Conference PaperPDF Available

Active Pneumatic Pulsation Damper for Peristaltic Pump Flow Loops

  • Danfoss Power Solutions, Neumunster, Germany


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 hemodynamic 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 pattern that resembles measured heart flow rate profiles. The nonlinear 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. Copyright © 2016 by ASME Country-Specific Mortality and Growth Failure in Infancy and Yound Children and Association With Material Stature Use interactive graphics and maps to view and sort country-specific infant and early dhildhood mortality and growth failure data and their association with maternal
Proceedings of the 2016 Bath/ASME Symposium on Fluid Power and Motion Control
September 7-9, 2016, Bath, England
Matthias Liermann
Fluid-Mechatronics Lab
Department of Mechanical Engineering
American University of Beirut
Lebanon 1107-2020
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
Category C
Motion controlled pump
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
gear, roller, membrane, progressive cavity, piston
impeller, centrifugal, etc
pump action
principle `
pump drive
brushed, brushless DC, AC, stepper, linear
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-
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
fluid in
fluid out
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.
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
test section of flow loop
FIGURE 4: Model schematic of flow loop with active compliance
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 ˙
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:
dt (mAcvTA) = ˙pAVA+pA˙
The work done by the fluid is
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
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
where the change of chamber volume is the difference between
out- and inflowing volumetric flow rates
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=293K,the mass flow rate equation can be written as
for u>0 and pA
TSfor u>0 and pA
for u0 and pex
TAfor u0 and pex
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 pApLcaused 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:
πd4,Re <2300
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 pLpBdue to acceleration of the fluid is
modeled by a simplified inductance term as
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
Gp=1.619 ·105m3
s4+3519s3+6.316 ·106s2+2.218 ·106s+3.536 ·108
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 ·103m2
b critical pressure ratio 0.21 -
c valve sonic conductance 0.45 L
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
T0standard temperature 293 K
TSpressure source temp. 293 K
γheat capacity ratio 1.4 -
µ/mathrmg dynamic viscosity of air 1.8127 ·105Pa s
µldyn. viscosity of water 0.001 Pa s
ωVnatural freq. of valve 2513 rad
ρ0standard density of air 1.185 kg
ρldensity of water 1000 kg
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
E(s)=3.018 ·104+1789s.(12)
Fig. 6 shows the closed loop frequency response that shows
a very promising bandwidth of around 10Hz =63 rad
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-
FIGURE 5: Frequency response of plant with input signal uand
output signal QL
Magnitude (dB)
Phase (deg)
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 5.3·105m3
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
time [s]
0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2
flow [l/min]
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]
measured flow rate of peristaltic pump
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]
Measured heart flow rate in aorta
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]
Measured flow rate of peristaltic pump
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-
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]
time [s]
0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5
flow [l/min]
time [s]
0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5
pressure [MPa]
time [s]
0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5
height [mm]
FIGURE 9: Simulation results showing flows, pressures and
heights of liquid levels in chambers while control is switched
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 3.55s.
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]
10 Qref
time [s]
0 0.5 1 1.5 2 2.5 3 3.5 4 4.5 5
u/umax [-]
time [s]
3.5 4 4.5 5
pressure [MPa]
time [s]
3.5 4 4.5 5
height [mm]
FIGURE 10: Simulation results showing flows, pressures and
heights of liquid levels in chambers while control is switched
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
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.
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... They are integral part of research on cardiovascular diseases and cures. There is a variety of design concepts for hemodynamic flow loops as discussed in [1]. The types of systems can be categorized in terms of flow source and type of flow modulation, see Fig. 1. ...
... In recent research at the American University of Beirut, a flow loop has been devised using a peri-M control M control FIGURE 1. Three categories for concepts of simulating physiological flow delivered from human heart [1] staltic pump and two in-house developed pulsation dampers for static flow conditions [2,3]. A further development of this setup is aimed at producing physiological flow with as little extra cost and complexity as possible. ...
... The new approach of the presented system is the combined use of one device to passively cancel parasitic pulsatility by an appropriate compliance chamber and to use that same compliance chamber to inject a desired pulsatility. A preliminary simulation study suggested that such a system may be realized [1]. This paper presents an experimental setup, an improved model and some measurement results. ...
Conference Paper
Full-text available
Hemodynamic flow loops are widely used for research on causes and cures of cardiovascular diseases. They replicate physiological blood flow pulsatility in vitro. Many different pump types exist for such flow loops. The variety of the flow loop types shows the lack of one concept that satisfies all requirements, which are ease of handling and sterilization, flexible and accurate realization of various profiles, low shear rate exerted on fluid, low amount of circulating fluid. This paper experimentally proves the concept of a new type of pulse damping/pulse generating device that can be used for flow loops that are operated with a peristaltic pump. The pulse generating device fulfills the double function of damping the undesired pulsatility of the peristaltic pump and injecting a desired pulsatility that replicates the flow profile delivered by the heart. The injection of the desired pulsatility is achieved by modulation of air pressure in a damping device. The experimental results show that it is possible to achieve the dual function in one device. An electromagnetic flow sensor provides the feedback for the air pressure control and a high-response flow control valve controls the pressure in the pulse damper/generator. The response time and accuracy of the sensor proved to be critical for achieving the objective. With the limitations of the relatively cheap components used for this functional prototype, the mean error in the flow rate signal could be kept below 10% for a simulated adult pulse rate of 60bpm.
... J o u r n a l P r e -p r o o f regulation valves, or diaphragms/elastomers [33,34]. ...
... Pulse dampeners are widely used to mitigate the interrupted pulsatile flow patterns which are a consequence of rollers in peristaltic pumps [34,36,40,41]. As part of this work, we compared the endogenous function of the peristaltic pump against our lab-built pulse dampeners and a commercial pulse dampener (Fig 2, SFig 1,2). ...
Full-text available
Endothelial cells (ECs) are the primary cellular constituent of blood vessels that are in direct contact with hemodynamic forces over their lifetime. Throughout the body, vessels experience different blood flow patterns and rates that alter vascular architecture and cellular behavior. Because of the complexities of studying blood flow in an intact organism, particularly during development, the field has increasingly relied on in vitro modeling of blood flow as a powerful technique for studying hemodynamic-dependent signaling mechanisms in ECs. While commercial flow systems that recirculate fluids exist, many commercially available pumps are peristaltic and best model pulsatile flow conditions. However, there are many important situations in which ECs experience laminar flow conditions in vivo, such as along long, straight stretches of the vasculature. To understand EC function under these contexts, it is important to be able to reproducibly model laminar flow conditions in vitro. Here, we outline a method to reliably adapt commercially available peristaltic pumps to study laminar flow conditions. Our proof-of-concept study focuses on 2-dimensional (2D) models but could be further adapted to 3-dimensional (3D) environments to better model in vivo scenarios, such as organ development. Our studies make significant inroads into solving technical challenges associated with flow modeling and allow us to conduct functional studies towards understanding the mechanistic role of shear forces on vascular architecture, cellular behavior, and remodeling in diverse physiological contexts.
... In this case, the output flow generated by the pump is highly pulsatile, which is undesirable if a fine control of the pressure/flow needs to be achieved. This behaviour can be attenuated by introducing an inline pulsation damper [15], although this adds complexity and cost to the system. These issues, when coupled with poor resolution in flow-rate control and hysteresis, make them unsuitable for applications such as medical actuators that require great precision. ...
With significant research focused on integrating robotics into medical devices, sanitary control of pressurizing fluids in a precise, accurate and customizable way is highly desirable. Current sanitary flow control methods include pinch valves which clamp the pressure line locally to restrict fluid flow; resulting in damage and variable flow characteristics over time. This paper presents a sanitary compression valve based on an eccentric clamping mechanism. The proposed valve distributes clamping forces over a larger area, thereby reducing the plastic deformation and associated influence on flow characteristic. Using the proposed valve, significant reductions in tube damage (up to 96%) and flow-rate error (up to 98%) were found, when compared with a standard pinch valve. Additionally, an optimization strategy presents a method for improving linearity and resolution over the working range to suit specific control applications. The valve efficacy has been evaluated through controlled testing of a water jet propelled low-cost endoscopic device. In this case, use of the optimized valve shows a reduction in the average orientation error and its variation, resulting in smoother movement of the endoscopic tip when compared to alternative wet and dry valve solutions. The presented valve offers a customizable solution for sanitary control of fluid driven actuators.
... It takes dedicated control algorithms to faithfully mimic the physiological waveform. 3,19,34 Hemodynamic wall shear stress modulates the biological response of the endothelium,. 8,32,42,44 Its timeaveraged component is not the only relevant factor in cellular response, but also the spatial, 14 and temporal variations. ...
Purpose Peristaltic pumps (PP) are favored in flow bioreactors for their non-contact sterile design. But they produce pulsatile flow, which is consequential for the cultured cells. A novel pulse damper (PD) is reported for pulsatility elimination. Methods The PD design was implemented to target static pressure pulsatility and flow rate (velocity) pulsatility from a PP. Damping effectiveness was tested in a macro-scale, closed-loop recirculating bioreactor mimicking the aortic arch at flow rates up to (4 L/min). Time-resolved particle image velocimetry was used to characterize the velocity field. Endothelial cells (EC) were grown in the bioreactor, and subjected to continuous flow for 15 min with or without PD. Results The PD was found to be nearly 90% effective at reducing pulsatility. The EC exposed to low PP flow without PD exhibited distress signaling in the form of increased ERK1/2 phosphorylation (2.5 folds) when compared to those exposed to the same flow with PD. At high pump flow without PD, the cells detached and did not survive, while they were perfectly healthy with PD. Conclusions Flow pulsatility from PP causes EC distress at low flow and cell detachment at high flow. Elevated temporal shear stress gradient combined with elevated shear stress magnitude at high flow are believed to be the cause of cell detachment and death. The proposed PD design was effective at minimizing the hemodynamic stressors in the pump’s output, demonstrably reducing cell distress. Adoption of the proposed PD design in flow bioreactors should improve experimental protocols.
Full-text available
Peristaltic pumps are used in a wide variety of applications due to their tightness, ease of mainte-nance and accurate delivery. Nevertheless, the use of peristaltic pumps is limited by their disad-vantages: short service life of the working body and uneven feed. This work provides an overview of the existing design solutions for pumps. The main advantages and disadvantages of the most common modern designs of peristaltic pumps are considered. The developed design solutions are presented. These solutions are designed to extend the service life of the elastic working body of the pump. These include a spiral hose design, where hose life is improved by reducing the number of cyclic compressions using just one roller. Another solution is to operate the pump with incomplete compression of the working element, which reduces the stress values and thereby prolongs the ser-vice life of the working element. The special shapes of protrusions in the compression area were developed in order to compensate the decrease in flow caused by the operation of the pump with incomplete compression of the working member. The paper provides an overview of solutions to reduce the uneven flow of a peristaltic pump. The simplest of these is the use of multiple parallel channels. In other designs, the elimination of flow pulsations is achieved with a pneumatic damper. There is also a constructive solution, in which a special algorithm of actuation of five squeeze ele-ments is used for uniform supply, each of which compresses only its own section of the pump working body. Based on the analysis, it is shown that in order to eliminate the disadvantages of per-istaltic pumps the various methods are used. Nevertheless, those methods need further improve-ment.
Conference Paper
Cardiovascular disease is modern-day plague with a vast number of lives claimed, and an enormous socio-economic cost incurred. Hemodynamics of the cardiovascular system play an important mechanistic role in disease development. For instance, atherosclerotic plaque depositions are often correlated with regions of turbulent flow patterns and disturbed hemodynamic shear stress. A simplified, rigid, in vitro, flow model of a real-size aortic arch is described. The flow in the arched vessel is attached and healthy at the outer curvature, while it is separated and disturbed at the inner curvature wall, which is an ideal setting to study cardiovascular disease. Endothelial cells can be cultured on the lumen of the aortic arch model under controlled flow conditions and extracted from the inner and outer curvature walls for biochemical signaling studies. The flow velocity field in the model is characterized using particle image velocimetry PIV which allows for the estimation of the wall shear stress. This helps in correlating the underlying hemodynamics to the biomechanical response of the endothelium.
Full-text available
It is a known fact that blood flow pattern and more specifically the pulsatile time variation of shear stress on the vascular wall play a key role in atherogenesis. The paper presents the conception, the building and the control of a new in vitro test bench that mimics the pulsatile flows behavior based on in vivo measurements. An in vitro cardiovascular simulator is alimented with in vivo constraints upstream and provided with further post-processing analysis downstream in order to mimic the pulsatile in vivo blood flow quantities. This real-time controlled system is designed to perform real pulsatile in vivo blood flow signals to study endothelial cells' behavior under near physiological environment. The system is based on an internal model controller and a proportional-integral controller that controls a linear motor with customized piston pump, two proportional-integral controllers that control the mean flow rate and temperature of the medium. This configuration enables to mimic any resulting blood flow rate patterns between 40 and 700 ml/min. In order to feed the system with reliable periodic flow quantities in vivo measurements were performed. Data from five patients (1 female, 4 males; ages 44-63) were filtered and post-processed using the Newtonian Womersley's solution. These resulting flow signals were compared with 2D axisymmetric, numerical simulation using a Carreau non-Newtonian model to validate the approximation of a Newtonian behavior. This in vitro test bench reproduces the measured flow rate time evolution and the complexity of in vivo hemodynamic signals within the accuracy of the relative error below 5%. This post-processing method is compatible with any real complex in vivo signal and demonstrates the heterogeneity of pulsatile patterns in coronary arteries among of different patients. The comparison between analytical and numerical solution demonstrate the fair quality of the Newtonian Womersley's approximation. Therefore, Womersley's solution was used to calculate input flow rate for the in vitro test bench.
Full-text available
Numerical models of the mitral valve have been used to elucidate mitral valve function and mechanics. These models have evolved from simple two-dimensional approximations to complex three-dimensional fully coupled fluid structure interaction models. However, to date these models lack direct one-to-one experimental validation. As computational solvers vary considerably, experimental benchmark data are critically important to ensure model accuracy. In this study, a novel left heart simulator was designed specifically for the validation of numerical mitral valve models. Several distinct experimental techniques were collectively performed to resolve mitral valve geometry and hemodynamics. In particular, micro-computed tomography was used to obtain accurate and high-resolution (39 μm voxel) native valvular anatomy, which included the mitral leaflets, chordae tendinae, and papillary muscles. Three-dimensional echocardiography was used to obtain systolic leaflet geometry. Stereoscopic digital particle image velocimetry provided all three components of fluid velocity through the mitral valve, resolved every 25 ms in the cardiac cycle. A strong central filling jet (V ~ 0.6 m/s) was observed during peak systole with minimal out-of-plane velocities. In addition, physiologic hemodynamic boundary conditions were defined and all data were synchronously acquired through a central trigger. Finally, the simulator is a precisely controlled environment, in which flow conditions and geometry can be systematically prescribed and resultant valvular function and hemodynamics assessed. Thus, this work represents the first comprehensive database of high fidelity experimental data, critical for extensive validation of mitral valve fluid structure interaction simulations.
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This study presents a numerical simulation of cardiovascular response in the heart failure condition under the support of a Berlin Heart INCOR impeller pump-type ventricular assist device (VAD). The model is implemented using the CellML modelling language. To investigate the potential of using the Berlin Heart INCOR impeller pump to produce physiologically meaningful arterial pulse pressure within the various physiological constraints, a series of VAD-assisted cardiovascular cases are studied, in which the pulsation ratio and the phase shift of the VAD motion profile are systematically changed to observe the cardiovascular responses in each of the studied cases. An optimization process is proposed, including the introduction of a cost function to balance the importance of the characteristic cardiovascular variables. Based on this cost function it is found that a pulsation ratio of 0.35 combined with a phase shift of 200° produces the optimal cardiovascular response, giving rise to a maximal arterial pulse pressure of 12.6 mm Hg without inducing regurgitant pump flow while keeping other characteristic cardiovascular variables within appropriate physiological ranges.
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Since most complications related to the operation of prosthetic heart valves is due to disturbances of flow, its hydrodynamic characterization is a useful aid in the design of new prostheses. Simulations of pulsatile flow in cardiac prostheses began nearly 40 years ago, through the development of different mock human circulatory systems, improving the clinical results interpretation. A new design of a pulse duplicator system was developed at Polytechnic School of USP to study prosthetic heart valves. To present the conception of a new mock circulatory system for hydrodynamic simulations of cardiac prosthetic valves and the assembly plan of an experiment whose focus is the test of mitral prosthesis. Its conception is based on the state-of-art's review of these studies and the experience got with the previous mock circulatory systems, particularly the one used in the Instituto Dante Pazzanese de Cardiologia, São Paulo, SP, Brazil. In this design, an electric servomotor controlled by computer emits, through a hydraulic piston, a pulse to the left ventricular chamber model, where the heart valves are accomodated. To characterize, in the future, the dynamic operation of mitral prosthetic valves, an experimental setup was mounted to provide measurements of volumetric flow, instantaneous pressure and velocity fields on these valves. Optical access is conveniently provided on the design, making possible the use, in the future, of a LDA system. In order to improve the analysis of hydrodynamic shear stress and prediction of haemolysis, the experimental results may be used to regulate a numerical model using 'Computational Fluid Dynamics' (CFD).
Full-text available
A pulsatile flow pumping system is developed to replicate flow waveforms with reasonable accuracy for experiments simulating physiological blood flows at numerous points in the body. The system divides the task of flow waveform generation between two pumps: a gear pump generates the mean component and a piston pump generates the oscillatory component. The system is driven by two programmable servo controllers. The frequency response of the system is used to characterize its operation. The system has been successfully tested in vascular flow experiments where sinusoidal, carotid, and coronary flow waveforms are replicated.
Full-text available
A hydromechanical simulator of the human left heart and systemic circulation system has been built for evaluating the dynamic actions of mitral valve prostheses. The facility includes a transparent and compliant model ventricle pumping into a simulation of the systemic circulation impedance. Transparent materials permit detailed high speed photography of valve action from the left ventricular aspect. This action has been analyzed by digitizing the valve open area and computing the variation of area with time. Additionally, pressures across and flows through the valves have been measured. Comparison has been made between the maximum observed valve open areas and the areas computed using the appropriate pressure and flow measurements substituted in clinically used equations.
Full-text available
A computer-controlled pump for use both in the study of vascular haemodynamics and in the calibration of clinical devices which measure blood flow is designed. The novel design of this pump incorporates two rack-mounted pistons, driven into opposing cylinders by a micro-stepping motor. This approach allows the production of nearly uninterrupted steady flow, as well as a variety of pulsatile waveforms, including waveforms with reverse flow. The capabilities of this pump to produce steady flow from 0.1 to 60 ml s-1, as well as sinusoidal flow and physiological flow, such as that found in the common femoral and common carotid arteries are demonstrated. Cycle-to-cycle reproducibility is very good, with an average variation of 0.1 ml s-1 over thousands of cycles.
Conference Paper
Pulsatile pressure/flow wave forms reproduction of blood in mechanical circulatory systems are still an open topic. Regarding the periodic behavior of pulsatile hemodynamics, a repetitive control algorithm was adopted as a potential methodology for rotary blood pumps. The developed algorithm was tested on a mock system including an oxygenator, a resistance, and a compliance. The post-oxygenator pressure served as the feedback of the control system. Initially, a model of the whole system was developed in order to use repetitive control algorithm. Then the performance of the developed algorithm was evaluated in three different scenarios. The experimental results indicated that the proposed method was able to accurately reproduce any pattern of pulsatile pressure. Moreover, it demonstrated an acceptable robustness in terms of model uncertainty and nonlinearity.
This exciting new reference text is concerned with fluid power control. It is an ideal reference for the practicing engineer and a textbook for advanced courses in fluid power control. In applications in which large forces and/or torques are required, often with a fast response time, oilhydraulic control systems are essential. They excel in environmentally difficult applications because the drive part can be designed with no electrical components, and they almost always have a more competitive power–weight ratio than electrically actuated systems. Fluid power systems have the capability to control several parameters, such as pressure, speed, and position, to a high degree of accuracy at high power levels. In practice, there are many exciting challenges facing the fluid power engineer, who now must have a broad skill set.
The ability to reproduce wave forms of the human arterial blood flow in vitro is essential for the study of vascular haemodynamics. The arbitrary flow generator described in this article is designed for this application. The flow generator consists of a progressive cavity pump driven by a computer-controlled servo motor. The chosen pump type offers advantages such as good linearity, high load tolerance and mechanical robustness. The flow generator is capable of producing noninterruptible standard and physiological wave forms. Standard wave forms include sinus, triangular, square and pulse. Three physiological wave forms are predefined: carotid, femoral and aorta. The performance of the flow generator is evaluated in detail. Constant flow up to 5 l/min is generated with high accuracy (<1%) and low ripple (−50 dB). The output is virtually insensitive to an increased flow resistance. A back pressure of 0.5 bar results in an output loss of less than 2.5 ml/min. The flow generator has an output bandwidth (−3 dB) of 11 Hz and a maximum slew rate of 2210 ml/s2. No significant difference (p<0.001) was found in the motor regulation during load. The achieved performance makes the flow generator very suitable for its application. © 2000 American Institute of Physics.