebral joint
configuration for given flexion and extension postures
of normal spines was evaluated on osteoporotic spine.
The osteoporotic lumbar spine (L1-L5) was modeled
as an open-ended kinematic chain. An optimization
algorithm with physiological constraints was
employed to determine the intervertebral joint
configuration. Intervertebral movements were
measured from sagittal computerized radiographs of
nine subjects as reference. The model is validated by
calculating the mean difference between radiograph
measurements of intervertebral rotation in the
sagittal plane and the values predicted by the inverse
kinematic model. It is concluded that the inverse
kinematic model needs to be customized for
osteoporotic spine in order to be clinically useful for
predicting intervertebral movements of osteoporotic
lumbar spine when radiograph or invasive
measurements are undesirable.
Introduction
Knowledge of intervertebral movements of the lumbar
spine is clinically useful in the assessment of spinal
disorders (such as instability) and treatment outcome. In
order to measure the intervertebral movements, many
approaches have been reported, such as radiographic,
electro-optical, and electromagnetic techniques.
However, radiographic techniques are complicated, and
have the inherent health risk of repeated radiation
exposure. Surface measurements using markers or
sensing devices are subjected to large error due to the
deformation of underlying soft tissues disguising the true
vertebral movement [1].
An inverse kinematic model was developed in a previous
study for the subject with normal lumbar spine [2].
However, whether this model is applicable to
osteoporotic spine is unknown due to the deformities of
vertebral bodies, degenerated intervertebral discs, and
possible different kinematic mechanisms. More attention
is being paid to osteoporosis of the elderly as our
population continues to age. The purpose of this study
was to evaluate the validity of using an inverse kinematic
model to determine the intervertebral joint movements of
osteoporotic spine, with given knowledge of the total
range of flexion and extension movements and the
positions of the spinous processes.
Methods
Nine volunteers (3 men, 6 women, 73±4 yrs old) with
different level of bone mineral density (2 osteopenia, 5
osteoporosis, 2 severe osteoporosis) agreed to participate.
The subjects were requested to take lateral radiographs in
three postures: neutral upright, full flexion, and full
extension. The osteoporotic lumbar spine was modeled as
a planar multilink chain system, as shown in Fig.1. Each
intervertebral joint has three degree of freedom (DOF),
i.e. one rotation and two translations. Global and local
coordinate systems are illustrated in Fig. 1.
Figure 1: The lumbar spine is modeled as an open-ended
kinematic chain. The joints are located at the centers of
intervertebral discs (Only the rotational joints are shown
in the figure, the x-, y- translational joints are omitted to
be concise).
The following information was assumed to be known:
(a) The positions of the most posterior parts of the
spinous processes. Clinically, these bony landmarks can
be easily palpated through the skin surface and their
positions can be predicted from surface measurements
[3]; (b) The total movements of the osteoporotic lumbar
IFMBE Proc. 2005;9: 257

Sports and rehabilitation biomechanics
spine; (c) The geometry of the vertebral bodies and the
length of each link of the kinematic chain.
The joint angles θ i , the x i , y i translations of the vertebrae
were the state variables ω i (i = 1, 2, …, 5). The positions
of the most posterior parts of spinous processes S i were
expressed as follow,
S i = f(ω i ) (1)
where ω i = θ i , x i , y i T . To derive ω i with a given S i , the
inverse of equation would be required. It denotes,
ω i = f -1 (S i ) (2)
There was more than one solution for a given set of S
values. The inverse kinematic problem was solved by the
general equation as follow,
where J is the Jacobian matrix, J + the pseudo-inverse of
J, and I the identity matrix. The potential function P was
minimized to optimize the solution. The following
potential functions were employed: (a) The error in
predicting the total movement was minimized; (b) The
intervertebral rotation and translations are constrained
within the physiological limits; (c) The Jacobian
determinant was minimized. An iterative refinement
algorithm was used to improve the solution.
Results
The predicted intervertebral joint movements is
satisfactory in some cases, but not in the others. Table 1
provides the mean absolute differences between obtained
by radiographic measurement and those predicted by the
inverse kinematic algorithm. The absolute prediction
error, which was defined by the differences between the
measured and predicted values, is presented separately
for each kinematic parameter and for each intervertebral
joint. Generally, the relative error of rotation is smaller
than that of translation, which is the same result as the
model applied on normal spine.
(3)
L5/S L4/5 L3/4 L2/3 L1/2
θ 6.68 2.90 3.59 1.09 3.19
x 2.45 1.95 6.37 5.29 4.26
y 0.75 0.69 1.01 0.86 1.32
Table 1: Mean absolute prediction error measured and
predicted kinematics parameters (rotation angle θ,
translation x, y) of the 5 intervertebral joints
The mean absolute error of intervertebral rotation ranged
from 1.09º to 6.68º, which was larger than that in the
prediction of normal spines, ranging from 1.0º to 1.6º.
Although the inverse kinematic model was evaluated as
inaccurate in predicting translational movements of
normal spine, the prediction of osteoporotic spine are
even worse and the absolute error ranged from 0.86 mm
to 6.37 mm, comparing with normal spine cases ranging
from 0.4 mm to 1.7 mm. The results implied that the
osteoporotic spine might behavior differently from
normal spine. It suggested that the inverse kinematic
model could not be applied directly on the osteoporotic
lumbar spine to predict the intervertebral movement. The
sources of prediction error include the error of link
lengths caused by the morphological difference between
the osteoporotic and normal vertebral bodies, and the
degenerated discs. Therefore, the kinematic model of
osteoporotic spine needs to be further improved to take
vertebrae deformity and disc degeneration as potential
functions.
Conclusions
An inverse kinematic model for predicting intervertebral
movements is evaluated on osteoporotic spine. The
pioneering results indicated that the model is not
applicable on osteoporotic spine for predicting
intervertebral movement. In order to utilize inverse
kinematic techniques to predict intervertebral
movements, other physiological constraints related to
osteoporosis have to be imposed on the model. In
addition, it is observed that there is a negative correlation
between the T-score and the total range of rotation of
lumbar spine. It is guessed that the larger disc height in
osteoporotic spine provides larger deformable range. This
proposition should be validated by further studies.
References
[1] PEARCY, M.J., and HINDLE, R.J. (1989): ‘New
Method for the Non-invasive Three dimensional
Measurement of Human Back Movement’, Clin.
Biomech., 4, pp. 73-79.
[2] SUN L.W., LEE R.Y.W., LU W., and LUK D.K.
(2004): ‘Modelling and Simulation of the Intervertebral
Movements of the Lumbar Spine Using an Inverse
Kinematic Algorithm’, Medical & Biological
Engineering & Computing, 42(6), pp. 740-746.
[3] BRYANT J.T., REID J.G., SMITH B.L., and
STEVENSONL.M. (1989): ‘Method for Determining
Vertebral Body Positions in the Sagittal Plane Using Skin
Markers’, Spine, 14(3), pp. 258-265.
Acknowledgements
The authors would like to acknowledge the funding
support of the Hong Kong Research Grant Council
(Competitive Earmarked Research Grant CERG PolyU
5251/04E).
Discussion
IFMBE Proc. 2005;9: 258

Sports and rehabilitation biomechanics
DYNAMIC CONTROL OF THE TRUNK IN OBESE SUBJECTS
C. Loong 1 , S. Yeung 1 , S. Kwong 1 , N. O'Dwyer 2 , R. Lee 1
1 Rehabilitation Sciences, Hong Kong Polytechnic University, Hunghom, Hong Kong
2 Exercise and Sport Science, The University of Sydney, Sydney, Australia
rsrlee@polyu.edu.hk
Abstract
Obesity and low back pain are becoming public health
concerns all over the world. However, most studies
about obesity were related to health risks such as
cardiovascular diseases and diabetes rather than low
back pain. The purpose of this study was to examine
the dynamic control of trunk in sitting in response to
unexpected, sudden perturbation. Fifty four obese
and non-obese subjects were recruited in this study.
They were subjected to an anteroposterior force at the
chest while in sitting. The force was suddenly
released, and the resulting trunk movements were
measured by gyroscopic sensors. It was generally
observed that obese subjects adoped a flexed sitting
posture after perturbation while non-obese subjects
adopted an extended posture. Obese subjects did not
exhibit a large amount of sway immediately after
perturbation.
Introduction
Obesity has been shown to be related to increase in the
lordosis of lumbar spine and the lumbosacral angle [1].
Control of body posture was found to be affected by
obesity and an increase in the sway of the trunk was
observed during quiet standing [2], which would lead to
poor dynamic balance [3]. It was also found that there
were differences in the gait characteristic between nonobese
and obese subjects [4]. Obesity was found to affect
the movements of trunk and functional performance [4].
However, most previous study examined whole body
posture rather than the dynamic control of spine.
Information about dynamic control of the spine would be
essential as instability of spine is an important cause low
back pain [5]. The purpose of this study was to examine
the dynamic control of the trunk in sitting in response to
unexpected, sudden perturbation.
Methods
Fifty four subjects, aged from 45 and 65, were recruited
for this study. Middle-age subjects were examined
because the prevalence of obesity was found to be the
highest in this age group [6]. They did not have any
spinal pathology, rheumatological disorders, fractures,
dislocations, spinal surgery or low back pain within the
last 12 months that requires medical attention or
treatment. They were divided into two groups according
to their body mass index (BMI) which was defined as
weight divided by square of height (i.e. kg/m 2 ) – (1)
Obese group: BMI ≥ 23; (2) Non-obese group: BMI

Sports and rehabilitation biomechanics
Figure 1. Sagittal movements of the lumbar spine after an
unexpected perturbation. (thin line-obese subjects, thick
line- non-obese subjects)
Discussion
The present study showed that the mechanical response
of the trunk to perturbation generally followed a second
order response. The number of body sways and the
time required to reach a stable posture after sudden
release in the obese subjects were found to be less than
those of non-obese subjects. However, the literature
[2] demonstrated that obese teenagers exhibited increased
body sway during quiet standing. There might be a
difference in the mechanical behaviour between obese
teenagers and middle age adults. The loading conditions
of two situations were also different. Our study involved
dynamic perturbation whereas the earlier work examined
quiet standing. Another finding of this study was that the
final posture of obese and non-obese subjects were very
different. This may indicate that the two groups of
subjects employed different strategies in maintaining
balance.
[7]WHO Expert Consultation (2004): ‘Appropriate
body-mass index for Asian populations and its
implications for policy and intervention strategies’,
Lancet, 2004 Jan 10, 363(9403), pp. 157-163.
Acknowledgements
The authors wish to thank the Hong Kong Polytechnic
Internal Competitive Research Grant in providing
funding of this research project (ICRG A-PF26).
Conclusions
This study showed that obesity significantly affect the
dynamic control of the trunk during sitting. Spinal
instability and motor control deficits are important causes
of low back pain, and thus further research should be
conducted to examine the biomechanical mechanisms
which would explain the changes in sitting posture and
postural sway after perturbation
References
[1]Ridola C., Palma A., Ridola G., Sanfilippo A.,
Almasio PL., Zummo G. (1994): ‘Changes in the
lumbosacral segment of the spine due to overweight in
adultes’, Ital J Anat Embryol., 99(3), pp. 133-143.
[2]Bernard PL., Geraci M., Hue O., Amato M., Seynnes
O., Lantieri D. (2003), ‘Influence of obesiy on postural
capacities of teenagers’, Ann Readapt Med Phys., 46(4),
pp. 184-190.
[3]Goulding A., Jones IE., Taylor RW., Piggot JM.,
Taylor D. (2003): ‘Dynamic and static tests of balance
and postural sway in boys: effects of previous wrist bone
fractures and high adiposity’, Gait Posture, 17(2), pp.
136-41.
[4]Spyropoulos P., Pisciotta JC., Pavlou KN., Cairns
MA., Simon SR. (1991): ‘Biomechanical gait analysis in
obese men’, Arch Phys Med Rehabil., 72(13), pp. 1065-
1070.
[5]Byl NN., Sinnott PL.. (1991): ‘Variations in balance
and body sway in middle-aged adults: subjects with
healthy backs compared with subjects with low-back
dysfunctions’, Spine, 16, pp. 325-330.
[6]Schoenborn CA., Adams PF., Barnes PM. (2002 Sep
6): ‘Body weight status of adults: United States, 1997-
1998’, Adv Data, 330, pp. 1-15.
IFMBE Proc. 2005;9: 260

Sports and rehabilitation biomechanics
A TECHNIQUE FOR EVALUATING INFANTS’ MUSCLE ACTIVITY
USING SURFACE EMG
N. Östlund*, S. Håkansson**, U. Edström*, J.S. Karlsson*
*Department of Biomedical Engineering & Informatics, University Hospital, Umeå, Sweden
**Department of Paediatrics, Umeå University, Umeå, Sweden
E-mail: stefan.karlsson@vll.se
Abstract: There is a lack of reliable indicators of
the prognosis in infants with obstetric lesion of
the brachial plexus. The aim of the present study
was to develop a method for evaluating muscle
activity in newborn infants using surface
electromyography (EMG). In healthy infants the
normal EMG patterns of brachial muscles
activated by the Moro reflex were investigated,
in order to evaluate the proposed method. The
results show that it is possible to obtain signals
with sufficient quality for on-set muscle activity
detection.
Introduction
The incidence of obstetric brachial plexus palsy
(OBPP) is in the range 0.5-2 per 1000 live births,
and has not decreased during the past 20 years [1].
The mechanism of injury is usually caused by
traction to the brachial plexus during delivery.
Although most cases (75- 90 %) resolve without
surgical intervention [2] there is a need for refined
objective criteria that can herald spontaneous
recovery or winnow out cases eligible for microsurgical
neuronal repair [1].
Electromyography (EMG) is a well-established
procedure for study of the muscle function, and
seems to be ideal tool for diagnostic purposes.
Using invasive EMG technique it may be
difficult to correlate muscle activity to the clinical
status of the infant [1]. Surface EMG technique
detects the superimposition of many different motor
unit action potentials generated by a large number
of active motor units, i.e., the interference signal.
The repeatability has been shown to be higher and
long-term monitoring becomes possible. Other
potential advantages of the superficial method are
the elimination of noxious input and risk of
infection.
The objective of this study was to investigate
surface EMG pattern in healthy infants, with the
perspective of applying this method as a
diagnostic/prognostic tool in the management of
OBPP.
Materials and Methods
Subjects: Fifteen healthy infants at age 1-4 days
were participating in the study. The parents
approved their infant’s participation in the study
after been given information in both written and
oral form. The Ethical Committee, Faculty of
Medicine, Umeå University, Umeå, Sweden
approved the study.
Surface EMG: For all subjects EMG signals
were recorded, with surface electrodes, from the
biceps brachii, triceps brachii, and palmar portion
of the thenar muscles on both left and right arms
and hands. The skin was first cleaned with alcohol.
The electrodes (see Fig. 1) were each made of two
sintered Ag/AgCl pellets that had been cast in
silicon rubber and for every measurement a
conductive gel was used. Because the electrode
consisted of two pellets a bipolar recording could
be obtained by using only one electrode. The
recording area of the electrodes was 2x28 mm 2 and
the centre-to-centre distance between the two
Ag/AgCl pellets was 5 mm.
Figure 1. A schematic drawing showing the locations of the
measuring electrodes (filled circles) and reference
electrodes (unfilled circles) and a close-up drawing of a
measuring electrode.
IFMBE Proc. 2005;9: 261

Sports and rehabilitation biomechanics
Fig. 2 for an example. Motion artefacts and
electrodes with poor connection did not affect the
ability to interpret the EMG signals at group level.
The proposed method shall be further evaluated
in infants with OBPP.
References
Figure 2. EMG recorded from a healthy infant during a
Moro reflex.
Two reference electrodes were used and placed
at the seventh cervical vertebrae (C7) on both the
subject and the examiner.
Experimental protocol: To standardise the
motor stimulus, the Moro reflex was used. The
reflex was elicited by first flexing the neck
followed by an extension of the neck with the head
falling backwards into the hands of the investigator.
During the performance of the reflex the head was
guided to remain in the mid-line. The reflex was
elicited five times and all infants were recorded on
video during the procedure. Later, the video
sequences were used to pick one reflex sequence
from each infant that were used for the analysis.
Data acquisition: The surface bipolar EMG
signals were sampled at 2 kHz using a 20-bit
converter with input range ± 2.5 V using our
custom made wireless acquisition system [3]
(CMMR > 100dB, system input noise < 2 µV RMS ,
input impedance >10 GΩ). Test contractions before
the Moro-reflex were also made to secure good
electrode-skin contact. Before sampling, the EMG
signals were amplified (fix gain 50) and to remove
aliasing also low-pass filtered at 430 Hz. After
sampling, signal was decimated to give a sampling
frequency of 1 kHz and the signals were then
digitally band-pass filtered at 25 to 350 Hz to
remove movement artefacts in the low-frequency
region. Notch filters were used at 48 to 52 Hz and
148 to 152 Hz.
[1] GERT VAN DIJK J., PONDAAG W., MALESSY J.A.
(2001): ‘Obstetric Lesions of the Brachial
Plexus’, Muscle Nerve, 24, pp. 1451-1461.
[2] BAGER B.(1997): ‘Perinatally Acquired Brachial
Plexus Palsy – a Persisting Challenge’, Acta
Paediatr., 86, pp. 1214-1219.
[3] KARLSSON J.S., BÄCKLUND T., EDSTRÖM U.
(2003): ‘A New Wireless Multi-Channel Data
System for Acquisition and Analysis of
Physiological Signals’, Proc of 17th Int. Symp.
on Biotelemetry, Brisbane, Australia.
Results and Discussion
The general aim of this study was to develop a
method for non-invasive monitoring and follow up
of OBPP with the ultimate objective to improve
diagnostic and prognostic tools for optimal therapy.
The muscles investigated were chosen to reflect
different nerve branches so that for children with
OBPP it would be possible to estimate the extent of
the detriment.
The custom made electrodes worked well
during the whole study and EMG-signals of good
quality were obtained from the different sites, see
IFMBE Proc. 2005;9: 262

Biomechanics
APPLICATION OF RESISTANCE TESTING FOR IDENTIFYING SHUNT
RESPONDERS IN IDIOPATHIC NORMAL PRESSURE HYDROCEPHALUS
A. Marmarou 1
1 Department of Neurosurgery, Virginia Commonwealth University Medical Center, Richmond, VA,
United States
Abstract
The resistance to outflow (R o ) of CSF is considered to be
the impedance of flow offered by the CSF absorption
pathways. Several methods can be used to measure the
resistance for example in the Katzman test a pump
introduces mock CSF fluid or saline at a known rate
through a needle placed in the lumbar subarachnoid
space. The outflow resistance as defined by the Katzman
infusion test, is the difference in the final steady state
pressure reached and the initial pressure divided by the
infused flow rate. In distinction, the bolus method for
estimating R o involves injecting a known volume into the
lumbar subarachnoid space at a fixed rate. The advantage
of the bolus method is that it also provides a measure of
the brain compliance as defined by the Pressure Volume
Index. For unknown reasons, R o values determined by
the infusion method (Katzman) are higher than that of the
bolus technique and it is important to distinguish between
the reported thresholds for each method when the values
are used for predicting shunt responsive INPH. The
utility of this method in identifying shunt responders will
be presented and an explanation of the difference
between infusion and bolus methods and the
mathematical implications of CSF model development
will be described.
amarmaro@vcu.edu
IFMBE Proc. 2005;9: 263

Biomechanics
Using the Pulsatility of Blood and CSF Flows to Probe the Biomechanical
State of the Craniospinal System
A New Methodology for Noninvasive Measurement of Intracranial Compliance and Pressure
Noam Alperin, PhD
University of Illinois at Chicago/ Radiology, Chicago, USA
Alperin@uic.edu
Abstract:
The pulsation of the cerebrospinal fluid (CSF) has
fascinated investigators of the intracranial physiology
since it has first been documented by invasive CSF
pressure measurements. Advances in dynamic
Magnetic Resonance Imaging (MRI) now enable
accurate characterization of the CSF flow dynamics
and improve our understanding of the origin for CSF
pulsation. This, in turn, has lead to the development of
a noninvasive method to measure important
biomechanical properties such as intracranial
compliance and pressure by MRI. Currently,
intracranial compliance is measured by injection (or
withdrawal) of a known volume of fluid into the CSF
spaces and recording the resulted change in pressure.
The MRI based methodology utilizes the small changes
in volume and pressure that occur naturally with each
heart beat due to the pulsatile blood flow into the
intracranial compartment. The volume change is
measured from the difference in volumes of blood and
CSF that enter and leave the intracranial compartment
during the cardiac cycle. The pressure change is
derived from the CSF pressure gradient waveform
calculated using the Navier Stokes relationship. Mean
ICP is then derived from the linear relationship
between elastance (inverse of compliance) and pressure.
The principles and the implementation of the MRImethod
will be presented. In addition, validation
studies to date with nonhuman primate animal model,
computer simulations, healthy human subjects and
patients will be presented.
The neurophysiologic basis of the MRI technique:
In a closed system such as the intracranial compartment
(See Fig. 1), pressure and volume are related. A change in
pressure due to increase (or decrease) in volume is
determined by the overall mechanical elastance of the
compartment. The relation between intracranial volume
and pressure has been studied extensively using invasive
techniques in animals and in humans. Ryder et al. [1] and
others [2-4] studied the pressure-volume relationship by
injection of fluid into the CSF space and measuring the
change in pressure. Marmarou et al determined that the
pressure-volume curve is a mono-exponential curve and
therefore, the ratio of pressure and volume changes during
the cardiac cycle, dP/dV, is a linear function of the
pressure [2].
Method:
Similarly, the MR-ICP method provides a measure of
elastance from the ratio of maximum pressure and volume
changes that occur with every heart beat. Intracranial
pressure is then derived through the linear relationship
between elastance and pressure. The small volume change
that occurs with each cardiac cycle is analogous to the
injected volume used in the volume-pressure response test.
Volume and pressure changes occur because of the
pulsatile nature of blood flow. During systole, the
intracranial volume increases because volumetric flow into
the cranial vault exceeds volumetric outflows (arterial
inflow exceeds venous and CSF outflows). During
diastole, volumetric outflow is larger and the intracranial
volume returns to its initial volume.
Fig. 1: The craniospinal flow-volume-pressure
model used in the derivation of ICP by MRI.
Validation studies:
The volume of the entire intracranial (IC) space is about
1500mL, the change in that volume during the cardiac
cycle is on the order of one mL, less than 0.1%! Therefore,
reproducible and accurate measurements of IC volume
change (ICVC) pose a challenge. The validity of the ICVC
measurement was assessed from studies in patients with
pathologies that affect the ICVC in a predicted manner,
and the inherent accuracy was assessed with a specially
designed craniospinal flow phantom [5]. These studies
were necessary since currently there is no alternative
method that could be used as a reference. The average
IFMBE Proc. 2005;9: 264

Biomechanics
maximum ICVC value measured in the phantom by MRI
was within 5% of the ICVC value measured independently.
The high reliability by which ICVC is measured by MRI is
attributed to the excellent temporal response of the cine
phase contrast MRI technique, which enables accurate
measurements of changes in volumetric flow rates.
The intracranial pressure change during the cardiac cycle is
derived from change in the CSF pressure gradient. The
relationship between time varying change in pressure and
in pressure gradient was evaluated experimentally with a
nonhuman primate [6] and theoretically with
computational fluid dynamics (CFD) [7]. The experimental
validation required the use of a large nonhuman primate
(baboon) because hydrodynamics is scale-dependent and
important fluid dynamic parameters such as the size of the
spinal canal and heart rate in baboons are similar to those
of humans.
The reproducibility of the pressure change measurement by
MRI is therefore determined by the reproducibility of the
peak-to-peak amplitude of the CSF pressure gradient
measurement. The CSF pressure gradient measurement
reproducibility was assessed from repeated MRI scans of
healthy subjects. A measurement variability of 8% was
found in these studies [6]. The MRI-derived intracranial
compliance and pressure (MR-ICP) are derived from the
ratio of the pressure and volume changes. Therefore the
overall MR-ICP measurement reproducibility is the square
root of the sum of the square of individual fractional
standard deviations, i.e., the fractional SD of the volume
and the pressure change measurements, which are currently
approximately 8% each. Therefore the overall MR-ICP
measurement variability is approximately 10%.
Direct comparison between MRI derived elastance index
and mean ICP value measured invasively at the time of the
MRI study in 5 patients demonstrated a linear relationship,
as expected, with high degree of correlation (R 2 = 0.96)
[6]. The false positive rate of the method was determined
by measurements in healthy subjects. A total of 71
simultaneous measurements of ICP were performed in
twenty three young adults (20 males, range 20 to 39, mean
age 25 +/- 5 years) with no known neurological problems.
ICP values range from 3.5 to 17.1 mmHg; most
measurements were between 7 and 9 mmHg. Wide
distribution of ICP values measured invasively in normal
human subjects has been previously reported: from 2 and
up to 18 mmHg [8]. Based on the current view that ICP
value of 20 mmHg is a critical threshold for elevated ICP
[10], no false positives were measured with MRI (i.e., a
zero-false positive rate).
Discussion
The MR-ICP measures intracranial compliance and
pressure based on neurophysiologic and fluid mechanics
principles. The potential role of the MR-ICP method is,
however, different than invasive ICP monitoring. Since the
MRI study provides a “snapshot” measurement of ICP
(analogous to noninvasive blood pressure measurement) its
potential role is mainly for diagnostics. As a diagnostic test
it may improve the understanding and treatment of several
neurological problems (e.g., hydrocephalous, hemorrhagic
strokes, Chiari Malformations, head trauma, and possibly
dementia). The method was recently applied to study the
effect of decompression surgery in Chiari Malformations
patients, and for the first time, it demonstrated the role of
intracranial compliance in the pathophysiology of this
poorly understood disorder [10].
This work may also contribute to a noninvasive bed-side
ICP monitoring by demonstrating what parameters needs
to be measured in order to estimate ICP noninvasively;
alternatively it may provide a reference for devices that are
sensitive to changes in IC compliance and pressure.
REFERENCES
[1] Ryder HW, Espey FF, Kimbel FD, Penka EJ,
Rosenauer A, Evans JP. The mechanism of change in
cerebrospinal fluid pressure following an induced
change in volume of the fluid space. J Lab Clin Med
1953; 41: 428-435.
[2] Marmarou A, Shulman K, La Morgese J.
Compartmental analysis of compliance and outflow
resistance of the CSF system. J of Neurosurgery
1975; 43: 523-534.
[3] Sklar FH, Elashvili I. The pressure-volume function of
brain elasticity. J of Neurosurgery 1977; 47: 670-679.
[4] Szewczykowski J, Sliwka S, Kunicki A, Dyko P,
Korsak-Sliwaka J. A fast method of estimating the
elastance of intracranial system. J of Neurosurgery
1977; 47: 19-26.
[5] Alperin N, Kadkhodayan Y, Loth F, Yedavalli R. MRI
measurements of intracranial volume change: A
phantom study. Proc. Intl. Soc. Mag. Reson. Med. 9
2001; 3: 1981.
[6] Alperin NJ, Lee SH, Loth F, Raksin PB, Lichtor T.
MR-Intracranial Pressure (ICP): A method to measure
intracranial elastance and pressure noninvasively by
means of MR Imaging: Baboon and human study.
Radiology 2000; 217: 877-885.
[7] Loth F, Yardimci MA, Alperin N. Hydrodynamic
modeling of cerebrospinal fluid motion within the
spinal cavity. J of Biomechanical Engineering 2001;
123: 71-79.
[8] Ayer JB. Cerebrospinal fluid pressure from the
clinical point of view. Assn. Res. Nerv. Mental Dis
1924; 4: 159-171.
[9] Marmarou A, Eisenberg HM, Foulkes MA. Impact of
ICP instability and hypertension on outcome in
patients with severe head trauma. J. Neurosrg. 1991;
75: s59-s66.
[10] Sivaramakrishnan A, Alperin N, Surapaneni S,
Lichtor T. Evaluating the Effect of Decompression
Surgery on CSF Flow and Intracranial Compliance in
Patients with Chiari Malformation. Neurosurgery.
Volume 55(6) 1344-1350 (2004)
IFMBE Proc. 2005;9: 265

Biomechanics
ESTIMATION OF CSF OUTFLOW CONDUCTANCE - REPEATABILITY AND
PRECISION
N. Andersson 1, 3 , J. Malm 2 , T. Bäcklund 1, 3 , A. Eklund 1, 3
1 Department of Biomedical Engineering and Informatics, Umeå University Hospital, Umeå, Sweden
2 Department of Clinical Neuroscience, Umeå University, Umeå, Sweden
3 Centre for Biomedical Engineering and Physics, Umeå University, Umeå, Sweden
nina.andersson@vll.se
Abstract
A new apparatus for performing infusion tests in a
standardized, automated way with real time analysis of
measurement precision was developed. The purpose of
this study was to evaluate the repetitive as well as
individual precision of measurements of outflow
conductance (C out ) on a hydrodynamic model and present
preliminary patient data. C out was measured on five steel
pipes of different length and diameter (n=6), as well as
repeatedly on 3 patients with suspected or treated
Idiopatic Adult Hydrocephalus Syndrome. We conclude
that the repetitiveness of the C out measurement with the
new infusion apparatus is high, and that the 95% CI of
individual C out measurements reflect the precision of the
parameter in an important and valuable way.
Introduction
The brain and spinal cord are protected from physical or
chemical injury by cerebrospinal fluid (CSF). CSF also
nourishes the central nervous system 1 . A disturbance to
the CSF system can lead to the development of Idiopatic
Adult Hydrocephalus Syndrome (IAHS) with main
clinical features like gait disturbance, urinary
incontinence and cognitive decline. Patients with IAHS
are treated by surgically placing a shunt system which
passes CSF in a silicon tube from the ventricles of the
brain to the abdomen. Selecting patients suitable for
shunt surgery is difficult and finding good predictive tests
has long been an important issue to research in the area 2 .
One predictive test is the infusion test, where the
dynamics of the CSF system is investigated. We have
developed a new apparatus for performing infusion tests
based on constant pressure levels. The apparatus uses
standardized and automated protocols and includes a real
time estimate of the precision of C out . The purpose of this
study was to evaluate the apparatus concerning statistical
precision in measuring C out , for individual measurements
as well as for repetitiveness. The evaluations were
performed on a hydrodynamic model simulating the CSF
system, and on preliminary patient data.
Methods
The study was divided into two parts; the first was
performed on a hydrodynamic model, and the second on
a patient material. An automated protocol designed for
investigations of patients with suspected or treated IAHS
was applied to the hydrodynamic model as well as 3
patients. On the model five steel pipes of different length
and inner diameter represented the outflow conductance
of the system. Measurements were performed without
and with the addition of physiological pressure
fluctuations. For the patients, two measurements of C out
were performed during the same session. C out was
determined using linear regression between six constant
pressure levels and their corresponding flows.
Results
The mean C out ± 95% CI of one steel pipe when no
pressure fluctuations were added was 17.88 µl/(s kPa).
Mean estimated CI in these individual investigations was
0.18 µl/(s kPa) (n=6). With the addition of pressure
fluctuations C out for that pipe was 16.50 µl/(s kPa) with
mean CI= 2.05 µl/(s kPa) (n=6). The outflow
conductance based on total measurement time from all
six repetitions on that pipe was C out pipe = 16.42 ± 0.57
µl/(s kPa). A general linear model for C out between 3.32
and 32.1 µl/(s kPa) for all five pipes, with pipes as a
factor, resulted in a repeatability of ± 1.05 µl/(s kPa)
(95% CI, n=30).
Table 1 shows the results from the patient investigations.
Table 1: C out from repeated patient investigations.
Discussion
Model
IFMBE Proc. 2005;9: 266

Biomechanics
To estimate the limitations of the model and the infusion
apparatus, C out was first measured without the addition of
pressure fluctuations. Repeated measurements on one
steel pipe gave an individual mean CI = 0.18 µl/(s kPa)
which is small as compared with precision when
fluctuations were added. This indicates that the infusion
apparatus in it self is stable and has a precision good
enough for performing C out investigations. In a C out
interval betweeen 3.32 to 32.1 µl/(s kPa)
when fluctuations were added, an overall CI of 1.05 µl/(s
kPa) (n=30) was achieved. This shows that the
repetitiveness of the investigation is high.
On the hydrodynamic model all individual C out ± CI
include C out pipe of their corresponding pipe. This indicates
that the individual CIs do not overestimate the precision
of the measurement, but can be viewed as an indicator of
the interval in which the true C out is. The fact that CI was
smaller for C out pipe than for individual C out for all pipes
also indicates that the precision of the C out parameter
increases with measurement time.
Patients
The preliminary patient investigations point towards a
good repeatability between measurements also in the
clinical setting. Regarding the precision of an individual
C out investigation, the CIs from the repeated patient
investigations overlap in all three cases. Thus we find it
reasonable to believe that the 95% CIs of the linear
regression between net flow and pressure reflects the
precision of the C out investigation in a good way. A
property which we believe to be very important and
valuable when the C out parameter should be weighted
together with other criteria in deciding whether as to
shunt a patient or not.
Conclusions
We conclude that the ability of the new infusion
apparatus to measures C out in a repetitive manner is high.
The study also indicates that the individual 95% CI
calculated in real time during each model/patient
investigation does reflect the precision of the current
measurement in a valuable way.
References
1. BERGSNEIDER M (2001): 'Evolving concepts of
cerebrospinal fluid physiology', Neurosurgery Clinics of
North America, 12(4), pp. 631-638.
2. VANNESTE J A (2000): 'Diagnosis and management
of normal-pressure hydrocephalus', J. Neurol, 247, pp. 5-
14.
IFMBE Proc. 2005;9: 267

Biomechanics
COCHLEAR AQUEDUCT PATENCY IN TYMPANIC MEMBRANE
DISPLACEMENT MEASUREMENT FOR INTRACRANIAL PRESSURE
ASSESSMENT
S. Shimbles 1 , K. Banister 1 , C. Dodd 1 , D. Mendelow 1 , I. Chambers 1
1 Newcastle General Hospital, Newcastle upon Tyne, UK
i.r.chambers@ncl.ac.uk
Abstract
We have previously reported [1] on the association
between tympanic membrane displacement (TMD) and
intracranial pressure (ICP). This relationship is based
upon a patent cochlear aqueduct and different methods
can be used to help with this assessment. We have
compared three different methods based upon values
obtained when the patient is supine and sitting.
All three methods showed a significant association
between TMD and ICP with relatively tight confidence
intervals but predictive limits are wide. Using the
comparison method of assessment, at 100nL the
variability of the ICP distribution is such that it could
only be predicted to be between approximately -20 and
+40mmHg with 95% certainty.
Although there is clearly a strong relationship between
TMD and ICP such accuracy limits the use of the
technique. Even with different methods of assessing
cochlear aqueduct patency the variability of the response
makes it difficult to make clinical decisions.
Introduction
A reliable method of non-invasive ICP measurement
would be of significant benefit in the management of
patients with CSF circulation disorders and could aid the
decision of shunt introduction or replacement. The use of
TMD measurements may provide an indirect index that
can be related to ICP and earlier work has shown that
stimulation of the acoustic reflex can induce a TMD that
was related to ICP[2]. An aural stimulus above the
acoustic reflex threshold (ART) will cause a contraction
of the stapedius muscle; this results in a movement of the
tympanic membrane (figure 1). The kinematic chain
between the stapedius and the tympanic membrane
involves the ossicles. One of these, the stapes, rests on
the oval window of the cochlear. This is a flexible
membrane; consequently, the pressure of the
intracochlear fluid determines the initial position of the
stapes, and hence the size and direction of the movement.
If there is fluid communication, via the cochlear
aqueduct, between the intracochlear and intracranial
spaces then TMD measurements may be able to provide
an indirect estimate of ICP.
Our previous work has shown a strong negative
correlation between TMD measured in this way and
invasively measured ICP [1] that was consistent with this
theory. However, intersubject variability appeared to be
too great for the technique be clinically useful. Patency of
the CA is fundamental to the technique and its
assessment is therefore central to the validity of the
method. Marchbanks calculated the ratio of the change in
maximum inward displacement between sitting and
supine to the maximum inward displacement in the
supine position. If this ratio was greater than 0.1 for all
intensities tested in a given ear the aqueduct was assumed
to be patent.
Applying this technique from data collected from healthy
volunteers the percentage of successful tests was low
when compared with histological studies of cochlear
aqueduct patency [3]. Nevertheless a strong correlation
existed between TMD and ICP in these patients but the
predictive value was low. If an improved method of
patient selection could be created this might increase the
potential use of the technique.
Methods
Data from thirty-six patients who underwent invasive
measurement of ICP as part of their clinical management
was studied. Local Ethics Committee approval and
informed consent were obtained prior to patient
recruitment. Patients had simultaneous TMD
measurements using the MMS analyser, with stimulus
intensities up to 20dB above the ART subject to a
specified safety limit of 110dB. The TMD measurements
were repeated at two monthly intervals for up to twelve
months.
IFMBE Proc. 2005;9: 268

Biomechanics
For each paired TMD / ICP measurement three different
methods were used to establish whether the cochlear
aqueduct was patent (i.e. test valid). These were:
1. Marchbanks maximum inward displacement
ratio method
2. A new algorithm based upon the subsequent
TMD measurements
3. A simple postural difference test
Our new algorithm used all the TMD data available from
each patient. Each patient had up to 6 visits where TMD
readings were obtained at up to three stimulus intensities.
Therefore for each ear tested up to 18 sets of readings
were obtained where comparison between the sitting and
supine positions was possible. Comparison was made on
the basis of the mean value of the Vm (figure 1).
The simple postural test involved observing the change in
TMD with a change in posture from sitting so supine –
increased ICP in the supine position would be expected to
produce a lower TMD measurement; if this was observed
then the patient’s CA was considered to be patent.
The ICP values obtained from each of the selected
patients were compared with Vm obtained at the ART +
10dB. Where both left and right CAs were classified as
patent the mean Vm value was used. Correlation and
regression analysis was performed.
Results
A total of 36 patients were recruited, ages ranging from 5
to 76 (median 31). Twenty two were female and 14 male.
ICP readings ranged from -12mmHg to 49 mmHg in the
sitting position and -22mmHg to 37mmHg supine. The
values of TMD (Vm) ranged between -780nL and
+532nL.
Each of the three methods identified a different number
of patent CAs. Methods 1 and 2 produced a similar result
(39% and 44% respectively) whilst method three
classified only 20% as patent. Regression analysis found
a strong correlation between Vm and ICP for the data
obtained from all three methods (p < 0.01 in all cases).
Figure 2 shows the results from the new algorithm, these
are similar, but not identical, to the two other methods.
The relatively tight 95% confidence intervals show that
there is strong association but the prediction limits are
wide. At a TMD value of 100nL the CI shows a tight
distribution of the standard error of mean ICP. However
the variability of the ICP distribution is such that it could
only be predicted to be between approximately -20 and
+40mmHg with 95% certainty.
Discussion
The assessment of CA patency carried out on each patient
in our initial study led to the exclusion from the analysis
of more than half of the patient group. Thirteen out of
twenty-nine patients were judged to have at least one
patent cochlear aqueduct and this is lower than
expected [3].
Three very different methods of assessing cochlear
aqueduct patency have been used. Although a very
simple postural test excludes much more data than other
methods it still demonstrated a strong relationship
between TMD and ICP. However it only identified one
patient as having both CAs patent compared to seven and
six using Marchbanks method and our method.
Conclusions
Although there is a strong relationship between TMD and
ICP the variability of the response is such that, even with
different methods of assessing cochlear aqueduct
patency, it does not produce accurate information that can
be relied upon to make clinical decisions.
References
[1] Chambers IR, Shimbles S, Dodd C, Banister K and
Mendelow AD (2004):
‘Clinical comparison of tympanic membrane
displacement with invasive ICP measurements’, Twelfth
International Symposium on Intracranial Pressure and
Brain Monitoring. Hong Kong, 2004, p54
[2]Madan S, Burge DM, Marchbanks RJ (1998)
Tympanic membrane displacement testing in regular
assessment of intracranial pressure in eight children with
shunted hydrocephalus. J. Neurosurg. 88:983-995
[3] Wlodyka J (1978): ‘Studies on cochlear aqueduct
patency’, Ann Otol Laryngol 87:22-28
Acknowledgements
This study was supported by a grant from Action
Medical, The Garfield Weston Foundation.
IFMBE Proc. 2005;9: 269

Biomechanics
COMPUTERISED ASSESSMENT OF CSF DYNAMICS IN HYDROCEPHALIC
PATIENTS
P. Smielewski 1 , M. Czosnyka 1 , Z. Czosnyka 1 , J. Pickard 1
1 Department of Clinical Neurosciences, University of Cambridge, , UK
ps10011@medschl.cam.ac.uk
Abstract
Hydrocephalus, as a disorder of CSF circulation, is
mostly treated with mechanical shunt implants. In
order for the shunt to be effective certain conditions
regarding the mechanical properties of the CSF
system have to be met. The paper presents a summary
of the authors over 10 years experience in using
computer aided approach to investigation of the
disease in patients before and after shunting.
Introduction
Hydrocephalus manifests with excessive accumulation of
fluid within the brain. Implantation of hydrocephalus
shunt is a standard way of management of
communicating hydrocephalus. As shunting is purely
mechanistic treatment, which radically affects pressurevolume
compensation, the hydrodynamics of patient's
own compensation should be ideally examined before a
shunt is implanted. Testing of CSF dynamics, although
invasive, may help with the decision about surgery. It
also provides a baseline information for further
management of shunted patient, when complications,
such as shunt blockage, under-, and over-drainage, arise.
In such cases, physiological measurement may aid the
decision about shunt’s revision.
Methods
The methodology employed in Cambridge includes
constant rate infusion test and overnight ICP monitoring
(Figure 1).
Figure1: During constant rate infusion test mock CSF
fluid is being infused into the CSF space (e.g Ommayer
reservoir, shunt pre chamber or lumbar space) and
pressure response is recorded.
Interpretation of the infusion test is based on the
mathematical model of CSF pressure volumecompensation,
introduced by A. Marmarou [1] and
modified in later studies (Avezaat and Eijndhoven [2],
Frieden and Ekstedt [3]) – Figure 2. The following
parameters of the model are identified: resistance to CSF
absorption (Rcsf), coefficient of compliance (E),
pressure-volume index (PVI), sagital sinus pressure (Pss)
and the CSF formation rate. The provision is also made
in the model for using only one needle for both infusion
and measurements (not recommended practice but often
unavoidable). To facilitate the in-vivo testing of shunts
the CSF space model has been extended with the shunt
branch so that the shunt resistance and its critical pressure
can also be estimated.
Figure 2. Model of the CSF space identified during the
infusion test. Part of the model on the right hand site is
only present for patient with shunts in situ.
If the results of the infusion test are inconclusive
overnight monitoring of ICP is performed. ICP signal is
being analysed for existence of spontaneous waves and
time trends of indices describing the CSF compensatory
reserve are calculated. In addition ABP and sometimes
transcranial Doppler is also being recorded providing
information about the vascular component of the disease
(CO2 reactivity and autoregulation) [4].
Data acquisition and analysis is performed using in-house
build software ICM+ [5]. The software collects the ICP
(and sometimes also ABP and FV) data and performs online
analysis producing time trends of parameters
reflecting brain compliance. Special tool is provided to
analyse the behaviour of ICP mean and ICP pulse wave
amplitude time trends during the constant infusion period
(Figure 3).
In order to facilitate in-vivo shunt testing, the software
contains an extensive database of shunts with their flow
characteristics measured in vitro in the shunt laboratory
[6]. This allows the software to estimate the critical
pressure for each individual case. For suspected
overdrainage a tilting test is performed (Figure 4).
IFMBE Proc. 2005;9: 270

Biomechanics
Figure 3: Example of the infusion test recording and
analysis. The top-left chart also shows fitting of the
model derived curved of ICP increase during infusion.
The resistance to CSF outflow can be estimated from the
difference between the plateau pressure during infusion
and the resting pressure. However, in many cases strong
vasogenic waves or an excessive elevation of the pressure
above the safe limit of 40 mmHg do not allow the precise
measurement of the final pressure plateau. Computerized
analysis, on the other hand, produces results even in
difficult cases when the infusion is terminated
prematurely (i.e., without reaching the end-plateau). In
addition to the resistance to CSF outflow it provides
estimates of the elastance coefficient or pressure-volume
index, cerebrospinal compliance and the CSF formation
rate. And for shunted patients it also provides estimates
of the shunt physical characteristic which then compared
against the database of in-vitro lab results provides
conclusive information about the shunt performance.
Conclusions
Physiological monitoring can be useful in the
management of hydrocephalus. Specialised computerised
infusion test helps doctors to exclude patients from
unnecessary shunting, to evaluate shunt functioning invivo,
and to detect vascular components of
hydrocephalus.
Figure 4: Example of the infusion test performed on a
patient with shunt in-situ. Overdrainage test was also
performed at the end. Critical shunt pressure threshold
and the ovedrainage pressure threashold are both shown.
The shunt was diagnosed as undedraining.
Results
Results of using the ICM+ powered constant infusion rate
test in different cases of CSF circulation pathology will
be presented and discussed. In total 1324 number of tests
have been performed in Cambridge from 1992-2004.
63% of tests were performed to assess CSF dynamics
before shunting and 37% to assess shunt function. In a
subgroup of patients whose more detailed clinical data
were available, severity of symptoms (Stein-Langfitt
score) significantly correlated with Rcsf (R=0.21;
p

Biomechanics
RESTING PRESSURE AND ACCURACY OF CEREBROSPINAL FLUID
INFUSION TEST
A. Eklund 1 , N. Andersson 1 , J. Malm 2
1 Biomedical Engineering and Informatics, Umeå University Hospital, Umeå, Sweden
2 Department of Clinical Neuroscience, Umeå University, Umeå, Sweden
anders.eklund@vll.se
Abstract
This study investigates the variation in resting pressure
and how it affects the accuracy of determined
cerebrospinal fluid outflow resistance in infusion tests.
Introduction
In 1965 Hakim and Adams 5 showed that patients with
the typical symptoms of hydrocephalus were improved
after shunting regardless of whether they had an
increased intracranial pressure (ICP) or not. This showed
that the etiology of hydrocephalus was not as simple as
just an increased pressure, and a new syndrome called
Normal Pressure Hydrocephalus (NPH, also named
IAHS) was identified. To diagnose these patients the
interest for more advanced cerebrospinal fluid (CSF)
dynamic investigations, as compared with simple ICPmonitoring,
emerged. Techniques used today are lumbar
constant rate infusion tests with modifications 6 2 , bolus
injections 8 , constant pressure infusions 3 and lumboventricular
perfusion methods 1,4 .
One purpose of these methods is to measure the outflow
resistance (R out ) or outflow conductance (C out ) of the CSF
system. The basic equations for steady state calculation
of R out and C out are
a lumbar infusion test with a constant pressure method.
The investigation started with a determination of P rest .
For IAHS patients P rest was determined as the mean value
over five minutes, taken after 25 minutes of rest in the
supine position. For the shunted patients P rest was taken
after approximately 10 minutes of rest. ICP was then
regulated by means of a computer controlled peristaltic
pump to six different elevated pressure levels. The mean
infusion rate needed to sustain each pressure level was
recorded. For each patient C out was determined as the
regression slope of the net flow as a function of
corresponding pressure levels. Using the linear
relationship the regression resting pressure, P regr , was
determined as the zero-flow crossing of the pressure axis.
Variation of ICP at rest
Three IAHS-patients underwent 18 hours of continuous
ICP registration with a Codman ICP-express. Data was
divided into five minute segments and mean ICP on each
segment was calculated. A histogram over the ICP
variation was constructed.
Results
In the infusion tests there was a significant difference
between P rest and P regr in both groups. For patients
investigated for IAHS P regr was lower (Paired t-test,
p=0.012) and for patients with shunts P regr was higher
than P rest (p=0.001).
Were P lev is a steady state ICP level produced by a
net infusion rate I. P rest is the resting ICP under zero
infusion. The resting pressure is thus an important
parameter for determining the outflow characteristics of
the CSF system. In this study we aim to investigate the
variability of P rest in two ways. The first estimates the
possible error in P rest during a CSF dynamic
investigation, and the second evaluates the natural
variability of ICP over a longer time period. We
furthermore conduct a theoretical inaccuracy analysis for
infusion test methods and evaluate the influence of P rest
on the precision of R out .
Methods
Resting pressure vs regression resting pressure
Forty-five patients investigated for suspected IAHS
(n=21) or for control of shunt function (n=24) underwent
Histogram over the 18 hour long registration of ICP with
patients in the supine position. P rest measured as mean
over five minutes varied with confidence intervals
between ± 3.3 to ± 7.0 mm Hg.
Discussion
IFMBE Proc. 2005;9: 272

Biomechanics
In spite of being such an important parameter, the
methods for determining resting pressure are rarely well
described. With a Gauss approximation the uncertainty of
R out , denoted by ∆R out , are described by
Focusing on the last quadratic term in equations and we
see that there is a dependency on the resting pressure. In
this study we showed that the variation and uncertainty in
P rest can be substantial. The variation is among other
things due to effects from leakage when placing the
needles and taking CSF samples prior to start of the
investigation. This is compensated for by waiting until
ICP is stable before measuring P rest . The waiting period
differs between studies and is typically described as
“After a steady state CSF pressure was obtained”. To
evaluate the approximate time needed for P rest to return to
its true value, the time constant (t csf ) for the CSF system
can be considered. Close to the resting pressure t csf is
calculated from R out and compliance at rest, and is in the
magnitude of 7 to 16 minutes. It will take three times the
size of t csf to compensate 95% of the difference between
the disturbed initial pressure and the true resting pressure.
In this study the difference between P regr and P rest for the
IAHS patients might follow from a too short time period
between placing the needles and measuring P rest . The
difference in shunted patients is probably a consequence
of a lowered ICP prior to the investigation, due to
shunting in the up-right position.
Furthermore, it is well known from the work of Lundberg
7 and others that there are continuous wavelike variations
in ICP, due to vasomotion. These will make P rest vary
with time, depending on when the resting
pressure measurement is taken and how long time is used
for mean value estimation. This was also indicated by the
histograms over the P rest variation in our three patients.
Even if we assumed that ∆P lev =0 and ∆I=0 a ∆P Rest =6 mm
Hg will give a ∆R out =4 mm Hgml/min with I=1.5
ml/min.
4. Gjerris F, Borgesen SE, Schmidt K, et al: , in
Rowan SL (ed): Intracranial presssure
VII: Springer-Verlag, 1986, pp 411-416
5. Hakim S, Adams RD: J Neurol Sci 2:117-126,
1965
6. Katzman R, Hussey F: Neurology 20:534-544,
1970
7. Lundberg N: Acta Psychiatr Scand 36(Suppl
149):1-193, 1960
8. Marmarou A, Shulman K, Rosende RM:
journal of neurosurgery 48:332-344, 1978
Conclusions
This study showed that uncertainty in determined resting
pressure can influence the determined outflow parameter
substantially. A thorough evaluation of the accuracy and
precision of these types of measurements is important. To
perform evidence-based medicine and research this is
essential. By including a precision analysis along with the
estimated output parameters the clinician or researcher
can appraise the result in a better way.
References
1 Borgesen SE, Gjerris F, Srensen SC: Acta
Neurologica Scandinavica 57:88-96, 1978
2. Czosnyka M, Batorski L, Laniewski P, et al:
acta neurochirurgica 105:112-116, 1990
3. Ekstedt J: Journal of Neurology,
Neurosurgery, and Psychiatry 40:105-119, 1977
IFMBE Proc. 2005;9: 273

Biomechanics
INTRACRANIAL PRESSURE MEASUREMENT VIA LUMBAR SPACE
Niklas Lenfeldt MSc.E.P., BM. 1,3 , Lars–Owe D. Koskinen M.D., Ph.D. 1 ,
Aina Ågren-Wilsson M.D. 1 , A. Tommy Bergenheim M.D., Ph.D. 1 ,
Jan Malm M.D., Ph.D. 1 and Anders Eklund Ph.D. 2,3
1 Department of Clinical Neuroscience
2 Department of Biomedical Engineering and Informatics
3 Centre for Biomedical Engineering and Physics
University of Umeå, Umeå, Sweden
niklas.lenfeldt@neuro.umu.se
Abstract: There is still no conclusive evidence that
intracranial cerebrospinal fluid (CSF) pressure
measured via the lumbar space (ICP CSF−LS )
accurately reflects intracranial pressure (ICP). Still,
lumbar measurements become more common when
investigating CSF hydrodynamics. To prove the
correctness of assessing ICP via the lumbar space,
intraparenchymal ICP (ICP IP ) and ICP CSF−LS were
simultaneously measured in ten patients suffering
from Idiopathic Adult Hydrocephalus Syndrome
(IAHS). The results showed that changes in ICP IP
were simultanously accompanied by equal changes
in ICP CSF−LS . Hence, assessing CSF hydrodynamics
via lumbar space is a viable method.
Conclusions
Our investigation shows conclusively that changes in
ICP can reliably be traced by gaining access to the
lumbar space and measure the CSF pressure. Thus, the
lumbar CSF hydrodynamic investigation is well suited
to study intracranial pressure relations in
communicating hydrocephalus.
References
1. Lundkvist, B., et al., Cerebrospinal fluid
hydrodynamics after placement of a shunt with
an antisiphon device: a long-term study.
Journal of Neurosurgery, 2001. 94(5): p. 750-6.
Introduction
CSF hydrodynamics investigations are becoming more
common when selecting hydrocephalus patients for
surgery and performing these investigations via the
lumbar space significantly simplifies the procedure.
However, this requires that ICP truly can be assessed
via the lumbar space.
We present a study that verifies the usefulness of the
lumbar CSF hydrodynamics investigation by comparing
ICP CSF−LS to ICP IP in patients suffering from IAHS.
Materials and Methods
The present study was based on ten patients and the
ICP IP measurement was performed by a catheter tip
transducer inserted close to the roof of the right
ventricle. The procedure of measuring the ICP CSF−LS has
previously been described [1].
Pressure data from both devices were averaged over one
second and recorded on a PC by a software that also
provided one minute means of ICP IP and ICP CSF−LS .
A comparison of the data was performed using a
univariate General Linear Model (GLM).
Results
The results of the GLM show that the correlation
between ICP IP and ICP CSF−LS is highly significant (R 2
=0.999, P

Biomechanics
NON-NEWTONIAN EFFECTS ON WALL SHEAR STRESS IN A
HUMAN AORTA WITH COARCTATION AND DILATATION
R. Gårdhagen*, J. Svensson*, D. Loyd* and M. Karlsson**
*Department of Mechanical Engineering, Linköping University, Linköping, SWEDEN
**Department of Biomedical Engineering, Linköping University, Linköping, SWEDEN
rolga@ikp.liu.se
Abstract: Steady state, laminar blood flow is
simulated in a realistic, patient specific geometry
of a human aorta with a coarctation and
a post-stenotic dilatation. WSS from Non-
Newtonian flows, with a shear thinning viscosity
according to the Carreau model, is compared
with corresponding values from Newtonian
flows.
Non-Newtonian effects are found in the
dilatation, especially for lower velocities. The
average WSS is, however, relatively unaffected,
while the local WSS is highly influenced
by non-Newtonian effects and by the
rotation of the flow in the aortic arch.
Introduction
Atherosclerosis is one of the major causes of decease
in the modern society and even though risk factors
like smoking and fatty food are well known and easy
to observe, they do not tell where in a vessel a lesion
is likely to occur. Besides, atherosclerosis also
strikes people with wholesome habits. The genesis
of atherosclerosis in a vessel is, however, highly affected
by the local wall shear stress (WSS), which is
a frictional load on the vessel wall due to the motion
of the blood.
Accurate and easy methods to estimate WSS
would facilitate the work to identify risk sites, make
diagnoses, plan interventions and follow up treatments.
On the other hand, WSS is a very complex
parameter depending on shear rate (velocity gradients)
at the wall and viscosity, which in turn can be
a function of the shear rate itself. Hence, in this context
”accurate and easy” might be difficult to combine.
The blood viscosity is constant (high) at low
shear rates, constant (low) at high shear rates and
changes almost linearly for shear rates in between.
In this work different steady state, laminar blood
flows in a human aorta with a coarctation and a post
stenotic dilatation are simulated in order to investigate
whether non-Newtonian effects are present and
to see if an easy expression containing volume flow
rate, vessel radius and viscosity gives a reasonable
estimation of the WSS.
Materials and Method
Patient specific data come from a juvenile. Images
of the aortic arch and upper thoracic region of the
aorta are acquired with a GE Signa Horizon magnetic
resonance imaging (MRI) scanner at the University
Hospital in Linköping. At the end of the arch
the patient has a coarctation and a post-stenotic dilatation.
Computational fluid dynamics (CFD) is
applied to simulate the blood flow passing the lesions.
This requires a geometry, which is constructed
from the MRI images and comprises the region from
the upper arch to the mid thoracic region, Figure 1.
Coarctation
Dilatation
Figure 1: Geometry of aortic arch and thoracic region
with coarctation and post-stenotic dilatation used for
blood flow simulations.
Normally, the flow in the arch has a clockwise
rotation (CWR) [1] and the influence of this is investigated
by simulating CWR, counter-clockwise rotation
(CCWR) and no rotation (NR) of the flow in the
arch. The rotating velocity component is set to 30%
of the axial velocity component. Three mass flow
rates are simulated, 0.04, 0.05 and 0.06 kg/s, corresponding
to maximum inlet velocities of approximately
0.5, 0.6 and 0.7 m/s, respectively.
The flow is governed by the Navier-Stokes equations
consisting of the continuity equation, Equation
1 and three momentum equations, Equation 2.
∇·V = 0 (1)
ρ∇·(V V ) =−∇p +∇·τ (2)
V is the velocity vector, ρ the density, p the pressure
and τ the stress tensor. The density is set to
1060 kg/m 3 . The viscosity in the non-Newtonian
simulations is determined according to the Carreaumodel,
Equation 3, which gives a shear thinning behaviour
[2].
n−1
µ = µ ∞ + (µ 0 − µ ∞ )
[1 + (λ˙γ)
2]
2
(3)
According to [2] λ = 3.313 s, n = 0.3586, µ 0 = 0.056
Ns/m 2 and µ ∞ = 0.00345 Ns/m 2 . At low shear
rates the viscosity is µ 0 and at high shear rates (approximately
grater than 100 s −1 ) the constant viscosity
is µ ∞ , which is also the value used for the
Newtonian simulations. τ and ˙γ are related to the
second invariant of the tensor ˙γ ij according to Equations
4 and 5, where u i,j is the derivative of the i : th
velosity component with respect to the j : th spatial
direction.
√
1
[∑ ∑ ]
˙γ ij = (u i,j + u j,i ), ˙γ = ˙γ ij ˙γ ji (4)
2
And thus
i
τ = µ (˙γ) ˙γ ij (5)
j
IFMBE Proc. 2005;9: 275

Biomechanics
At the inlet the flow is specified as a fully developed
laminar velocity profile. The vessel wall is rigid and
represented by a no-slip condition and at the outflow
a condition to ensure conservation of mass is used.
Non-Newtonian effects are characterized by a local
non-Newtonian importance factor, I L , proposed
by [3]. It relates the apparent viscosity, µ, of the
flow to the constant Newtonian viscosity, µ ∞ , as
I L = µ/µ ∞ .
The WSS is studied as average values in cross
sections of the vessel and is easily obtained from the
simulation results. It is also determined according to
Equation 6, which is valid for laminar flow in (long)
straight circular pipes (called Hagen-Poiseuille flow).
It gives average values in a cross section with radius
r, where the volume flow rate is Q.
WSS HP = 4µ ∞Q
πr 3 (6)
Results
Figure 2 (left) shows cross sectional area of the vessel,
average WSS (in cross sections along the flow)
for Newtonian and non-Newtonian CWR flow with
ṁ = 0.04 kg/s and WSS HP . Figure 2 (right) shows
the non-Newtonian importance factor I L for CCWR
and CWR flow, ṁ = 0.04 kg/s and ṁ = 0.06 kg/s.
The x axis is a streamline coordinate in the flow direction.
The inlet corresponds to s = 0.
WSS [Pa]
5
4
3
2
1
0
WSS
WSS N
CCW m4
WSS nN
CW m4
CCW m6
2
CW m6
1
0
6
4
2
0.04 0.06 0.08 0.1 0.12 0.14 0.16
WSS HP 0.04 0.06 0.08 0.1 0.12 0.14 0.16
s [m]
s [m]
Area [cm 2 ]
Figure 2: (Left) Newtonian and non-Newtonian average
WSS compared with WSS HP for, ṁ = 0.04 kg/s. (Right)
Local non-Newtonian importance factor I L for ṁ = 0.04
kg/s and ṁ = 0.06 kg/s with CCWR and CWR. The
bottom curve in both diagrams shows the cross sectional
area of the vessel.
Figure 3 shows local WSS in a cross section
in the dilatation for Newtonian (WSS N ) and non-
Newtonian (WSS nN ) flow, CCWR (left) and CWR
(right).
WSS [Pa]
3
2.5
2
1.5
1
0.5
WSS [Pa]
0
0 90 180 270 360
θ [°]
N
nN
I L
WSS [Pa]
3
2.5
2
1.5
1
0.5
I L
WSS [Pa]
0
0 90 180 270 360
Figure 3: WSS N and WSS nN in a cross section at s =
0.096m for CCWR (left) and CWR (right) with ṁ = 0.04
kg/s.
Discussion
According to Figure 2 (left) the average WSS in a
θ [°]
6
4
2
N
nN
Area [cm 2 ]
cross section of the vessel is moderately affected of
the non-Newtonian effects, while the difference between
WSS from the simulations, denoted WSS sim ,
and WSS HP is remarkably large. It is obvious that
the latter strongly underestimates the WSS. Both
WSS sim and WSS HP indicates a maximum in the
constriction (located at about s = 0.06m) but the
fluctuations in the dilatation downstream are not
captured by WSS HP . The tendency is similar for all
other mass flow rates and rotations and the higher
mass flow rate the larger is the underestimation with
WSS HP .
In Figure 2 (right) the average of the local
non-Newtonian importance factor indicates non-
Newtonian effects in the dilatation, i.e. between
s = 0.07m and s = 0.14m. The effect is larger when
the mass flow rate, and hence, the velocity is low.
Although the average WSS is almost unaffected
by non-Newtonian effects Figure 3 clearly shows an
influence on the real, local WSS, where the difference
reaches about 25%. It is also clear that the
rotation in the aortic arch affects the WSS, since regions
with high and low values in the cross section
differs significantly, as does the maximum value in
the cross section. The average values of the WSS in
Figure 3 (left), i.e. CCWR, are 0.54 and 0.56 Pa for
Newtonian and non-Newtonian flow, respectively. In
Figure 3 (right) corresponding values for CWR are
0.59 and 0.61 Pa, which are the same as in Figure
(2) left at s = 0.096m. The large variations that obviously
exist in a cross section possibly indicate that
average values in general do not give correct information
of the WSS.
It should be noted that these simulations are for
steady flow and thus when taking unsteady effects
into account, the mass flow rate and hence the velocity
will be even lower and non-Newtonian effects
will certainly be even more important, but also vary
during a heart beat.
Conclusion
Non-Newtonian effects exist for the studied flows and
are more significant at lower velocities. WSS HP does
not give an appropriate measure of the WSS. In general,
average WSS in a cross section is possibly a
too unprecise measure due to large variations in each
cross section. The rotation of the blood flow in the
aortic arch does not affect the average WSS in the
vessel, but is of crucial importance for the local WSS,
e.g. where a maximum or minimum will occur.
Referenses
[1] Yang G. Z. Mohiaddin R. H. Firmin D. N. Kilner,
P. J. and Longmore D. B. Helical and retrograde
secondary flow patterns in the aortic arch studied
by three-directional magnetic resonance velocity
mapping. Circulation, 88:2235–2247, 1993.
[2] Y. I. Cho and K. R. Kensey. Effects of nonnewtonian
viscosity in a diseased arterial vessel.
part 1: Steady flows. Biorheology, 28:241–262,
1991.
[3] Johnston P. R. Corney S. Johnston, B. M. and
D. Kilpatrick. Non-newtonian blood flow in human
right coronary arteries: steady state simulations.
Journal of Biomechanics, 37:709–720,
2004.
IFMBE Proc. 2005;9: 276

Biomechanics
ERRORS IN APPLANATION TONOMETRY RELATED TO REFRACTIVE
SURGERY
P. Hallberg 1, 4, 5 , K. Santala 2 , T. Koskela 3 , O. Lindahl 4, 5 , C. Lindén 2 , A. Eklund 1, 5
1 Dept. of Biomedical Engineering & Informatics, Umeå University Hospital, Umeå, Sweden
2 Dept. of Clinical Science, University Hospital/Ophthalmology, Umeå, Sweden
3 Koskela eye clinic, Umeå, Sweden
4 Dept. of Applied Physics and Electronics, University of Umeå, Umeå, Sweden
5 Center for Biomedical Engineering and Physics, University of Umeå, Umeå, Sweden
per.hallberg@vll.se
Abstract
Laser surgery is a common technique for correction of
sight. The laser uses ablanation to change the shape of the
cornea. The modifications of the cornea give rise to errors
in measurement of the intra ocular pressure, IOP. The
aim of the study was to evaluate effects in IOP
measurement from laser surgery in an in vitro model.
Two different tonometers were used, Goldmann
applanation tonometer, and applanation resonance
tonometer. The study showed a significant
underestimation of the IOP after surgery and also a
dependency of measurement technique.
Introduction
Excimer laser surgery is a common technique to correct
myopia, hyperopia combined with astigmatism. Two
methods that are frequently used are Photo Refractive
Keratectomy (PRK) and Laser Insitu Keratomileusis
(LASIK). Both methods are based on modification of the
cornea to correct vision and seems to have effect on the
accuracy of intra ocular pressure, IOP, estimation. The
modifications of the cornea imply a reduction of the
cornea thickness and a change of the corneal curvature.
Both factors are possible sources to measurement errors,
[1-3] but the individual influence has not been made
clear.
Glaucoma is a group of diseases associated with optic
nerve damage and visual field loss. It is one of the
leading causes of blindness. One of the major risk factors
for glaucoma is an elevated IOP [4]. All treatment, so far,
is aimed at reducing IOP. A recent study showed
considerable beneficial effects of IOP reduction and that
it significantly delayed glaucoma progression [5]. It have
been shown that the IOP is underestimated after a laser
correction for myopia [6], a underestimation that could
result in a delayed diagnosis for glaucoma.
The aim of this study was to investigate the possibility to
evaluate effects of IOP measurements after PRK
treatment in an in vitro pig eye set-up using Goldmann
applanation tonometry, GAT, and a recently developed
sensor based on mechanical resonance technique,
Applanation resonance tonometer, ART [7-10].
Methods
Cadaver eyes from approximately 6-month-old Landrace
pigs were enucleated immediately after the pigs were put
to death at the abattoir (Ullånger, Sweden). To keep the
eye in place during measurement, the eye was moulded
into a petri dish with agar gel (15g/dm 3 ). To be able to
perform GAT measurements on the porcine eye a fixture
for the petridish was attached on a biomicroscope (Haag-
Streit slitlamp, Bern, Switzerland) that put the eye in
similar position as a human eye in a sitting patient
situation. A calibration study on porcine eyes showed that
GAT was functional for IOP exceeded 13 mm Hg [11].
The ART probe was mounted on a servomotor-controlled
lever and applied to the cornea of the eye with controlled
velocity of 7 mm/s. Measurement of the contact area was
based on a resonator sensor element made of lead
zirconate titanate (PZT). A feedback circuit processed the
signal from the element and powered the PZT in order to
sustain the oscillations in the resonance frequency. At the
end of the PZT element a force transducer (PS-05KD,
Kyowa, Tokyo, Japan) was mounted to receive the
contact force and at the opposite end of the PZT, a
bakelite piece was applied for contact against the cornea.
The contact surface of the piece was convex with a radius
of curvature, r =7 mm and diameter=4 mm. When the
contact piece touches the cornea of the eye the resonance
frequency will change. The amount of change depends on
the contact area between the sensor and cornea [8-10,
12]. The sensor system output signals are the contact
force and the shift of the oscillation frequency from
unloaded to loaded condition, sampled continuously
(1000Hz) during the applanation. The IOP was
determined from the equation
where dF C was change in contact force and df was change
in resonance frequency, corresponding to a change in
contact area. [9] The vertical position, L, of the sensor
was measured with an inductive position transducer
placed at the lever.
The central corneal thickness, CCT, was measured with a
Pach-pen (ultra sonic technique) and for the refractive
surgery a SCHWIND-Multiscan excimer laser was used.
IFMBE Proc. 2005;9: 277

Biomechanics
Six eyes were moulded into petri-dishes with agar. A
winged thin walled cannula was introduced through the
side of the eyeball into approximately the middle of the
vitreous chamber and connected to a saline column for
pressurise. All measurement was performed at IOP VC =30
mm Hg. Measurement order was: Central corneal
thickness, CCT, IOP GAT , IOP ART , and then PRK laser
treatment. Ten repetitions on each measurement. The first
measure sequence was done on unchanged cornea, the
second sequence was done on cornea with the epithelia
removed by the eximer laser. The following sequences
were measured after laser correction. Six eyes were
corrected for totally 25 dioptres (D) in steps of 5, 10 and
10 D. Blinking was simulated with saline, applied with a
sweep of a soft goat-hair brush.
Results
IOP was measured with GAT and ART (n=10, six eyes)
at 5 treatment levels. The first sequence was performed
on unchanged eye with the epithelium intact. Treatment
=1 represent measurement on eyes with the epithelium
removed by the excimer laser, but no correction.
Subsequent treatment, level 2, 3, and 4 represent PRK
correction of totally 5D, 15D, and 25D respectively. No
measurements with GAT were practicable on eye number
1 at treatment level 3 and 4, and eye number 2 at level 4.
Figure 1. Treatment against IOP GAT and IOP ART for six
eyes. Error bars shows the SD. Treatment = 0 represent
unchanged eye, treatment =1 represent the epithelium
removed from the cornea. Treatment = 2, 3, and 4 were
IOP measurements after laser corrections of 5, 15, and
25 D.
Discussion
This is the first study, to our knowledge, that investigates
the relationship between repeated PRK treatment on eyes
and errors in IOP reading with applanation method
tonometers. It was found that both the ART and the GAT
was sensitive to treatment. When the epithelium was
removed it did not influence the IOP GAT readings, but the
IOP ART dropped from 29 mm Hg to 26 mm Hg, This
indicated that a change in tissue is an impact factor for
the ART and it is not unreasonable to relate such a
change to a change in acoustic impedance of the corneal
surface. From a biomechanical view it’s interesting to
analyse if the there was any differences between GAT
and ART after the removal of the epithelium. ART
readings dropped from 26 mm Hg to 20 mm Hg after a
correction of 25 D. Corresponding value for GAT was a
drop from 32 mm Hg to 19 mm Hg, i.e. more than twice
as much as the ART. Mean decrease of the IOP was 2.5
mm Hg of ART and 5.2 mm Hg of the GAT
measurements. The study showed that GAT and ART
underestimated the IOP after PRK surgery. Excluded the
change when the epithelium was removed, ART
underestimated 6 mm Hg after correction of 25 D, which
is small in this context. Consequently the ART was less
sensitive than GAT for laser corrected eyes.
Conclusions
This study indicate that corneal keratectomy effect the
IOP readings different depending on tonometer use.
With respect to measurement error in IOP assessment the
ART is less sensitive to PRK treatment than Goldmann.
References
1. Sommer, A., Intraocular pressure and
glaucoma. Am J Ophth, 1989. 107(2)
2. Heijl, A., et al., Reduction of intraocular
pressure and glaucoma progression... Arc of Ophth,
2002. 120(10)
3. Whitacre, et al, Sources of error with use of
Goldmann-type tonometers. Surv of Ophth, 1993. 38(1)
4. Whitacre, et al, The effect of corneal thickness
on applanation tonometry. Am J of Ophth, 1993. 115(5)
5. Mark, H., Corneal curvature in applanation
tonometry. Am J Ophth, 1973. 76(2)
6. Eklund, A., Resonator sensor technique for
medical use... in Dept of Rad sci.
7. Eklund, A.et al, A resonator sensor for
measurement of intraocular pressure... Phys Meas, 2000.
21(3)
8. Eklund, A., et al., Applanation resonance
sensor for measuring intraocular pressure... Inv Ophth &
Vis Sci, 2003. 44(7):
9. Eklund, A., et al., Evaluation of applanation
resonator sensors for intraocular pressure
measurement... MBEC, 2003. 41
10. Hallberg, P. et al, Comparison of Goldmann
applanation- and Applanation resonance tonometry...
JMET, in press.
11. Hallberg, P., et al., Applanation resonance
tonometry for intraocular pressure in humans. Phys
Meas, 2004. 25
12. Schmidt, T., Zur applanationtonometri an der
spaltlampe. Ophthalmologica, 1957. 133:
Acknowledgements
The authors thank Tomas Bäcklund, Dept. of Biomedical
Engineering for skillful technical assistance and
measurement support.
IFMBE Proc. 2005;9: 278

TRANSMURAL MYOCARDIAL STRAIN DISTRIBUTION -
THEORETICAL RESULTS AND IN VIVO DATA
Biomechanics
K. Kindberg and M. Karlsson
Division of Biomedical Modelling and Simulation, Department of Biomedical Engineering,
Linköping University, Linköping, SWEDEN
katki@imt.liu.se
Abstract: Transmural distributions of myocardial
strain are presented as means for seven
animals and are compared to the estimated
and the true strains of an analytical model of
the left ventricular deformation. The model
gives good estimates of the animal strains.
Introduction
Biomedical models are useful when attempting to
understand the mechanisms underlying the actions
of the normal organs in the body, diagnosing disease
or to predict the results of new surgical methods or
medicine. In order to create a trustworthy model,
comparisons with real data need to be done. Here the
wall strain of a cylindrical model of the left ventricle
of the heart is compared to the strain of the ovine
pumping heart.
Materials and Methods
Surgical Preparation and Data Acquisition: The surgical
preparations and the data acquisition have previously
been reported in [1] and hence only a brief review
will be done here. Seven adult Dorsett hybrid sheep
were anesthetized and 13 subepicardial radiopaque
markers were surgically implanted to silhouette the
left ventricular, LV, chamber. In addition, three transmural
columns of four beads each were implanted into
the lateral LV wall in a region equally-spaced between
the papillary muscles. The columns were placed normally
to the epicardial tangent plane and three 0.7 mm
diameter beads in each column were evenly spaced between
endo- and epicardium. The fourth bead of each
column was larger, 1.7 mm in diameter, and was sewn
onto the epicardium above each column.
Eight weeks after surgery, biplane videofluoroscopic
images of all radiopaque markers were acquired at 60
Hz. Analog LV pressure and surface lead ECG signals
were recorded in digital format on each individual
video image. Data from the two two-dimensional views
were merged to yield the three-dimensional coordinates
of the centroid of each marker every 16.7 ms.
Analytical Model: A deformable thick-walled cylinder
is used as a model of the left ventricle. The model
is in the reference state described by cylindrical coordinates
(R, Θ,Z) and the deformed coordinates (r, θ, z)
are given by
√
α(R2 − R1 2 r =
) + r1
2 λ
θ = φR +Θ+βZ
z = ωR + λZ
where the undeformed inner radius is R 1 =2cm,the
undeformed outer radius is R 2 = 3 cm and the deformed
inner radius is r 1 =1.65 cm. The parameters
are axial extension ratio λ =0.8, torsion parameter
β = 0.2, transverse shear parameters φ = 0.1 and
ω =0.3 and compressibility α = 1 (incompressible).
The model has previously been described in [2, 3]. The
deformation is chosen to resemble the true deformation
at end systole, ES, taking end diastole, ED, as
reference. The analytical strain of a deformed cylinder
is presented in [4].
The strain is estimated in a dense region of assumed
beads, made of three transmural columns with six
beads equally spaced between epi- and endocardium.
The three beads on the epicardium form an isosceles
triangle with base 1.04 mm and the other two sides
0.96 mm long.
Cardiac Finite Strains: A local orthogonal cartesian
coordinate system is used at the position of the
bead array when computing the strain, on the real
data as well as on the model. The coordinate directions
are circumferential, X 1 , longitudinal, X 2 , and
radial, X 3 . The longitudinal direction is apex-to-base
and the radial direction is endo-to-epicardium. The
X 1 − X 2 -plane is tangential to the epicardium.
For each beat, the reference state is taken at ED.
A material point which in the undeformed state has
the coordinate X, moves to the position x in the deformed
configuration, which here corresponds to the
ES state. The material gradient of the position field
∂x
∂X ,
is the local deformation gradient tensor F =
and the ( Lagrangian strain tensor E is computed via
E = 1 2 F T F − I ) , where I is the identity tensor.
Strains are interpolated along the centroid of the bead
columns at 1% increments of wall depth from the epicardium
to the most subendocardial bead.
The deformation gradient tensor F is on the analytical
model estimated with a finite element, FE,
method using a bilinear-quadratic element which was
implemented by following [2]. In addition, an approach
based on polynomial least squares fitting [5], is used
to compute the deformation gradient tensor on the
analytical model as well as on the real data.
Statistics: For each animal, three consecutive beats
were sampled. The mean strain components of the
three beats were taken as that animal’s strain values.
All values are given as mean±SE for n = 7 animals.
Results
In Figure 1, the transmural analytical strains and
the strains estimated with both of the two methods
are plotted. Figure 2 shows the mean transmural distributions
of the strains at ES, using ED as reference,
for seven animals.
IFMBE Proc. 2005;9: 279

Biomechanics
Discussion
Comparison with real data is needed for determining
the accuracy of a model. Here the mean transmural
strains at ES of seven animals, as well as the strains
of the cylinder model are presented. The analytical
model gives good estimates of the ovine transmural
strains, but slightly exaggerated in magnitude.
The simulated bead array with approximately 1
mm column separation and six beads per column is
denser than the real data. This is to show how accurate
the estimates can be. A sparser array leads to
larger errors, [2, 5].
References
Figure 1: Dashed: FE strains, dotted: strains from the
polynomial method, solid: analytical strains.
[1] A. Cheng, F. Langer, F. Rodriguez, J. C. Criscione,
G. T. Daughters, D. C. Miller, and N. B. Ingels Jr.
Transmural cardiac strains in the lateral wall of the
ovine left ventricle. Am J Physiol Heart Circ Physiol,
2004. In press.
[2] A. D. McCulloch and J. H. Omens. Nonhomogeneous
analysis of three-dimensional transmural
finite deformation in canine ventricular myocardium.
JBiomech, 24(7):539–48, 1991.
[3] K. Kindberg and M. Karlsson. Cardiac wall strain
variations due to compressibility. In: Proceedings
of NSCM-17, Stockholm, Sweden, 2004.
[4] A.J.M.Spencer.Continuum mechanics. Longman
mathematical texts. Longman, London, 1980.
Figure 2: The mean±SE transmural distributions of
the strains at end systole for n = 7 animals.
[5] K. Kindberg, M. Karlsson, N. B. Ingels Jr, and
J. C. Criscione. Non-homogeneous strain from
transmural myocardial beads for cardiac hemodynamics.
In manuscript.
IFMBE Proc. 2005;9: 280

Biomechanics
SPRING-DAMPER MODEL FOR PROSTATE TISSUE
N. Norén 1, 2 , B. Andersson 1, 2 , A. Bergh 3 , B. Ljungberg 4 , O. Lindahl 1, 2
1 Applied Physics and Electronics, Umeå University, Umeå, Sweden
2 Center for Biomedical Engineering and Physics, Umeå University, Umeå, Sweden
3 Department of Medical Bioscience, Umeå University, Umeå, Sweden
4 Department for Surgical and Perioperativ Science, Umeå University, Umeå, Sweden
niklas.noren@tfe.umu.se
Abstract
New non-invasive methods that can detect prostate
cancer are highly needed. In this study creep
measurements were conducted on four prostate slices
containing both cancer and non-malignant areas. A model
consisting of a linear spring in series with two Voigt
elements was fitted to the experimental data and the
models ability to describe the creep was evaluated. The
coefficients of explanation, R 2 , from all the fits were
found to be between 0.98-0.99. This indicates that the
model has a good ability to describe the creep of prostate
tissue.
Introduction
In the European Union and the USA, prostate cancer
is the most common cancer for males. In 2003, 220900
males were diagnosed with prostate cancer in the US [1].
In Sweden 41.3 percent of all cancer cases in males
reported 2003 was prostate cancer, which translates into
9035 cases [2]. Prostate cancer is generally diagnosed by
a high blood PSA (prostate specific antigen) level, rectal
palpation and ultrasound examination of the prostate
followed by histological examination of prostate
biopsies. Screening solely based on palpation misses 30-
50 percent of prostate cancer in comparison to screening
based on transrectal ultrasound and/or PSA 3]. Evidence
is also pointing to that it is foremost the smaller tumours,
which are easier to treat, that are missed [3]. Therefore,
there is a need for new knowledge about prostate tissue
and improved non-invasive methods to detect prostate
tumours in a reliable and easy way.
All biological tissues are viscoelastic 4].
Viscoelastic materials show a combination of elastic and
viscous effects. Creep, relaxation and hysterisis are all
viscoelastic phenomena's [4]. Creep is when a body is
subjected to a constant load and the body continuous to
deform over time [4]. To describe viscoelastic materials
mechanical models, which consists of springs and
dashpots, are often used [4].
The goal of this study was to develop a model
consisting of springs and dampers that can describe the
creep of prostate tissue.
Methods
A counter-balance system was used in the
experiments. It consisted of a gramophone pick-up arm
mounted on a metal box. The pick-up arm was able to
pivot vertically around its mounting axis and the
movement was controlled by a stepping motor placed in
the metal box. At the front end of the pick-up arm a probe
with a force sensor (5S-05KC, Kyowa, Japan) was
attached. The probe had a spherical tip with a diameter of
5 mm and a curvature of 0.125 mm -1 . At the back end of
the pick-up arm counterweights could be attached which
made it possible to control the load the probe exerted on a
sample. A position sensor made out of a steel rod and a
conductor was used to measure the probe tips position. A
LabVIEW6i® (National Instruments, USA) program
installed on a Toshiba laptop with a DAQCard-AI-16XE-
50 (National Instruments, USA) was used to sample the
data.
Creep measurements were conducted, after approved
by an ethical committee, in collaboration with a
pathologist and urologist on four prostate tissue samples.
The samples were slices of whole prostates about 10 mm
thick that were cut out by a pathologist. The four slices
came from different prostates. Before measuring, the
prostate slices were pinned at the edges to polystyrene.
The polystyrene with the prostate slice were then fixed to
an x-y positioning table. Measurements were done on 2-
3 points on each slice. 5 measurements, with a time
interval around 5 minutes in between, were conducted on
each point. The tissue was carefully bathed with saline
solution between every measurement.
The force exerted on the tissues was 0.04 N for all
measurement which was calibrated using a scale (BL310
Sartorius AG Göttingen, Germany). The sampling time
was 60 seconds and the sampling frequency 1000 Hz.
After the measurements were conducted the tissue
slices fixed in formalin, embedded in paraffin and
sectioned. Using morhpometrical methods the tissue
composition (epithelium, glandular lumina, stroma and
prostate stones) under the measured points was
determined, both in malignant and non-malignant areas.
The model chosen to be tested was a generalized
Voigt model consisting of one linear spring in series with
two Voigt elements.
A Voigt element consists of a linear spring in parallel
with a linear damper. The models creep functions was
IFMBE Proc. 2005;9: 281

Biomechanics
calculated and fitted to the experimental data. The fitting
was done with a Gauss-Newton algorithm with linesearch
implemented in MATLAB (Comsol AB, Sweden). The
coefficient of determination, R 2 , was calculated for each
fit.
Results
An example of a fit which is typical can be seen in
figure 1. Generally the coefficient of explanation was
between 0.98-0.99 for all fits.
Discussion
The high coefficient of explanation (R 2 =0.98-0.99)
for the fits conducted shows the models ability to
describe the creep of prostate tissue. A problem though is
the high standard deviation in the parameter values found
from some of the fits done to data gained at the same
measurement point. This variation cannot be explained
by varying tissue composition as for the spread in
parameter values between different measurement points.
The findings that there was a remaining impression
between following measurements could be an
explanation. It indicates that the tissue has not regained
its properties and thus not responding in a similar manner
between following measurements. The choice to have
about a 5 minute time interval was in beforehand thought
to be enough for the tissue to regain its shape and
mechanical properties, but this was obviously not the
case. One could speculate that preconditioning the sample
with a suitable number of impressions could be the
answer.
Conclusions
Figure 1: Model fitted to experimental data
Figure 2 shows a boxplot of spring parameter k 0 .
The model is able to describe the creep in prostate
tissue well. In the future more measurements on prostate
tissue will be conducted and correlations between
parameter values and tissue composition will be tested.
References
[1] American Red Cross, Prostate Cancer Statistics 2003
from ACS - American Cancer Society, Internet site
address: http://prostate-help.org/castats.htm
[2] Socialstyrelsen, Internet site address: http://
www.socialstyrelsen.se/NR/rdonlyres/86E6C7D0-4650-
43DA-82DF-FAC11CDB54C5/3019/20044 210.pdf
[3]Svensk Urologisk förening, Internet site address:
http://www.urologi.org/sota/STA026/sta026.htm
#Prostatacancer
[4] Fung Y.C. (1993): ‘Biomechanics: Mechanical
Properties of Living Tissues’, (Springer-Verlag, Berlin-
Heidelberg-New York), pp. 41-7
Figure 2: Boxplots of parameter k 0
The parameter values found from fittings done to
data gained from measurements on the same point had a
standard deviation that varied between 0.5-20 percent of
the mean value.
Variations in parameter values between fittings done
to data from different points were also found.
Comparisons between following measurements on
the same spot was done. It was found that there was a
remaining impression due to earlier measurements. This
impression grew larger with each measurement and was
up to 20 percent of the total impression depth.
The total impression also decreased, up to 7 percent,
between the first and last measurement.
Acknowledgements
The study was supported by EU objective one,
Northern Sweden.
IFMBE Proc. 2005;9: 282

Biomechanics
MODELLING OF PERIPHERAL BLOOD FLOW DYNAMICS: COMPARISON
WITH FINGER PHOTOPLETHYSMOGRAPHY SIGNALS
U. Rubins 1 , J. Spigulis 1
1 Institute of Atomic Physics and Spectroscopy, University of Latvia, Riga, Latvia
uldis.rubins@lu.lv
Abstract
Two models of human blood circulatory system have
been discussed – Y-branched artery model and combined
3-element Windkessel model. These models have been
simulated in Matlab. The measured skin blood volume
changes from human’s finger have been compared with
the computer simulation models. The program estimated
4 parameters for Y-branched model and 6 parameters for
combined model using the least-square minimization.
Introduction
The blood volume changes induced by the heart
contractions are propagating along the arteries as pulse
waves. The peripheral pulse waveform depends on
various factors such as heart and respiration function,
arterial wall distension, branching and narrowing sites.
Each inhomogenity produce reflected wave that distort
the initial wave. The pulsations may considerably change
when the pressure wave reaches the periphery.
Blood volume changes are proportional to the blood
pressure changes in peripheral vessels. The peripheral
volume pulsations can be effectively detected from the
living tissues by non-invasive reflection
photoplethysmography (PPG). For instance, the results
obtained by PPG fingertip sensor devices with special
signal filtering and averaging algorithms [1] may be
useful for verifying the model calculations.
Detailed analysis of the PPG signal waveform reveals
physiological parameters of human’s vascular system if
the input (ascending aorta) pressure waveform is known.
Figure 1: Single Y-branch model
The effective reflecting site in a human arterial system is
a bifurcation of aorta with left and right Illac arteries [2].
Aorta from the output of the heart’s left ventricle till
bifurcation is assumed as elastic tube (L 0 = 40 cm) with
branching at the end.
Combined arterial model is shown in Figure 2. This
model includes Y-branch model and 3-element
Windkessel [3], who has applied to subclavian, brachial
arteries and its subarteries and arterioles. If pressure at
model’s input p(t) is known, pressure at distal end p 1 (t)
can be evaluated by solving differential equation (2):
(2),
where R 1 – input arterial resistance, R 2 – peripheral
resistance, C – arterial compliance.
Methods
A single Y-branch model of systemic vessel branching is
shown in Figure 1. According to this model, pressure
wave is partially reflecting from the branching and fully
reflecting from the root of the vessel, so changing the
initial wave (1):
(1),
where p 0 – initial pressure wave, α – the reflection
coefficient at branching point, L 0 – length of the vessel,
c 0 – wave propagation velocity.
Figure 2: Mixed model
The aortic pressure waveform was assumed to be
Gaussian:
IFMBE Proc. 2005;9: 283

Biomechanics
where b – Gaussian bandwidth, τ - time delay.
(3),
Results
Figure 3 shows the simulation results. The input Gaussian
pressure function contains two parameters: b and τ; the
branched model contains two parameters: α and c 0 and
the mixed model contains additional two parameters: a 1 =
R1/R2 and a 2 = R1·C.
A single period PPG (SPPPG) signal was calculated from
the measured data using the special signal averaging
algorithm [1]. The model’s output pressure waveform has
been fitted to SPPPG data. The parameters in branched
model and mixed model have been calculated iteratively
by means of the least-square minimization (Table 1).
A: Branched model pressure fitted to SPPPG data
Table 1. Evaluated parameter values in simulated models
Model Y-Branched Mixed
b 8 6.5
tau 0.15 0.1
alpha 0.2 0.2
c0 4.3 m/s 4.2 m/s
a1 - 0.02
a2 - 0.16
A: Pressure simulation in the branched model
B: Pressure simulation in the mixed model
Figure 3. Model simulation results: initial pressure
(dotted) and simulated pressure at the periphery.
B: Mixed model pressure fitted to SPPPG data
Figure 4. Model simulation results: SPPPG data (circles)
and calculated pressure (solid line).
Discussion
We have developed a combined human’s arterial model
by joining the Y-branched artery model and the
Windkessel model. We have revealed that 3-element
Windkessel model can represent the output pressure,
if input pressure is known, without the flow data. This
allows to compare calculated blood pressure at system’s
output with the PPG-measured volume changes in
periphery.
Conclusions
This work demonstrates a simple method that allows to
calculate arterial system’s parameters from the measured
volume changes in the periphery. The non-invasive
reflection SPPG method seems to be well suited for
experimental tests of arterial flow models.
References
[1] J. Spigulis, R. Erts, U. Rubins. Micro-circulation of
skin blood: optical monitoring by advanced
photoplethysmography techniques. Proc. SPIE 5119:
219-225, 2003.
[2] C. Caro, T. Pedley, R. Schroter, W. Seed. The
mechanics of the circulation. Oxford University Press,
NY-Toronto, 1978.
[3] M. Yoshigi, G. D. Knott, B. B. Keller. Lumped
parameter estimation for the embryonic chick vascular
system: a time-domain approach using MLAB. Computer
Methods and Programs in Biomedicine. vol: 63 issue: 1,
p. 29-41, 2000.
IFMBE Proc. 2005;9: 284

Biomechanics
THE EFFECT OF MEMBRANE MOVING PATTERN ON THE BLOOD FLOW
IN A SAC-TYPE VENTRICULAR ASSIST DEVICE
Faramarz Firouzi 1 , Nasser Fatouraee 2 , and Siamak Najarian 3
1,2 Biological Fluid Dynamics Laboratory, Biomedical Engineering Faculty,
Amirkabir University of Technology (Tehran Polytechnic), Tehran, IRAN, 15914
3 Department of Mechanical and Industrial Engineering,
Concordia University, Montreal, Quebec, Canada H3G 1M8
1 FaraFirouzi@yahoo.com, 2 Nasser@aut.ac.ir , 3 SNajaria@me.concordia.ca
Abstract: Nowadays, long-term implantable
ventricular assist devices (VADs) are widely
demanded. To reduce hemolysis and thrombus
formation, the implantable sac-type devices producing
pulsatile flow are suitable. In this paper two different
models of sac-type VADs have been numerically
simulated. In model 1, the movement of the elastic
membrane is assumed to be simple. In model 2, in
order to investigate the effect of the shape of
membrane on blood flow, movement is considered to
have wavy shape. The results demonstrate the effect of
the membrane movement pattern on the blood flow
pattern inside the chamber is negligible.
Introduction
In spite of recent noticeable achievements in science
and technology, heart diseases make too many people to
die. VADs help patients with myocardial dysfunction or
end-stage heart failure to survive [1].
There are many researches related to numerical
analysis of passing blood flow through VADs producing
continuous flow [2]. The pulsatile flow simulation within
a VAD needs to consider moving wall conditions, and this
complicates the numerical modeling. The research in this
area is still progressing.
Hochareon et al. used a physical modeled of the
LionHeart LVAD [3]. They conducted a series of
experiments to determine the influence of diaphragm
motion on the flow pattern where they found a wave-like
pattern for the diaphragm motion during filling phase.
In this paper, a VAD named HeartSaver which is sactype
device was investigated numerically. In order to
understand the relationship between moving pattern of
membrane and produced flow pattern, two different
models were chosen. The motion of membrane was
assumed to be elliptic in the first model and wavy in the
second one. Flow within the models was simulated using
finite element method with segregated approach and the
mixed Eulerian-Lagrangian formulation. For domain
remeshing the spine method was used.
Methods
Mathematical model: Membrane was assumed to be
the only moving wall. Blood was assumed to be a
homogeneous, incompressible, and Newtonian fluid with
density of 1060 kg/m 3 and viscosity of 0.004866 Pa.s [4].
Governing equations used were conservation of mass and
momentum.
Boundary conditions: No-slip boundary condition and
zero tangential velocity were applied to the stationary
walls and inlet and outlet boundaries, respectively.
The displacement and velocity of the moving wall
were assumed to be prescribed at all the time and had
values such that the outflow and inflow profiles were like
the natural heart. The displaced volume and volumetric
flow rate cycle are shown in Figure 1. The cycle period
was 0.7s.
(a)
(b)
Figure 1: a) Displaced blood volume, b) Volumetric flow rate for natural
heart [5].
The maximum Reynolds number and Womersely
number were 3288 and 13.58 respectively.
Solution method: A deforming mesh approach was
employed and this was coupled with a segregated-type
iterative procedure. Nodes in chamber region were
allowed to move along the y-axis only. Galerkin Finite
Element technique was applied to the transient governing
equations. The FIDAP code (Fluent Inc.) was employed to
solve the models.
Geometry and grid of models: The geometry was
based on Mussivand [6]. The location of inflow and
outflow unidirectional valves was similar to Mussivand
[6].
The mesh was generated orthogonal to the y-axis
direction. The 4-node quadrilateral elements with the total
number of 20550 nodes were used.
Results
Grid independent verification: For model 1 different
mesh was generated (12800, 15400, 20550, and 29460
IFMBE Proc. 2005;9: 285

Biomechanics
nodes). A mesh with 20550 nodes was shown to be a good
choice.
Periodic results: The results of several successive
cycles were compared. A little difference between the
results of third and forth cycle was recognized (about
0.5%) so, results of the third cycle was considered.
Validation: For procedure validation, a oscillatory
flow through parallel surfaces was modeled. Numerical
results were compared with theoretical results obtained
from Currie [7], which showed a maximum of 6%
deviation. The results were considered good enough to our
simulations.
Flow pattern: Some of numerical analysis results of
blood flow for the two different models are shown in
Figure 2. The results demonstrate that during a complete
cycle, there are vortices in chamber, inflow and outflow
region. Vortex 1 generates at the beginning of diastole
phase and moves toward outflow port, and it occupies a
wide area at the entry of outflow port, at the end of systole
phase.
as the velocity of fluid driven decreases, a lot of local
vortices occur that may be so useful because they cause to
move fluid and prevent stagnation point and platelet
aggregation, and sweep within the chamber.
Two models are acting almost similar in the value,
location, and moment of maximum velocity and the
number of local vortices within a period of cycle.
Stress distribution along membrane: It can be inferred
that change of membrane movement pattern does not
affect vertical stress variations (Figure 3).
The results shows that for both models the maximum
shear rate does not exceed 2000 s -1 , and the maximum
value of viscosity term in vertical total stress is 15 Pa. It is
also seen that the pressure term is the more important
parameter of vertical total stress and viscosity term effects
are small. So, one can conclude that the flow pattern on
model 1 is accurate enough and the effect of the
membrane movement pattern on the blood flow pattern
inside the chamber is negligible.
Model 1 Model 2
Figure 2: Streamlines at phase 51.4º of cycle. The same streamline level
sets are drown for both models
The start of clockwise vortex 2 is the beginning of the
systole phase and it grows and moves toward the aorta
valve. Flow patterns comparison show that the two
models have no difference at the beginning of systole
phase. In the middle of systole phase, the higher speed of
moving wall opposite of inflow port in model 2 causes the
vortex 5. At the end of systole phase, flow pattern in
model 1 and 2 differs a little with each other. Moreover,
the total pattern of inflow and outflow ports is the same in
all of the models. The only difference can be observed
within chamber. Flow pattern, along the diastole phase, in
the two models, is exactly the same at inflow port and
chamber except the region next to outflow port.
Blood stress on membrane: In order to better
understand the effect of fluid dynamic onto moving wall,
forces and stresses applied by fluid to moving wall is
investigated. Fluid total stress vector for each element is
as follows:
t σ n σ = − δ + µ u u
i
= which ( )
ij
j
ij
p
ij i, j
+
j,
i
The amount of applied vertical stresses by adjacent fluid
elements to moving wall elements is given in Figure 3 for
different phases of pulse cycle. The maximum variation of
vertical stress along the moving wall for model 1 and 2
are 243 and 342 Pa, respectively.
Conclusion
Flow pattern: It is concluded that in the two models,
there is a particular disturbance in flow pattern. Moreover,
Model 1 Model 2
Figure 3: Total applied vertical stresses by adjacent fluid elements to
moving wall elements for different moments.
References
[1] BRONZINO J.D. (2000) "The Biomedical
Engineering Handbook”, (CRC Press LC).
[2] BURGREEN G.W., ANTAKI J.F. (2001)
“Computational Fluid Dynamics as a Development
Tool for Rotary Blood Pump”, Artificial Organs, Vol.
25, No. 5, pp. 336-340.
[3] HOCHAREON P., MANNING K.B., FONTAINE
A.A., DEUTSCH S. (2003) “Diaphragm Motion
Affects Flow Patterns in an Artificial Heart”, Artif.
Organs, Vol.27, No.12, pp. 1102-9.
[4] FIROUZI F., KATOOZIAN H., FARHANIEH B.
(1997) “Fluid Flow Analysis in Bifurcation and its
affect on intimal thickening”, Faculty of Mechanics,
Sharif Univ. of Tech.
[5] FARREL A.P. (2001) “Encyclopedia of life
sciences–circulation in vertebrates”, (Nature
Publishing Group).
[6] MUSSIVAND T. (1996) “Electrohydraulic
Ventricular Assist Device”, US Patent, No. 5569156.
[7] CURRIE I.G. (1992) “Fundamental Mechanics of
Fluids”, (McGraw-Hill).
IFMBE Proc. 2005;9: 286

Biomechanics
BONE MINERAL DENSITY IN RELATION TO FRACTURAL
ANALYSIS BY BIOMECHANICAL PROPERTIES
M. Mokhtari-Dizaji * , M.R. Dadras * , B. Larijani ** , G. Torkaman *** , Kazem-nejad A ****
* Department of Medical Physics, Tarbiat Modarres University
** Metabolism and Endocrine Research Center, Tehran Medical Sciences University
***Department of Physiotherapy, Tarbiat Modarres University
****Department of Biostatistics, Tarbiat Modarres University, Tehran, Iran
mokhtarm@modares.ac.ir
Abstract: Dual energy x-ray absorptiometry
(DXA) has been suggested for the assessment of
fracture risk. In this study, bone mineral density
measurements were conducted on rabbit’s bone
using commercially clinical instruments. The
bones were then mechanically tested by threepoint
bending test to obtain stiffness, ultimate
strength and energy expenditure until fracture.
The results of correlation analysis show that
there are relatively high correlation between
BMD and mechanical parameters. We conclude
that BMD important for the mechanical
properties of bones.
Introduction
Bone density or mass is a major factor in
determining bone strength [1]. Bone strength may
be assessed by DXA and fractal analysis [2]. Bone
strength is the ultimate indicator of bone quality,
but unfortunately is not directly measurable in vivo.
Based on theoretical and experimental analysis,
mechanical behavior in bone depends also on the
structure features and density of the tissue [3,4].
In this study, Rabbit’s tibia and femur bones were
investigated using DXA and mechanical
measurements. The objectives of this work were to
investigate the relationships between the
mechanical properties and clinical DXA parameter
in rabbit’s bone.
MATERIALS AND METHODS
A total of 22 three-month old white rabbits
weighted 1851±271gr were anaesthetized by
intrapertoneum Ketamin hydrochloride (10%) and
Xylazin hydrochloride (2%) injection (3:4, 0.7
mg/kg). The BMD of femur (n=22) and tibia (n=22)
in three regions (up: 1/3 of length, down: 2/3 of
length and middle: ½ of length) were measured and
averaged using DXA (Lunar Corp, USA). The
measurement protocol for the femur with a 15*12
cm 2 window was used. The animals were killed
with Ether facility protocol. The tissue surrounding
the femur and tibia was left intact.
Biomechanical studies consisted of compression
and three-point bending tests (Zwick-477514,
Germany). Each specimen was loaded to failure at a
rate of 1mm/min using displacement control. The
stiffness of bone (N.mm -1 ) was determined from the
initial strength line portion of the load-deformation
as the highest point of curve, and the energy
expenditure until fracture (N.mm) as the area under
the curve (Figure 1).
Figure1: Load – deformation curve
Pearson correlation coefficients were used to
express relationships between the parameters. The
relationship between various mechanical properties
and BMD was examined using linear regression.
Results
The bone mineral density (g/cm 2 ) values and
results of the mechanical tests of the femur and the
tibia bones are presented in Table 1. The BMD
were measured in three regions (up, down and
middle) and averaged for the tibia and the femur
IFMBE Proc. 2005;9: 287

Biomechanics
bones. The fracture loads had variation from 143N
to 247N with a medium 194N for femur and from
129N to 201N with a medium161N for tibia bone.
The correlation and regression functions between
the mechanical parameters and bone mineral
density are presented in Table 2 and Table 3.
Table 1: Mean ±SD of BMD values and mechanical
data in the femur and the tibia bones.
Parameters Femur(N=22) Tibia(N=22)
present study, all mechanical parameters were
assessed by a destructive invasive test and
compared to bone mineral density. In Figure 2, the
plot shows the correlation between the BMD with
the stiffness (interval confidence 95%).
.34
.32
.30
.28
BMD
(gr/cm 2 )
Stiffness
(N/mm)
load of
fracture: F max
(N)
Energy
expenditure
until fracture
(N. mm)
.277±.031 .236±.026
90.31±18.36 52.98±13.99
194.11±30.11 161.32±22.17
243.17±39.39 276.21±33.82
BMD
.26
.24
.22
.20
.18
20
40
STIFFNES
60
80
Figure 2: The correlation plot of between BMDstiffness
Conclusion
100
120
140
Table 2: Pearson correlation coefficients and
significant level of BMD and mechanical
parameters of rabbit’s bone.
Parameters
BMD F max Stiff-
Ness
Energy
absorption
BMD 0.41 * 0.56 * 0.40 *
F max 0.87 * 0.14
Stiffness -0.16
* Correlation is significant at the 0.01level (2-
tailed).
Table 3: results of regression analysis and
significant level of BMD (g/cm 2 ) and mechanical
parameters {F max (N), stiffness (N. mm -1 ) and
Energy expenditure until fracture (N. mm)} of
rabbit’s bone
Regression function
Sig.
.001stiffness+6.627*10 -5 F max +0.273= BMD .000
Discussion
Several previous studies have shown significant
correlation between mechanical strength of femoral
neck and bone mineral expressed as QCT mass (2).
Other researcher studied cortical tibia bone as a
material and found poor correlation between QCT
density values and bone mineral strength [4]. In the
We first it justified to conclude that BMD
important for the mechanical properties of bones.
Further investigations are in progress to evaluated
correlation between acoustic parameters and BMD
–mechanical properties.
References
[1] WU C, Hans D, He Y, Fan B, Nieh C F, Augat
P, Richards J and Genat H K. (2000): ‘Prediction of
bone strength of distal forearm using radius bone
mineral density and phalangeal speed of sound’,
Bone, 26, pp. 529-533.
[2] Toyras J, Nieminen M T, Kroger H and Jurvelin
J S. (2002): ’Bone mineral density, ultrasound
velocity and broadband attenuation predict
mechanical properties of trabecular bone
differently’ Bone, 31, pp. 503-507.
[3] Toyras J, Kroger H and Jarvelin J S. (1999):
‘Bone properties as estimated by mineral density,
ultrasound attenuation and velocity’, Bone, 25, pp.
725-731.
[4] Synder S M and Schnider E. (1991),’Estimation
of mechanical properties of cortical bone by
computed tomography’, J. Orthop, Res, 9, pp. 424-
431.
IFMBE Proc. 2005;9: 288

Biomedical signal processing
HEART RATE VARIBILITY IN
FAMILIAL AMYLOIDOTIC POLYNEUROPATHY
U. Wiklund* , **
* Department of Biomedical Engineering & Informatics, University Hospital, and
**Department of Radiation Sciences, Umeå University, Umeå, Sweden
E-Mail: urban.wiklund@vll.se
Abstract
What are the challenges when applying more or
less sophisticated methods for signal analysis on
patients with severe abnormalities in their regulation
of the cardiac pacemaker? Among all patients being
investigated, treated and followed-up at the
University Hospital in Umeå, Sweden, there is a
small group of individuals with the hereditary
disease Familial amyloidotic polyneuropathy (FAP),
where pronounced disturbances in the autonomic
nervous system are common. This paper gives a brief
review of problems and challenges encountered
during the analyses of heart rate variability signals
recorded from patients with FAP.
Introduction
Autonomic dysfunction is common in patients with
FAP, and can be identified early in the course of the
disease by analysis of short-term heart rate variability
(HRV) [1]. Analysis of HRV is a non-invasive
technique that is based on the beat-to-beat fluctuations
in heart rate, determined from the time intervals
between two successive heartbeats in the ECG.
A striking finding in many tilt-table recordings in
FAP patients is an almost total absence of HRV, except
from some very low-amplitude noise-like fluctuations,
and a nearly constant mean heart rate. The most typical
HRV recording, however, shows a marked reduction in
respiratory related fluctuations in HR, both during
spontaneous and controlled breathing, but the mean
heart rate often shows a significant increase after
passive tilt from supine to the upright position.
These two different patterns in HRV indicate the
presence of a severe cardiac autonomic dysregulation,
which clearly discriminates the majority of FAP patients
from healthy subjects. This is in contrast to many other
clinical studies, where the differences in HRV between
patients and healthy controls may be rather subtle, and
only are apparent after statistical comparisons of group
averages. However, even if there appear to be almost no
HRV in many FAP patients, by applying different tools
for signal processing to these data, new insights
regarding both the underlying physiological process and
the mathematical methods may be obtained, as
discussed in this paper.
Familial amyloidotic polyneuropathy
FAP is a rare disease, but there are local clusters of
individuals with FAP in the northern costal region of
Sweden. At present, approximately 150 patients in this
region are living with the disease. Other “clusters” of
FAP can be found in the northern part of Portugal and in
Japan. FAP is a hereditary form of systemic amyloidosis
characterised by deposits of an insoluble protein —
amyloid — in various organs and tissues, such as the
nervous system, heart, kidneys and the gastrointestinal
tract. The initial symptom of the disease is usually a
painful sensory motor polyneuropathy starting in the
lower extremities. Autonomic nervous dysfunction is
also common, with symptoms such as orthostatic
hypotension, gastrointestinal motility disturbances,
cardiac conduction disturbances, and urinary bladder
dysfunction [2]. The onset of FAP occurs in adult age,
the disease has a progressive course, and is fatal with a
survival of approximately 10 years after the first
symptoms appear. Liver transplantation is the only
treatment to halt the progression of FAP.
Heart rate variability
Many analytical and statistical methods have been
presented for analysis of HRV. The method most
frequently used in the studies on Swedish FAP patients
is power spectral analysis by autoregressive (AR)
algorithm.
The HRV power spectrum is normally divided in
three different spectral regions [3]. The very lowfrequency
(VLF) component with an upper limit of 0.04
Hz is attributed to several physiological variables such
as thermoregulatory fluctuations in vasomotor tone and
fluctuations in the renin-angiotensin system. The lowfrequency
(LF) component (0.04-0.15 Hz) seems to be
related to baroreceptor mediated blood-pressure control,
but these oscillations have also been associated with
central autonomic outflow in some patient groups. The
high-frequency (HF) component (above 0.15 Hz) is
normally reflecting respiratory related fluctuations in
HR.
Arrhythmia or cardiac autonomic modulation?
The analysis of HRV is based on the assumption that
the activity in the autonomic nervous system is the
major source for the triggering of heartbeats. ECG
IFMBE Proc. 2005;9: 289

Biomedical signal processing
abnormalities are common in Swedish FAP patients, at
least among patients older than 55 years [4]. Therefore,
all data series are screened for the occurrence of ectopic
beats or complex patterns in the beat-to-beat heart rate
of non-neural origin. The tool used for this purpose is
Poincaré graphs, i.e., plots of each R-R interval against
the subsequent value [5]. Moreover, patients with
arrhythmia often have spectral peaks at other
frequencies in the HRV spectrum than in the
corresponding respiratory power spectral density.
No low-frequency component in the HRV spectrum?
The shape of the AR power spectrum is smooth,
with a relatively low number of spectral peaks, which
often can be associated with physiological mechanisms.
The selection of model order is not obvious. If the
model order is too low, then there is a poor resolution of
spectral peaks. A model order that is too high may
introduce spurious peaks that are not relevant to the
measured data.
Several FAP patients have a HRV power spectrum
with a very small or even absent peak in the LF region.
This could be a similar finding as in quadriplegic
patients, where the absence of LF peaks has been
reported. The autonomic dysfunction associated with
FAP may also have caused a shift in the frequency of
the LF component, similar to what has been found in
diabetic patients [6]. This could imply that the LF peak
is masked by a stronger VLF component. In our
experience, reasonable estimates of spectral components
are only obtained if high-order AR models are used. In
general, the power spectra are determined using
AR(30)-models of linearly-detrended sequences of twominute
duration.
Very high-frequency oscillations
The presence of HRV in the very high frequency
(VHF) region (above 0.4 Hz) has been reported in heart
transplant patients, and has been associated with a lack
of parasympathetic control of the sinus node [7]. VHF
peaks have also been found in FAP patients [8]. In some
FAP patients, as in several healthy subjects, the VHF
peaks reflects harmonics to the respiration frequency.
However, in several patients the VHF peaks were not
associated with any harmonics to respiration. In at least
two FAP patients the amplitude of the VHF peaks
increased successively after passive tilt to the upright
position, and in one patient this was followed by the
onset of a cardiac arrhythmia.
VHF oscillations may be an indicator of vagal
dysfunction also in FAP patients. However, it could also
be an early sign of disturbances in the atrium, such as
several foci for triggering of heartbeats, or within the
cardiac conduction system, which might develop into
more severe cardiac arrhythmia as the progress of the
disease continues.
Multivariate data analysis
The results from the analysis of HRV are
summarised by a number of variables, such as the mean
heart rate and power in different spectral regions,
calculated for selected segments from different
procedures (supine, upright and controlled breathing).
Principal component analysis (PCA) has been used to
project the multidimensional data into two dimensions,
in order to disclose and visualise groupings of data
points [9]. PCA enhanced the differences in HRV
between FAP patients, as compared to the univariate
statistical analysis.
.
References
1. OLOFSSON B.O., SUHR O., NIKLASSON U.,
WIKLUND U., BJERLE P., and BECKMAN A. (1994):
‘Assessment of autonomic nerve function in
familial amyloidotic polyneuropathy - a clinical
study based of spectral analysis of heart rate
variability’, Amyloid 1, pp. 240-246
2. ANDO Y., and SUHR O.B. (1998): ‘Review:
Autonomic dysfunction in familial amyloidotic
polyneuropathy (FAP) ’, Amyloid: Int. J. Exp. Clin.
Invest. 5, pp. 288-300
3. TASK FORCE (1996): ‘Heart Rate Variability :
Standards of Measurement, Physiological
Interpretation, and Clinical Use’. Circulation, 93,
pp.1043-1065.
4. HÖRNSTEN R, WIKLUND U, OLOFSSON BO, JENSEN
SM, SUHR OB.(2004): ‘Liver transplantation does
not prevent the development of life threatening
arrhythmia in familial amyloidotic polyneuropathy,
Portuguese type (ATTR Val30Met) patients’.
Transplantation 78, pp.112-116.
5. WOO M.A., STEVENSON W.G., MOSER D.K.,
TRELEASE R.B, and HARPER R.M. (1992). ‘Patterns
of beat-to-beat variability in advanced heart
failure’. Am Heart J 123, pp. 704-10.
6. KAMATH MV, FALLEN EL.(1993): ‘Power spectral
analysis of heart rate variability: a noninvasive
signature of cardiac autonomic function’. Crit Rev
Biomed Eng 21, pp. 245-311.
7. TOLEDO, E., PINHAS, I., ARAVOT, D. AND
AKSELROD, S. (2003): ‘Very high frequency
oscillations in the heart rate and blood pressure of
heart transplant patients’, Med. Biol. Eng. Comput.
41, pp. 432-438
8. WIKLUND U, HORNSTEN R, SUHR OB, AKSELROD
S. (2004): ‘Very high frequency heart rate
fluctuations in patients with severe autonomic
dysfunction’. Proc. 10th Mediterr. Conf. of the
IFMBE, July 31-Aug 5, Ischia, Naples, Italy.
9. WIKLUND U, OLOFSSON-BO, SUHR OB,
NIKLASSON U. (2002): ‘Heart rate variability
patterns in familial amyloidotic polyneuropathy’.
Proc. 4th Int. Workshop On Biosignal Interpret.,
June 24-26th, 2002, Como, Italy, pp. 151-154.
IFMBE Proc. 2005;9: 290

Biomedical signal processing
A WAVELET-BASED TECHNIQUE FOR BASELINE WANDER CORRECTION
IN ECG AND MULTI-CHANNEL ECG
A. Khawaja 1 , S. Sanyal 1 , O. Dössel 1
1 Institute of Biomedical Engineering, University of Karlsruhe, Karlsruhe, Germany
ak@ibt.uni-karlsruhe.de
Abstract
In this paper, a new offline method for automatic baseline
drift correction in Electrocardiogram is presented. It is
based on Discrete Wavelet Transform (DWT) and
analyzing high scale Approximation Coefficients (AC). A
set of 650 noisy ECG signals was created by mixing
different artificially generated noise-free ECGs and
baseline wanders. By applying different mother Wavelets
on each noisy signal, twelve stage DWT decomposition
was carried out and twelve filtered ECGs were
reconstructed by canceling the highest level AC at each
stage. The similarities between initially generated
baseline and canceled AC, as well as between the
corresponding noise-free and reconstructed ECGs were
examined every time by means of Correlation technique.
The results from all 650 signals were considered in order
to find the suitable Wavelet and AC level. The highest
correlations, better than 99.9% for baseline and 99.99%
for filtered ECG, were found with the ninth scale
approximation coefficients when using Daubechies11 or
Symlet12 as prototype wavelet. The algorithm was
applied on various MIT-BIH and Multi-channel ECG
signals. Furthermore, the baseline elimination results
were considered to be very promising.
Introduction
In Multi-Channel ECG, Baseline variation may be caused
by coughing, breathing with large chest movement,
variations in temperature and bias in the instrumentation
and amplifiers as well as poor contact and polarization of
the electrodes [1].
The frequency components of the Baseline wander are
usually below 0.5 Hz in rest ECG [2]. Several methods
have been used in the literature to remove the base line
wander. Some of them used high-pass FIR filters with
fixed cut off frequencies [3] [4], whereas another group
has applied a time varying filtering technique letting the
cut off frequency be controlled by low frequency
properties of the ECG [5]. Given that ECG and Baseline
wander frequency spectra usually overlap, it is not
possible to remove the latter with a linear filter without
distorting the ECG components [2].
Another group used third order approximation called
cubic spline [6]. This is a non-linear approach whose
performance depends on the PR intervals (knots)
determination accuracy. It is degraded as the knots
become more separated (low heart rate) [7]. Adaptive
filter has been also proposed, but it still modifies the ST
components [8]. To solve this problem, Cascade adaptive
filter method has been developed [9]. This filter needs a
QRS detector to remove baseline wander preserving the
QRS correlated ECG components (in particular the ST
segment) [7]. Another group used Short Time Fourier
Transform (STFT) technique and time varying filtering to
remove baseline wander [10]. Attaining the optimal time
and frequency resolutions at the same time was the
limitation of this method.
Here, we present a new method to eliminate the
interference of baseline wander using Discrete Wavelet
Transform (DWT).
The ECG signal is characterised by a cyclic occurrence of
patterns with different frequency contents. Due to this
nature of ECG, Wavelet Transform proves to be the best
analysis tool for it. In DWT, time-scale description of a
digital signal is achieved by means of Digital Filters. The
original signal is passed through a series of high and low
pass filters and the outputs are subsequently sub-sampled
by two, producing a set of Approximation and Details
coefficients. The Approximations represent the slowly
changing low-frequency components of the signal,
whereas, the Details reflect the rapidly changing highfrequency
components.
As already stated, Baseline drift in ECG signal is caused
by low-frequency interference (below 0.5 Hz). Hence, its
effect is more pronounced at higher level Approximation
Coefficients. In our algorithm, we go on decomposing the
ECG signal till twelve levels and try to find out which
Approximation level resembles the Baseline wander most
closely.
Methods
We begin with the artificial generation of thirteen noisefree
ECG signals, in the range of 60 to 180 Pulses per
minute (sampling frequency 1 KHz), with the help of inbuilt
commands in Matlab. At the same time, a set of fifty
sinusoidal signals was created, with frequencies ranging
from 0.01-0.5 Hz, to simulate the baseline wander.
Thereafter, a set of 650 test signals were generated by
mixing the artificial ECGs with artificial baseline wander
signals in one to one correspondence.
Each of the 650 Mixture Signals, or noisy ECGs, were
decomposed into 12 DWT levels using different Mother
Wavelets, namely Coiflets (first till fifth order), Symlets
(first till twelfth order) and Daubechies (first till twelfth
order). At every subsequent level, the previous level
IFMBE Proc. 2005;9: 291

Biomedical signal processing
approximation coefficients were decomposed using the
half-band high and low pass filters.
After each level of decomposition, two subsequent
reconstructions were followed immediately before
proceeding to the next level. The reconstruction strategy
in either case is just the reverse of the decomposition
strategy. The first reconstruction is meant to have a
filtered ECG free from the current level approximation
coefficients (CLAC), whereas, the second reconstruction
computes the pure morphology of the CLAC. The first
reconstruction is carried out with CLAC to be all zeros.
On the contrary, second reconstruction is performed with
all coefficients (Details coefficients of the previous
levels, to be specific) other than CLAC to be all zeros.
At each level, the similarity (S1) between the first
reconstruction and the original noise-free ECG as well as
the similarity (S2) between the senond reconstruction and
the corresponding baseline wander signal were computed
by applying the Cross-Correlation technique.
For each of the 650 noisy ECGs, we build two 12x29
matrices containing S1 and S2 values respectively. The
rows denote the DWT decomposition level and each
column represents a different mother wavelet.
To choose the best suitable Mother wavelet and the
appropriate scale (i.e. DWT decomposition level) over
the whole range of ECG signals and baseline wander
under consideration, we calculated M1 and M2 matrices.
M1 is the mean of S1 matrices and M2 is the mean of S2
matrices for all the 650 different noisy ECGs. Finally we
considered the elements in M1 that are greater than
99.99% and we matched them with the corresponding
elements in M2.
Results
The effect of Baseline wander is found to be most
pronounced at the ninth level Approximation.
Figure 1: Single Channel from a Multi-channel ECG
A: ECG Signal Corrupted With baseline wander
B: Ninth level Approximations using Daubechies11
C: Reconstructed ECG
signals, are Daubechies11 and Symlet12. The similarity
percentages represent the Cross-Correlation coefficents in
M1 and M2.
Discussion
This method is able to eliminate the Baseline drift
without any distortion of ST segment as observed with
conventional high pass filters. Moreover, it can be
applied equally well to short and long duration ECG
signals. We tested the performance of a conventional
high pass filter (second order Butterworth filter with 0.5
Hz cut off frequency) and our technique on a generated
ECG having 0.1 Hz Baseline wander. The conventional
filtered ECG showed only 97% similarity to the freenoise
ECG, whereas our wavelet-based technique showed
greater than 99%.
References
[1] RANGARAJ, M. R. (2002): 'Biomedical Signal
Analysis', (J. Wiley and Sons, New York).
[2] LAGUNA P., JANE R., and CAMINAL P., 'Adaptive
Filtering of ECG Baseline Wander', IEEE, 1992.
[3] J.A. VAN ALSTE, T. S. SCHILDER, 'Removal of
Baseline wander & Power line interference from the ECG
by an efficient FIR filter with a reduced number of taps',
IEEE Trans. Biomed. Engg, val BME-32, No.12, Dec
1985, pp1053-1060.
[4] I.I. CHRISTOV, I.A. DOTSINSKY, I. K.
DASKALOV, 'High Pass filtering of ECG signals using
QRS elimination'. Med. &Biol. Engg. & Computing,
March 1992, pp 389-337.
[5] L.SORNMO.'Time varying digital filtering of ECG
baseline wander'. Med& Biol. Engg. & Computing,
September 1993, pp 305-508.
[6] MEYER, C.R., and KEISER H.N.,'Electrocardiogram
baseline noise estimation and removal using cubic splines
and state-space computation techniques', Computers and
Biomedical Research, vol. 10 pp. 459-470. 1977.
[7] JANE R., LAGUANA P., THAKOR N.V., and
CAMINAL P. (1992): 'Adaptive Baseline Wander
Removal in the ECG: Comparative Analysis With Cubic
Spline Technique', IEEE.
[8] THAKOR, N.V., ZHU, Y. 'Applications of Adaptive
filtering to ECG analysis: noise cancellation and
arrhythmia detection', IEEE trans. Biomed. Eng., vol. 38,
n. 8, pp. 785-794. 1991.
[9] LAGUNA P., JANE R., Meste O., Poon P., Caminal
P., RIX H., and THAKOR N.V.,'Adaptive filter for
event-related bioelectric signals using an impulse
correlated reference input: Comparision with signal
averaging techniques' IEEE Trans. On Biomedical
Engineering, vol. 39 No 10. 1992.
[10] PANDIT S.V., 'ECG Baseline Drift Removal
Through STFT', IEEE, 1997.
The mother Wavelets, which yielded greater than 99.99%
similarity between the filtered and noise-free ECGs and
99.9% similarity between the baseline and ninth level AC
IFMBE Proc. 2005;9: 292

Biomedical signal processing
Evaluation of an Efficient Method for Handling Ectopic Beats in HRV
K. Solem 1 , P. Laguna 2 and L. Sörnmo 1
1 Signal Processing Group, Dept of Electroscience, Lund University, Lund, Sweden
2 Aragon Inst. for Engineering Research, Zaragoza University, Zaragoza, Spain
kristian.solem@es.lth.se
Abstract: The problem of analyzing heart rate
variability in the presence of ectopic beats is revisited.
Based on the IPFM model and the closely related
heart timing signal, a new computationally very
efficient technique is introduced which corrects for
the occasional ectopic beats. From actual heart
rate data, the results show that the new technique
is associated with a much lower computational
complexity (almost 3000 times faster) than the
original heart timing approach, while producing
similar performance. This also implies that the power
spectrum and related clinical indices obtained by the
new technique are more accurately estimated than by
other methods.
1. Introduction
The presence of ectopic beats perturbs the impulse
pattern initiated by the sinoatrial node, and implies that
RR intervals adjacent to an ectopic beat cannot be used
for heart rate variability (HRV) analysis. Since ectopic
beats may occur in both normal subjects and patients with
heart disease, their presence represents an important error
source which must be dealt with before spectral analysis
can be performed. If not dealt with, the analysis of an
RR interval series containing ectopic beats results in a
power spectrum with spurious frequency components.
The heart timing (HT) signal was recently suggested
for characterization of heart rate variability [1]. This
signal is based on the well-known integral pulse
frequency modulation (IPFM) model for the generation
of normal sinus beats [2], characterizing HRV in
terms of a modulation function m(t). The definition
of the HT signal has later been extended to also
account for the presence of occasional ectopic beats [3].
In terms of spectral distortion, the results showed
that the HT-based correction produced one order of
magnitude lower error than did interpolation-based
correction techniques. While producing excellent
results, the HT-based correction is associated with
heavy computations which, for example, in the analysis
of Holter recordings, may become prohibitive. The
present paper introduces a correction method which
drastically reduces the computational demands of the
method presented in [3], while introducing no significant
deterioration in performance.
2. Methods
The heart timing signal d HT (t) is at each heartbeat
occurrence time t k defined as, d HT (t k ) = kT 0 −
t k , where T 0 denotes the mean RR-interval length [1].
The HT signal is closely related to the IPFM model
and its modulating signal m(t). Using the HT signal,
the modulating signal m(t) can be estimated in order
to produce the HRV power spectrum. If the Fourier
transform of m(t) and d HT (t) are denoted with D m (Ω)
and D HT (Ω), respectively, it can be shown that [1]
D m (Ω) = jΩD HT (Ω), (1)
where Ω = 2πF . Hence, the desired spectral
estimate D m (Ω) can easily be computed once the Fourier
transform of the heart timing signal, D HT (Ω), is known.
In the description below, we assume that sinus beats
occur at the times t 0 , t 1 , . . . , t K , and that one ectopic
beat occurs at time t e . The time t e is not included in
the series t 0 , t 1 , . . . , t K , and the sinus beat immediately
preceding the ectopic beat occurs at t ke and the sinus beat
immediately following at t ke+1.
Heart Timing Representation: In order to compensate
for the presence of an ectopic beat, the above definition of
d HT (t k ) is modified by the introduction of a parameter s
according to [3].
d HT (t k ) =
{
kT 0 − t k k = 0, . . .,k e ,
(k + s)T 0 − t k k = k e + 1, . . .,K.
(2)
The parameter s can be viewed as a jump in the resetting
of the integral in the IPFM model, and may be obtained
by LS estimation described in [3]. An estimate of T 0 is
obtained by t K /(K + ŝ).
Computationally Efficient Representation: A different
approach to deal with ectopic beats is to observe that an
ectopic beat shifts the occurrence times of the following
normal heart beats. By estimating the time shift δ, the
presence of an ectopic beat can be accounted for by
d HT δ
(t k ) =
{
kT 0 − t k k = 0, . . . , k e ,
kT 0 − t k + δ k = k e + 1, . . .,K,
and δ estimated according to [4]
(3)
ˆδ = t ke+1− 2t ke + t ke−1. (4)
An estimate of T 0 is obtained by (t K − ˆδ)/K.
IFMBE Proc. 2005;9: 293

Biomedical signal processing
3. Database
The database consists of 132 ECG episodes selected
from the European ST-T database, previously studied
in [3]. The ectopic beat composition of the 132 episodes
is as follows: 91 episodes contain one ectopic beat,
28 contain two, 5 contain three, 4 contain four, 2
contain eight, and 2 contain ten ectopic beats. Each
ECG episode is divided into three overlapping, fourminute
segments: A, B, and C. Segments A and C are
ectopic-free, whereas segment B contains the ectopic
beat(s). Segment A contains the four minutes preceding
the ectopic beat(s), segment B is centered around the
ectopic beat(s), and segment C contains the four minutes
following the ectopic beat(s).
The database is studied using the evaluation approach
introduced in [3], where it was suggested that the spectral
characteristics of the ectopic-free segments A and C can
be compared to the corrected segment B assuming that
the heart rate is stationary once ectopy has been removed.
Three parameters, ∆AC, ∆AB, and ∆BC, are defined,
where ∆AC denotes the difference in spectral power
between segments A and C, and so on. Moreover, the
power is divided into two subbands: a low frequency
(LF) band (0.04–0.15 Hz) and a high frequency (HF)
band (0.15–0.40 Hz). Assuming stationarity during the
segments A, B, and C, ∆AC is expected to be close to
zero in both frequency bands, and following correction
of segment B, ∆AB and ∆BC should be close to that of
∆AC.
4. Results
The performance of the δ estimator in (4), based
on d HT δ
(t k ), is compared to that of the minimum LS
estimator of s based on d HT (t k ), in terms of HRV power
spectral differences. The HRV power spectra of the
ectopic-free segments A and C is computed using the
definition of the HT signal, whereas the power spectrum
of segment B requires that either the δ estimator or the
s estimator is used.
Table 1 presents the results when all 132 ECG episodes
are analyzed. The performance of the two different
estimators is almost identical for both the LF and HF
bands of ∆AB and ∆BC, and is comparable to the
variation of ∆AC.
In order to compare complexity of the two estimators,
the number of floating point operations (flops) was
studied, see Table 2. The results show that the
s estimator requires almost 3000 times more flops than
the δ estimator. It is also noted that the number of flops
used by the δ estimator is deterministic since, in contrast
to the s estimator, it is independent of where the ectopic
beat occurs.
5. Conclusions
Table 1. HRV power spectral differences related to the s
and δ estimators. Values are given in mean±std in the
unit ms −2 .
∆AC
Estimator LF HF
s
δ
49±808 –22±190
∆AB
Estimator LF HF
s –32±644 –26±153
δ –32±651 –30±136
∆BC
Estimator LF HF
s 80±498 4±158
δ 81±501 8±145
Table 2. Flop statistics for the s and δ estimators, when
using the HT signal for the 132 ECG episodes.
Estimator Mean Std
s 22514 26359
δ 8 7
the contribution of a new HT-based method. The
performance was compared to the original method, which
is based on the heart timing signal [3], and was found
to be the same when performance is measured in power
spectral terms. However, the new method requires
dramatically less computations and is therefore much
better suited for implementation in systems for long-term
ECG analysis.
References
[1] MATEO J, LAGUNA P. (2000): ’Improved heart rate
variability signal analysis from the beat occurrence
times according to the IPFM model’ , IEEE Trans
Biomed Eng, 47, pp. 997–1009.
[2] HYNDMAN BW, MOHN RK. (1975): ’A model
of the cardiac pacemaker and its use in decoding
the information content of cardiac intervals’ ,
Automedica, 1, pp. 239–252.
[3] MATEO J, LAGUNA P. (2003): ’Analysis of heart
rate variability in the presence of ectopic beats using
the heart timing signal’ , IEEE Trans Biomed Eng,
50, pp. 334–343.
[4] SOLEM K, LAGUNA P, SÖRNMO L. (2004):
’Handling of ectopic beats in heart rate variability
analysis using the heart timing signal’ , In Proc. Med.
Conf. Med. Biol. Eng. (MEDICON), Ischia, Italy pp.
(CD–ROM).
The present paper sheds new light on the problem
of ectopic beat correction in HRV analysis with
IFMBE Proc. 2005;9: 294

Biomedical signal processing
ACCURACY LIMITATIONS OF THE DOPPLER ULTRASOUND
SIGNAL IN RELATION TO FHR VARIABILITY ANALYSIS
J. Wrobel, J. Jezewski, K. Horoba, A. Gacek
Institute of Medical Technology and Equipment, Department of Biomedical Informatics
118 Roosevelt Str., Zabrze, Poland
januszw@itam.zabrze.pl
Abstract: This work was aimed at estimation of the
influence of commonly used Doppler ultrasound
method on the clinical assessment of the fetal state,
assuming that at present this is based on calculated
automatically indices describing instantaneous fetal
heart rate (FHR) variability. The method has been
developed which enables comparison of variability
indices obtained from ultrasound method with their
reference values calculated from direct fetal electrocardiogram.
The results obtained reveal that ultrasound
method is not able to provide the enough
accuracy of beat-to-beat interval determination,
which is required for reliable quantitative evaluation
of FHR variability. This limitation causes a decrease
of the indices, but may not have serious consequences
for the detection of fetal distress signs.
Introduction
Information on fetal heart activity is a basis for fetal
wellbeing assessment. Present-day fetal monitors
provide this information in a form of fetal heart rate
(FHR) signal using Doppler ultrasound method. Accuracy
of FHR measurement at a level of 1% ensured by
the ultrasound method is sufficient for visual analysis
of a strip-chart record [1]. However, the fetal state assessment
based on visual interpretation only is characterized
by low inter- and intraobserver agreement. One
of the most important signs of physiology in FHR record
is continuous fluctuation in time of beat-to-beat
intervals. There are various mathematical indices used
for quantitative evaluation of two types of FHR variability
called short and long term variability [2]. Definitions
of these indices have been created in 1970s
basing on beat-to-beat intervals precisely determined
from direct fetal electrocardiogram (FECG). As the
ultrasound method has became the standard approach
to fetal monitoring FHR variability indices have been
straight implemented without any adaptive changes.
But we should consider limitations of the ultrasound
approach with regard to accuracy of the fetal heart rate
determination on a beat-to-beat level. Commonly used
autocorrelation technique is able to recognize heart
beat events even in a signal of such complex shape like
the Doppler envelope, but the precision of the heart
beats localization is lower than using fetal electrocardiogram
[3]. Additionally, autocorrelation function
tends to average the determined successive cycles. The
aim of this study was to estimate what is the influence
of ultrasound method on correct clinical assessment of
the fetal state, assuming that at present this assessment
is computerized with FHR variability indices calculated
automatically. This aim required the variability indices
obtained from ultrasound method to be compared to
their reference values derived from fetal electrocardiogram.
Methods
Measurement station has been based on laptop PC
with PCMCIA data acquisition card – DAQCard-AI-
16XE (National Instruments). This card has eight analog
inputs and A/D converter which can operate with
maximal sampling frequency of 200 kHz. FHR signals
were recorded using the MT-430 (Toitu) fetal monitor.
Fetal heart beats are detected by analysis of the envelope
of ultrasound wave reflected from moving parts of
fetal heart (valves or walls). Simple peak detection can
provide incorrect values of cardiac cycles due to the
complex and unstable shape of the envelope signal. Figure
1 shows a fragment of ultrasound envelope with
marked particular events of cardiac cycle.
Figure 1: Determination of T RR duration from the Doppler
ultrasound envelope (US) using peak detection
method. Cardiac cycle events: atrial wall contraction –
Atc, mitral valve opening and closure – Mo and Mc,
aortic valve opening and closure – Ao and Ac
IFMBE Proc. 2005;9: 295

Biomedical signal processing
Various durations of cardiac cycle can be obtained
depending on which event is selected as representing a
given cycle. The proper value is only one – calculated
from FECG signal. Due to this, for the recognition of
consecutive heart beats in ultrasound signal the autocorrelation
technique considering full shape of analyzed
signal is commonly applied. Distance between
two consecutive peaks of autocorrelation function corresponds
to the interval between two consecutive R-
waves. Values of T RR intervals are transformed into
instantaneous values of FHR expressed in beats per
minute accordingly to the equation:
FHR i [bpm] = 60000/T RRi [ms]
Direct FECG was recorded by means of a typical
spiral electrode. Amplifier with a selectable gain between
500 and 2000 V/V was applied. Band-pass filter
at 0.05 Hz ensured suppression of low-frequency noise
components and elimination of an isoline drift. The
upper frequency of 300 Hz overcame the aliasing
problem. Recorded analog signals were sampled with
2 kHz, which ensured a required accuracy for reference
T RR intervals. The virtual instrumentation software
for measurement system was implemented using
LabVIEW environment. The reference FHR signal
was determined off-line on a basis of FECG signal using
two independent QRS detection methods: energy
sensitive technique and crosscorrelation analyzing the
shape of signal. The reference T RR interval was accepted
if the difference of R-wave locations between
the methods did not exceed 0.5 ms.
Results
Traces were collected from 5 women during the labour,
total length was 185 minutes. For the acquisition
methods comparison based on variability indices, very
important was to mark a lack of T RR interval if at least
one of the corresponding intervals has been missed in
one of FHR signals being compared. Thanks to that,
for both signals a given index was calculated using the
same number of intervals in every one-minute period.
Table 1: Descriptive Statistics of the Indices Errors
distributed in one-minute periods, only 12 of 185 periods
were discarded. In order to unify the comparison
procedures, the index definitions have been modified to
determine FHR variability in one-minute periods. Values
of indices and their relative errors were calculated
for the ultrasound method where the reference values
were obtained from FECG signal (Table 1).
Conclusions
The errors of indices determined using the ultrasound
method always take negative values, which
means that these indices are lower than the reference
indices calculated for FECG signal. This is mainly the
effect of correlation techniques applied in ultrasound
channel which causes the averaging of T RR values. So,
differences between consecutive T RR intervals and in
consequence values of variability indices decrease.
Taking into account the obtained errors values we
can put long term indices in distinctive groups. The best
indices (having the lowest sensitivity to the measurement
method) are: de Haan’s with error of –2 % (SD =
±7.4 %) as well as Yeh’s, Organ’s and Dalton’s with
error about –5 % (SD = ±5 %). The next group includes
Zugaib’s and Dawes’ indices with an error reaching –10
% (SD = ±10 %). Huey’s index whose error equals –
33.4 % (SD = ±8.9 %) is outside of any expectable
range. Short term indices can be ordered from the worst
to the best: de Haan’s, Van Geijn’s, Huey’s, Dalton’s
Zugaib’s and Yeh’s. Their error decreases from about –
40 %, with 7 % step for successive indices to value of –
5 %, dispersion of errors is very large.
The Doppler ultrasound method is not able to provide
the FHR signal of the accuracy required for reliable
quantitative evaluation of FHR variability – particularly
short term variability – based on the indices calculated
automatically. However, this limitation prevents fetal
distress signs from being unnoticed since the values of
the variability indices are decreased.
Acknowledgment. Scientific work financed from the
State Committee for Scientific Research resources in
2003÷2005 years as a research project No.4T11F00225.
References
Limit for signal-loss for the one-minute period has
been established at 20 % of number of intervals. The
average length of interval equals from 400 to 500 ms,
so 20 % threshold corresponds to the range of 24 to 20
lost beats. Because the lost intervals were not regularly
[1] DAWES G. S., VISSER G. H. A., GOODMAN J. D .S.,
REDMAN C. W. G. (1981): ‘Numerical analysis of
the human fetal heart rate: The quality of ultrasound
records’, Am. J. Obst. and Gynecol., 141, pp. 43-52.
[2] KUBO T., INABA J., SHIGEMITSU S., AKATSUKA T.
(1987): ‘Fetal heart variability indices and the accuracy
of variability measurements’, Am. J. Perinat., 4,
pp. 179-186.
[3] JEZEWSKI J., WROBEL J., HOROBA K., MATONIA A.,
KUPKA T. (2002): ‘Estimation of beat-to-beat accuracy
of fetal heart rate data obtained via Doppler ultrasound’,
Proc. of EMBEC’02, Vienna, XII 2002,
pp. 1536-1537.
IFMBE Proc. 2005;9: 296

Biomedical signal processing
OPEN-LOOP MODEL OF EQUINE HEART CONTROL
J. Holcik*, P. Kozelek*, M. Jirina*, J.Hanak**, M.Sedlinska**
* Inst. Biomedical Engineering, Czech Technical University, Kladno, Czech Republic
** Equine Clinic, University of Veterinary Science and Pharmacy, Brno, Czech Republic
holcik@ubmi.cvut.cz
Abstract: The paper describes a mathematical
model and results of simulation experiments
explaining relationship between RR and QT
intervals in equine ECG signals. The simulation
results support a hypothesis that electrical
processes in equine heart are controlled by
different mechanism than in human. In particular
it means that the ventricular activity of equine
heart is more influenced by a vagal neural system
when compared with human heart.
Introduction
Publications of sudden cardiovascular death
mortality rate in horses associated with general
anaesthesia estimate it is about hundred times greater
than in human population. Such a great difference in
the mortality rate led our team to study electrical,
mechanical as well as control processes running in
equine heart and made us look for causes of the
problem.
First findings dealing with differences in
relationship between RR intervals (describing control
of SA node) and QT intervals (describing properties
of spreading electrical excitation trough tissue of
myocard ventricles) in equine heart were published in
[3]. Experimental results mentioned in [3] show that
equine ECG records exhibit much more
heterogeneous and complicated relationships between
QT and RR intervals compared with human ECG.
The most typical case is represented by prolonged QT
intervals during shortening of RR intervals and only
after that their values return to the initial point - this
kind of behaviour has not been found in human ECG
signals yet.
The last case is completely beyond our
expectations and there is no relevant explanation of
the mentioned fact in any available publications.
Therefore, it has been necessary to confirm or
exclude some hypothesis about causes of the
observed phenomenon either experimentally or by
means of computer simulation.
In [4] we described and applied very simple
mathematical open-loop model of the control of
electrical processes in equine heart. However, the
model uses only implicit stimulation of ventricular
tissue by parasympathetic nerves, so it was difficult to
analyze particular contributions of both neural
branches to stimulation of all parts of the
myocardium.
Materials and Methods
The structure of the developed model can be
described by an open loop block diagram depicted in
Fig.1. The model expresses separate influence of both
neural branches on SA node and ventricular walls.
Figure 1: Block diagram of the model
The blocks SYM and PSYM express level of an
activity of the sympathetic and parasympathetic
systems according to Fig.2 and the following equation
⎧
bs,p
⎪as,pNE
+ bs,p
if NE
> −
Ns ,p = ⎨
as,
p
(1)
⎪
⎩ 0 otherwise
Figure 2: Sympathetic and parasympathetic
transform functions
The paths representing generation of RR interval
sequence do not contain any additional subsystems
(delay and amplifying) because the model input is
able to define the basic reference level with sufficient
precision required for the simulation experiments. On
the other hand, the blocks in branches defining QT
intervals represent time delay and level of influence of
both neural branches, as it is necessary for calculation
of the QT sequence. The delay blocks do not represent
only delay following from the finite velocity of
IFMBE Proc. 2005;9: 297

Biomedical signal processing
RR[s] →
QT[s] →
2.2
2.1
2
1.9
1.8
1.7
1.6
1.5
0.56
0.55
0.54
0.53
0.52
0.51
0.5
0.49
0.56
0.55
0.54
0.53
0.52
0.51
0.5
0.49
260 280 300 320 340 360 380 400 420 0.48
t[s] →
spreading electrical excitation along nerve fibres, but
also inertia of the nerves around heart ventricles.
The formulas describing relationships among
RR, QT and levels of neural activity N s and N p , resp.
are based on additive relations
RR(t)
= RRsau − ksR
Ns
(t) + k pR Np
(t) (2a)
and
QT(t) = QTk
− k sQ N s (t − τsQ
) +
(2b)
+ k N (t − τ )
pQ
where RR sau is a basic heart period of sine node, N s
and N p represent sympathetic and parasympathetic
activity levels, OT k is a basic length of QT interval at
neural ventricular blocade, k sR , k pR , k sQ and k pQ are
multiplicative parameters of the neural branches and
finally τ sQ and τ pQ are delays in sympathetic and
parasympathetic neural branches at heart ventricles.
Input signal N E represents response of the neural
feedback to impulse stimulation. It is described as
⎧ ⎡ ⎛ 2π
π ⎞ ⎤
⎪A.
⎢sin⎜
(t − t1)
− ⎟ + 1⎥,
⎪ ⎣ ⎝ T 2 ⎠ ⎦
N E = ⎨ if t1
≤ t ≤ t1
+ T
(3)
⎪
⎪
⎩
0, otherwise,
where T is a duration of the input impulse and t 1
represents its lag after some reference starting point.
p
RR (experimental)
RR (simulated)
QT (experimental)
QT (simulated)
experimental data
simulated data
0.48
1.5 1.6 1.7 1.8 1.9 2 2.1 2.2
RR[s] →
Figure 3: Experimental and simulated RR and QT
sequences and their state-space representation for
optimised parameters - k sQ =3.0 s, k pQ =-3.5 s ,
τ sQ =8.1 s, τ pQ = 6.9 s , T = 40 s and A = 0.034
pQ
QT[s] →
Results
Optimization of model parameters (k sQ , k pQ, τ sQ
τ pQ, T and A) has been done for 8 most representative
sets of experimental data and the optimised
parameters were used for simulation calculations.
Results of simulations were compared to the
measured real data (Fig.3).
Discussion
The results of optimization calculations show
very close, nearly linear, dependency between both
values of delay in the sympathetic and
parasympathetic branches and similarly between
values of gain in both branches. These dependencies
make the model less sensitive to the most appropriate
parameters for the model. On the other hand, all
obtained results explain the occurrence of the
prolonged QT intervals during the shortening of RR
intervals. This behaviour is caused by a significant
influence of vagal neural system on ventricular
activity of equine heart.
Conclusion
The described model provides a good tool for
verifying hypotheses about background of electrical
and control processes in equine heart. The future work
should invest more effort into development of new
models which will be more sensitive to the particular
values of delay and gain in both branches representing
neural control of ventricular activity in equine heart.
Acknowledgement
The research was granted by the project of the
Grant Agency of the Czech Republic No.102/04/0887
„Methods and Technical Tools for an Analysis of
Sudden Cardiovascular Death in the Horse“
References
[1] JOHNSTON G. M. et al. (1995): Confidential
enquiry of perioperative equine fatalities
(CEPEF-1): preliminary results. Equine Vet J.,
27, pp.193-200.
[2] JOHNSTON G. M. (1995): The risks of the game:
the confidential enquiry into perioperative equine
fatalities. Br Vet J., 151, pp.347-50.
[3] HOLCIK J., MUSIL J., HANAK J., and SEDLINSKA
M.: ‘Pilot Study of the RR and QT Interval
Relationship in Equine Electrocardiograms‘, CD
ROM IFMBE Proc. of MEDICON 2004 and
HEALTH TELEMATICS - X Mediterranean
Conf. on Med. and Biol. Eng. and Comput,
Ischia, Italy, 2004, 4p.
[4] HOLCIK J., KOZELEK P., HANAK J., and
SEDLINSKA M.: ‘Mathematical Modelling as a
Tool for Recognition of Causes of Disorders in
QT/RR Interval Relationship in Equine ECG.
Proc. of PRIA2004, St. Peterburg, Russia,
part.III, p.688-691.
IFMBE Proc. 2005;9: 298

Biomedical signal processing
ADAPTIVE MULTICHANNEL FILTER FOR HEART BEAT DETECTION
F. Ragnarsson*, N. Östlund* , ** and U. Wiklund* , **
* Department of Biomedical Engineering & Informatics, University Hospital, and
**Department of Radiation Sciences, Umeå University, Umeå, Sweden
E-Mail: urban.wiklund@vll.se
Abstract: In this study an adaptive multichannel
filter for detection of heartbeats in ECG signals was
implemented and evaluated. The adaptive
multichannel filter uses a spatial and temporal finite
impulse response filter to extract the heartbeat
events from the ECG signal. The filter coefficients
are adaptively updated with an independent
component analysis (ICA) algorithm to maximize the
super-gaussianity of the output signal. At the filter
output the heartbeat events occur as distinct peaks
and are detected with a threshold detector. The
multichannel filter was tested with eight channels of
recorded ECG signals. Although the noise level was
very high in some of the channels, 99% or more of
the heartbeat events were extracted in recordings
from three healthy subjects.
Introduction
Realtime analysis of the beat-to-beat fluctuations in
heart rate requires that the time instance for every
heartbeat is detected, and that there are as few missing
and false detections as possible. However, automatic
detection of heartbeats could be difficult when the ECG
signal is mixed up with noise. Many of the proposed
algorithms for heartbeat detection are sophisticated
single-channel detectors that can handle relatively high
noise levels. It is, however, natural to assume that a
multichannel detector would have a better performance
then a single-channel detector, because the redundancy
of the system is expected to be higher when data from
many channels are analysed simultaneously. Recently,
an adaptive multichannel filter was developed for
analysis of electromyographic signals [1], and the aim
of this study was to apply this filter for extraction of
heartbeat events in noisy ECG signals.
Materials and Methods
Data acquisition: Eight ECG channels were
recorded at 500 Hz using a wireless multichannel data
acquisition system. Data were recorded from three
healthy male subjects (age 27-32). The ECG was
measured bipolar and the electrodes were placed on the
chest. During the recording intermittent noise was
generated by high levels of skeleton muscular activity.
Disturbances were generated by removal and replacement
of electrodes. In addition, touching the
electrodes rapidly generated high-amplitude transients.
Multichannel filter: The basic idea is to design a filter
where the output signal has distinct peaks corresponding
to the time instants where the QRS complexes occur,
and are close to zero elsewhere. In that signal, most of
the data points have an amplitude value close to zero
and the peaks have a significantly larger value and
occur as a marked tail in the histogram, i.e., has a supergaussian
distribution. The aim with the filter is to
maximize the super-gaussianity of the output signal.
In order to maximize the super-gaussianity, the
adaptive multichannel filter uses both spatial and
temporal filtering. The time filtering is performed using
individual finite impulse response filters on each input
channel i according to:
z i = h i * x i (1)
where x i is the input signal and h i = [h i1 h i2 ... h ip ] is the
filter kernel.
With M input channels the spatial filtering is
accomplished as
y = g 1 z 1 + g 2 z 2 + ... + g M z M (2)
Substitution of g i h i with w i the filter equation yields:
y = w 1 * x 1 + w 2 * x 2 + ... + w M * x M (3)
Obviously, the output signal y is a linear combination of
time delayed input signals. The filter coefficients were
determined adaptively for blocks of length N=3000 with
an ICA algorithm. In this study the FastICA [3]
algorithm was used to maximise the super-gaussianity
of the output y, with skewness as cost function.
Pre- and post-processing: In order to stabilise the
algorithm, pre- and post-processing was added. The
input signals were high-pass filtered in order to suppress
base line drift. Also, the signals were normalised to zero
mean and unit variance.
Since the analysis was performed on blocks of data,
an algorithm for time-alignment of successive blocks of
the filter output was implemented [3].
The ICA-algorithm outputs a set of solutions where
the super-gaussianity is maximized for all individual
solutions with the constraint that the solutions are
uncorrelated. The number of output signals is equal to
the number of input channels times the filter length.
IFMBE Proc. 2005;9: 299

Biomedical signal processing
In general, the most super-gaussian output signal
gives a solution where the QRS-complexes are extracted
and the noise is suppressed. However, a method based
on unsupervised clustering was used to increase the
robustness of the algorithm for selection of the “best”
solution [3].
Heart-beat detection: Time instants for individual
heartbeats were determined from the selected output
channel using an threshold detector.
Evaluation of performance: All heartbeat event
times were manually confirmed and errors corrected
based on visual inspection of the ECG data.
All calculations were performed using the Matlab
software package (Mathworks, Natick, Mass.).
Results
Figure 1 shows a part of one ECG recording, where
severe muscular disturbances are present in all six input
channels. The time instants for the spikes in the output
signal before and after noise was present indicate that
the algorithm was able to suppress the noise.
The results in another part of the recording with
other disturbances are shown in Figure 2.
A thresholding algorithm was used to determine
time instants for individual heart beats based on the
output signal from the multichannel filter. As shown in
Table 1, a low number of misclassifications of heart
beats were made in the recording from subject 3. All
misclassifications occurred in a single subsegment
where spikes were present in one channel, most likely
due to communication problems between the wireless
data acquisition system and the computer.
Figure 2: In this part of the recording noise was added
by touching and removing electrodes. All six input
channels were used to determine the output signal
shown at the bottom.
Discussion
In this study we have implemented an adaptive
multi-channel filter for detection of time instants of
heart beats. Although the noise level was very high in
some of the channels, 99% or more of the heartbeat
events were extracted in recordings from three healthy
subjects.
Acknowledgements
The study was supported by grants from the
Swedish Research Council (number 2003-4833). The
development of the system is performed within the
Centre for Biomedical Engineering and Physics at
Umeå University, Sweden.
References
Figure 1. In this figure two input channels are shown,
where the ECG was disturbed by electrical activity from
the chest muscles. The bottom tracing shows the output
signal, where six ECG channels with muscular noise
were used as inputs.
Table 1. Performance of the adaptive channel filter
Subject total correctly false missed
number detected detections beats
of beats
1 415 415 0 0
2 674 674 0 0
3 582 578 5 4
1. ÖSTLUND, N., YU, J. AND KARLSSON, J.S. (2005):
‘Adaptive spatio-temporal filtering of multichannel
EMG signals', Submitted
2. HYVÄRINEN, A.: ‘The fast ICA package for Matlab',
http://www.cis.hut.fi/projects/ica/fastica/
3. RAGNARSSON, F. (2005): 'Adaptive multichannel
filter for ECG analysis', Masters thesis, Umea
University.
IFMBE Proc. 2005;9: 300

Biomedical signal processing
OTOACOUSTIC EMISSIONS TIME FREQUENCY MAPPING BASED ON
THE ENSEMBLE CORRELATION AND HILBERT-HUANG TRANSFORM
A. Janušauskas*, A. Lukoševicius*, V. Marozas*, and L. Sörnmo**
*Kaunas Technology University/Biomedical Engineering Institute, Kaunas, Lithuania
** Signal Processing Group/Lund University, Lund, Sweden
artjanu@ktu.lt
Abstract: This paper presents an application of the
Hilbert-Huang transform (HHT) and ensemble
correlation for high-resolution time-frequency
mapping of transient evoked otoacoustic emission
signals (TEOAE). The algorithm decomposes
TEOAE responses into intrinsic modes using
empirical mode decomposition followed by
extraction of the correlated signal parts, Hilbert
spectrum calculation for each mode, and mapping
extracted instantaneous frequencies and amplitudes
onto the time-frequency plane. The results show
good correspondence between audiometric
thresholds better than 20–30dB HL and the presence
of otoacoustic emissions at these frequencies.
Therefore, high-resolution audiogram thresholds
worse than 20–30dB HL may be predicted from
TEOAE signal records.
Introduction
Transient evoked otoacoustic emission is a tiny
acoustic signal recorded in the outer ear canal in
response to a short acoustic stimulus. The TEOAE
contain information about a large part of the inner ear.
Studies have shown that otoacoustic emissions can be
generated in subjects with a mean hearing level better
than 20–30dB HL [1,2].
Analysis of TEOAE time-frequency properties is
also of interest due to their relation to cochlear
mechanisms. Transient otoacoustic emissions
represent a highly non-stationary signal, and various
time-frequency distributions have been used for
TEOAE time-frequency mapping including the shorttime
Fourier transform, the wavelet transform, the
Wigner-Ville distribution [3,4]. However, all these
methods have serious drawbacks related to poor
resolution of the short-time Fourier transform, the
presence of cross-terms of the Wigner-Ville
distribution, or reduced time-frequency resolution due
to a limited match of basis functions to the nonstationary
signal in the wavelet case.
In this paper, a technique based on the Hilbert-
Huang transform and the ensemble correlation method
(EC) is proposed for TEOAE signal extraction from
noise and for high-resolution time-frequency mapping
[2, 5]. This technique does not require basis functions
but is fully signal adaptive. It allows a very fine
frequency resolution while the time resolution depends
only on sampling frequency.
Materials and Methods
The method includes the following steps:
1) averaging of TEOAE signal realizations (upper path
in Fig. 1), as recorded by the ILO92 OAE recording
equipment in “raw” mode from adult subjects [2,4] into
12 sub-averages, extraction of the first 4 intrinsic mode
functions (IMF) for each subaverage using empirical
mode decomposition method (EMD), and ensemble
correlation (EC) function calculation for each group of
intrinsic modes; 2) averaging of the whole OAE
ensemble (lower path in Fig. 1), calculation of the first 4
IMF for the whole average, and application of Hilbert
transform for each IMF to get instantaneous frequencies
and amplitudes; 3) smoothing of instantaneous
frequencies, and time-windowing of Hilbert amplitudes
using EC information; 4) merging of pre-processed
instantaneous frequencies and amplitudes from 4 time
scales into one time-frequency plane and display of
results. The methods used in each step are described
below.
Figure 1: Algorithm for TEOAE time-frequency
mapping.
Empirical mode decomposition method is a
technique for decomposing the signal into a series of
intrinsic oscillatory modes, so-called "intrinsic mode
functions", that describe the instantaneous frequency of
the associated Hilbert spectrum [5]. The EMD method
is a fully data driven method consisting of a few
iterative steps: 1) local maxima and minima are
determined in the signal and used for the creation of
lower and upper envelopes using interpolation between
the maxima and minima, respectively; 2) the mean
value of the resulting envelopes is calculated for each
signal point and extracted from the signal. The steps are
repeated iteratively until the processed signal satisfy
certain IMF conditions: a) the numbers of extrema and
zero crossings are either equal or differ at most by one,
b) at any point, the mean value of the envelope defined
IFMBE Proc. 2005;9: 301

Biomedical signal processing
by the local maxima and the envelope defined by local
minima is zero. Then, the extracted IMF is subtracted
from the input signal and the residual is used as the
input for the next IMF extraction process.
A time-frequency spectrum is formed by merging
instantaneous frequencies and Hilbert amplitudes of 4
IMFs into 20Hz bins. Instantaneous frequencies of the
modes are calculated as the first derivatives of the
Hilbert transform phase and are smoothed in order to
obtain general trends of the signal time-frequency
pattern. The time-frequency spectrum is displayed as a
contour plot of isolines representing 0.1 from the
maximum amplitude.
The ensemble correlation function is a function,
which weights each signal sample of the averaged
signal in relation to the correlation across the ensemble
of signal realizations [2]. Signal parts weighted by EC
values lower than 0.6 were considered as noise and
excluded from the time-frequency spectra.
Results
The method was applied to a number of TEOAE
signals from subjects with audiograms having hearing
thresholds both better and worse than 25dB HL in
500Hz-6kHz frequency range. In the most cases, the
proposed method provided time-frequency spectra
with excellent correspondence between the audiogram
and the range of TEOAE presence. A representative
example is presented in Fig. 2. The subject had the
following audiogram: 5dB HL at 500Hz, 10dB HL at
1kHz, 30dB HL at 2kHz, 15dB HL at 4kHz and 15 dB HL
at 6kHz. Time-frequency spectrum of TEOAE signal
recorded from the same subject clearly shows lack of
signal components in the frequency range between
1,9–2,3kHz.
Discussion
The results show that the presented method provides
high frequency resolution of the investigated signals and
resulting time-frequency maps correlated very well with
corresponding audiograms. The results could be even
better if additional signal quality criteria would be used
to select good quality signals or retrocochlear hearing
losses would be excluded. Achieving 100% detection of
hearing thresholds of 30dB HL is impossible also from
physiological point of view since several studies
showed that TEOAE may or may not be generated when
hearing level is in between 20 to 30dB HL [1, 4].
The presented results indicate a close relationship
between TEOAE signal time-frequency properties and
cochlear mechanics. The method may thus be
successfully used for estimation of TEOAE signal
properties, prediction of hearing thresholds worse than
20-30dB HL at a specific frequency, building and tuning
of the cochlear models, etc.
Conclusions
The presented method makes it possible to
investigate important OAE physiological features with
high accuracy and resolution. The results show that a
high correlation exists between TEOAE signal timefrequency
properties and hearing levels.
References
[1] PRIEVE B.A. (1996): ‘Click and Tone-Burst-
Evoked Otoacoustic Emissions in Normal and
Hearing Impaired Ears’, J. Acousl. Soc. Am., 99 (5),
pp. 3077-3086
[2] JANUŠAUSKAS A., SÖRNMO L., SVENSSON
O., ENGDAHL B. (2002): ‘Detection of Transient
Otoacoustic Emissions and Design of Time
Windows’, IEEE Trans. Biomed. Eng., 49 (2), pp.
132-139
[3] TOGNOLA G., GRANDORI F., RAVAZZANI P.
(1998): ‘Wavelet Analysis of Click-Evoked
Otoacoustic Emissions’, IEEE. Trans. Biomed. Eng.,
45 (6), pp. 686-697
Figure 2: TEOAE time-frequency spectrum.
The number of 317 TEOAE time-frequency
distributions from subjects with audiograms having
thresholds both better and worse than 25dB HL in
500Hz–6kHz frequency range were visually
investigated. Full correspondence between OAE
presence and all audiogram thresholds better than
30dB HL was found in 246 cases (77,6%). In the other
64 cases, the proposed method failed at only one
frequency.
[4] JANUŠAUSKAS A., MAROZAS V., SÖRNMO L.,
SVENSSON O., HOFFMAN H.J., ENGDAHL B.
(2002): ‘Otoacoustic Emissions and Improved
Pass/fail Separation Using Wavelet Analysis and
Time Windowing’, Med. Biol. Eng. Comput., 39, pp.
134-139
[5] HUANG N., SHEN Z., LONG S., SHIH H.,
ZHENG Q., YEN N., TUNG C., LIU H. (1998):
‘The Empirical Mode Decomposition and Hilbert
Spectrum for Nonlinear and Non-stationary Time
Series Analysis. Proc. R. Soc. London., 454, pp.
903-995
IFMBE Proc. 2005;9: 302

Biomedical signal processing
WAVELET-BASED FEATURE EXTRACTION
FROM PREHENSILE EMG SIGNALS
I. R. Carreño* and M.I. Vuskovic**
*Universidad Pública de Navarra, Pamplona, Spain
**San Diego State University, San Diego, California, USA
irodriguez@unavarra.es
Abstract: A feature extraction method based on three
moments applied to three wavelet transform sequences
has been used for classification of prehensile surface
EMG patterns. The method has performed significantly
better than the three window short-time Thompson
transform applied to the same signals recorded from a
real subject. The classifier used to evaluate the two
approaches was a Mahalanobis-distance based
ARTMAP classifier presented elsewhere.
Introduction
The electromyographic signal (EMG), measured at the
surface of the skin, provides valuable information about the
neuromuscular activity of a muscle and this has been essential
to its application in clinical diagnosis, and as a source
for controlling assistive devices, and schemes of functional
electrical stimulation [1],[2]. Its application to control prosthetic
limbs has recently presented a great challenge, due to
the complexity of the EMG signals.
An important requirement in this area is to accurately
classify different EMG patterns for controlling a prosthetic
device. For this reason, effective feature extraction is a
crucial step to improve the accuracy of pattern classification
and many signal representations have been suggested.
Various temporal and spectral approaches have been
applied to extract features from these signals [1],[5],[6]. A
comparison of some effective temporal and spectral approaches
is given in [7], where the authors have applied
moments to short time Fourier transform (STFT) [5], and
short time Thompson transform (STTT) [8]. The later
transform has shown the best performance in case of short
temporal sequences.
The wavelet transform-based feature extraction techniques
(WT) have also been successfully applied with
promising results in EMG pattern recognition [3], [4].
The main goal of this work is to compare the WT with
the best spectral approach, the STTT. The two approaches
are used in conjunction with the first three moments that
were applied to the transformed sequences in order to further
reduce the dimensionality of the feature space, which
has proven to be very advantageous in the classification
stage. The evaluation of the two approaches was carried out
with an identical classifier, the Mahalanobis-based
ARTMAP (Adaptive Resonant Theory-based algorithm for
supervised incremental learning and classification)[9],
which has shown an excellent performance in classification
of EMG patterns [10].
Materials and Methods
Four-channel raw surface EMG signals from a healthy
subject were recorded at 1000 Hz sampling frequency. The
recording was done while the subject has repeatedly performed
the first four grasp types from the Schlesinger classification:
cylindrical grasp, precision grasp, lateral grasp
and spherical grasp (see Fig. 1).
Fig. 1. Four grasp types.
There were 180 grasp recordings evenly distributed
across the four grasps types. Two different EMG sequence
lengths were used: 200 ms and 400 ms. The 200 ms sequences
were obtained by truncating the recordings of 400
ms sequences. The shorter sequences are crucially important
in control applications.
Two features extraction methods were applied to these
signals: STTT and discrete WT.
The Thompson’s estimator of power spectral density
(multitaper method [8]) is known as the best estimator
which can simultaneously reduce the spectral leakage and
the spectral variance in the case of very short time sequences.
The STTT was applied to three 30% overlapping
windows. This approach was first suggested for feature
extraction from EMG signals in [6]. (The application of
STTT is detailed in [7]).
The DWT decomposes a signal into an approximation
signal and a detail signal. The approximation signal is subsequently
divided into new approximation and detail signals.
This process is carried out iteratively producing a set
of approximation signals at different detail levels (scales)
and a final gross approximation of the signal.
The detail D j and the approximation A j at level j can be
obtained by filtering the signal with an L-sample high pass
IFMBE Proc. 2005;9: 303

Biomedical signal processing
filter g and an L-sample low pass filter h. Both approximation
and detail signal are downsampled by a factor of 2.
This can be expressed as follows:
L−1 L−1
∑
D [ n] = g[ k] A [2 n− k], A[ n] = h[ k] A [2 n−k
],
∑
j j−1 j j−1
k= 0 k=
0
where A [ ] 0
n , n = 0,1,…N-1 is the original EMG sequence.
Sequences g[n] and h[n] are associated with wavelet function
ψ ( t)
and the scaling function ϕ ( t)
through inner
products:
gn [ ] = ψ(), t 2 ψ(2 t− n) , hn [ ] = ϕ(), t 2 ϕ(2 t−
n) .
Specifically we used the Daubechies wavelet of order
18 for applying 2-scale decomposition and obtaining the
wavelet coefficients (two details and one approximation).
The A and D sequences obtained as the result of DWT
are still massive in terms of the number of samples, which
contributes to large dimensionality of feature space. Besides,
the sequences have a high noise component inherited
from the original EMG signal. In order to reduce the dimensionality
and to smooth out the noise, we applied three
moments to transformed signals:
N −1
m
M = x S[ n], m = 0,1,2.
m
∑
n = 0
where x represents frequency in case of STTT and time in
case of DWT. Similarly S[n] represents either power spectral
density, or A and D sequences. Log transformation was
applied to the moments as it reduces skewness and kurtosis,
resulting in estimated probability density functions that
appear more like normal distributions. This nonlinear transformation
of features has improved the classification rate
significantly.
Results
The classification hit rate for one subject and two feature
extraction methods are shown in Tables 1 and 2. The
hit rates are averaged over 100 independent experiments,
with 50% random partition of samples into training and test
sets. The tables also show the average standard deviation of
hit rates for each sample length. As expected, shorter sequences
contributed to lower precision of classification and
to higher variation of results.
Table 1: Short-Time Thompson Transform (3 windows)
M 0 M 0 +M 1 M 0 +M 1 +M 2 STD
200 ms 88.37 % 82.04 % 79.40 % 4.33
400 ms 96.61 % 94.67 % 95.78 % 1.94
Table 2: Discrete Wavelet Transform (2D+1A)
M 0 M 0 +M 1 M 0 +M 1 +M 2 STD
200 ms 90.31 % 91.59 % 91.86 % 2.26
400 ms 98.43 % 98.56 % 98.72 % 0.95
The results suggest a clear advantage of DWT over the
STTT, which is known as the most effective spectral estimator
for short sequences. The standard deviations of the
hit rates are significantly lower at shorter sequences. In
addition, the DWT approach shows improvement of hit
rates for shorter sequences if all three moments are used;
which is consistent with the fact that more moments contribute
to more information.
Conclusions
The measurements performed in this study show the superiority
of the wavelet transformation over the traditional
spectral approach for feature extraction from prehensile
EMG signals. The method employed a simple DWT with
only three transform sequences, instead of the full wavelet
decomposition used in more complex wavelet packet transform.
This has eliminated the tedious feature reduction
procedures and PCA.
References
[1] HUDGINS B., PARKER P. and SCOTT R. N. (1993): ‘A
New Strategy for Multifunctional Myoelectric Control’,
IEEE Trans. on Biomed. Eng., vol. 40, No. 1, pp. 82-94.
[2] ENGLEHART K., HUDGINS B., PARKER P., STEVENSON M.
(1998): ‘Time-frequency representation for classification
of the transient myoelectric signal’, Proc. of the
20th Annual Int. Conf. on Eng. in Med. and Biol. Soc.
Chicago, Ill., p.2627 – 2630, vol.5.
[3] ENGLEHART K. (1998): ‘Signal Representation for Classification
of the Transient Myoelectric Signal’, Doctoral
Thesis. University of New Brunswick, Fredericton,
New Brunswick, Canada.
[4] ENGLEHART K., HUDGINS B., PARKER P. and
STEVENSON M. (1999): ‘Improving Myoelectric Signal
Classification using Wavelet Packets and Principle
Component Analysis’, Proc. of the 21st Annual Int.
Conf. of the IEEE on Eng. in Med. and Biol. Soc., Atlanta.
[5] HANNAFORD B. and LEHMAN S. (1986): ‘Short Time
Fourier Analysis of the Electromyogram: Fast Movements
and Constant Contraction’, IEEE Trans. On
Biomed. Eng., BME-33, pp. 2392-2397
[6] FARRY K. A., WALKER I. D., and BARANJUK R.G. (1996).
‘Myoelectric Teleoperation of a Complex Robotic
Hand’, IEEE Trans. On Rob. and Auto., 12, No.5.
[7] DU S. and VUSKOVIC M. (2004): ‘Temporal vs. Spectral
approach to Feature Extraction from Prehensile EMG
Signals’, The IEEE Int. Conf. on Information Reuse
and Integration (IEEE IRI-2004), Las Vegas, Nevada.
[8] THOMSON D. J. (1982): ‘Spectrum estimation and harmonic
analysis,’ Proc. of the IEEE, Vol. 70, pp. 1055-
1096.
[9] XU H. (2003), ‘Mahalanobis Distance-Based ARTMAP
Networks’, MS thesis. Dept. of Computer Science, San
Diego State University. San Diego, California.
[10] XU H. AND VUSKOVIC M. (2004): ‘Mahalanobis Distance-Based
ARTMAP Network,’ Int. Joint Conf. on
Neural Networks (IJCNN-2004), Budapest, Hungary,
July 25-29.
IFMBE Proc. 2005;9: 304

Biomedical signal processing
WHEEZE ANALYSIS AND DETECTION WITH
NON-LINEAR PHASE-SPACE EMBEDDING
C. Ahlstrom*, P. Hult* and P. Ask*
* Dept. of Biomedical Engineering, Linköping University, Linköping, Sweden
christer@imt.liu.se
Abstract: Wheezes are abnormal lung sounds
indicating airway obstruction, relevant for instance
in chronic obstructive pulmonary disease and
asthma. A wheeze detection method is presented,
based on concepts originally developed for nonlinear
time series analysis. A segment of the lung
sound signal is transformed into a trajectory in d-
dimensional phase space using Takens’ delay
embedding theorem. The trajectory is visualised in a
recurrence plot (RP), revealing structure in
apparently stochastic data. Feature vectors are
extracted from the RP and classified into wheezing
and normal segments. The method was tested on 11
patients with obstructive lung diseases where the
sensitivity and specificity was found to be 93% and
95%, respectively.
Introduction
Wheezes are additive, continuous adventitious lung
sounds probably caused by interaction between
fluttering airway walls and the gas moving through the
airways [1]. Wheezes are present in a wide variety of
diseases, and even if they can’t be used as a sole
discriminator, it is an interesting feature that provides
valuable information. Automatic detection offers better
objectivity and accuracy, especially in long-term
{ s s }
S = ,...,
(1)
1 , 2
s n
from which a sequence of N d-dimensional vectors a k
can be constructed using Takens’ delay embedding
theorem [5]. In this representation, the evolution of the
system can be described as a vector moving along some
trajectory in an abstract phase space (by reconstructing
the phase space of the signal, the geometric structure of
the multivariate dynamics is unfolded.). The vectors are
defined as
{ s s , s s }
a (2)
k
=
k
,
k + τ k + 2 τ
,...
k + ( d − 1)
τ
where τ is a delay parameter and d is the embedding
dimension. The techniques of average mutual
information and Cao’s method [6] were used to
determine τ and d, respectively.
Recurrence plots (RP) were introduced to
graphically display recurring patterns and nonstationarity
in time series. RPs can be used on rather
short time series and represent the recurrence of states
of a system (i.e. how often a small region in phase space
is visited). Unlike other methods such as Fourier,
Wigner-Ville or wavelets, recurrence is a simple
relation, which can be used for both linear and nonlinear
data [7]. An RP is a symmetric NxN matrix where a
point (i,j) represents the Euclidian distance between a i
and a j . Present tools for quantification of RPs operate on
binary data, so the matrix is calculated according to:
auscultation and analysis of the patient.
RP = Θ( ε − a − a )
Previous wheeze detection methods are mainly
based on various time frequency representations
combined with peak detection, thresholding or image
processing techniques [2]-[3]. Non-linear analyses of
healthy lung sounds suggest that lung sounds represents
a chaotic system [4]. Hence, in this study, a robust nonlinear
analysis approach is employed for automatic
location and identification of wheezes. The principles of
the method and the results are presented in this work.
Materials and Methods
Patients and data acquisition: 11 patients with
varying degrees of obstructive lung diseases were
enrolled in the study. The sounds were recorded using
Siemens EMT25C accelerometers, filtered via an 8 th
order Butterworth analogue filter (50-2500 Hz), and
amplified with a gain of 200. The signals were recorded
during resting conditions at the right posterior lower
lobe. Digitization occurred at 10240 Hz and 12-bits
whereupon the signals were digitally down-sampled to
5000 Hz. Time-dependence were obtained by analysing
segments of 256 samples.
Recurrence Quantification Analysis: A signal S can
be considered as a set of n scalar measurements
( i,
j)
i i j
(3)
where i,j=1,…,N, ε i is a cutoff distance, ║·║ is the
Euclidian norm and Θ(·) is the Heaviside function.
Different measures for RQA have been developed; they
are based on diagonal structures, vertical structures and
time statistics. Isolated recurrence points occur if states
are rare, if they do not persist for any time or if they
fluctuate heavily. Diagonal lines occur when a segment
of the trajectory runs parallel to another segment, i.e.
when the trajectory visits the same region of the phase
space at different times. Vertical (horizontal) lines mark
a time length in which a state does not change or
changes very slowly.
Classification: Ten feature vectors were extracted in
accordance with the RQA results. A simple discriminant
analysis was used to separate wheezing segments from
normal segments based on these features.
Periodicity histogram: Each pair of points is
separated in phase space by some distance r and in time
domain by some time ∆t. A space-time separation plot
visualises the relationship between these two distances,
and, when a fundamental frequency is present,
concentrates small r-values around time separation
values corresponding to the fundamental frequency [8].
IFMBE Proc. 2005;9: 305

Biomedical signal processing
Counting the number of distances r for each time
distance ∆t results in a periodicity histogram.
N
∑ − k
1
H ( k,
r)
= Θ( r − a i
− a i + k
) (4)
N − k i=
1
where k=0,…,(N-1) is a bin index, r is a chosen
neighbourhood radius and N is the number of phase
space vectors. The histogram is normalised since the
summation interval shrinks linearly with increasing k.
The fundamental frequency is calculated with a simple
peak detection algorithm operating on the periodicity
histogram.
Wheezing segment
Normal segment
calculated fundamental frequencies shows good
correspondence.
Wheezes have very prominent characteristics in
phase space, see Figure 1. The periodic structure of
wheezes is apparent as diagonal lines in the RP as well
as in the concentrated values at the fundamental
frequency in the space time separation plot.
Conclusions
This study shows that a nonlinear phase space
embedding reveal valuable information about wheezes
and that RQA provides good feature vectors.
Discriminant analysis is a simple, yet sufficient, method
for classifying respiratory sounds into normal and
wheezing segments.
Acknowledgements
The authors are much obliged to H. Pasterkamp, J.
Gnitecki (Resp. Acoustics Lab, Univ. of Manitoba,
Canada) and V. Gross (Dept. of Int. Med., Philips
University, Germany) for providing the data. This study
was supported by the Swedish Agency for Innovation
Systems (VINNOVA).
References
Figure 1: Left column illustrates a wheezing segment and right
column a normal segment. (a,b) shows the phase space with
τ=6 and d=3. (c,d) is the RP, (e,f) shows the space time
separation plot and (g,h) the normalised periodicity histogram.
All units are in samples (100 sample = 20 ms.).
Results and Discussion
The first minimum in the average mutual
information was located at τ=6 and was chosen as time
delay. Similarly the embedding dimension was set to
d=5 according to Cao’s method. Data from five
randomly chosen patients (39 wheezes) made up the
training set and the remaining six patients (35 wheezes)
constituted the test set. The classification resulted in a
sensitivity of 93% and a specificity of 95%. Visual
comparison between spectrogram analysis and the
[1] MESLIER N., CHARBONNEAU G. AND RACINEUX J.
L. (1995): ‘Wheezes’, Eur. Resp. J., 8, pp. 1942-8
[2] TAPLIDUO S. A., HADJILEONTIADIS L. J., KITSAS I.
K., PANOULAS K. I., PENZEL V., GROSS V. AND
PANAS S. M. (2004): ‘On Applying Continuous
Wavelet Transform in Wheeze Analysis’, Proc. 26th
Ann. Int. Conf. of the IEEE EMBS, San Francisco,
USA, 2004, pp. 3832-5
[3] WARIS M., HELISTO P., HALTSONEN S., SAARINEN A.
AND SOVIJARVI A. R. (1998): ’A new method for
automatic wheeze detection’, Tech. Health Care, 6,
pp. 33-40
[4] VENA A., CONTE A., PERCHIAZZI G., FEDERICI A.,
GIULIANI R. AND ZBILUT J. P. (2004): ‘Detection of
physiological singularities in respiratory dynamics
analyzed by recurrence quantification analysis of
tracheal sounds’, Chaos, Solitons & Fractals, 22, pp.
869-81
[5] MARCH T. K., CHAPMAN S. C. AND DENDY R. O.
(2005): ‘Recurrence Plot Statistics and the Effect of
Embedding’, Physica D, 200, pp. 171-84
[6] CAO L. (1997): ‘Practical Method for Determining
the Minimum Embedding Dimension of Scalar Time
Series’, Physica D, 110, pp. 43-50
[7] ZBILUT J. P., THOMASSON N. AND WEBBER C. L.
(2002): ‘Recurrence quantification analysis as a tool
for nonlinear exploration of nonstationary cardiac
signals’, Med. Eng. & Phys., 24, pp. 53-60
[8] TEREZ D. E., (2002): ‘Robust pitch determination
using nonlinear state-space embedding’, Int. Conf.
on Acoustics, Speech and Sign. Proc., Orlando,
USA, 2002, pp 345-8
IFMBE Proc. 2005;9: 306

Biomedical signal processing
Position approximation from acceleration data of an ischemic pig heart,
synchronized with the electrocardiograph signal
L. A. Fleischer*, L. Hoff*, O. J. Elle**, S. Halvorsen**, E. Fosse**
*Vestfold University College, Faculty of Science and Engineering, Horten, Norway
**Rikshospitalet University Hospital, The Interventional Centre, Oslo, Norway
Lars.A.Fleischer@hive.no
Abstract: The main focus in this article was to
measure the difference in heart motion associated
with coronary artery occlusion. We have been
monitoring the heart motion using a three-axis
accelerometer attached to the heart wall during open
thorax heart surgery. To interpret the heart
acceleration signal, each cardiac cycle was separated
using the Q and R wave of the electrocardiography
(ECG) signal. Each separated cycle of the
acceleration signal was then integrated to get an
approximation of the position. The integration was
done using the assumption that each point of a heart
during a cardiac cycle, starts and ends in the same
position. The result was then presented as an
average cardiac cycle.
Introduction
The work on this paper has been done as a part of
the Micro Heart project, which is collaboration between
the Interventional Center at Rikshospitalet University
Hospital in Oslo and the Institute for Microsystems
Technology at Vestfold University College in Horten.
This project is based upon an idea that originated at
Rikshospitalet. The idea is to use a 3-axis accelerometer
to monitor the heart motion [1]. The goal is to monitor
postoperative coronary artery bypass graft patients, to
detect regional myocardial ischemia or infarction.
This article presents data acquired using a hybrid 3-
axis accelerometer sutured on the apex of the left
ventricle, during open thorax heart surgery. The ECG
was also recorded in order to relate the acceleration
signal to the cardiac cycle. The sensor was placed in an
area of the heart where myocardial ischemia was
expected when the coronary artery was occluded.
This measurement setup has previously been shown
to be an adequate way to measure the heart wall motion
[2].
Methods
The ECG signal was used to locate the Q and R
wave. Since the ECG and the acceleration signal were
recorded synchronously, this method uses the R-wave to
separate each cardiac cycle. Since the heart rate of a pig
naturally fluctuates, each cardiac cycle will be of
different length. It was therefore necessary to adjust the
number of samples each cardiac cycle consists of. This
adjustment changes the sampling frequency, and it was
therefore necessary to calculate one based on the
average length of the cardiac cycles.
This method was based on the assumption that the
heart, in a cardiac cycle, starts and stops in the same
position. This assumption was used when integrating
from acceleration to a velocity and then to a position
approximation. This means that the average acceleration
was set to zero before integration, and the average
velocity was also set to zero before integrating to
position. This makes the velocity and position start and
stop with a zero. The integration method was by a
trapezoidal approximation. This resulted in several
cardiac cycles which were used to generate an average
one.
Figure 1: The average ECG cardiac cycle for the
normal period, early and late occlusion period.
Results
Data acquisition was done on the heart of a pig,
during open heart surgery. The sensor was sutured on
the heart wall in the apex of the left ventricle. The signal
from the acceleration sensor was recorded with a sample
frequency of 250Hz.
An average normal contition for the cardiac cycle
was generated using 319 cycles from a period of 4
minutes before the left anterior descending (LAD)
artery was occluded. The LAD was then occluded for 60
seconds, or 80 cardiac cycles. To get a better view of
how the ischemia developed during a 60 seconds period,
it was divided into two periods where the first one
contains 38 cycles and the second one contains 40. This
resulted in two average cardiac cycles for the occluded
period.
Figure 1 shows the average ECG signal for the
normal period and the first and last phase of the
occluded period.
IFMBE Proc. 2005;9: 307

Biomedical signal processing
From the figure we can see, that during the
occluded period, a reduction and almost an inversion of
the T wave occurring over time. This may be an
indication of ischemia [3].
The heart rate increased from 80bpm in the normal
period, to 82 and 83bpm in occlusion period one and
two respectively. This gives rise to a shortening of the
cardiac cycle when plotted against a time scale. Due to
this and since the graphs are synchronized on the R
wave peak, the latter part of the graphs in the Figures
are unaligned. This is visible in the P wave in Figure 1.
The result was a cardiac cycle with zero net movement.
If the sensor movements do not follow the assumption,
the result might lead to a big deviation, like the one in
Figure 3.
Conclusions
Because of the big deviation in the cardiac cycle it
is necessary to look into how to reduce this. It is
therefore necessary to look into how the respiration
affects the cardiac motion.
Since this article only looks at one axis in a three
axis system, future work will be done looking at all axes
together, to extract the total sensor movement during a
cardiac cycle. [4]
Figure 2: Average cardiac cycles of the
approximated position. The different graphs
represent normal period, early and late occlusion
period.
Figure 2 shows the position approximation of the
acceleration signal of one axis. Since the orientation of
the axis was unknown, the one with biggest change
during the occluded period was chosen. The R and T
waves have been included in the figure in order to relate
the graphs to the cardiac cycle.
Figure 3 shows the statistical variation of the
cardiac cycle position data in the normal period. It is
also plotted together with the occlusion period for
reference.
Discussion
In Figure 2, the latter occlusion period shows a very
different movement pattern compared to the normal
period. In the sections between 0 and 0.3 seconds the
movement pattern is similar except the distance among
the graphs is increasing. In the section between 0.4 and
0.5 seconds the distance is reduced by a negative
movement relative to the normal.
Figure 3 shows a large variation in the signal during
the normal period. This may be caused by the
respiration, since it was observed that it moves the heart
during open thorax heart surgery.
The assumption that each point of the heart during a
cardiac cycle starts and ends in the same position was
mentioned earlier. If the heart moves with the
respiration, the assumption mentioned earlier will lead
to this variation.
The assumption made the integral of the
acceleration and velocity zero. This was done by first
setting the mean acceleration and mean velocity to zero.
Figure 3: Standard deviation of the approximated
position during the normal period. The average
cardiac cycles for the occluded periods are also
plotted here for comparison.
References
[1]ELLE, O.J., HALVORSEN, S., GULBRANDSEN,
M.G., FOSSE, E. (2005): 'Early recognition of regional
cardiac ischemia using a 3-axis accelerometer sensor',
Physiol. Meas. 26. In press
[2]HOFF, L., ELLE, O.J., GRIMNES, M.J.,
HALVORSEN, S., ALKER, H.J., FOSSE, E. (2004)
'Measurements of Heart Motion using Accelerometers',
Proc. of 26th Ann. Internat. Conf., IEEE Eng. in Med.
and Biol. Society, San Fransisco, USA, 2004, p. 2049-
2051.
[3]KINNEY, R.M, PACKA, R.D., ANDREOLI, K.G.,
ZIPES, D.P., (1991):'Comprehensive Cardiac Care',
Mosby Year Book, seventh edition, 1991.
[4]GRIMNES, M., HOFF, L., HALVORSEN, S., ELLE,
O.J., ALKER, H.J., FOSSE, E. (2004): 'Velocity and
Position Approximations from Left Ventricular 3D
Accelerometer Data', Proc. of the IEEE 30th Ann.
Northeast Bioeng. Conf., Springfield, MA, USA 2004, p
25-26
IFMBE Proc. 2005;9: 308

Biomedical signal processing
BIOIMPEDANCE ANALYSIS OF CARDIAC FUNCTION USING
IN-VIVO EXPERIMENTS, MEASUREMENT AND SIMULATION
R. Gordon*, A. Haapalainen**, P. Kauppinen** and R. Land*
* Institute of Electronics, Tallinn Technical University, Tallinn, Estonia
** Ragnar Granit Institute, Tampere University of Technology, Tampere, Finland
rauno@elin.ttu.ee
Abstract: Invasive impedance measurement can
provide detailed information about work-load and
condition of the heart tissue. Non-invasive,
multichannel impedance measurement provides
information about cardiovascular dynamics in a
safe, easy to use manner. Both methods, while
promising, are still in development and share
common research needs. This paper describes how
international scientific work around these topics is
combined using in-vivo experiments, measurement
and simulation.
Introduction
12-lead impedance cardiography (ICG) is a method
for obtaining physical data about cardiovascular
function, without any invasive procedures. Previous
ICG methods have utilized only one or two channels,
giving unreliable results. Simulations and experiments
about impedance measurement sensitivity suggest that
multichannel approach should give more precise
information. However, more simulation and
experimental work is needed in order to provide solid
evidence about the reliability of the 12-lead ICG.
Invasive impedance cardiography can provide much
more detailed information about heart muscle condition
and cardiac workload. These parameters can be
monitored for example by cardiac pacemakers to
constantly assess and provide patient well-being.
Impedance measurements can be made between several
electrodes in different locations in the heart. Increase in
the intelligence of the electronics allows a less intrusive
device to extract more practical information. That goal
is achieved by simultaneous multichannel
multifrequency analysis.
Best locations for electrodes in experimental
measurement are found by analyzing results of
numerous electrode configurations through simulation.
On the other hand, precise measurement methods and
experiment set-up are needed to validate simulation.
Methods
A new virtual system for simultaneous multifrequency
measurement of electrical bioimpedance have
been proposed [1] for simultaneous measurement of
bioimpedance on four channels at several frequencies.
This virtual system was composed as a tool for
development of multichannel multifrequency
measurement devices that operate in invasive
experiment configurations. It utilizes eight generators at
frequencies from 100 Hz to 5 MHz that are multiplexed
to external electrodes within a biological sample.
Signals are detected with non-uniform synchronous
undersampling and analysis method. It allows
simultaneously to determine the complex impedance
“spectrum” at four independent pairs of electrodes
within the specimen or to simultaneously measure the
impedance of multiple regions. The system is very
useful for training personnel and planning the
experiments before the real biological samples become
available. This virtual measurement set-up can use
signal input from a real in-vivo measurement [2] in
progress or a simulation.
Simulation of ICG signals with invasive electrodes
has been conducted with Finite Difference Method on a
3D dynamic model of a heartbeat with 20 frames of
motion. Multi-frequency simulations are made possible
with implementation of frequency dependent models for
each represented tissue. Anatomical 3D model of the
heart has been obtained from segmented MRI slices. It
has been enhanced to remove patient specific
peculiarities and segmentation artefacts. Electrode
positions in the simulation are chosen to represent
possible pacing leads of a cardiac pacemaker.
Right
Ventricle
I excit
Left
Ventricle
Z [Ω]
Z [Ω]
Z [Ω]
Time
Figure 1. Cross-section of the first frame of 3D dynamic
model. Numbers 1-12 represent electrodes in the
simulation inside heart. Electrodes 1 and 4 are used for
excitation. Impedance variation is shown for electrodes
1, 2 and 3 in reference to electrode 4.
IFMBE Proc. 2005;9: 309

Biomedical signal processing
Methods for simulating and analyzing non-invasive
ICG measurement configurations with the application of
lead field theory and computer modelling have been
developed. The measurement properties of conventional
ICG were first studied with detailed torso models and
fundamental shortcomings in its methods emerged. ICG
was shown not to be specifically sensitive to detect
conductivity changes in the regions assumed to be the
source of ICG signal, and reported modified
measurement configurations are subject to additional
errors [3].
New configurations were derived based on the 12-
lead electrocardiography (ECG) electrode system,
which allows clinical utilization simultaneous with
routine ECG acquisition. Several thousand
configurations were analyzed with the lead field
approach and three different computer models of the
human thorax. As compared to results with
conventional ICG, clearly more selective measurements
of the structures of the cardiovascular system were
obtained. The simulation results suggest, that in order to
derive reliable information based on an ICG method, a
number of measurements should be taken with electrode
configurations possessing regional sensitivity [4].
Results
To verify simulation results of the non-invasive
electrode system, clinical experimentation with a
number of regionally sensitive configurations was
undertaken. A front-end unit for commercial bioimpedance
instrumentation was developed for this
purpose.
Several configurations were preliminarily tested on
healthy volunteers and patients undergoing valve
replacement. Parameters derived from the
measurements exhibited significant correlation with the
simulation data. Valvular disease, which disturbs
conventional ICG, was distinguished within the study
population and recorded 12-lead ICG signals evinced
waveforms and landmarks different from those of
conventional ICG. Certain configurations showing
resemblance to invasive data were noted. The
encouraging results suggest that reliable ICG
measurements producing more direct physiological data
can be obtained.
New approach for the multichannel ICG hardware
was invented in co-operation with the Tampere
University of Technology, Technical University of
Tallinn and Institute of Electronics and Computer
Science in Riga [5, 6]. New device for 12-lead ICG is
designed for multiplexed measurement from the
beginning. Its structure enables rapid channel switching
and measurement accuracy that is needed for relatively
complex signal analysis. With the co-operation of the
project partners, the new device is going to be in test
phase in summer 2005.
Discussion
Dynamic simulation of the whole thorax is needed
for non-invasive as well as invasive simulation in order
to predict the results obtained with measurements.
However, the complexity of the subject limits the
realization. As a start, dynamic model of the heart can
be placed into a torso model with the dynamics of the
most important vessels. Fluid dynamics of the blood are
included in some extension. With these operations, the
model should be able to give approximations of the
cardiac originated impedance data to non-invasive
electrodes. Inclusion of breathing dynamics should be
the next step in bringing the simulation closer to reality.
Developed instrumentation enables the collection of
interesting material, which can be used together with the
modelling data to improve the impedance measurement
method.
Conclusions
Work is currently in progress in all of these areas –
models of cardiac function are being perfected and
modeling methods refined, measurement-devices for
invasive and non-invasive experiments are being
finalized and will soon allow conducting experiments
using identical equipment by cooperation partners in
Finland, Estonia and USA.
References
[1] GORDON R., LAND R., MIN M., PARVE T., SALO R..
(2005): ‘A Virtual System for Simultaneous Multifrequency
Measurement of Electrical Bioimpedance’, Proc.
of the Joint Meeting of 5 th International Conference on
Bioelectromagnetism and 5 th International Symposium on
Noninvasive Functional Source Imaging within the Human
Brain and Heart, BEMNFSI’05, May 12-15, (to be
published)
[2] KINK A., RATSEP I., PARVE T.. (2003): ‘Impedance
Controlled Pacing Rate Limits in Cardiac Pacemakers -
Experimental Validation on Isolated Heart’, International
Journal of Bioelectromagnetism, Special Issue: Advances
in Electrocardiology, No 1, Vol. 5, 2003, pp.63.64.
http://www.ijbem.org
[3] KAUPPINEN PK., HYTTINEN JA., MALMIVUO JA..
(1999): ‘Sensitivity distributions of impedance
cardiography using band and spot electrodes analysed by a
three-dimensional computer model’, Annals of Biomedical
Engineering, 26, pp. 694-702
[4] KAUPPINEN PK, HYTTINEN JAK, KÖÖBI T,
MALMIVUO J. (1999): ‘Multiple lead recordings improve
accuracy of bio-impedance plethysmographic technique’
Medical Engineering & Physics, 21, pp. 371-375
[5] HAAPALAINEN, A., KAUPPINEN, P., HYTTINEN, J.,
MALMIVUO, J. (2004): ‘Instrumentation for 12-lead
ICG’, Proc. of the XII International Congress on Electrical
Bio-Impedance & V Electrical Impedance Tomography,
Vol 2, Gdansk, Poland, 2004, p. 379-381
[6] DASPTOOL project. http://dasptool.edi.lv
The work was supported by the EU 5 th FW project IST-2001-34552 DASPTOOL
and Grants of Estonian Science Foundation No 5892 and 5902.
IFMBE Proc. 2005;9: 310

Biomedical signal processing
CUT-OFF MECHANICAL INDEX FOR EFFECTIVE CONTRAST IMAGING
F. Conversano 2 , S. Casciaro 1,2 , R. Palmizio Errico 2 , E. Casciaro 1,2 , C. Demitri 2 , A. Distante 1,2
1
Institute of Clinical Physiology, National Council of Research, Lecce, Italy
2
Bioengineering Division, Euro Mediterranean Scientific Biomedical Institute, Brindisi, Italy
conversano@isbem.it
Abstract: Aim of this study was to evaluate contrast
enhanced ecographic signals arising from a new
hydrogel-based phantom at variable mechanical
index (MI) to define the cut-off MI for effective
imaging using low power technologies. Experiments
were performed using a phospholipidic microbubble
contrast agent (CA), whose signal intensity was
evaluated at variable ultrasound (US) emission
power (MI: 0.08 to 0.3) for 3 different CA dilutions
(1:30000, 1:40000, 1:80000). For each tested dilution,
our results suggested MI=0.2 as the optimal
compromise between a good contrast enhancement
achieved in fundamental B-mode and low acoustic
pressure employed, while for the harmonic modality
this compromise seems to be suitable only for the
lowest CA dilution, while for more diluted samples
the cut-off MI could probably be slightly higher.
Introduction
Acoustic pressure variations can change the nature
of contrast microbubble oscillation from linear to non
linear. This has been pushing towards the development
of contrast-specific harmonic modalities for the
detection of microbubbles within human body
structures. However, tissue itself can produce harmonics
that will be received by the transducer [1], with the
effect of reducing contrast between bubbles and tissue.
A strong effect of emission power on the onset of
tissue harmonic signals has been widely demonstrated
for different ultrasound (US)-based imaging modalities
[2-4]. Fortunately, a clear difference between harmonics
produced by tissues and those produced by bubbles was
also demonstrated [1]: tissue harmonics require a high
peak pressure, so they are only evident at high
mechanical index (MI).
Furthermore, safe use of microbubbles is in doubt
because, at sufficient acoustic pressures, the contrast
agent (CA) gas bodies, or bubbles derived from them,
can also provide nuclei for inertial cavitation [5].
The concern over potential cavitation-induced
bioeffects has led to the development of several indices
to describe the relative likelihood of cavitation, among
which MI, proportional to the peak of negative pressure,
is now recognized as the major determinant of the
response of microbubbles to US [1].
This work shows how to reduce the inertial
cavitation risk by limiting MI without loosing the
diagnostic benefits of a good contrast enhancement,
thanks to the latest technological innovations in the field
of signal processing and radiofrequency (RF) spectrum
analysis.
Materials and Methods
The phantom used in this study was a customdesigned
hydrogel-based phantom containing two 1-mm
diameter vessels. The sound propagation velocity
throughout the phantom matrix was similar to that one
in the human liver, while vessel walls were made of a
different hydrogel.
The flow circuit was fed by a 500-mL reservoir
filled with a saline-diluted suspension of an
experimental phospholipidic CA (supplied by Bracco
Research SA, Geneva, Switzerland), continuously
mixed by a magnetic stirrer and driven through the
circuit by a peristaltic pump (Peri-Star, WPI Inc., FL,
USA) providing a laminar, steady flow at 8 mL/min.
A linear array probe (LA 532, Esaote Spa, Florence,
Italy), positioned on the top of the phantom so that the
imaging plane resulted perpendicular to the vessels, was
used for insonification with 2.5-MHz US pulses. The
transducer was connected to a digital ecograph (Megas
GPX, Esaote Spa, Florence, Italy) externally linked to a
prototype for RF analysis (FEMMINA, developed by
Florence University), able to get the full raw signal of
the probe [6].
Three different CA dilutions (1:80000, 1:40000,
1:30000) were employed. We chose just these dilution
values because in a preliminary study we demonstrated
that the considered CA shows a strong linear
relationship (r=0.995) between signal intensity and CA
dilution in the range 1:80000-1:30000. The tested MI
values were 0.08, 0.1, 0.2, 0.3 and each of them was
employed with all the three dilutions.
Sequences of backscattered signals were digitally
stored as RF raw data and analyzed off-line by means of
Fortezza software (supplied by Florence University). A
square region of interest (ROI) was defined (side = 0.5
mm), covering exclusively the inner cross-sectional area
of one vessel.
Mean Fast Fourier Transform (FFT) curve was
calculated within the selected ROI for each acquired
data frame. Then fundamental and second harmonic
component values were extracted from the obtained
IFMBE Proc. 2005;9: 311

Biomedical signal processing
curves and averaged over the corresponding frame
sequence. Finally, these values were plotted versus MI.
Results
Figure 1 displays the measured curves of the
backscatter intensity of single FFT components plotted
versus MI, for each tested CA dilution.
For all dilutions, fundamental component increases
arising MI until 0.2, while, beyond this value, it remains
approximately constant to a sort of “plateau value”. The
second harmonic component, on the contrary, seems to
follow this trend only for the 1:30000 dilution case and
to be proportional to MI in the next two cases, with a
positive slope even in the last step of the curve.
a) Dilution = 1:30000
105
Average Backscatter Intensity (dB)
100
95
90
85
80
0,05 0,10 0,15 0,20 0,25 0,30
b) Dilution = 1:40000
105
Average Backscatter Intensity (dB)
100
95
90
85
Mechanical Index
fundamental
2nd harmonic
80
0,05 0,10 0,15 0,20 0,25 0,30
c) Dilution = 1:80000
105
Average Backscatter Intensity (dB)
100
95
90
85
Mechanical Index
fundamental
2nd harmonic
80
0,05 0,10 0,15 0,20 0,25 0,30
Mechanical Index
fundamental
2nd harmonic
Figure 1: Plot of average backscatter intensity of single
FFT components versus MI: a) CA dilution = 1:30000,
b) CA dilution = 1:40000, c) CA dilution = 1:80000.
Discussion
It is now widely accepted that US-induced
bioeffects, due to the presence of contrast microbubbles,
arise primarily via an inertial cavitation mechanism,
whose likelihood increases with acoustic pressure.
Our in vitro obtained results indicate the possibility
of an effective contrast imaging in fundamental B-mode
at low acoustic pressure (MI=0.2) and employing a
quite high dilution of the tested CA (1:80000).
A similar approach could also be suitable to
maximize the efficacy of harmonic B-mode imaging,
which requires the definition of a MI cut-off value not
so high to produce evident tissue harmonics but
sufficient to cause microbubble non linear oscillations.
We demonstrated that for the 1:30000 dilution case
such a cut-off value should be around MI=0.2, while for
more diluted suspensions it could be slightly higher.
Conclusion
Presented data suggest MI=0.2 as the optimal
compromise between good contrast enhancement
achieved in fundamental B-mode and low acoustic
pressure employed for every tested CA dilution.
On the other hand, for the harmonic modality a
similar indication is valid only for 1:30000-diluted CA,
while further studies are needed to determine the
corresponding cut-off MI for more diluted suspensions.
References
[1] BURNS P.N. (2002): ‘Instrumentation for Contrast
Echocardiography’, Echocardiography, 19, pp. 241-
258
[2] TIEMANN K., VELTMANN C., GHANEM A., LOHMAIER
S., BRUCE M., KUNTZ-HEHNER S., POHL C., EHLGEN
A., SCHLOSSER T., OMRAN H., BECHER H. (2001):
‘The Impact of Emission Power on the Destruction of
Echo Contrast Agents and on the Origin of Tissue
Harmonic Signals Using Power Pulse-Inversion
Imaging’, Ultrasound Med. Biol., 27, pp.1525-1533
[3] BECHER H., TIEMANN K., SCHLOSSER T., POHL C.,
NANDA N.C., AVERKIOU M.A., POWERS J., LUDERITZ
B. (1998): ‘Improvement of Endocardial Border
Delineation Using Tissue Harmonic Imaging’,
Echocardiography, 15, pp. 511-517
[4] THOMAS J.D., RUBIN D.N. (1998): ‘Tissue Harmonic
Imaging: Why Does It Work?’, J. Am. Soc.
Echocardiogr., 11, pp. 803-808
[5] CHEN W.S., BRAYMAN A.A., MATULA T.J., CRUM
L.A. (2003): ‘Inertial Cavitation Dose and Hemolysis
Produced In Vitro with or without Optison’,
Ultrasound Med. Biol., 29, pp. 725–737
[6] SCABIA M., BIAGI E., MASOTTI L. (2002): ‘Hardware
and Software Platform for Real-Time Processing and
Visualization of Echographic Radiofrequency
Signals’, IEEE Trans. UFFC, 49, pp. 1444-1452
IFMBE Proc. 2005;9: 312

Biomedical signal processing
COMPUTER AIDED FETAL MONITORING SYSTEM
USING NONINVASIVE ELECTROCARDIOGRAM
A. Matonia, T. Kupka, K. Horoba, J. Jezewski, P. Labaj
Institute of Medical Technology and Equipment, Department of Biomedical Informatics
118 Roosevelt Str., Zabrze, Poland
adamm@itam.zabrze.pl
Abstract: An alternative approach to conventional
cardiotocographic fetal monitoring is presented. It
relies upon analysis of bioelectrical signals recorded
from maternal abdominal wall. Due to strong interferences
present in abdominal signal advanced
methods of signal processing had to be developed to
extract fetal electrocardiogram and uterine electrical
activity signal. Presented monitoring system,
apart from classical analysis, enables analysis of
fetal electrocardiogram morphology and electrophysiological
properties of uterus. These features
could be useful for very early recognition of fetal
distress.
Methods
We have developed the system for acquisition and
analysis of bioelectrical signals recorded on maternal
abdomen (Fig. 1). The basic merits of the used recording
module are very low level of its own noise
measured with reference to input: < 0.5 µV (peak-topeak)
and large value of CMRR coefficient: 120 dB.
Introduction
Biophysical monitoring of fetus and mother during
pregnancy and labour is based on information on fetal
heart rate and uterine contraction activity. So far, it is
accomplished by mechanical approach. Applying the
Doppler principle of ultrasound waves aimed at the
fetal heart enables monitoring of the mechanical activity
of the fetal heart. Determination of instantaneous
fetal heart rate (FHR) relies on the detection of heart
beats based on the analysis of ultrasound beam reflected
from the moving valves or walls. Uterine contraction
activity (UC) is recorded with a help of tocodynamometric
transducer attached to maternal abdomen.
In the common opinion mechanical methods are
very subjective and do not provide credible information
on functioning of the organs being examined.
These disadvantages result from the fact that as the indirect
measuring techniques, they records the effects of
electric excitation i.e. fetal heart movements and mechanical
contractions of uterus. Considerably higher
accuracy and reliability can be obtained by recording
of the primary bioelectric signals [1]: fetal electrocardiogram
– FECG and electrohysterogram – EHG as an
electrical activity of uterine muscle. Consecutive cardiac
cycles can be determined more accurate by detection
of QRS complexes in FECG than by analysis of
reflected ultrasound beam of a very complex shape.
From the other hand, electrohysterography seems to be
more sensitive than mechanical tocography. Due to
limited accuracy and sensitivity the tocography is used
mainly to control the labour progress. The electrical
approach can provide more comprehensive information
for early recognition of disorder of central nervous
system or for detection of premature labour.
Figure 1: The system for bioelectrical signals analysis
IFMBE Proc. 2005;9: 313

Biomedical signal processing
The main aim of the system was to ensure the same
information as provided by the mechanical approach:
fetal heart rate and uterine activity signals. Analysis of
FECG comprises the initial filtration of low-frequency
interferences, suppression of interfering maternal electrocardiogram
(during this process the signal of maternal
heart rate is also determined), detection of the fetal
QRS complexes and thus calculation of the FHR values
expressed in beats per minute:
FHR i [bpm] = 60000 / T RRi [ms]
where: T RRi is the consecutive beat-to-beat interval determined
from FECG signal.
Maternal electrocardiogram is a very strong interfering
signal. Its amplitude of about 200 µV is much
higher than fetal electrocardiogram amplitude, which
reaches level of only 10 µV. The frequency range of
MECG overlaps the FECG range which makes the
suppression of MECG using the simple filtration impossible
[2]. The method of MECG suppression developed
by us is based on precise subtraction of averaged
maternal QRS complex in each abdominal channel.
Reference maternal QRS complexes are created and
then after scaling they are subtracted from successive
maternal QRS complexes in each abdominal signal.
The phenomenon of inaccurate pattern synchronization
can be eliminated by subtraction of the first derivative
of maternal QRS complex pattern.
Results
The system for acquisition and analysis of abdominal
signals provides the FHR signal and uterine activity
information. They are analysed in conventional
way including detection of acceleration/deceleartion
patterns and uterine contractions (Fig. 2).
and stepped with 1 min. In every step, samples are ordered
from the lowest to the highest value and the mean
value from 10 % of samples from a lower side is calculated.
The threshold value for contractions detection is
obtained by adding to basal tone the value equal to 25 %
of signal range in the analysed window. Contraction is
detected if its duration exceeds 30 s and its amplitude is
higher then double value of the threshold.
Measurement of electrical activity of fetal heart enables
to carry out assessment of the morphology of the
fetal QRS complexes, comprising measurement of an
amplitude and time dependences between individual
waves, mainly analysis of the ST segment, determination
of the ratio of T wave amplitude to the QRS complex
amplitude (Fig. 3) and correlation of the PR segment
with the value of the FHR signal.
Figure 3: Fetal QRS morphology analysis
Conclusions
The measurement of fetal heart rate, maternal heart
rate and uterine activity signals and their classical automated
analysis can be accomplished by the recording of
maternal abdominal signals. In addition, analysis of fetal
electrocardiogram signal morphology and electrophysiological
properties of uterus can be carried out.
Estimation of timing parameters of ST segments and
evaluation of T/QRS ratio changes can enable very early
detection of fetal distress.
Acknowledgment. Scientific work financed from the
State Committee for Scientific Research resources in
2004÷2006 years as a research project No.3T11E01726.
References
Figure 2: The window with fetal and maternal traces
Electrical uterine activity component is separated
from abdominal signal in selected lead, by means of
band-pass (0.1 to 3.5 Hz) filtration. The slow wave is
extracted from EHG signal by a calculation of consecutive
root-mean-square values in 60 s window
stepped with 3 s. The basal tone is a reference for determination
of contraction timing parameters. Its consecutive
values are determined in windows 4 min wide
[1] PIERI J.F., CROWE J.A., HAYES-GILL B.R., SPENCER
C.J., BHOGAL K., JAMES D.K. (2001): ‘Compact
long-term recorder for the transabdominal foetal and
maternal electrocardiogram’, Med. Biol. Eng. Comput.,
39, pp. 118-125.
[2] JEZEWSKI J., MATONIA A., KUPKA T., GACEK A.,
HOROBA K. (2002): ‘Suppression of maternal ECG
interference in abdominal fetal electrocardiogram’,
Proc. of the 12th Nordic Baltic Conf. on Biomed.
Eng. and Med. Physics, Iceland, VI 2002, 162-163.
IFMBE Proc. 2005;9: 314

Biomedical signal processing
CHARACTERIZATION METHODOLOGY FOR SOUND ABSORPTION OF
SYNTHETIC MATERIALS IN ULTRASOUND IN VITRO STUDIES
S. Casciaro 1,2 , R. Palmizio Errico 2 , F. Conversano 2 , E. Casciaro 1,2 , L. Ostuni 1 , A. Distante 1,2
1
Institute of Clinical Physiology, National Council of Research, Lecce, Italy
2
Bioengineering Division, Euro Mediterranean Scientific Biomedical Institute, Brindisi, Italy
3
Innovation Engineering Department, Lecce University, Italy
casciaro@ifc.cnr.it
Abstract: The use of contrast agents (CA) in
diagnostic ultrasound (US) is increasing and the
need to develop in vitro setups that allow to fully
understand their acoustic properties and to minimize
chemical and physical effects due to environmental
conditions is evident. A method to evaluat three
synthetic materials (Polyurethane, Airex © , ethylene
vinyl acetate (EVA © )) used to reduce the acoustic
artefacts is presented. The study is performed
insonifying a custom-designed tissue-mimicking
phantom having the synthetic materials at the
bottom. Polyurethane showed the highest absorption
intensity and it can be used to minimize environment
effects on the vessel cavity where ultrasonic
microbubble contrast agent flows.
Introduction
In recent years, the knowledge of the behaviour of
microbubble contrast agent has improved by using in
vitro experiments. The need to develop an in vitro
experimental setup that reduces to the maximum
additional physical, chemical and acoustical effects of
the environment on the CA often results fundamental
[1-5].
Here we discuss about the design of a new tissuemimicking
phantom developed and modulated in order
to be able to study the ultrasonic properties of CA in
almost all kind of human tissues and in different
vascular system conditions, minimising boundary
conditions interferences. In particular the purpose of this
study is to develop a method for evaluation of
backscatter properties of three synthetic materials
(Polyurethane, Airex © , ethylene vinyl acetate (EVA © ))
at a large frequency range, in order to establish what is
the best material is to be used to minimize the artefact
effects on the vessel cavity, i.c. the key Region Of
Interest (ROI).
Materials and Methods
The phantom was a custom-designed tissuemimicking
phantom, made of hydrogel with a sound
propagation velocity (1559,5 m/s) very close to the
human liver value (1560 m/s) [6]. It was 6 cm deep, 5
cm long and 8 cm wide. Three materials (Polyurethane,
Airex © and EVA © ) were positioned at the bottom of the
phantom.
An US transducer (LA 523, Esaote, Florence, Italy)
was positioned on the top of the phantom and connected
to a digital ecograph (Megas GPX, Esaote, Florence,
Italy) optically linked to a prototype for radiofrequency
(RF) analysis (FEMMINA, developed by Florence
University), able to get the raw signal of the probe
without hardware and software filtering of the ecograph
itself [7]. The received signals were digitized at 40
MHz.
The phantom was insonified at variuos transmit
frequencies (1.66, 2, 2.5, 3.3, 5, 7.5, 10, 13 MHz) and
the whole raw data were acquired in sequences and
stored in FEMMINA for further off-line analysis.
Quantitative off-line analyses were performed using
Fortezza algorithms developed by our group. Stored raw
data were used to reconstruct image data and to properly
choose the Region Of Interest (ROI). They were not
actually filtered, but only enveloped with the highest
possible cut frequency (20 MHz), in order to obtain the
absolute value of the signal.
Two ROIs were selected, respectively closer (ROI1)
and further (ROI2) from the material surface. ROI1 was
made up of 20 traces, 500 points high and 800 points far
from the material interface. ROI2 had the same
dimensions as ROI1 but its distance from the material
interface was 1300 points.
Figure 1: Schematic representation of the ROIs.
The intensity values of the defined ROIs were
calculated and averaged for acquired frames of each
sequence corresponding to a specific frequency. Values
were recorded in a Fortezza proprietary format file, then
converted in XLS format.
IFMBE Proc. 2005;9: 315

Biomedical signal processing
Results
Figure 2 displays the relationship between averaged
pixel intensity and transmit frequency by which
synthetic materials laid on the bottom of phantom were
insonified.
Figure 2: Plot of average pixel intensity versus
frequency in ROI1 and ROI2.
In both figures it is observed that each material has a
trend that is almost constant in studied frequency range.
The lowest intensity was showed by polyurethane.
EVA © .
Discussion
The described phantom is potentially able to study
ultrasound contrast agents reproducing almost all kind
of human tissues, and can easily modulate the vessel
sizes and different spatial and geometrical
configurations. Here we presented a method to
determine which was the best material and to establish
the optimal localization of vessel cavity to minimize the
acoustic effects on contrast agents flowing in it. Airex © ,
EVA © and Polyurethane have a different absorption
grade that is almost constant varying the frequency. The
best sound absorption property is showed by
Polyurethane in both ROIs.
Looking at figure 2 it is possible to note how for
Airex © and EVA © , the average pixel intensities increase
at bigger distance from the material surface itselfs
(ROI2). This is probably due to the lower resonance
frequency indicated by backscatter intensities at low
frequencies for these materials rather than polyurethane
(Fig. 3).
Conclusions
In conclusion, we developed and presented an
acoustic characterization method. About these three
materials, the polyurethane showed the best behaviour
as sound-absorbant material when used to cover the
Plexiglas bottom of the phantom in ultrasonic
applications, as it shows the minimal average pixel
intensity of reflected signals for any distance. Its
application allows to minimize the artefacts and to
investigate the backscatter properties of different
contrast agents without additional aspects due to the
environmental boundary conditions interferences.
By using this method, further characterization of
synthetic materials commercially available is possible in
order to improve the performance of phantoms, making
it suitable for use in flowing systems that simulate
different human organs and allowing more accurate
investigations and consequent measurements.
References
[1] FRINKING PJA, DE JONG N. (1998): ‘Acoustic
modeling of shell-encapsulated gas bubbles’
Ultrasound Med Bio ,24, pp. 523-533
[2] WARD M, WU J, CHIU JF. (2000): ‘Experimental
study of the effects of Optison concentration on
sonoporation in vitro’, Ultrasound Med Biol, 26,
pp. 1169-1175
[3] PORTER TR, OBERDORFER J, RAFTER P, LOF J AND
XIE F. (2003): ‘Microbubble responses to a similar
mechanical index with different real-time perfusion
imaging techcniques’, Ultrasound Med. Biol., 29,
pp. 1187-92
[4] SBOROS V, MORAN CM, ANDERSON T, PYE SD,
MACLEOD IC, MILLAR AM AND MCDICKEN WN.
(2000a): ‘Evaluation of an experimental system for
the in vitro assessment of ultrasonic contrast
agents’, Ultrasound Med. Biol., 26, pp. 105-11
[5] SBOROS V, MORAN CM, ANDERSON T, GATZOULIS
L, CRITON A, AVERKIOU M, PYE SD AND MC
DICKEN WN (2001): ‘An in vitro system for the
study of ultrasound contrast agents using a
commercial imaging system’, Phys. Med. Biol., 46,
pp. 3301-3321
[6] BAZZOCCHI M. 2001: ‘Ecografia’, II edition.
Idelson Gnocchi Editor
[7] SCABIA M, BIAGI E AND MASOTTI L (2002):
‘Hardware and software platform for real-time
processing and visualization of echographic
radiofrequency signals’, IEEE Trans UFFC, 49, pp.
1444-1452
Figure 2: Backscatter Intensity (BI) versus frequency
for ROI1 and ROI2.
IFMBE Proc. 2005;9: 316

Biomedical signal processing
HEART SOUND CONFIDENCE INTERVALS ESTIMATION BASED ON
THE IEFE COEFFICIANTS
Data
M.U. Rusha*, S. Hussain*, Ahmed Babekir Normalization **
*Department of Microelectronics and Computer Engineering, UTM-Skudai, Johor, Malaysia.
**Khartoum Developed Clinic Centre-Sudan
Email: rushacom@hotmail.com , hussain@mail.fke.utm.my
Abstract: This paper discussed the 95% confidence
interval of the mean for a normal and abnormal
heart sounds. The heart sounds were recorded
within one minute to capture the heart sound
signals. The time domain signals were transformed
into the frequency domain using the instantaneous
energy and frequency estimation (IEFE) technique.
The confidence interval computation is based on the
IEFE coefficients. The results show that we can
confidently differentiate between the normal heart
sound and the abnormal heart sound. The abnormal
heart is based on mitral regurgitation. The
experiments were conducted over 20 normal
patients and 12 patients with the mitral
regurgitation problem.
Introduction
Stethoscope is the traditional and valuable
diagnostic tool which has been used since 1861 [1].
One can describe the heart sound as two followed
sounds have simply the sounds of (dupp and lupp).The
normal cardiac cycle consists of the synchronized
activity of the atria and the ventricles. The heart sound
produced is due to the contraction and the relaxation of
the atrial and the ventricles which is known as the
systole and diastole. For the normal heart at the systole
and the diastole time durations, the heart sound
components should not be exist; but for the abnormal
heart bits there is an extra sounds will appear in the
systole and diastole time duration; these sounds known
as heart murmurs. The murmurs energy is quite high
and observable. During the auscultation process many
other noises can present as murmurs. Also
misplacement of the stethoscope can lead to a wrong
auscultation which leads to the wrong decision. The
data were all taken by the cardiologist in a suitable
environment. The normality and abnormality also has
been judged by the cardiologist.
Confidence interval for a parameter is a range
believed to contain the parameter with a specific level
of probability (“confidence”). It based on point
estimation and the spread of the sampling distribution
(standard error). When the sampling distribution is
approximately normal, the confidence interval is simply
plus or minus a specific number of standard errors
around the point estimate [2].
Methodology
There are three main modules in the heart
diagnostic system which is shown in Figure 1.0. The
data acquisition, feature extraction as well as the
confidence interval estimation algorithm with 95%
confident.
Data
Auscultation
IEFE Feature
Extraction
Data Acquisition
Confidence
interval estimation
Results
Figure 1.0: Show the system design
Heart sound samples are collected from 32 patients
(20 normal patients and 12 abnormal patients). Mitral
regurgitation has been taken as a sample of the heart
abnormality because it’s well known and most spread
over the patients.Mitral regurgitation is a progressive,
long-term disorder in which the mitral valve, which
separates the left upper chamber of the heart (atrium)
from the left lower chamber (ventricle), does not close
properly. This causes blood to leak (backflow or
regurgitation) into the left atrium from the left ventricle
during contraction of the heart (systole).
Data acquisition was achieved in Khartoum
developed clinic centre. Electronic stethoscope was
used to record the heart sound. The sampling rate is
2000Hz with stereo channel and one minute recording
time.
Feature Extraction
The heart sound features will be extracted as
Instantaneous Energy and Frequency Estimation (IEFE)
coefficients by taking the instantaneous energy and
instantaneous frequency of the heart sounds. Studies
have shown that different heart disease produces
different energy and frequency. Thus, it is appropriate
to take these unique features to represent every possible
occurrence of heart problem [3].
IFMBE Proc. 2005;9: 317

Biomedical signal processing
is:
Instantaneous Energy (IE), Ez of heart sounds
Ez z( n)
z *( n)
= c(
n)*
c(
n)
= (1)
Where z (n), c (n) is heart sound signal.
Instantaneous frequency (IF) of heart sound is given by:
fs
1 d
( n)
= [ φ(
n)
] (2)
2π
dn
In order to find a solution for the
differentiation of a discrete signal, one can use the
forward and backward finite differentiator (FFD) and
(BFD) so that he can get an unbiased and zero group
delay for linear FM signals. This we called as a central
finite difference, instantaneous frequency:
1
fs (n) = [ φ(n + 1) − φ(n − 1) ] (3)
4π
The IEFE applied to the transformed data to
represent the important feature for each cardiac cycle.
A 95% mean confidence interval has been taken to the
IEFE coefficients to find the existence probability of
the popular mean for such a cycle.
Confidence interval
The parameter of concern here is the mean. In
order to estimate the 95% confidence interval we first
estimate the mean [5]. Then we get the variance in
order to get the standard deviation
The standard error of y :
σ
σ y = (4)
n
Then 95% the confidence interval for y :
Results and Discussion
y = ±1.96σy
(5)
Table 1.0 shows the 95% confidence interval for
the Normal and the Mitral regurgitation heart sound.
The difference is obvious between the two cases as
shown in Table 1.0. The best result causes from IF
where the Lower Confidence Limit (LCL) for the
normal heart is 17.954 and the Upper Confidence Limit
(UCL) is 19.529 compared to the mitral regurgitation
(MR), the Lower Confidence Limit (LCL) is 14.683
and the Upper Confidence Limit (UCL) is
16.1753.However we can see that the confidence
interval overlaps for the raw data.
Conclusion
Instantaneous Frequency (IF) can separate between
the normal and the Mitral regurgitation patients by
using 95% confidence interval. This shows that the
heart sound can be represented and interpreted
statistically without the traditional ways of the
recognition.
HEART
UCL
LCL
SITUATION
NORMAL
IEFE 0.289966 0.233376
IF 19.529758 17.954884
IE 0.347615 0.308288
RAW DATA -0.038082 -0.04481
ABNORMAL(MITRAL RIGURGITATION)
IEFE 0.320162 0.275163
IF 16.175343 14.683075
IE 0.402304 0.349271
RAW DATA -0.039344 -0.041838
Table 1.0: Normal and Abnormal 95% confidence
interval.
References
[1] ARA G.TILKIAN, (1993). University of California.
MARY BOUDREAU CONOVER (1993).
“Understanding Heart Sounds and Murmurs with
an introduction to lung sounds.” Santa Cruz
California. In Saunders Company 92-49309.
[2] ROBERT JOHNSON (1992).” Elementary
Statistics” PWS-KENT Publishing Company
053492980X.
[3] BOUALEM BOASHASH, Seniour member of
IEEE1992. “Estimating and Interpreting the
Instantanious Frequency of a Signal-Algorithms
and Applications 1&2” Proceeding of the IEEE,
Vol 80,NO. 4 , pp 540-568.
[4] ALAN V.OPPENHEIM, Massachusetts Institute
of Technology.(1999).RONALDW.SCHAFER,
Georgia Institute of Technology(1999).JOHN
R.BUCK,University of Massachusetts
Dartmouth(1999).“Discrete-Time Signal
Processing”. Prentice Hall International,
Inc.0130834432.
[5] SIMON JACKMAN. “Statistical Inference
Confidence Intervals for Point Estimates for
Mean” Stanford University, political science 151B.
IFMBE Proc. 2005;9: 318

Medical physics
CLINICAL MR SPECTROSCOPY, PAST, PRESENT AND FUTURE - A
REVIEW
J. Hauksson 1
1 Radiation Physics Laboratory, Northern Sweden University Hospital, Umeå, SWEDEN
Abstract
Magnetic resonance imaging (MRI) has proven to be an
indispensable tool for both researcher and clinician alike.
However, MRI is often nonspecific in defining the
underlying pathology of disease. This is not surprising,
since MR images are based on the distribution of water
and/or fat in the anatomies which are imaged. Although
differences in relaxation times of water molecules with
different mobility are the most common underlying
physical mechanisms behind image contrast in MRI,
attempts to find direct relationships between relaxation
times and disease have been unsuccessful. By
comparison a high degree of diagnostic specificity can be
achieved with magnetic resonance spectroscopy (MRS)
because it can detect the biochemical changes that
accompany specific diseases. MRS is the medical name
for nuclear magnetic resonance spectroscopy (NMR)
applied to medicine. MRS applies the rigorous and
quantitative approaches of NMR spectroscopy to the
characterization, diagnosis and monitoring of disease.
This review lecture will describe the physical principles
of MRS as well as some of the obtainable biochemical
information. Some of the technical challenges will be
described, and an attempt will be made to give an
overview of where we are today, as well as what can be
expected in the near future. Since MRS has established
itself as a clinically useful tool primarily in brain,
prostate and the female breast, the examples given in this
talk will be primarily from these parts of the human body.
jon.hauksson@vll.se
IFMBE Proc. 2005;9: 319

Medical physics
IMPLEMENTING PATIENT SELF-EVALUATED RECTAL PROBLEMS IN
NTCP MODEL FOR LOCALIZED PROSTATE CANCER PATIENTS TREATED
WITH EXTERNAL BEAM RADIOTHERAPY
S. Åström 1
1 Radiation Sciences, Radiation Physics, Umeå, Sweden
sofia.astrom@vll.se
Abstract
The purpose of this project was to investigate whether
clinical data could be applied to an empirical model for
normal tissue complication probability (NTCP).
Introduction
Rectum complications are often the dose limiting factor
for treating prostate cancer patients with external beam
radiotherapy. Different models of the predicted biological
response (NTCP) together with clinical experience could
be an aid in comparing different dose plans. To test the
validity of the NTCP models used today there is a need to
test them against clinical data.
Methods
One hundred and one prostate cancer patients treated with
external beam radiotherapy (78 Gy) at Umeå University
Hospital between 1999 and 2002 were included in the
analysis. 35 of the patients had received treatment also to
the seminal vesicles. The patients had evaluated their
intestinal problems in a self-assessed questionnaire at the
start of treatment, at treatment end and 1 year after. This
study only analyzes the symptoms evaluated at the
beginning of treatment and 1 year after. The Lyman-
Kuther-Burman (LKB) NTCP model was used to
calculate the NTCP values with the aid of a PC-program
called Bioplan. Cumulative DVH:s were used as input
and parameters for milder complications was used. NTCP
values were also calculated for a couple of nonstandard
parameter sets.
The volume parameter n in the LKB-model is a measure
of the architecture of the tissue. The two extremes are
parallel and serial structure. A low n-value suggests the
tissue is more serial, whereas a higher n-value suggests
the tissue is more parallel.
vesicles had received about 20 % larger volume of
intermediate dose.
The NTCP values with standard parameters
overestimated the reported patient problems for the
groups with no or minor problems. The same
phenomenon occurred for the NTCP values with
nonstandard parameters. The results were better for the
group with more severe problems, but due to the small
number of patients with severe problems this result is
uncertain.
For more serious complications and softer endpoints the
rectum structure seems to be more serial. However for
milder complications the structure seems to be more
parallel. Due to the small number of patients with severe
problems and the many factors of uncertainty these
results are statistically insignificant.
Discussion
Most of the patients had no or minor problems. There
was no difference in the total irradiated volume between
the patients treated to or not treated to the seminal
vesicles. The patients with treatment to the seminal
IFMBE Proc. 2005;9: 320

Medical physics
QUALITY FACTOR VS. FIVE-POINT SCALE IN EVALUATION OF
CLINICALY ACCEPTABLE IMAGE QUALITY IN CHEST CT
O. Sveljo*, B. Reljin**, Z. Markovic***, M. Lucic*, O. Adjic*, M. Prvulovic*
*Institute of Oncology, Sremska Kamenica, Serbia and Montenegro
** Faculty of Electrical Engineering, Belgrade, Serbia and Montenegro
***Clinical Center of Serbia, Belgrade, Serbia and Montenegro
sveljo.olivera@onko.onk.ns.ac.yu
Abstract: Modern CT scanners offer different
modes of operations and a wide choice in exposure
parameters. Evaluation of various protocols in
visualizations of anatomical structures and overall
image quality should ensure optimal relation
between image quality and dose for individual
patient. There are different methods for assessment
of clinically acceptable image quality. In this paper
we compared some of this approaches.
Introduction
In computed tomography image quality has many
components and is influenced by many technical
parameters [1]. While image quality has always been a
concern for the engineering and physics community,
clinically acceptable image quality has become even
more of an issue as strategies to reduce radiation dose –
especially to pediatric patients – become a larger focus.
Clinically acceptable image quality has been usually
established according to independent assessment of
two or more radiologist. Some investigators [2] use
quality factor (QF) as criteria for comparing different
protocols and the other [3] use five-point scale in
assessment of clinically acceptable protocols. In this
study we compared four different CT protocols using
both methods.
Materials and Methods
The study includes 44 patients subdivided into two
groups according to their body weights. The first group
with 24 patients weighted more than 70 kg were
examined according to the standard protocol: with the
nominal tube current of 159 mAs and edge enhancing
filter (EEF). In the second group 20 patients weighted
less than 70 kg were examined with reduced tube
current of 146 mAs. After scanning, images were
reconstructed with smoothing filter (SF) Examination
protocols are shown in Table 1. All exams were printed
on film under the same conditions. Two experienced
radiologists, independently, rated general image quality
and anatomic details using the following marks: 1 –
unacceptable, 2 – substandard, 3 – acceptable, 4 –
above average, and 5 superior. The readers were
blinded to technical and clinical data. Anatomic details
criteria are taken from the European guidelines on
quality criteria for computed tomography [4]. (Table
Table 1: The acquisition and reconstruction protocol
parameters for both groups of examined patients
I group II group
kV
140 kV
Current 159 mA 146 mA
Thickness
8 mm
Pitch factor 1.5
Reconstruction filter edge enhancing
smoothing
2). Quality Factor QF has been estimated as percent of
criteria rated 3 and more. The two tailed p-values of

Medical physics
assessment for both readers Figure 1. There were no
statistically significant difference (p>0.05) between
protocols with different mAs and same reconstruction
filter. Assessment of the protocols with same mAs and
different reconstruction filter were statistically different
(p< 0.05) for both readers.
show clearly ranked protocols with statistical
verifications according to five-point scale. Although
protocol ranking according to QF has been similar as
protocol ranking according to five-point scale there
were no statistically significant differences in QF for
different protocols. In our study QF has been estimated
according to equally weighted criteria taken from [4]
for general chest exam. Relatively high value for QF
for all readers and all protocols are probably due to the
fact that all most in all exams first group of 4 criteria
are rated 3 and more. It can be concluded that all
considered criteria are not equally important for
assessment of clinically acceptable image quality e.g.
in estimation of QF anatomic details criteria taken from
European guidelines on quality criteria for computed
tomography should be considered with different
weighted factors.
Conclusions
Figure 1: Linear regression of the protocol assessment
for readers ML and OA according to five-point scale
Protocol assessment according to QF has been shown
similar protocol ranking Figure 2, but there were no
statistically significant differences in QF of any
compared protocols.
Five-point scale in assessment of clinically acceptable
image quality give only an information about overall
image quality for anatomical region of interest and
could be used in comparison of different CT protocols.
QF is dependent of visualization of important anatomic
details and gives more information about visualization
of important anatomic structures. Anatomical detail
criteria from European guidelines on quality criteria
for computed tomography could be used in estimation
of QF as criteria for comparision of different chest CT
protocols but those criteria are not of equal importance
for clinically acceptable image quality and should be
considered with different weighted factors in
estimation of QF.
Figure 2: Linear regression of the protocol assessment
for RS, ML and OA according to QF
Discussion
Objective method for assessment of clinically
acceptable image quality has been not established yet
[5]. Ranking of different CT protocols according to
overall image quality with five-point scale are
relatively simple method but does not provide a closer
look in visualization of particular anatomic details. On
the other hand QF estimation takes in consideration
visualization of important anatomic details. Our results
References
[1] KALENDER, W.A. (2000): 'Image Quality', in
KALENDER, W.A.: 'Computed Tomography',
(Publicis MCD Verlag, Munich), pp. 83-118
[2] JURIK AG, JESSEN KA, HANSEN J.
(1997):’Image quality and dose in computed
tomography’ Eur. Radiol. 7, pp.77-81
[3] RUBIN GD, LEUNG AN, ROBERTSON VJ,
STARK P. (1998): ‘Thoracic Spiral CT: Influence of
Subsecond Gantry Rotation in Image Quality’
Radiology 208, pp. 771-6
1. [4] BONGARTZ G, GOLDING SJ, JURIK A,
LEONARDI M, GELEIJENS J, JESSEN KA, et al.
(1997): ‘Quality criteria for computed tomography’
Working Document EUR 16262, (European
Commission, Luxembourg)
[5] COHNEN M, FISHER H, HAMACHER J, LINS
E, KOTER R AND MODDER U. (2000): ‘ CT of the
Head by Use of Reduced Current and Kilovoltage:
Relationship between Image Quality and Dose
Reduction’ AJNR 21, pp. 1654-60
IFMBE Proc. 2005;9: 322

Medical physics
CALCULATING DOSE OUTPUT AT OFF-AXIS POSITIONS IN PHOTON
BEAMS USING A LATERALLY BEAM QUALITY DEPENDENT PENCIL
KERNEL MODEL
J. Olofsson 1 , T. Nyholm 1 , A. Ahnesjö 1 , M. Karlsson 1
1 Inst. for Radiation Sciences, Umeå University, Umeå, Sweden
Abstract
We have generalized a photon pencil kernel dose
calculations model to include explicit modeling of offaxis
softening (OAS). This is achieved by varying the
kernel characteristics according to typical shifts in beam
quality depending on the lateral position. The only
measured input data needed for each individual photon
beam is the quality index TPR 20/10 .
Methods
Calculations were made, both including and excluding
the effect of OAS, and compared to measured results.
Comparisons were performed at 5, 10, and 20 cm depth
for four different photon beam qualities delivered by a
Varian Clinac 2300C/D and a Siemens Primus
accelerator. In total the dose was measured in 264 points
located up to 18 cm from the central axis in a number of
different fields of varying size and position. In all cases
the results were normalised to the reference geometry for
each beam quality, to reflect the absolute dose delivery
per monitor unit (MU).
jorgen.olofsson@vll.se
Results
Results that include the effect of OAS show significantly
smaller deviations from measurements as compared to
when it is excluded (improvements up to 5% were
observed). In general the deviations including OAS are
within ±1%, but some variations in the amplitude of the
effect of OAS among the investigated beams seem to
increase the deviations in some cases.
Conclusions
A pencil kernel model implemented to take lateral beam
quality variations into account has the potential to
significantly reduce systematic calculation errors at offaxis
positions as compared to a standard implementation
based solely on a central axis beam quality description.
IFMBE Proc. 2005;9: 323

Medical physics
THE USE OF CARCINOGENESIS RISK ESTIMATION MODELS FOR
RADIOTHERAPY
A. Daşu 1 , I. Toma-Daşu 1 , J. Olofsson 1 , M. Karlsson 1
1 Department of Radiation Sciences, Umeå University, Umeå, Sweden
Abstract
There is an increased interest in predicting the risk
for secondary cancers from radiotherapy in order to
be used as a complementary criterion for the selection
of plans in addition to the estimation of the possible
deterministic effects. These attempts must however
take into consideration the specific features of
radiation treatment (fractionation, non-uniform dose
distributions etc.). We have explored several possible
methods for estimating the risk of cancer following
radiotherapy in order to investigate the influences of
the fractionation and the non-uniformity of the dose
to the irradiated organ. The results suggested that
dose inhomogeneity must be taken into account
through the use of the dose volume histograms as it
plays an important role in predicting the risk for
secondary cancer. Comparisons of the results with
clinical data also indicated that a linear risk model
might not be appropriate and that the competition
between cell killing and the induction of DNA
mutations has to be taken into account for more
realistic risk estimations. Furthermore, more reliable
parameters could be obtained if this competition is
also included in analyses of epidemiological data from
radiotherapy applications.
alexandru.dasu@radfys.umu.se
IFMBE Proc. 2005;9: 324

Medical physics
CORRECTION FOR SCATTER AND SEPTAL PENETRATION IN 123I BRAIN
SPECT IMAGING - A MONTE CARLO STUDY
A. Larsson 1 , M. Ljungberg 2 , S. Jacobsson Mo 3 , K. Å Riklund 3 , L. Johansson 1
1 Department of Radiation Sciences, Radiation Physics, Umeå University, Umeå, Sweden
2 Department of Medical Radiation Physics, Clinical Sciences, Lund University, Lund, Sweden
3 Department of Radiation Sciences, Diagnostic Radiology, Umeå University, Umeå, Sweden
Abstract
Photons which scatter in the patient, scatter in the
collimator or penetrate the septa in the collimator and
become registered after scattering, deteriorate the contrast
in SPECT imaging. For many radionuclides the only
significant of the effects above is scatter in the patient,
but some radionuclides emit high energy photons for
which the collimator effects are important. One such
radionuclide is 123 I which for example is used for brain
SPECT as a tool for diagnostics of Parkinson’s disease.
Both scatter and septal penetration have a negative effect
on the accuracy of quantitative SPECT measurements.
The effects can however be corrected for, and in this
work two different correction techniques are evaluated,
one that operates on the 2D projection images, and one
that is incorporated in the iterative reconstruction process.
The method operating on the projection images is
transmission dependent convolution subtraction (TDCS)
and two versions of this method are evaluated. The
method incorporated in the iterative reconstruction uses
the effective source scatter estimation (ESSE) for
modeling scatter, and collimator detector response (CDR)
which includes geometric and penetration components.
The corrections are evaluated for 123 I brain SPECT for
two different types of studies, one study of the dopamine
transporters in the striatum and one study of the regional
cerebral blood flow. The images are produced using a
newly developed Monte Carlo code added to the program
SIMIND which can include interactions in the collimator.
anne.larsson@vll.se
IFMBE Proc. 2005;9: 325

Medical physics
ERROR ESTIMATION IN DOSE CALCULATION FOR VERIFICATION
PURPOSES
T. Nyholm 1
1 Dept. Radiation Sciences, Umeå University, Umeå, Sweden
Abstract
Introduction
Modern radiotherapy includes an increasing complexity
in the given fields. Fields can be shaped almost arbitrary
with the MLC, and the energy fluence map can be varied
with IMRT. From a verification perspective this
development implies an increased need for verifications,
and also that traditional verification methods becomes
more difficult or unfeasible. An independent calculation
tool (ICT) for MU verifications, can be an effective
method to catch errors, not only from the TPS, but also
introduced in for example the R&V system. When using
an ICT one will find the occasions with mismatch
between the TPS and the independent calculation. To be
able to investigate these deviations in a uniform way at a
department, there will be a need for specified levels
where different actions/investigations are meaningful. A
single, fix action level is not an optimal solution; the
action level should instead dynamically reflect the
uncertainty in the ICT calculation. Generally the
uncertainty of a dose calculation is lower in geometries
“close” to the reference geometry than in for example an
IMRT treatment with the calculation point located offaxis.
tufve.nyholm@radfys.umu.se
Methods
Using a large set of measured data we have developed an
empirical method for estimation of the uncertainty in
energy fluence per MU calculation performed with a
published model. Also a semi-analytical method for
estimation of uncertainty in the dose per energy fluence,
for a pencil dose deposition model, has been developed.
Both the calculations methods and the uncertainty
estimation methods have been tested versus a set of 300
measurements in irregular MLC fields at three depths.
Results
A good agreement was found between calculated and
measured doses. No deviation larger than 2% was
observed. Predicted uncertainty and observed deviations
was positively correlated.
IFMBE Proc. 2005;9: 326

Medical physics
INTRA-OPERATIVE MONITORING WITH NEEDLE ELECTRODES IN
CONJUNCTION WITH SURGERY,INCLUDING ELECTROSURGICAL UNIT
(ESU).A HAZARD ? !
P. Fällmar 1 , T. Winkler 1 , R. Flink 1
1 Clin.Neurophys, University hospital, Uppsala, Sweden
Abstract
Intra-operative Neurophysiological Monitoring has
become very important to avoid post-operative
neurological deficits.
This report will highlight the risk of induced energy from
the Electrosurgical Unit (ESU) in combination with
needle electrodes during surgery.
3 patient cases, where needle recordings led to small
lesions, will be presented.
When the needle and the cable are exposed to the radio
frequency field (RF), they will transfer induced energy.
The shape of the needle will result in a high current
density at the tip, inducing high temperature locally. The
current will also give rise to an electrochemical process
where metal ions are released. A lesion caused by the
heating effect will heal without any remaining defects,
however the metal-ion deposits will be visible for a long
time.
Therefore, the quality of the material in the needle
electrode is of great importance. By using needle
electrodes made from stainless steel there is an added
risk, which is due to the current alloy in the needle, that
will give a high potential risk for long-term effects, i.e.
allergic reactions. If needles made of platinum is used,
the same amount of energy is present, as well as identical
increases of temperature, but no metal-ions are emitted.
It is not possible to completely avoid induced energy into
the tissue of the patient when ESU is used, therefore it is
important to use surface electrodes in all cases where it is
possible
peo.fallmar@akademiska.se
IFMBE Proc. 2005;9: 327