Dosimetric Characteristics of Clinical Photon Beams

Dosimetric Characteristics of 

Clinical Photon Beams 

Jatinder R Palta PhD 

University of Florida 

Department of Radiation Oncology 

Gainesville Gainesville, Florida

Disclosures 

�Research �Research development grants from Philips 

Medical Systems, Elekta Oncology 

Systems Systems, and Sun Nuclear Associates Associates. 

�NIH research award. 

�B �Bankhead kh d CColey l research h award. 

d

Learning g Objectives j 

�Understanding dosimetric properties of 

clinical photon beams. 

�Understanding gpphysical y pparameters 

that 

affect dosimetric properties of clinical 

photon p beams. 

�Understand the need for accurate 

characterization of clinical photon beams 

in a treatment planning system.

Photon Beam Delivery Systems 

S band Linear Accelerators 

X band Linear Accelerators 

Medical Linear 

Accelerators: 

�� Accelerate electrons in ppulses lses 

to kinetic energies from 4 to 25 

MeV. 

�Use �Use non-conservative 

microwave RF fields in the 

frequency range from 103 MHz (L 

band) to 104 band) to 10 MHz (X band), with 

the vast majority running at 2856 

MHz (S band). 

� Some provide beams only in 

the low megavoltage range (4-6 

MV), while others provide both 

photons and electrons at various 

megavoltage l energies. i A typical i l 

modern high-energy linac can 

provide 2-3 photon energies.

Sources of radiation that determine dosimetric 

characteristics of clinical photon beams 

Indirect (headscatter) 

Flattening filter 

Monitor Chamber 

Collimator jaws 

Electron 

Contamination 

MLC 

Charged particle 

contamination dose 

Secondary 

electrons 

Source 

Direct 

Output radiation or 

Incident radiation 

Primary dose 

SScatter tt ddose 

�Direct Radiation (Focal 

Radiation) 

�Photon radiation generated 

at the target g that reaches 

patient without any 

intermediate interactions. 

�Indirect Radiation (Extra- 

focal Radiation): 

�Photon radiation with a 

history of interaction/scattering 

in the head of the treatment 

unit with the flattening filter filter, 

collimators, or other structures 

in the treatment head . 

�Contaminant 

electrons/positrons 

� secondary electrons and 

positrons released from 

interactions with either the 

treatment head or the air 

column . 

AAPM TG74 Report

Sources of Direct and Indirect 

Radiation 

� A Monte Carlo study (Chaney et al., Med. Phys. 21,1994) 

� Siemens MD2, 6MV 

Direct 

Indirect

Characterizing Dosimetric Properties 

of Clinical Photon Beams 

� Beam penetration 

�� Normalized depth dose (NDD) or tissue phantom ratio 

(TPR). 

� Beam Output 

� Total output ratio: S c,p, in-air output ratio: S c, phantom 

scatter factor: S p. 

�� Cross Cross-beam beam profile 

� Isodose distribution. 

�� Attenuation factors for beam modifiers 

� hard wedges, compensators, trays, etc. 

With the ultimate goal of ensuring that computerized treatment 

With the ultimate goal of ensuring that computerized treatment 

plans accurately reflect the dose received by patients

Beam Penetration 

D 

�� d 

d f Q� 

d 

d, 

s, 

f , Q � 

NDD 

f 

D 

dref 

where d is the depth of measurement on the 

central axis of the phantom, s is the field size at 

the surface of the phantom, f is the sourcesurface-distance, 

Q is the quality q y of the clinical 

photon beam, and Dd and Ddref are dose at 

depth d and dref respectively. 

TPR data can be determined from measured 

NDD as follows: 

Water 

s 

d 

�� 

2 

f � d � �� 

S �� s �� 

�� 

TPR�d, s � � � 

d , Q � NDD d, 

s, 

f , Q � � � 

f dref 

� 

� � � 

p 

S 

p 

dref 

�s� d

Normalized Depth Dose Data 

Energy Dependence 

SSurface f region i 

BBuildup ild region i 

15 MV 

6MV 6 MV 

FS = 10 x 10 cm 2 

TCPE region


Field Size Dependence 

This depth corresponds to range 15 MV Photon Beam 

of the highest energy 

contaminant charged particles 

4x4 

16x16


Wedge/Open Comparison 

6 MV (W/O) 

15 MV (W/O) 

FS=10x10cm2 

FS = 10 x 10 cm2


Wedge/Open Comparison 

Minima


Field sizes: 

6x6, 10x10 

and 20x20 

cm2 cm2 6 MV 

- Siemens 

-- Varian 

.Elekta . Elekta 

18 MV 

These data from 

Radiological 

Physics Center 

show that all NDD 

for both 6 and 18 

MV photon beams 

at depths of 5 cm 

and 15 cm for 

different field sizes 

have a 

maximum i % %σ of f 

0.5% and this 

increases to 0.7% 

at a depth of 20 cm cm.

Monte Carlo Calculated Photon 

Beam Spectra 

•The spectral shapes 

are somewhat similar 

•The differences at the 

high-energy end are 

caused by the 

differences in the 

mean incident electron 

energies and their 

spread 

Sheikh-Bagheri & Rogers, Med. Phys., 

29, 2002

Monte Carlo Calculated Average Energies 

•The average energies 

for the same nominal 

accelerating potential 

are somewhat similar 

•The average energies 

decrease at off-axis 

distances for all clinical 

beams 

• more pronounced difference 

at higher g energies g 

Sheikh-Bagheri & Rogers, Med. Phys., 

29, 2002

Beam Penetration for Irregularly- 

Shaped Fields 

Concept p of Equivalent q Square: q 

�The equivalent field is defined as 

that standard (square or circular) 

field which has the same centralaxis 

depth dose characteristics as 

the given non-standard field. 

“Day’s Rule”: 

S 

�� r 

r � 

S 1 

��r 

� 

��r 

� 

r � S 1� 

e 

� 

� � � � � r � e 

� 

� 

Water 

S(r) = the central axis scatter in a field of radius r, S∞ = the central axis scatter in 

fi field ld of f infinite i fi it radius, di λ iis a scaling li parameter, t and d μ iis a di dimensionless i l shape h 

parameter. They computed equivalent square fields for a complete set of 

rectangular fields using a value of λ=0.26 cm-1 and μ=0.5. 

s 

f 

d

Equivalent square d 

Sterling g Formula: 

(Sterling et.al., Brit. J. Radiol. 37, 544 (1964)) 

2LW 

S �� 

�� 

4 A / P 

L �W 

Assuming, λ = 0.26 cm cm-1., 1., and μ = 0.5 

L/2 W/2 

SLW ( , ) 4 Dxydxdy ( , ) 

� � � 

L / W 

0 0 

1 2 3 4 5 

SLW ( , ) / S (10,10) 1.000 0.993 0.982 0.969 0.958 

L 

W s

KLEIN- NISHINA CROSS SECTION 

FOR THE COMPTON INTERACTION 

d e� 

d� 

Ψ 

� 

2 2 

0 2 

sin 

r �� 

h� 

' �� 

�� 

h� 

h� 

' 

� � � � � 

2 � h� 

� � h� 

' h� 

PHOTONS SCATTERED AT AN ANGLE, Ψ 

�� 

� � 

� 

PHOTONS SCATTERED INTO A UNIT 

SOLID ANGLE, Ω 

SOLID ANGLE AVAILABLE PER 

UNIT ANGLE 

d� 

d d�� 

� 2� sin � 

Based on the kinematics of Compton interaction, the average 

p g 

energy of scattered photons is less than 1Mev and is independent 

of the incident energy.

Measurement of Normalized Depth 

DDose ddata 

Follow AAPM TG Report p # 106 recommendations: 

� Use 4-5 mm diameter ion chamber for depth 

beyond 1cm. 

� Use parallel plate or extrapolation chamber to 

measure data near the surface. 

�� Di Diodes d and d di diamond d ddetectors t t are appropriate i t 

as long as data measured with these detectors 

is scoss cross-referenced e e e ced to data measured easu ed with t aan 

ion chamber. 

� Prone to radiation damage and non-linear response.

Is depth ionization data depth dose? 

YES!!! 

With the caveat, 

�� TCPE exists at the point of measurement 

measurement. 

� the energy spectrum of incident photons does not change with the 

depth. 

�� fluence across the detector remains the same same. 

These conditions are met at depths beyond the range of 

contaminant charged particles 

However at shallow depth, The contaminants and 

secondary electrons have energy spectra that change 

rapidly with depth. 

� Results in a variation of ~10% in restricted mass stopping power 

ratio data for water and air. 

�Translates into a spatial uncertainty of less than 1.5 mm in dose in the 

build up region

S c 

c 

S 

p 

�� s 

� 

Beam Output 

S 

c, 

p 

S 

c 

�� s�� 

�� c 

f 

f 

(Derived) 

S c,p 

10 cm 

Water 

c

In-air output Ratio 

Elekta: 4-18 MV clinical photon beams beams.

Monte Carlo Calculations of In- 

Simulation Geometry 

(Varian 2100EX) 

Ai Air Output O t t Ratio R ti 

(BEAMnrc code) 

O o 

/tex/rof/clxyro 

1.05 

1.00 

0.95 

0.90 

In-Air Output Ratio 

6 MV measured 

6 MV calc calculated lated 

18 MV measured 

18 MV calculated 

0 5 10 15 20 25 30 35 40 45 

Side of square field /cm

Energy spectrum of head scattered photons 

Mean Energy:0.5 MeV 

(Varian 2100C.)

Energy spectrum of head scattered photons 

Mean Energy:0 Energy:0.5 5 MeV 

(Varian 2100C.)

In-air output Ratio 

e: Elekta, , s: Siemens, , and v: Varian 

(for clinical photon beams ranging from 6-25 MV.

Flattening Filter 

Beam Modifier 

(internal wedge) 

Lower Collimator 

Beam Modifier 

(external wedge) 

Monitor Back Scatter 

Monitor Chamber 

Upper Collimator 

Tertiary Collimator 

(Cerrobend Block 

or Varian MLC) 

Machine MBS Publication 

Varian Clinac 1800 1-5% Kubo, Med. Phys. 

16, 295 (1987) 

Therac 20 7.5% Hounsell, , P.M.B. 

43, 445 (1998) 

Elekta SL15

MMeasurement t of f In-Air I Ai Output O t t Ratios R ti 

• Mini phantom p 

– Water-equivalent materials. 

– 4g/cm2 diameter and 10g/cm2 depth to maintain lateral 

CPE and eliminate contaminant electron electron. 

• For small segment fields (c

Cross Beam Characteristics 

� Affected by the radially symmetric conical high Zmaterial 

flattening g filter, , which 

� Flattens the beam by differentially absorbing more photons in the 

center and less in the periphery 

� unwanted consequence of flattening the beam is the differential 

change in beam quality at off-axis points points. 

� hardens the beam 

� Cross beam flatness is defined as: 

D 

F � 100� 

D 

max 

max 

� D 

� One flattening filter for each clinical photon beam results in a 

compromise of beam flatness characteristics of small and large 

fields. 

�� Fl Flattening tt i filt filters are ddesigned i d tto give i a gradually d ll iincreasing i radial di l iintensity. t it 

This is referred to as “horns” on a cross-beam profile 

� Cross beam profiles may not be radially symmetric due 

to non circular focal spot. p 

� Therefore, cross-beam data is characterized by a set of two 

orthogonal dose profiles measured perpendicular to the beam’s 

central axis at a given depth in a phantom 

� 

D 

min 

min

Cross Beam Profile 

6 MV Photon Beam, Depth of 5.0 cm, Field size of 4x4, 10.4x10.4, and 21x21 cm 2 . 

The flatness of photon beams is extremely sensitive to change in energy of the 

incident beam. A small change in the penetrative quality of a photon beam results in 

very large change in beam flatness.

Cross Beam Profile 

6 MV Photon Beam, , Field Size of 10.4x10.4 cm2, , Depths p of 1.5, , 5.0, , 10.0, , 15.0, , 

and 25.0 cm. 

The field flatness changes with depth. This is attributed to an increase in scatter to 

primary dose ratio with increasing depth and decreasing incident photon energy off 

axis

Effect of Electron Steering 

on Beam Flatness 

Symmetric Tilted Displaced

Effect of a Dipole p Magnet g on Exit 

Energy Spread 

Beam 

Radial Displacement 

Radial Divergence

Cross Beam Symmetry 

S 

� 100 � 

Area 

left 

� 

Area � 

left 

Dosimetry and beam 

steering system 

Area 

Area 

right 

right

Isodose Distribution 

30 cm X 30 cm 

18 MV X-ray beam

6 MV 

Isodose Distributions 

(20 X 20 Cm2 (20 X 20 Cm ) 2 ) 

18 MV 

Note contaminant electrons contribute to dose outside the field 

at shallow s a o depths. dept s The e magnitude ag tude aand d eextent te t oof dose outs outside de 

the geometric edge of a field at shallow depths increases with 

beam energy.

Isodose Distributions 

(20 X 20 Cm2 (20 X 20 Cm 18 MV) 

2 , 18 MV) 

Note Contaminant electrons contribute to dose outside the 

field at shallow depths. p The magnitude g and extent of dose 

outside the geometric edge of a field at shallow depths 

increases even more in the presence of beam modifiers.

Cross Beam Measurements 

Diameter 

Penumbra 

20%~80% 

Wh What t iis th the affect ff t of f 

detector size? 

�Incorrect 

measurement of 

penumbra region 

Diode CC04 CC13 

0.8x0.8 

mm 2 4 mm 6 mm 

4.0 mm 6.1 mm 7.2 mm

Detector Size Effect on TPS 

Commissioning 

Impact of 

Treatment Planning detector size 

System 

effect on dose 

Commissioning di distribution??? 

t ib ti ??? 

Yan G et. al., Med. Phys (35)., 2008

Extraction of True Profile

IMRT QA results: DTA 2%/2 mm 

CC13 

CC04 

Deconvolved

Measurement of Attenuation 

FFactors t for f Beam B Modifiers M difi 

� The attenuation factor for a beam modifier is defined as 

the ratio of the dose rate at the point of calculation for a 

given field with and without the modifier in place. 

� Attenuation factors for devices such as block trays, y , accessories 

etc. are often assumed to be independent of field size, depth and 

SSD. 

� These factors should be measured at a depth well beyond the 

maximum range of electron contamination 

� The attenuation devices that are in contact with the 

patient skin (immobilization apparatus, table top, etc.) 

req require ire additional considerations 

considerations. 

� These devices not only attenuate the incident beam but they 

introduce scatter radiation that increase the scatter to primary 

ratio within the patient patient. 

� It is best to include such attenuation devices as a part of the patient 

in 3DRTPS

Measurement of Wedge Factors 

� The WF is defined as the ratio of the dose rate 

at t the th reference f depth d th for f a wedged d d field fi ld tto th that t 

for the same field without a wedge modifier . 

�� The field size dependency of the WF originates from 

a wedge-induced increase in head scatter. 

� the field size dependence p of the WF is correctly y 

accounted for by in-air output ratios (Sc) wedge 

specifically measured for wedged fields 

�These �These data should be measured with the chamber axis 

perpendicular to the gradient direction of the wedge 

�Two sets of measurements should be made with the wedge 

in opposite orientations to ensure the correct placement of 

the chamber

Characterizing Clinical Photon 

Ahnesjo et al al., PMB 1999 

BBeams iin 3DRTPS

Approaches to Dose Computation Algorithms 

Reconstitute water data 

Calculate inhomogeneity 

corrections to water data 

Data measured in water 

and in air 

Parameterize water data 

Calculate dose directly 

based on beam and 

phantom configurations 

“Correction 

Correction” ” based “Model Model” ” based 

methods methods 

Figure 8.9,The Modern Technology of 

Radiation Oncology; J. Van Dyk

Correction vs. Model Based Methods 

Correction Based Model Based 

Measured data used as basis for 

Dose Computation. 

Measured data used to setup 

description of treatment beam. 

Require measurements with buildup Require a parameter to estimate size 

cap in air or in a mini-phantom. of photon source at target. 

Require lots of data. Generating 

functions used to reduce size of 

data set for convenient clinical use 

(i.e. less storage space). 

Require more time for tuning of 

model parameters. 

p 

Patient dose distribution obtained by Patient dose distribution obtained by 

first computing Dose in water from 

generating function, then correcting 

for tissue heterogeneity, patient 

contour, t and d bbeam modifiers. 

difi 

computing beam and beam transport 

(i.e. beam interactions in treatment 

head and in patient) directly.

Accuracy Goal in Dose Calculations 

•Required q accuracy y (overall ( treatment < 5%): ) 

Ahnesjo et al., PMB 1999


Beams in 3DRTPS 

�� MUST model the following features realistically: 

� Finite size of source (& penumbra) 

�� EExtra-focal t f l radiation di ti (primary ( i collimator, lli t fl flattening tt i filt filter) ) 

� Beam spectrum (& change in spectrum with position) 

� Beam intensity variation across field (e.g., beam horns) 

� Transmission through secondary collimators 

�� SScatter tt outside t id fi field ld ( (related l t d tto extra-focal t f l radiation) di ti ) 

� MLC, blocks, block tray 

�� Dynamic wedge wedge, fixed wedge wedge, compensators (beam 

hardening)


Caveats: 


� Almost all photon dose computation with convolution models 

assumes kernel invariance, which requires the photon dose kernel to 

be constant with spatial locations in the calculation phantom phantom. 

� However, in clinical treatments, patient inhomogeneities, as well as beam 

divergence and polychromaticity, cause kernel variation in various ways. 

� Modeling of charged particle contaminants is at best an 

approximation of real clinical situation 

�� Modeling of indirect radiation as a single or multiple analytical 

source functions, modeling of off-axis softening with a simple 

parametric fit, source size, etc. are best effort estimates of physical 

processes


Caveats (continued): 


� One can always use a set of beam modeling parameters 

to get the best agreement between the computed and 

measured beam data in a phantom phantom. . 

�However, that would not be a sufficient condition for robust and 

accurate beam modeling . 

� The value or function used to describe a parameter 

should have some physical meaning. 

�each parameter used in the dose calculation algorithm should 

model the physical reality it represents even if there is less than 

perfect agreement between measure and computed data. 

� The observed differences often reflect limitations of the dose 

computation algorithm

Benchmark Dataset 

(D (Developed l d under d NIH iinitiative) iti ti ) 

A collaborative effort involving Sun Nuclear Associates; the 

contractor, t t and d consultants lt t from: f the th University U i it of f Florida; Fl id 

the RPC at M.D. Anderson Cancer Center; the University 

of Iowa; and the Vassar Brothers Hospital. p 

� Already measured a complete set of data on the new 

generation of Elekta (Synergy), Siemens (Oncor) and 

Varian (Trilogy) linear accelerators 

�Measured data are comprehensive in beam geometries to 

validate dose computation for any clinical situation. 

�data are sufficient in spatial resolution and were validated by 

independent measurements 

� This benchmark datasets will be sufficient for the TPS 

� This benchmark datasets will be sufficient for the TPS 

companies to compare the accuracy of their dose 

modeling for treatment delivery

Summary 

� The dosimetric properties of a clinical photon 

beam are characterized by: 

� Its ability to penetrate a tissue-like medium (water) 

� its change g in dose output p with field size 

� Its cross beam behavior 

� Its attenuation through modifying devices (e.g., wedge, 

compensator etc.) etc ) 

� The dosimetric properties of clinical photon 

beams from linacs depend on the photon energy 

fluence distribution emanating from the treatment 

head, on the geometry of the linac, and on the 

radiological properties of the medium with which it 

interacts.

Summary 

�It is quite evident that all modern clinical 

li linear accelerators l t (li (linacs) ) of f a particular ti l 

commercial make produce beams of very 

similar characteristics 

�High quality benchmark data have already been 

acquired by comprehensively characterizing 

single linacs of each make. 

�These benchmark data thoroughly describe the 

characteristics of photon beams so that 

treatment-planning companies and clinics 

throughout the United States can use it to 

examine the accuracy of dose-calculation 

algorithms.

Dosimetric Characteristics of Clinical Photon Beams

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