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Proceedings of the 2016 Bath/ASME Symposium on Fluid Power and Motion Control
FPMC2016
September 7-9, 2016, Bath, England
FPMC2016-1707
ACTIVE PNEUMATIC PULSATION DAMPER FOR PERISTALTIC PUMP FLOW
LOOPS
Matthias Liermann
Fluid-Mechatronics Lab
Department of Mechanical Engineering
American University of Beirut
Lebanon 1107-2020
Email: matthias.liermann@aub.edu.lb
ABSTRACT
A novel concept of an active pulsation damper is described
that cancels parasitic flow pulsatility of peristaltic pumps and is
able to inject desired pulsatility signatures such as physiological
heart beat. Peristaltic pumps avoid contact between the moving
parts of a pump and the operating fluid. They are used for clean
or sterile fluids as well as for highly aggressive fluids, whenever
it is important to isolate the fluid from the environment. The
background application for the proposed active pulsation damper
is the simulation of hemodynamic flow.
The paper presents a novel pulse damper concept that allows
the use of roller or peristaltic pumps as primary pumps for hemo-
dynamic flow loops. The problem with peristaltic pumps is that
they exhibit a high parasitic pulsatility that needs to be canceled
before a desired pulsatility can be injected. The active pulsation
damper does this and is also used to superpose a desired flow pat-
tern that resembles measured heart flow rate profiles. The non-
linear dynamic equations of a test system with active pulsation
dampers are established and linearized to allow a first analysis of
the achievable bandwidth. Simulation results of the closed loop
system are presented based on the non-linear equations.
1 Introduction
According to the World Health Organization WHO, cardio-
vascular diseases are the main cause of death worldwide [1]. A
large scientific community focuses on understanding the causes
and developing possible treatments for these ailments. Pumps
that replicate hemodynamic flow and pressure waveforms are
central to this type of research. They have been developed since
the mid 20th century and are today an integral part of labora-
tory research. Another class of systems simulating hemodynamic
flow is used for clinical treatment. They are used on patients,
for example, as mechanical circulatory support systems or for
dialysis. It has been found that natural pulsatility enhances per-
fusion [2, 3]. It is important for those systems to be extremely
reliable and to have low-cost disposables. They can be based on
roller pumps, bulb pumps [4] or more recently impeller pumps.
This paper gives a brief overview over both types of systems but
focuses on those that are used in laboratory research for testing
of implants and other experimental research such as flow visual-
ization studies. The presented novel concept for a hemodynamic
flow loop uses a peristaltic pump and an actively controlled com-
pliance chamber that cancels parasitic pulsatility and injects a
desired pulsatility profile. It is intended for precise laboratory
setups, where an accurate reproduction of a desired flow wave-
form is important. An application is, for example, to reproduce
flow and pressure signatures in various flow phantoms of the car-
diovascular system based on apriori measured flow profiles from
real patients. The requirements for such a device include: ac-
curate and reproducible volume flow waveform, wide range of
flow rates (even reversible flow), ease of programming, capabil-
ity of producing continuous flow, prevention of entrainment of
gas bubbles and cavitation, and low shear forces.
1 Copyright c
2016 by ASME
1.1 Hemodynamic Flow Devices in Literature
The literature provides a rich palette of existing devices that
address these requirements. Fig. 1 can serve as a guide to cat-
egorize each device presented in the following literature review.
Three categories can be distinguished based on how the pulsatile
flow is created. In the first category the pulsation is created just
like in the real heart: A volume contracts and expands, while the
displaced fluid is delivered through an arrangement of (check)
valves. The variable volume can be realized as an elastic bulb
that is pressurized from outside or by a motion-controlled pis-
ton/cylinder. In the second category of devices, a pump delivers
a constant flow, while a second device is used (either valve or
pump) that modulates the flow to inject the desired pulsatility. In
the third category a single pump is used, and through means of
motion control of the pump, both, the average baseline flow and
the pulsatility are realized.
Category A
Reciprocating pump & valves
Category B
Continuous pump + flow
modulator
Category C
Motion controlled pump
M
control
M
control
FIGURE 1: Categories of pulsatile pump systems
Each system has at its core a pump that can be categorized
by its pumping action principle, drive system, and its mode of
control to achieve the desired pulsatility. Fig. 2 lists various em-
bodiments found for these functions.
Category A A large area of need for pulsatile blood pumps is
for testing of heart implants. The design of such a device has
to make a compromise between the need to accurately simulate
the hemodynamic flow characteristics and being practicable for
routine laboratory use. Systems for the characterization of heart
valves have to simulate the contraction of the heart muscle. This
motion has been emulated in many setups with the bulb pump
principle, where the bulb resembles the heart chamber and is
caused to contract through application of outside pressure, see
left column of Fig. 1. In the Yoganathan-FDA system, pressur-
ized air was used while the Aachen pulse duplicator used pres-
surized liquid. Other setups have simulated the heart contraction
directly through piston pumping action (Sheffield pulse duplica-
tor) [5]. Similar to the Aachen pulse duplicator, the commer-
cial ViVitro Pulse Duplicator system developed by ViVitro Labs
Inc. in Canada also uses pressurized liquid applied to the outside
of a model heart to cause its pumping action [6]. It goes back to
a device developed in 1979 by Scotten et al to assess mitral valve
prostheses [7]. A commercial version is available since 2007 and
is today being used in many research projects. Other examples
for commercial pulsatile blood pumps are the Harvard Apparatus
Pulsatile Blood Pump or the BDC Laboratories HDT-500 Pulse
Duplicator. Nevertheless many research groups still build their
own test benches on the same principle with small variations, for
example [8, 9].
In many cases, such as in flow visualization studies in the
peripheral arterial system, it is not necessary to simulate the con-
tracting nature of the heart ventricle. The above described sys-
tems that use the bulb pump principle always require artificial
valves (mitral and aortic). A model heart is expensive to replace.
In addition, even though the pumping action is mechanically sim-
ilar to the natural heart, the downstream pulsatility is not neces-
sarily accurate, since it is also dependent on the compliance of
the downstream system. This compliance needs to be tuned with
devices such as the Vivitro device. When performing research
on parts of the cardiovascular system downstream of the aortic
valve, other flow loops are simpler that reproduce the pulsatility
not by a true mechanical representation but by controlled opera-
tion of one or more pumps.
Category B Many systems employ two pumps, where one is
used to provide a baseline flow, while the second is used to super-
pose a pulsatile flow, see middle column of Fig. 1. In [10] a roller
pump is used to provide a baseline flow, while the pulsatile flow
is simulated by a piston pump. Roller pumps or peristaltic pumps
are positive displacement pumps that cause minimum damage to
blood cells and can easily be cleaned, because the fluid is con-
fined by a tube. However, they are very pulsatile themselves,
which is undesired for the purpose of mimicking heart flow. In
section 3 of this paper, a measurement of such pulsatile flow is
presented. They have nevertheless been used. Law et al. [11] use
a modified roller pump with controlled stepper motor. The im-
plementation is limited to a specific output flow rate, and despite
modifications to the pump, pulsations from the pumping action
could not be eliminated. A gear pump is employed to generate
the baseline flow in [12]. Gear pumps cause significant noise in
the flow and pressure signal and should not be used for studies
with blood cells. In [13–15] a centrifugal pump is used to provide
the baseline flow because of its low noise signature compared to
a roller pump. A piston pump superposes the pulsatile flow. Var-
ious patents have been published that define features of pulsatile
pumps [16]. In [17] a pulsatile membrane pump device is de-
scribed that is actuated by predefined cam mechanisms that are
driven by a constant speed motor. Such a non-controlled system
2
displacement
gear, roller, membrane, progressive cavity, piston
hydrodynamic
impeller, centrifugal, etc
pump action
principle `
pump drive
system
electro-mechanic
brushed, brushless DC, AC, stepper, linear
fluidic
hydraulic/pneumatic
pulsation control position/angle/displacement control, velocity/flow control, force/pressure control, cam follower
parameter embodiment
FIGURE 2: Design parameters and embodiment of pulsatile pump
is possible when the baseline pump is of the displacement type
and therefore the flow pulsations are only affecting the down-
stream.
Category C Some systems try to emulate the pulsatile flow
with only a single pump which is electronically or mechanically
controlled in a way to regenerate the complete flow pattern, see
right column of Fig. 1. In [18] a pump is described that uses a
progressive cavity pump with a motion control system. Such a
pump exhibits a low noise signature compared to other displace-
ment pumps, such as roller pumps, but is also more complex
and less simple to clean and sterilize. Servo-driven piston pumps
can combine the advantage of simple components that can eas-
ily be cleaned, minimum pump noise and accurate controllabil-
ity. A double piston system is described in [19] that provides
non-interrupted pulsatile flow. It consists of two pistons that are
connected to a common shaft. While the first one extends and
delivers fluid to the system in a controlled manner, the other one
is retracted and charged with fluid from a reservoir. The piston
motion is controlled to realize a desired pulsatility. Piston pumps
do not cause shear load on blood particles and allow good repro-
duction of the pulsation profile.
In more recent studies, there have been attempts to achieve
pulsatile flow by means of speed modulation of an impeller
pump. However, it has been shown experimentally and theo-
retically that at near-healthy arterial pressure pulsation patterns,
strongly regurgitant flow occurs in the pump with several detri-
mental effects. It causes the impeller to operate in regions of
inferior efficiency, which increases energy consumption; further-
more, high shear levels that result from the impeller working in
regurgitant flow causes blood cell damage. To avoid regurgita-
tion with the impeller speed modulation method, the pulsation
profiles have to be properly planned and controlled [20,21].
1.2 Active Compliance Chamber
Very few studies, such as [10] (top middle in Fig. 1) use
a roller pump as a baseline pump because roller pumps exhibit
high parasitic pulsatility. This pulsatility has to be canceled first
using compliance chambers before a desired pulsatility can be
superposed by another pump. The system presented in this paper
uses a roller or peristaltic pump with a special compliance cham-
ber that is used to cancel parasitic and inject desired pulsatility
at the same time, see Fig. 3. This has not been done in previous
studies.
pSpV
control
p1Q1
fluid in
fluid out
pA
hA
FIGURE 3: Proposed solution: active compliance chamber
The compliance chamber is a flow-through type, also re-
ferred to as inline type pulsation damper. The setup is very sim-
ple, easy to clean, and the amount of fluid in the loop is small.
The pressure in the compliance chamber is modulated to accel-
erate and decelerate the downstream flow. A pneumatic control
valve is connected to the compliance chamber. The valve is con-
trolled by a controller that realizes the flow and pressure pulsatil-
ity based on an internal reference and measured signals of flow
and pressure in the downstream tube.
In the following, the paper describes the mathematical
model of a flow loop with active compliance chambers. The
frequency response of the system allows preliminary assessment
about the achievable bandwidth of the pulse control based on a
linearized model. Simulation results are presented that are ob-
tained with a nonlinear simuation model and a simple propor-
tional controller.
3
2 Mathematical Model
The flow loop with two active compliance chambers is illus-
trated in Fig. 4. The system consists of five main components:
compliance chambers A&B, pneumatic high response control
valve, flow loop test section and the pump. For the purpose of
analysing the dynamics of the active compliance chambers, and
their ability to inject desired oscillating flow patterns, it is not
important to include the pump in the model as long as it can
be assumed that it is a displacement pump with high input and
output impedance. That means that the output flow rate is lit-
tle affected by pressure pulsations in the system. The peristaltic
pump is therefore excluded from the following model, but in-
cluded later in the analysis. The pump delivers the same amount
of fluid that it consumes and introduces parasitic pulsatility into
the system. This pulsatility can be treated as a disturbance and
can later be superposed to the flow produced by the active flow
chambers.
pSpex
u
QPRHLH
pApB
hAhB
QL
mAmB
pL
A
test section of flow loop
control
pref
pA
FIGURE 4: Model schematic of flow loop with active compliance
chambers
The air volume of the compliance chamber can be modeled
with the assumption that no heat exchange takes place during fast
changes in pressure. With this assumption the change of internal
energy ˙
Uis equal to the enthalphy influx ˙
Hminus the work done
by the air volume ˙
W.
˙
U=˙
H−˙
W(1)
The change of internal energy depends on the change of mass mA
and temperature TA. Eliminating the temperature from ter term
by use of the ideal gas law gives:
˙
U=d
dt (mAcvTA) = ˙pAVA+pA˙
VA
γ−1(2)
The work done by the fluid is
˙
W=pA˙
VA,(3)
and the enthalpy influx is
˙
H=˙mAcpTAif ˙mA<0,
˙mAcpTAin if ˙mA>=0.(4)
where TAis the chamber temperature and TAin is the temperature
of the incoming gas from the valve. Assuming adiabatic flow
through the valve, the incoming temperature is
TAin =TSpA
pSγ−1
γ
(5)
In the following TqA is used as the temperature of the fluid enter-
ing or leaving the chamber. Combining Eqs.(2-4) and isolating
for the pressure gradient gives
˙pA=γ˙mARTqA
VA
−γpA˙
VA
VA
,(6)
where the change of chamber volume is the difference between
out- and inflowing volumetric flow rates
˙
VA=QL−QP.(7)
Chamber B is modeled accordingly.
The control valve modulates the mass flow in and out of
the chambers. It is nonlinearly dependent on the upstream and
downstream pressures, upstream temperature and valve opening.
Several ways exist to model this flow. Applying the model de-
scribed in the standard ISO 6358 [22] is useful when the respec-
tive model parameters are given in the valve data sheet, namely
the sonic conductance cand the critical pressure ratio b. Us-
ing the density of air at standard reference conditions ρ0and
T0=293◦K,the mass flow rate equation can be written as
˙mA=
ucρ0pSqT0
TSs1−pA
pS−b
1−b2
for u>0 and pA
pS>b,
ucρ0pSqT0
TSfor u>0 and pA
pS≤b,
ucρ0pAqT0
TAs1−pex
pA−b
1−b2
for u≤0 and pex
pA>b,
ucρ0pAqT0
TAfor u≤0 and pex
pA≤b.
(8)
4
The equation depends on the flow direction given by the valve
opening uand the ratio between downstream and upstream pres-
sure. The mass flow rate of chamber B is obtained similarly.
The flow loop test section is usually not very long (<1m)
and can be modeled with a hydraulic resistance and inductance.
It is assumed that the capacitance of the line is insignificant com-
pared to the capacitance of the compliance chambers and is there-
fore not modeled. The combination of resistance and inductance
can be modeled as a dynamic system of first order. The pressure
drop pA−pLcaused by resistance to the flow QLdepends on
whether it is laminar or turbulent. It can be modeled with the
Darcy-Weisbach equation [23, 24] as:
pA−pL=(QL128µll
πd4,Re <2300
Q2
L
0.3164·8·ρll
Re0.25d5π2,Re ≥2300 (9)
Eq. 9 is a coarse approximation. It does not attempt to describe
laminar- turbulent flow transition and is assuming smooth pipes
and stationary flow conditions. For the purpose of this study this
simplification is sufficient.
The pressure drop pL−pBdue to acceleration of the fluid is
modeled by a simplified inductance term as
pL−pB=lρl
A
˙
QL(10)
The parameters of the model are listed in Table 1.
3 Performance Analysis and Simulation results
The above described model is of at least 5th order. The state
variables are: the pressures pA,Band liquid volumes VA,Bof each
chamber, and the volumetric flow rate in the test section QL. All
other variables are algebraic variables. The valve spool dynam-
ics were neglected because the desired closed loop bandwidth is
much lower than the valve natural frequency, which is estimated
ωV=400Hz. The system can be linearized using the parameters
listed in Table 1. Because of the symmetry of the system it seems
appropriate to cancel the pressure in chamber B as a state. In the
linearized model the change of pressure in chamber B is always
equal in value but opposite in sign to the change of pressure in A.
The simplified linear model Gpcan be represented as a transfer
function with valve signal u/umax[−]as input and QL[m3/s]as
output.
Gp=1.619 ·105m3
s
s4+3519s3+6.316 ·106s2+2.218 ·106s+3.536 ·108
(11)
Represented as frequency response plot, the model is depicted
in Fig. 5. The plot shows only the relevant frequency range up
TABLE 1: List of model parameters
Symbol Comment Value Unit
A area of chamber 7.85 ·10−3m2
b critical pressure ratio 0.21 -
c valve sonic conductance 0.45 L
bars
h height of chamber 0.1 m
l length of test section 1 m
pSsource pressure 0.15 MPa
pex exhaust pressure 0.066 MPa
R ideal gas constant of air 287 J
kg·K
T0standard temperature 293 ◦K
TSpressure source temp. 293 ◦K
γheat capacity ratio 1.4 -
µ/mathrmg dynamic viscosity of air 1.8127 ·10−5Pa s
µldyn. viscosity of water 0.001 Pa s
ωVnatural freq. of valve 2513 rad
s
ρ0standard density of air 1.185 kg
m3
ρldensity of water 1000 kg
m3
ζVdamping ratio of valve 0.7−
to 15 Hz =100 rad
s. In this range the system appears simply as a
low damped second order system with a high resonance at around
1.2Hz =7.5rad
s. The damping ratio is ζ=0.02. The simplicity
of the model is a pleasant surprise.
The control loop is closed by taking the measurement of the
pulsatile flow QLand comparing it to the desired pulsatile flow
Qref. The difference e=Qref −QLis amplified by the controller
and applied as input uto the valve. A proportional-derivative
control is suitable for such a system. To prove the concept of the
system, a simple controller has been determined as
Gc=U(s)
E(s)=3.018 ·10−4+1789s.(12)
Fig. 6 shows the closed loop frequency response that shows
a very promising bandwidth of around 10Hz =63 rad
s.
The controller is tested in simulation with a nonlinear imple-
mentation of the model in Dymola using the equation based mod-
eling language Modelica. This model includes, beyond the men-
5
Magnitude (dB)
-120
-100
-80
-60
-40
100101102
Phase (deg)
-225
-180
-135
-90
-45
0
Frequency (rad/s)
FIGURE 5: Frequency response of plant with input signal uand
output signal QL
Magnitude (dB)
-8
-6
-4
-2
0
2
100101102
Phase (deg)
-90
-45
0
45
Frequency (rad/s)
FIGURE 6: Frequency response of closed loop with input signal
Qref and output signal QL
tioned equations in the previous section, a 2nd order dynamics
model for the motion of the valve spool (ωV=400Hz,ζV=0.7),
a limitation of the valve spool velocity, and the saturation of the
valve spool opening at 100%. A graphical representation of this
model is shown in Fig. 7.
The reference flow rate is a typical aortic flow at a heart rate
of 85bpm [25]. The average flow rate is 5.3·10−5m3
s=3.18 L
min .
The peristaltic pump is set to deliver this average flow rate. The
flowrate pulsatility of a peristaltic pump was measured by PIV
in [26, 27] and is given as input to the simulation. The pulsatility
pS
pA
pV
pA
QL
QP
uA
pB
pSpV
pB
QP
uB
time [s]
0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2
flow [l/min]
0
5
10
measured heart flow rate in aorta
time [s]
0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2
flow [l/min]
0
5
10
measured flow rate of peristaltic pump
control
-1
FIGURE 7: Graphical representation of Dymola/Modelica model
of hemodynamic flow loop
of the peristaltic pump is on a much higher frequency band and
can therefore be absorbed in the compliance chamber, while the
desired pulsatility can be injected by the control loop. Fig. 8
shows the two input signals used in the simulation.
time [s]
0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2
flow [l/min]
0
5
10
Measured heart flow rate in aorta
Qref
mean flow
time [s]
0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2
flow [l/min]
0
5
10
Measured flow rate of peristaltic pump
QP
FIGURE 8: Desired flow rate pulsatility [25] and pulsatility de-
livered from peristaltic pump
3.1 Simulation Results
Two simulation results are shown in the following. The first
simulation result shows the damping capability of the compli-
6
ance chamber, when the control is switched off, see Fig. 9. The
top diagram shows the pulsatile flow delivered from the peri-
staltic pump QP. The second diagram shows the flow in the test
section QL. It starts from zero initial conditions and is acceler-
ated to the average flow rate delivered by the peristaltic pump.
The passive compliance of the damping chambers smoothes out
the flow ripples effectively. After 3.5 s steady state flow condi-
tions can be observed. The 3rd and 4th diagram show the devel-
opment of pressures in the chambers and the height of the water
line. In both signals the small pertubations caused by the flow
ripples of the peristaltic pump can clearly be observed. At steady
state conditions there is a certain pressure drop that corresponds
to the resistance of the test section.
time [s]
0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5
flow [l/min]
0
5
QP
time [s]
0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5
flow [l/min]
0
5
10
QL
time [s]
0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5
pressure [MPa]
0.095
0.1
0.105
pA
pB
time [s]
0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5
height [mm]
48
50
52
hA
hB
FIGURE 9: Simulation results showing flows, pressures and
heights of liquid levels in chambers while control is switched
off.
The second simulation result, where the control is switched
on, is seen in Fig. 10. The plots show larger detail in the tempo-
ral resolution. The simulation time is chosen between 3.5−5s.
At this time, the large scale transients that could be seen in Fig. 9
have vanished. The top diagram compares the reference flow
Qref with the actual flow in the test section QL. The pump flow
QPis also shown. The top plot shows a good reproduction of the
desired flow pattern with some phase lag and overshoot. The sec-
ond diagram shows the valve input signal. The valve cannot open
more than 100%, therefore it is evident that the control demands
higher flow gains than the valve can produce. The size of the
valve corresponds to an available valve in our lab (Festo MPYE-
5-M5-010-B), which is planned to be used for experimental test-
ing in future work. It is the smallest valve of its product range,
time [s]
3.5 4 4.5 5
flow [l/min]
0
5
10 Qref
QP
QL
time [s]
0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5
u/umax [-]
-1
0
1u
time [s]
3.5 4 4.5 5
pressure [MPa]
0.05
0.1
0.15
pA
pB
time [s]
3.5 4 4.5 5
height [mm]
45
50
55
hA
hB
FIGURE 10: Simulation results showing flows, pressures and
heights of liquid levels in chambers while control is switched
on.
so the saturation effect that is seen in the simulation results does
not represent a general limitation of the working principle. Even
with the saturation, the match between desired and achieved pul-
satility is very promising. The 3rd and 4th diagram show the
pressure and waterline height variations in the chambers. It can
be seen that the average value of pressure in chamber A is still
above the average pressure of chamber B. The valve, however,
modulates the instanteneous pressure diference back and forth in
order to superpose the desired pulsatility. The heights of the wa-
terlines in chambers A and B shown in the bottom diagram are
slightly offset compared to Fig. 9. This is caused by a mismatch
between the average flows delivered by the peristaltic pump and
the reference flow profile. It can also be caused by the control er-
ror. The level of the compliance chambers should be monitored
by a sensor and a super-ordinate control loop should ensure that
neither is depleted.
4 Conclusion
This paper presents a proof-of-concept for a novel active
pneumatic damper concept that allows the use of peristaltic
pumps for the simulation of heart flow patterns in laboratory se-
tups. The peristaltic pump is the pump of choice for testing with
sterile or aggressive fluids because no pump components are in
contact with the fluid. However, due to its high pulsatility it is
rarely used and the literature shows that preference is often given
to other pumps that are much more difficult to clean and steril-
ize. For hemodynamic flow simulation, a peristaltic pump has to
be coupled with a compliance chamber that can cancel the flow
7
pulsations and another device that modulates a desired pulsatil-
ity. Examples for such systems have been found in the literature.
It was the research question of this study, whether the latter two
functions, the elimination of parasitic pulsatility and the gener-
ation of a desired flow pattern, could be combined in a single
device. A simulation study was used to answer this question.
An active pulsation damper is described. It has a characteris-
tic passive compliance that is effective in cancelling the parasitic
pulsatility of the pump. At the same time it is connected to a
controlled pressure supply by which its internal pressure can be
adjusted. A control loop is used to modulate the internal pressure
based on measurement of the actual flow in the test section. The
paper presents the nonlinear system equations and the linearized
model that is the basis of a proportional-derivative control de-
sign. The control is implemented in simulation with the non-
linear model. The passive pulsation cancelling properties of the
system are demostrated, as well as the ability to follow the heart
flow rate pattern for a typical 85bpm pulse. It is seen that the
valve is slightly undersized compared to the sizes of the cham-
bers and geometry of the test section. But it is also evident that
the bandwidth of the control loop is high enough to reproduce
the desired pulsatility.
The work is ongoing to implement the proposed concept.
The main anticipated challenge will be the measurement of the
the actual flow in the test section. A significant lag added in
the measurment is going to decrease the achievable bandwith of
the closed loop control. It will be tested as well, whether the flow
pulsations can simply be added by feed-forward control with pre-
generated valve signals and a superordinate control of the levels
of the compliance chambers.
5 Acknowledgements
The author would like to thank Dr. Ghanem Oweis for the
measurement of the peristaltic pump pulsatility.
Funding The author gratefully acknowledges the funding by
the American University of Beirut, University Research Board
for its support to conduct this research.
REFERENCES
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Organization. [updated 2013 Mar; cited
2013 Jun 22]. [Online]. Available:
http://www.who.int/mediacentre/factsheets/fs317/en/
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