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Supervisors:
Kim H. Parker, Phd
Frans N. van der Vosse, Phd
-
Characterising pressure and flow in the
Jarvik 2000 � Heart and the study of its
acoustic properties as a possible noninvasive
assessment of its operating conditions
Arjen van der Horst, 0506054,
Rapport no. BMTE 06.24
Department Biomedical Engineering,
Tissue engineering and Biomechanics ,
Eindhoven university of technology

Contents 2
Contents
1 Introduction 3
2 Material and Methods 5
2.1 Device components . . . . . . . . . . . . . . . . . . . . . . . . . . . 5
2.2 Flow and pressure characteristics of the Jarvik 2000 . . . . . . . . . 6
2.3 Rotary pump acoustics . . . . . . . . . . . . . . . . . . . . . . . . . 9
2.3.1 Thrombosis . . . . . . . . . . . . . . . . . . . . . . . . . . . 10
2.3.2 Frequency shift due to heart contractions . . . . . . . . . . . 10
2.4 Experimental setup . . . . . . . . . . . . . . . . . . . . . . . . . . . 10
2.4.1 H-Q relationship . . . . . . . . . . . . . . . . . . . . . . . . 11
2.4.2 Phase delay . . . . . . . . . . . . . . . . . . . . . . . . . . . 12
2.4.3 Inlet blockage . . . . . . . . . . . . . . . . . . . . . . . . . . 12
2.5 Clinical measurements . . . . . . . . . . . . . . . . . . . . . . . . . 13
2.5.1 Best microphone position . . . . . . . . . . . . . . . . . . . 13
2.5.2 Frequency shift due to heart contraction . . . . . . . . . . . 13
3 Results 14
3.1 Flow and pressure characteristics . . . . . . . . . . . . . . . . . . . 14
3.1.1 H-Q relationship . . . . . . . . . . . . . . . . . . . . . . . . 14
3.1.2 Phase delay . . . . . . . . . . . . . . . . . . . . . . . . . . . 15
3.2 Acoustics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16
3.2.1 Inlet blockage . . . . . . . . . . . . . . . . . . . . . . . . . . 17
3.3 Clinical results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 18
3.3.1 Best microphone position . . . . . . . . . . . . . . . . . . . 18
3.3.2 Frequency shift due to heart contractions . . . . . . . . . . . 20
4 Discussion and conclusions 22
5 Acknowledgements 23
2

1 Introduction 3
1 Introduction
The main goal of decades of research in mechanical cardiovascular support devices has been to
ameliorate poor survival and improve quality of life of patients with congestive heart failure. This
research was stimulated by the increasing prevalence of congestive heart failure. Heart failure
affects about 5 million Americans, which is more than 2 percent of adults in the United States.
With 550,000 new cases of heart failure diagnosed in the United States each year and an estimated
direct annual cost of $ 29.6 billion this is a serious problem [1].
In the 1980’s the introduction of cyclosporine established heart transplantation as the most effective
therapy for end-stage heart disease. But due to a shortage of donor hearts an alternative
was needed to bridge the waiting period for a donor heart. The use of ventricular assist devices
(VAD’s) has been considered controversial because such an approach results in redistribution of
donor hearts at additional cost. Nevertheless, proponents of this strategy argue that bridging for
transplantation is justifiable because it results in technological improvement and an unique insight
into the mechanics of heart failure.
In 1964, the National Heart, Lung and Blood Institute began to sponsor the development of
mechanical devices for circulatory support. This resulted in several circulatory support systems
which were designed for short-term use in patients with end-stage heart failure. Initially there were
problems like thomboembolism, infection and low survival rates [2]. But in 1985 a multicenter
clininal evaluation of left ventricular assist devices (LVAD’s) as a bridge of transplantation showed
that these devices had a survival rate to transplantation of 65 percent compared to 50 procent
for medically treated patients [3]. On basis of this evaluation and other studies [4, 5] the Food
and Drug Administration (FDA) approved the use of left ventricular assist devices as a bridge to
transplantation in 1994. Vented electric devices were approved for same purpose in 1998 [2].
Nowadays there are two categories of LVAD’s: positive displacement and rotary pumps. Examples
of positive displacement systems are: HeartMate (Thoratec Laboratories Corp., Pleasanton, CA,
USA) and the Novacor N1000PC (World Heart, Ottawa, ON, Canada). Examples of rotary pump
based systems are: HeartMate II (Thoratec Laboratories Corp.,Pleasanton, CA), the DeBakey
VAD (MicroMed Technology, Inc., Houston TX) and the Jarvik 2000 (Jarvik Heart, Inc., New
York, NY). The advantages of the rotary LVAD’s compared with positive displacement LVADs are:
small size, relatively low risk of thombus formation and ease of implantation. On the other hand,
positive displacement LVAD’s give more physiological waveforms with bigger pressure and flow
amplitudes, especially when there are no pressure changes due to heart contractions. Although
several studies suggested potentially deleterious effects associated with diminished blood flow
pulsatility [6, 7], including increased thrombogenicity, other studies have not [8, 9, 10, 11]. No
deleterious physiological effects of loss of blood flow pulsatility have been shown in long-term
animal studies [8, 12].
So far LVAD’s have been mainly used to as a bridge to transplantation, but due to increased
reliability and a lack of donor hearts, LVAD’s are increasingly used as a bridge to myocardial
recovery or for permanent circulatory support [13]. As LVAD’s are used for longer periods, the
device durability becomes more important. Furthermore, the ability to detect mechanical failure
at an early stage could be advantageous. That is why an easy, fast, cheap and non-invasive method
would be preferable. The aim of this preliminary study is to investigate whether the sound of an
axial flow LVAD: The Jarvik 2000 � Heart (Jarvik Heart Inc., New York, NY) can be used as a
diagnostic tool. As mentioned, the Jarvik 2000 is an axial flow pump and is used in the United
States as a bridge to transplantation under a FDA-approved clinical investigation. In Europe, the
Jarvik 2000 has a CE Mark certification for both bridge to transplant and lifetime use.
The subjects in this study can be categorized as follows: characterisation of pressure and flow,
acoustics and clinical measurements of the Jarvik 2000. In the section on the characterisation
of pressure and flow the relation between pressure head and flow and the phase delay between
3

1 Introduction 4
pressure and flow will be discussed. In the section on acoustics, the sound of the Jarvik 2000
and the influence of the inlet blockage is investigated. In the last section, sound measurements on
patients are described. Here the frequency spectrum of the experimental and clinical measurements
are compared. The best place on the chest is to record the sound is investigated. Finally, a new
method to gain insight in the contractility of the heart is explained and discussed.
4

2 Material and Methods 5
2 Material and Methods
The Jarvik 2000 � Heart (Jarvik Heart Inc., New York, NY) is an intraventricular LVAD which is
designed to give partial to complete assist to the failing heart. The pump is implanted through a
left thoracotomy or median sternotomy. The pump is positioned in the apex of the left ventricle
through a circular incision and secured with a silicone-polyester sewing cuff. This is shown in
Figure 2.1. An outflow graft is extended from the outlet of the pump to either the ascending
or descending aorta. The ascending aortic anastomosis is the preferred configuration as it avoids
retrograde flow in the aortic arch and blood stagnation at the level of the aortic root [14]. A power
cable is brought to a controller via the abdominal wall or via a skull mounted pedestal located
posterior to the left of the ear. This controller regulates the rotation frequency of the pump and
provides visible and audible signals.
Figure 2.1: Position of the Jarvik 2000 in the thorax (left) and a picture the Jarvik 2000 with
sewing cuff (right). Also shown is a cut-away pump showing the impeller (lower left). Reproduced
with permission of Jarvik Heart, Inc., New York, NY.
2.1 Device components
The Jarvik 2000 � Heart consists of a blood pump, a Hemashield outflow graft, a Dacron insulated
percutaneous power cable, a pump speed controller and a battery for direct-current power supply.
The pump weighs 90 g, 2.5 cm in diameter and has a displacement volume of 25 cm 3 . A great
advantage of the small size of the pump is that it can be implanted in small adults and even in
children. The only moving part of the pump is an impeller which is located at the center of the
titanium housing. The impeller consists of a neodymium-iron-boron magnet suspended by two
ceramic bearings. On the surface of the impeller there are two titanium blades. The impeller is
rotated by a brushless direct-current motor which is located within the housing. The blood flow,
created by the rotation of the two blades of the impeller, is directed to the outflow graft by four
stator blades. The purpose of stator blades is to recover the rotation energy imparted to the blood
by the rotor. In Figure 2.2 the impeller and stator blades are visualized. The intraventricular
position of the pump obviates the need for an inlet cannula, which avoids the unfavorable effects of
negative pressure on the blood (eg. hemolysis and platelet adhesion and destruction). To further
minimize the formation of thrombus all blood-contacting surfaces are made of smooth titanium
[15, 16]. The Jarvik 2000 Heart normally comes with a controller with 5 different speed levels:
8000, 9000, 10000, 11000 and 12000 rpm. The pump always tries to maintain the prescribed
number revolutions per minute. With the highest rate of rotation, the pump can provide in excess
of 11 L/min [17]. For this study Jarvik Heart Inc. was so kind to make one of their Jarvik 2000
pumps available. This pump is unwelded so alterations can be made more easily.
5

2 Material and Methods 6
Figure 2.2: The impeller and the stator blades of the Jarvik 2000. Reproduced with permission
of Jarvik Heart, Inc., New York, NY.
2.2 Flow and pressure characteristics of the Jarvik 2000
Figure 2.3 shows the amount of flow which can be produced under different pressure heads and
different rotation speeds. It shows that there is approximately a linear relationship between flow
(Q) and the pressure difference between inlet and outlet (H). So at a certain rotating frequency
and constant pressure head, the pump will create a continuous flow. In this study this effect is
checked for the available pump. The experiment is explained in section 2.4.1.
Figure 2.3: Adapted from Macris et al[17]. in vitro flow/pressure curves with Jarvik 2000 blood
pump. The pump was tested in 3.3 · 10 −6 m 2 /s glycerol/water mixture under steady-state conditions.
When the heart beats the pressure is, of course, not constant. The flow through the LVAD will
thus change according to the pressure flow relationship shown in the Figure 2.3. So if the heart
contracts there is an increase in inlet pressure and the pressure difference between the inlet and
outlet will therefore decrease (the outlet pressure is of course higher than the inlet pressure) and
the flow will therefore increase. When the heart is beating more vigorously, the flow changes will
be larger. Because of an increase in flow (due to decrease in pressure head) the aortic pressure
will also increase. The effect of this pressure-flow character of the pump on the (compliant)
aortic pressure can than be approximated according to the Windkessel model. The variation of
the aortic Windkessel pressure (PW k) is determined by the difference between inflow (Qin) and
outflow (Qout) [18]:
dPW k(t)
dt
= dPW k(t) dVW k(t)
=
dVW k dt
Qin(t) − Qout(t)
C
(2.1)
6

2 Material and Methods 7
with,
C = dVW k
dPW k
Here C is the compliance of the entire arterial tree and is assumed to be constant. When the
outflow is described as the resistive relationship:
Qout(t) = PW k(t) − P∞
R
With P∞ the asymptotic pressure of the diastolic exponential decay and R is the effective resistance
of the peripheral systemic circulation. Substitution of equation 2.3 into equation 2.1 leads to:
dPW k(t)
dt
= Qin(t)
C − dPW k(t) − P∞
RC
When all the flow to the aorta Qin(t) is produced by the LVAD at a certain speed this Qin(t) can
be written in terms of the linear pressure head-flow relationship of Figure 2.3:
(2.2)
(2.3)
(2.4)
Qin(t) = Qmax + (PLV (t) − PW k(t))a (2.5)
Here, Qmax is the flow with no pressure head and a is the slope of the H-Q curve. Substitution of
equation 2.5 in equation 2.4 results in:
dPW k(t)
dt
= Qmax + (PLV (t) − PW k(t))a
C
− dPW k(t) − P∞
RC
And by using an integrating factor and applying the periodic boundary condition PW k(0) =
PW k(T ) = P0, the following equation for PW k is obtained.
Here P0 is
PW k(t) = e
P0 =
� T
0
a 1 −( C + RC )t
� t
e
a ( C
0
e
a ( C
1 + RC )t′
�
Qmax + PLV (t ′ )a
+
C
P∞
�
dt
RC
′ + e
1 + RC )t′
�
Qmax + PLV (t ′ )a
+
C
P∞
�
dt
RC
′
a 1 −( C + RC )t P0
Due to possible loose coupling between the impeller rotation and the blood flow and the inertia of
the blood in the pump there might occur a phase lag between the rise in left ventricular pressure
and an increase in flow. Recent clinical measurements of the pressure in the left ventricle and
femoral artery suggest this phase lag is about 87 ms [19]. The measurements are shown in Figure
(2.6)
(2.7)
(2.8)
7

2 Material and Methods 8
2.4. To obtain this result the pulse wave propagation from the left ventricle to the femoral artery
is estimated to be 50 ms [19]. This was based on an estimated aortic path length of 50 cm and
estimated pulse wave velocity for a forty year old patient. It would be interesting to measure
this phase lag in vitro so the phase lag between pressure and flow caused by the pump can be
investigated under more controlled conditions. Therefore, an experimental set-up is created to
investigate this phase delay. This set-up is explained in section 2.4.2.
Figure 2.4: Adapted from Bowles et al. [19]: Pressure versus time relationship for left ventricular
pressure (lower trace) and right femoral artery pressure (upper trace) using the Jarvik 2000 LVAD
with the outflow anastomosed to the descending aorta. 1 square = 10 mmHg (vertical) and 200
ms (horizontal). The recorded phase lag between the peaks of the two traces was 137 ms. Note
that left ventricular pressure variation was propagated into the systemic arteries in spite of the
left ventricular pressure being insufficient to open the aortic valve.
Figure 2.5: Adapted from Bowles et al. [19]. Decrease in pulse pressure due to increasing pump
speeds.
Figure 2.5 shows that, even though the heart is contracting, increasing the pump rotation frequency
will decrease the pulse pressure in the aorta. This can be understood as follows: Because increasing
8

2 Material and Methods 9
the pump speed means an increase in flow, the blood volume in the left ventricle will be less.
Therefore the heart muscle will be stretched less and according to the Starling principle will
contract less vigorously. This will then result in a smaller amplitude of the left ventricular pressure
wave and therefore will decrease the pulse pressure in the aorta.
2.3 Rotary pump acoustics
There have been several studies which investigated the nature of the sound of axial flow devices
[20, 21, 22, 23]. These suggest that there are several sources of sound created by a rotary pump.
The most significant source of vibrations is the number of impeller blades which pass a certain
point. So in the case of the Jarvik 2000, that would be the two impeller blades times the rotation
frequency [23], but due to the curled impeller blade design (See Figure 2.2) there are not two
specific crossing sites. Therefore it is expected that the rotation frequency will also be clearly
present in the frequency spectrum. There is a specific crossing site with the stator blades at the
end of the impeller blades near the stator. This will, of course, depend on the clearance between
the stator and impeller blades, but because there is only three millimeter clearance it is likely to
have a demonstrable effect [23]. This effect will be more clear when turbulence flow will hit the
stator blades. The Reynolds number can give an indication of the presence of turbulence.
Re = ρvh
µ
Here is ρ the density of the fluid, v the mean velocity of the fluid, h the clearance between the
impeller and the casing and µ the dynamic viscosity. h is about two millimeter, ρ is 10 −3 kg ·m −3 ,
µ is 10 −3 P a · s and v is calculated as follows:
v = Q
A ≈
Q
2π R h =
(2.9)
0.5 · 10−4 2 π 5 · 10−3 = 0.80 m/s (2.10)
2 · 10−3 Q is the flow through the clearance with surface A and R is the radius to the edge of the impeller.
At normal operating conditions, 10000 rpm and a pressure head of about 75 mmHg, the flow will
be about three liters per min. The Reynolds number will therefore be:
Re = 10−3 0.80 2 · 10 −3
10 −3 = 1600 (2.11)
The flow will therefore be likely to be turbulent when considering the rotation of the impeller.
This turbulence will increase the acoustic effect of the impeller blades crossing the stator blades.
Due to the four stator blades and two impeller blades and taking the symmetry of the impeller
into account these impeller-stator blade crossings will introduce noise at four times the rotating
frequency. But due to the placement of the microphone, the time it takes for the noise to reach
the microphone may become important. The crossing of the stator blade which is furthest from
the microphone will be detected later than the one which is closer. In this case the time it takes
the sound to reach the microphone will be so small that a sampling rate of about 150,000 s −1 is
needed, which much bigger than the sampling rate used (see section 2.4). So it is expected that
the most energetic frequencies in the frequency spectrum are at the rotation frequency and the
first and third harmonic.
9

2 Material and Methods 10
2.3.1 Thrombosis
Although the incidence of thrombosis due to LVADs has decreased, it still is a significant problem.
A non-invasive way to detect thrombus formation on the blood path of an axial impeller LVAD
would be highly desirable. A possible way to detect thrombus blocking the entrance of an axial
flow pump, and therefore influencing the function of the pump, is to measure the sound which is
created by the pump. Trunzo et al. [20] showed that turbulence created by struts placed in front
of the inlet of a fan will influence the sound it makes. More specific, they showed that these struts
will have effect on the on the energy at the rotating frequency and/ or the harmonics [20]. Another
possible effect of blocking a part of the inlet could be an asymmetrical flow profile which can cause
an imbalance of the impeller. This could be detected by vibrations of the bearings, which will
result in an extra peak in the frequency spectrum. The effect of inlet blockage is simulated with
an in vitro experiment which is explained in section 2.4.3.
2.3.2 Frequency shift due to heart contractions
Another property of an axial flow pump, which can be detected by measuring sound, is the change
in pressure head. As said before, the Jarvik 2000 will try to maintain the prescribed speed
of rotation. However, the reaction of the pump when exposed to pressure head changes is not
instantaneous. Therefore the rotation speed will increase with a sudden decrease of pressure head.
The controller will then decrease the rotation speed to the prescribed level. So in theory it should
be possible to detected the pressure head changes due to the heart contractions by measuring the
frequency shift of the rotating frequency. This can be used to quantify the ability of the heart
muscle to contract. This is of course very important for patients who have the Jarvik 2000 as a
bridge to recovery, as amplitude of frequency shift may correlate to recovery. It seems therefore
useful to explore if there is such a frequency shift due to heart contractions. This is done by
clinical follow-up measurements explained in the section Clinical measurements (2.5.2).
2.4 Experimental setup
Figure 2.6 shows the basic experimental setup used in this study. Minor alteration are made to
this set-up to do the different experiments which are explained in the next sections. The set-up
consists of two reservoirs which contain filtered water. The water is filtered to prevent calcium
formation on the impeller blades. The pump is located in between the two reservoirs and connected
with silicon tubes. Reservoir 1 (right) and 2 (left) respectively have inner dimensions (l × d × h):
39.1 × 23.6 × 29.6 cm and 16.9 × 20.0 × 24.6 cm. Reservoir 2 has an overflow to create a constant
water level. This overflow is located at a height of 17.9 cm above the bottom of the reservoir.
Silicone tubes are used to direct the water from the overflow back to reservoir 1.
Rubber rings are wrapped around the pump and the pump clamped within a perspex tube (see
the left side of Figure 2.7) which is placed between the two reservoirs. This construction is chosen
to protect the ceramic bearings from mechanical shock. Two taps are placed on the perspex tube
to extract air bubbles from the perspex tube. The reservoirs can be lifted up and down to create
pressure difference between the inlet and outlet of the pump. The pressure difference is thus
gravity based. A tap is placed between the outlet of the pump and reservoir 2. When the pump is
not running, this tap prevents backward flow from reservoir 2 to reservoir 1. To record the sound
of the pump an ordinary stethoscope is adapted to produce an analog electrical output using a PC
microphone. The stethoscope is the Sprague stethoscope made by Prestige Medical (Northridge,
CA).
The stethoscope has one big bell and one smaller bell. For the in vitro experiments the small bell
is used. The main reason for this is that, in the in vitro experiments, the small bell fitted better
10

2 Material and Methods 11
Figure 2.6: Sketch of the experimental set-up used. The arrows indicate the direction of flow
Figure 2.7: Left: Picture of the Jarvik 2000 located in the perspex container. Right: Picture of
the Stethoscope and the recording Walkman.
onto the small exterior of the pump and therefore increased the signal-noise ratio. In section
2.5.1 the differences between the small and big bell are further explained. The acoustic signal is
recorded with the Sony MZ-RH10 Hi-MD Walkman Digital Music. The sampling rate is 44100
Hz. The right side of Figure 2.7 shows the walkman and stethoscope. The stethoscope is clamped
and placed as close as possible to the pump, which is covered by a sheet of latex with coupling
gel. This position is chosen because in this way the stethoscope is closest to the pump. Every
1
measurement has a duration of two minutes which gives a spectral resolution of 120 = 8.33 · 10−3
Hz.
2.4.1 H-Q relationship
To obtain the relationship between pressure head and flow, for this particular pump, the closed
circulation described above is opened. The overflow from reservoir 2 is collected in a reservoir. This
overflow is then measured by timed collection. This is done for ten different pressure heads at five
different pump speeds. Every measurement is repeated three times. So a total of 150 measurements
are made. In every measurement the flow created by the pump is measured for 30 seconds. The
water level of reservoir 1 is decreasing during these 30 seconds when the pump is running. Therefore
instead of a steady state there is a quasi-steady state situation. For physiological flow rates this
11

2 Material and Methods 12
quasi-steady state is justified because due to the size of reservoir 1 only minor changes (1-3 mmHg)
in the water level are expected to occur. The results are shown in section 3.1.1
2.4.2 Phase delay
To measure the pump related phase delay between pressure and flow the experimental setup of
Figure 2.6 is changed. This setup is shown in Figure 2.8. One flow and two pressure measurements
are made and recorded simultaneously. The pressure is measured two centimeters upstream and
downstream of the pump. The pressure measurements are made with a pressure transducer (Gaeltec,
CTC Series, Isle of Man). The flow measurement is done 16 centimeter downstream of the
outlet of the pump. The flow is measured with a flow transducer (Transonic, Ithaca, NY, USA).
Instead of reservoir 2, of the original set-up, a compliance reservoir is connected to the outlet of
the pump. The connection is made with latex tubing. This latex tube is submerged in water to
enable the flow transducer measurements. Because of the major decrease in wave speed in the
latex tube compared to the silicone and perspex tubing, the probe to measure the flow is placed
as close to the start of the latex tube as possible. To create the pressure wave, the tap before
the inlet of the pump is suddenly closed and openend again. This will create a negative-block
shaped pressure wave. This is of course not a physiological waveform, but keeping in mind this is
a preliminary study this is enough to give more insight into the resulting phase lag.
Figure 2.8: Sketch of the experimental set-up used for the phase delay experiment. The arrows
indicate the direction of flow
2.4.3 Inlet blockage
To simulate inlet blockage, obstructions are placed just before the pump inlet. The obstructions
consist of a plastic cap, which fit tightly into the perspex container. An opening is made in these
plastic caps to create obstructions of: 1 1 3
4 , 2 and 4 the original cap area. Figure 2.9 shows a picture
of the three caps used. Because of the symmetry plane of the Jarvik 2000 the obstructions are
placed at only one side of that symmetry plane. The obstructions are therefore placed in two
configurations (see Figure 2.9). For each obstruction five measurements are made. All measurement
are made at 10000 revolutions per minute. This is chosen because this is the most common
rotation speed in patients.
12

2 Material and Methods 13
Figure 2.9: Left: A representation of the two configurations used. Right: A picture of the caps
used to simulate inlet blockage.
2.5 Clinical measurements
Clinical measurements are made to investigate whether the acoustic results from the in vitro experiments
can actually be applied to patient data. Therefore two kind of clinical measurements
are made: First, the sound of the Jarvik 2000 is measured in one patient to obtain more insight
into te difference between the big or small bell and the importance of the position of the stethoscope.
Furthermore, to explore the possible frequency shift due to heart contractions follow-up
measurements are made on a patient recovering from a Jarvik 2000 implantation.
2.5.1 Best microphone position
To get more insight into the importance of the position of the stethoscope and which bell should
be used to obtain the best results, measurements are made on a male patient. According to the
stethoscope manual the small bell is better for high frequencies and the big bell is better for low
frequencies. But it is unknown what this means for the sound measurements of the Jarvik 2000 as
those frequencies are much higher than associated with a normal heart beat. The measurements
are made at three places: Position 1 is on the second rib below the left nipple, position 2 is on the
same rib but at an angle of 45 degrees to the left arm of the patient and position 3 on the third
rib below the left nipple. This is shown in Figure 2.10. On both locations on the second rib below
the nipple two measurements are made; one with the small bell and one with the big bell. On the
position on the third rib below the left nipple only the measurement with the small bell is made.
The results are shown section 3.3.1
Figure 2.10: The different measurement positions on the chest.
2.5.2 Frequency shift due to heart contraction
To measure the expected changes in rotation speed the sound of the Jarvik 2000 is measured in
a patient who is recovering from a Jarvik 2000 implantation. This is done because, in the period
when the heart is recovering from the implantation of the Jarvik 2000, it will be beating less
vigorously than normal. Therefore, there is a good chance that follow-up measurements will show
13

3 Results 14
the difference between less and more vigorous contractions on the rotation speed. So it is expected
that just after the implantation the changes in rotation speed are less compared to days after the
implantation. To compare the results and check whether the changes in rotation speed are caused
by changes in blood pressure, measurements are made of the blood pressure and heart rate. The
measurements are made on the position with the small bell at position 2.
3 Results
3.1 Flow and pressure characteristics
3.1.1 H-Q relationship
The flow (Q) is measured for 10 different pressure heads (H). The results are shown in Figure 3.1.
For physiologically representative pressure differences the relation between H and Q is approximately
linear. The reason why the relation is not linear for small pressure heads is because of the
quasi-steady behavior of the inlet pressure. This quasi-steady behavior, caused by the decreasing
water level in reservoir 1, exacerbates with higher flows. For the physiological representative
pressure heads the obtained results are similar to the results of Macris et al. [17]
Q (L/min)
10
9
8
7
6
5
4
3
2
1
0
0 20 40 60 80
H (mmHg)
100 120 140 160
Figure 3.1: Flows (Q) created under different pressure heads (H) as function of pump rotation
speed. The fluid used in this experiment was filtered water at room temperature.
14

Inlet Pressure (mmHg)
Flow (m 3 /s)
−50
3 Results 15
3.1.2 Phase delay
In Figure 3.2 the results of the phase delay experiment are shown. The top graph shows the inlet
pressure. The mean pressure is negative just before the inlet of the pump. This shows the ability
of the pump to create negative pressure, which is of course unfavourable in the heart. The pressure
wave is not exactly a block due to damping and reflection of waves. This damping and reflections
are increased on the pressure measurement after the outlet of the pump. The time it takes for the
pressure wave to cross the pump and the 2 cm before and after the pump is about 6.3 ms. This
is measured looking at the point where the blocking wave starts. This is indicated with the two
red straight lines. The flow wave arrives about 40 ms after the pressure wave at the outlet. The
flow is measured 14 cm behind the site where the pressure is measured. Because the wave speed
is unknown it is only known that the delay is less than 40 ms. Compared to the cardiac cycle this
is of course very small.
0
−100
1 1.5 2 2.5 3 3.5 4
Outlet Pressure (mmHg)
50
40
30
20
10
1 1.5 2 2.5 3 3.5 4
1
0.5
0
−0.5
1 1.5 2 2.5
time (s)
3 3.5 4
Figure 3.2: Inlet pressure, Outletpressure and Flow vs. time showing the effects of a transient
blockage of the inlet flow caused by closing and opening the valve upstream of the pump.
15

3 Results 16
3.2 Acoustics
The left side of Figure 3.3 shows the spectrum of the sound of the Jarvik 2000. The rotation speed
is approximately 10000 revolutions per minute. The large peak at 170 Hz, the one which belongs
to the rotating frequency, is clearly visible. The harmonics are also clearly visible. This is also
visualized in the right side of Figure 3.3. One of the striking things is the area of the peak at
the 7 th harmonic (8 × rotation frequency). The peak at 270 Hz is also unexpected. This is most
likely due to pump-container interactions.
area
10 4
10 3
10 2
10 1
10000 rmp
10
0 5 10 15 20 25
0
Harmonic
Figure 3.3: Left: Spectrum of sound of the Jarvik 2000 made at a rotation speed of 10000 revolutions
per minute and a pressure head of 74 mmHg. Right: The area under the different harmonic
peaks.
The left side of Figure 3.4 shows 10 measurements made in vitro with slightly different stethoscope
positions. It shows that slight changes in microphone positions have a large impact on the
frequency spectrum. The right side shows the ratio between the area under the peaks of the harmonics
and the area under the ground frequency (rotation frequency). This shows that the ratio’s
between the peaks are really different and that there is not just an offset between the different
measurements. Without the minor changes in microphone position the spectra were all more or
less the same.
Area
10 5
10 4
10 3
10 2
10 1
10000 rmp
10
0 5 10 15 20 25
0
Harmonic
area / area of rotation frequency
10 0
10 −1
10 −2
10 −3
10 −4
10000 rmp
10
0 5 10 15 20 25
−5
Harmonic
Figure 3.4: Left: Ten measurements with slight differences in stethoscope position. Right: The
area under the peaks divided by the area under the rotation frequency.
16

3 Results 17
3.2.1 Inlet blockage
In this section the results from the inlet blockage experiments will be discussed. Figure 3.5 shows
the frequency spectrum of a 3
4 blockage (right) and no blockage (left). It is impossible to compare
the height of area under the peaks because the differences in spectrum caused by differences in
microphone position is too large. Comparing the two spectra in Figure 3.5 it becomes clear that
the only major difference is the amount of harmonic content in the high frequencies (more than
2000 Hz). Because of the influence of the microphone position it again is hard to draw conclusions.
3
Figure 3.5: Left: The frequency spectrum without blockage. Right: 4 blockage in placed in
configuration 1.
Figure 3.6 shows for 1
2 blockage that there are not many major differences between the frequency
spectrum of the two different configurations. Only for the low frequencies does the spectra show
some significant differences. These differences are not seen for the spectra of 1 3
4 and 4 blockage.
These differences can be caused by external sources or might be caused to the different microphone
positions and therefore no conclusions can be drawn.
Figure 3.6: Left: 1
2
ration 2.
blockage in placed in configuration 1. Right: 1
2
blockage in placed in configu-
17

3 Results 18
3.3 Clinical results
Figure 3.7 shows the frequency spectrum of the Jarvik 2000 measured in a patient. The small bell
is used at position 2, which is explained in section 2.5.1. The rotation frequency is clearly visible.
It can be seen that the rotation frequency is about 185 Hz, which corresponds to controller level 4,
and the harmonics of this frequency are also clearly visible. The main reason for the peaks being
not as sharp as with the experimental set-up is because of the heart is contracting and therefore
the rotation frequency will change due to pressure head changes (see section 3.3.2). Compared to
the results from the experimental set-up, the fifth harmonic appears to be stronger in the clinical
measurements. Furthermore, there are two peaks at frequencies that are not harmonics of the
rotation frequency: these peaks are at 125 and 240 Hz. Because this is a preliminary research,
the origin of these peaks is not investigated further. It is interesting to mention that the patient
measured had pain on the chest which was most likely to be caused by contact of the Jarvik 2000
with the ribs. This interaction might be an explanation for the extra peaks in the spectrum.
Figure 3.7: The frequency spectrum of the sound of the Jarvik 2000 measured in a patient.
3.3.1 Best microphone position
As described in section 2.5.1 to gain more insight in the best position to measure the sound of
the Jarvik 2000 in patients, measurements are made on three different locations. The results are
shown in the graphs below.
Figure 3.8 shows the difference between the big and small bell of the stethoscope. It is clear
that the small bell is better for measuring of the high frequencies. The big bell measures the low
frequencies better. The transition point seems to be a about 500 Hz with frequencies below the
500 Hz better measured with the big bell and higher frequencies with the small bell. For example,
the big bell will be more applicable for tracking the rotation frequency.
Figure 3.9 shows that on both positions on the second rib below the nipple the sound measurements
are sufficient to see the rotation frequency and harmonics. Furthermore, position 2 seems to be
somewhat better than position 1. This is especially true for the higher frequencies. A striking
difference is that the peak at 240 Hz is much clearer at position 2. This was for both the small
and big bell. When moving to the third rib (Figure 3.10) the signal attenuates considerably.
18

3 Results 19
Figure 3.8: Left: Big bell at position 2. Right: Small bell at position 2.
Figure 3.9: Left: Big bell at position 1. Right: Big bell at position 2.
Figure 3.10: Left: Small bell at position 1. Right: Small bell at position 3.
19

3 Results 20
3.3.2 Frequency shift due to heart contractions
Figures 3.11 and 3.12 show the results obtained from the patient who is recovering from the
Jarvik 2000 implantation. Measurements are made on the day of the implantation just after the
implantation (day 1) and one day after (day 2). The heart rate and blood pressures at the time
when the sound was measured are summarized in Table 1.
Day 1 Day 2
Diastolic pressure 50 73
Systolic pressure 50 81
Mean pressure 50 77
Heart rate 136 125
Table 1: Blood pressure and Heart rate on day 1 and day 2.
Here, the heart rate is measured only once, so it is just an indication of the actual heart rate at
a specific time. But again because of the preliminary nature of this study this is justified. As
expected the pulse pressure on the day of the implantation is much smaller than the pulse pressure
on the day after.
Figure 3.11 is a contour plot of the sound spectrum as a function of time and frequency. It clearly
shows the frequency shift caused by the pressure heads on the second day. The mean frequency
of the peak is at about the measured heart rate. The fact that the pulse pressure is only eight
mmHg shows the relative sensitivity of these frequency changes. Especially when comparing the
results of day 1 and day 2 (see Figure 3.13 and 3.12) the difference between no pulse pressure and
a pulse pressure of eight mmHg is obvious.
freq. (Hz)
185
180
175
170
165
160
155
10 10.5 11 11.5
time (s)
12 12.5 13
Figure 3.11: Contour plot of the power spectrum as a function of time and frequency on day 2.
20

freq. (Hz)
170
165
160
155
150
145
140
3 Results 21
freq. (Hz)
170
165
160
155
150
145
140
10 10.5 11 11.5 12 12.5 13 13.5
time (s)
Figure 3.12: Contour plot of the power spectrum as a function of time and frequency on day 1.
10 10.5 11 11.5 12 12.5 13 13.5
time (s)
freq. (Hz)
185
180
175
170
165
160
155
10 10.5 11 11.5
time (s)
12 12.5 13
Figure 3.13: The contour plots of day 1 (left) and day 2 (right) next to each other
21

4 Discussion and conclusions 22
4 Discussion and conclusions
This is a preliminary study to investigate several aspects of the Jarvik 2000. Although the preliminary
nature of this study did not allow it to draw quantitative conclusions, most results seem
promising and can be used as the basis for further research. Furthermore, the results shown here
are probably also applicable for other rotary pumps like the HeartMate II (Thoratec Laboratories
Corp.,Pleasanton, CA) and the DeBakey VAD (MicroMed Technology, Inc., Houston TX).
To obtain more physiological results for the H-Q relationship it will be better to use actual blood.
In this study this is not so important, but if the actual pump in the body is modeled (with for
example the Windkessel equation) the H-Q behavior of the Jarvik 2000 with blood must be known.
The result of the phase delay experiment which resulted in a delay of 40 ms obviously has to be
verified. More physiological waveforms have to be used. It seems that the clinically determined
87 ms was an overestimation. 87 ms was already presented as very small compared to the heart
cycle so the phase lag which is created by the Jarvik 2000 seems to be almost negligible compared
to the heart cycle.
This study also shows that it is possible to determine the rotation speed very accurately noninvasively.
This was the case for both in vitro and in vivo measurements. This is of course a
very important result in itself. Because the peak at the rotation frequency is much bigger than
all the other components of the spectrum, this frequency is also very easy to detect. As said
before, the peak at 270 Hz in vitro could be caused by the set-up but must, of course, be further
investigated. The peaks in the spectrum of the patient data, which cannot be attributed to the
rotation frequency should be further researched. The relatively large influence of minor changes in
microphone position is of course very inconvenient. The microphone position - frequency spectrum
relationship requires further investigation.
The inlet blockage does not seem to have a major effect on the audio signal recorded. If there are
changes in the frequency spectrum the changes are very subtle. There seems to be more frequency
content in the higher frequencies but this will be hard to quantify. Because of the influence of the
microphone position it is impossible to see any differences in the ratio between the peaks of the
harmonics like those that Trunzo et al. showed.
Another possible problem which can be detected by acoustic analysis is thrombosis on one of
the impeller blades. This will cause an imbalance of the impeller. This imbalance will cause
vibrations in the bearings which can be recorded by measuring the sound the Jarvik 2000 creates.
The housing of the Jarvik 2000, which was made available for this study was unwelded to enable
a solid mass to be attached to one of the impeller blades to cause an imbalance. The effect of
creating such an imbalance on the acoustic properties of the device should be subject of further
studies.
The best position measurements clearly shows that the small bell is better for the high frequencies
and the big bell better for the low frequencies. Furthermore, it showed that both in position 1
and 2 the rotation frequency and its harmonics are clearly visible. To confirm the best position
more measurements have to be made on different patients; both men and women and patients
with specific features such as obesity.
The results with the frequency shift due to heart contractions are most promising. Furthermore,
this method has the potential to become a diagnostic tool in a short time. For a diagnostic
tool this frequency shift due to pressure head transients has to be quantified. A possible way to
quantify this frequency shift is to create an experimental set-up in which different waveforms can
be created. When this quantification is completed this method has the potential to be a fast, easy
and cheap diagnostic tool.
22

5 Acknowledgements 23
5 Acknowledgements
First and foremost I like to thank professor Kim H. Parker for giving me the opportunity to do this
exciting project on the Jarvik 2000 at Imperial College London. I also wish to express my gratitude
for him always being there, willingly answering all my questions and for all the interesting/fun
stories (un)related to the subject. I also like to thank Dr. Christopher T. Bowles for introducing
me to the field of ventricular assist devices, for all his help and valuable advice. Finally, I wish to
emphasize the kindness of Dr. Robert Jarvik for providing a Jarvik 2000 for this research.
23

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