Conference Paper

Design of a blood pump for a wearable artificial kidney device

DOI: 10.1109/ECCE.2012.6342308 Conference: Energy Conversion Congress and Exposition (ECCE), 2012 IEEE
ABSTRACT
The aim of a European project Nephron+ is design of a wearable artificial kidney device (WAKD). This paper is focused on design of the corresponding ultralow-hemolysis continuous-operation blood pump. Accurate specifications and operating principle of the pump are determined; a first, non-optimal configuration of an electro-mechanical actuator which will be used to pump the blood is designed; its prototype is presented along with the corresponding driving electronic circuit; finally, based on the measurements, the actuator is optimized, and the final design and first results are presented.

Full-text

Available from: Yves Perriard, Dec 03, 2014
IEEE TRANSACTIONS ON INDUSTRY APPLICATIONS, VOL. 49, NO. 5, SEPTEMBER/OCTOBER 2013 2053
Design and Optimization of A Blood Pump for
A Wearable Artificial Kidney Device
Miroslav Markovic, Michael Rapin, Marc Correvon, and Yves Perriard
Abstract—The aim of the European project Nephron+ is the
design of a wearable artificial kidney device. This paper is focused
on the design of the corresponding ultralow-hemolysis continuous-
operation blood pump. Accurate specifications and operating
principle of the pump are determined. A first nonoptimal configu-
ration of a linear electromechanical actuator which will be used to
pump the blood is designed. Its prototype is presented along with
the corresponding driving electronic circuit. Finally, based on the
measurements, the actuator is optimized, and the final design and
first experimental results are presented.
Index Terms—Actuators, efficiency, hemolysis, medical control
systems, microcontroller, pumps.
I. INTRODUCTION
T
HE kidneys serve the body as a natural filter of the blood,
as they remove wastes which are diverted to the urinary
bladder [1]. In producing urine, the kidneys excrete wastes
such as urea and ammonium, and they are also responsible for
the absorption of water, glucose, and amino acids, keeping the
body’s internal equilibrium of water and minerals.
A. Dialysis
In the case of kidney disease or failure (which is referred
to as nephropathy), an artificial replacement should be made,
which means that an external device should remove waste from
the blood. This artificial filtering process is called dialysis. It
is regarded as a “holding measure” until a renal transplant can
be performed (or, sometimes, as the only supportive measure
in patients for whom a transplant would be inappropriate).
However, this “holding” function should often last for a long
time, as the need for a kidney is much higher than the offer: In
Switzerland in 2009, the mean waiting time was 700 days [2].
Manuscript received October 2, 2012; revised November 22, 2012; accepted
January 3, 2013. Date of publication May 3, 2013; date of current version
September 16, 2013. Paper 2012-EMC-509.R1, presented at the 2012 IEEE
Energy Conversion Congress and Exposition, Raleigh, NC, USA, September
15–20, and approved for publication in the IEEE T
RANSACTIONS ON INDUS-
TRY APPLICATIONS by the Electric Machines Committee of the IEEE Industry
Applications Society.
M. Markovic and Y. Perriard are with the Integrated Actuators Labora-
tory (LAI), School of Engineering (STI), Ecole Polytechnique Federale de
Lausanne, 1015 Lausanne, Switzerland (e-mail: yves.perriard@epfl.ch).
M. Rapin and M. Correvon are with the Swiss Center for Electronics and
Microtechnics (CSEM), 2002 Neuchatel, Switzerland (e-mail: michael.rapin@
csem.ch; marc.correvon@csem.ch).
Color versions of one or more of the figures in this paper are available online
at http://ieeexplore.ieee.org.
Digital Object Identifier 10.1109/TIA.2013.2260851
The main type of the dialysis is the hemodialysis. It con-
sists in removing the waste by circulating blood outside the
body through an external filter, called a dialyzer, that contains
a semipermeable membrane. The patient’s blood is pumped
through the blood compartment of a dialyzer. The blood flows
in one direction through the membrane, and a special liquid,
called dialysate, flows in the opposite direction. The countercur-
rent flow of the blood and dialysate maximizes the concentra-
tion gradient of solutes between the blood and dialysate, which
helps to remove more waste from the blood.
The main problem with dialysis is that the patient needs to
be placed in the hospital and that it is typically performed three
times per week, with 4 h for each treatment. In addition, each
treatment requires 120 L of dialysate. All this is, of course,
related to a huge cost of treatment (typically, in Switzerland,
80 000 euros per patient per year).
B. Envisaged Artificial Kidney
In order to significantly improve the patient’s quality of life
and reduce the costs, the goal of the European project Nephron+
[3], which s tarted in April 2010 and will last for four years, is to
design a wearable device to replace the bulky dialyzer. Its main
property will be the continuous blood filtering, which is much
more beneficial for t he patient’s health, with a significantly
reduced quantity of dialysate. The participants in the project
are, among many others, the Swiss Center for Electronics and
Microtechnics (CSEM) from Neuchatel, Switzerland, and the
Ecole Polytechnique Federale de Lausanne, Switzerland.
The device is to be supplied from batteries and carried by
the patient. Hence, the device can be regarded as a wearable
artificial kidney. This will be the first device of this type at the
market.
C. State of the Art
Many different studies have been done on wearable or im-
plantable rotary blood pumps for ventricular assistance devices,
which may be either centrifugal [4], [5] or axial [6]–[8].
They show that the required blood flow and pressure drop are
different from the requirement of a blood pump needed for
a hemodialysis system. Indeed, the blood flow for a kidney
device is usually more than 30 times lower. Also, the required
pressure in standard utilization is two times higher. Due to these
important differences, the mentioned rotary pump does not suit
for our use. Even if the rotary blood pump may reach the
required pressure drop for an artificial kidney, it would require a
0093-9994 © 2013 IEEE
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2054 IEEE TRANSACTIONS ON INDUSTRY APPLICATIONS, VOL. 49, NO. 5, SEPTEMBER/OCTOBER 2013
Fig. 1. Principle of operation of the artificial kidney device.
higher rotational speed, which would significantly increase the
hemolysis rate.
Lee et al. studied a valveless pulsatile blood pump [9]: It
is shown that this kind of pump provides better results than
conventional units in terms of hemolysis, which is an extremely
important characteristic in continuous extracorporeal blood
circulation.
Other candidates for the pump are also analyzed. The elec-
troosmotic pumps can generate the flow rate which is too small
for the presented application, and in addition, they require a too
high voltage (typically 100 V). The syringe-type pumps may
generate clotting problems.
Finally, the references which served as inspiration for the
presented project are [10]–[14]. They present various config-
urations of pumps, for blood and other liquids. Reference [15]
presents the design of a pump for a portable renal replacement
system. The principle is different from that which will be
presented in this paper: Several cams sequentially compress
fingers, which compress flexible tubes, thus eliminating valves.
The pump volume of 150 cm
3
and the achieved flow rate of
100 mL/min are similar to the performance which will be
presented in this paper.
II. A
RTIFICIAL KIDNEY AND ITS BLOOD PUMP
A. Principle of Operation
The principle of operation of the artificial kidney device is
shown in Fig. 1. It is a hemodialysis machine, in which the
patient’s blood is continuously, using a blood pump, circu-
lated through a semipermeable extracorporeal membrane and
returned to the patient (blood circuit in red in Fig. 1).
The opposite side of the same membrane is washed with
an electrolyte solution (dialysate) containing the normal con-
stituents of plasma water. Diffusion (osmotic pressure) and con-
vection (hydrostatic pressure) are the two mechanisms causing
a flow of water and dissolved substances from the blood to the
dialysate compartment.
A special pump is used to circulate the dialysate (dialysate
circuit in blue in Fig. 1). The undesired molecules are removed
from the dialysate using a filter with nanoparticles in a sorbent
unit.
The aim of this paper is to show the procedure of the design
of the blood pump. For cost reasons, the pump must be reusable
without excessive maintenance.
B. Blood Pump Specifications
The mean blood pressures in arteries and in veins are P
a
=
100 mmHg and P
v
=40mmHg, respectively (750 mmHg =
100 kPa). In addition, the semipermeable membrane will be
opposed to the blood flow, which means that it will gener-
ate a pressure drop. The corresponding value is estimated to
P
m
= 450 mmHg. Hence, the blood pump should compen-
sate the total pressure drop, which is therefore estimated to
P
p
= P
m
+ P
v
P
a
= 390 mmHg =52 kPa. Note that the
positive difference P
a
P
v
will “help” the pump during its
operation.
The maximal necessary flow rate of the blood is estimated to
200 mL/min, which means that Q
p
=3.3 × 10
6
m
3
/s. The
necessary mechanical power to be provided by the pump is
hence P
p
Q
p
=0.172 W. If we suppose the efficiency of 10%,
the necessary electrical power to be consumed by the pump i s
approximately 1.7 W. To complete the specifications, the pump
volume and mass should not overpass 140 cm
3
and 400 g,
respectively.
C. Blood Pump Additional Requirements
Apart from the “mechanical” specifications concerning the
pumping of blood, some other considerations should be taken
due to the specificity of the blood. At first, blood is a compli-
cated fluid. It contains red blood cells (RBCs) which contain
the hemoglobin.
The RBC can be easily destroyed by one of the follow-
ing three factors: mechanical force, sudden variation of the
pressure, or the temperature. Concerning the first two factors,
their critical values are not easy to estimate. Concerning the
temperature, it should not overpass 45
C[16].
The process of RBC destruction is called hemolysis: When
an RBC is destroyed, the hemoglobin is freed. It means that
the increased concentration of hemoglobin means the increased
hemolysis. The normalized index of hemolysis (NIH) is a
clinical measure of hemolysis measured as the concentration of
hemoglobin in the blood. Its value should not overpass 100 mg
per 100 L of pumped blood.
In addition, if the blood stays immovable during a short
time, there is a risk of thrombosis (formation of blood clot).
Hence, the geometry of the pump should be made taking
this fact into consideration: It should not contain “corners” in
which the blood can stay without moving. All the mentioned
requirements are too difficult to take into account during the
pump design. Only the measurements on the pump prototype
will give responses if those requirements are met.
Finally, the pump should be “bloodtight” which means that
no blood can escape and (equally important) that the air cannot
enter. An air bubble can have catastrophic consequences for the
patient [17].
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MAR KOVI C et al.: DESIGN AND OPTIMIZATION OF A BLOOD PUMP FOR A WEARABLE ARTIFICIAL KIDNEY DEVICE 2055
Fig. 2. Pump configuration and its operation in four steps [13].
III. PUMP CONFIGURATION AND OPERATION
A. Pump Configuration
After a preliminary study, it is decided that an optimal pump
configuration for this application is a linear peristaltic pump
with two elastic tubes, as shown in Fig. 2. As it is peristaltic,
it enables to know the blood flow rate without a sensor which
significantly improves the system reliability and cost.
The pump contains four actuators (A, B, C, and D) which
move in the vertical direction. Using two tubes, when the
actuator is to switch from one to another end (extreme) vertical
positions, the elastic force of the closed tube will help the
movement toward the position where another tube should be
closed.
It is also important to point out that the actuator couples A
and C, and B and D, operate simultaneously and in the opposite
direction: When one from the couple goes up, another goes
down and vice versa.
In the closed state, the tube opening will be zero (tube
completely pinched). In the open state, the tubes will not be left
completely without external force: A remaining force will keep
the tube opening equal to 50% of its diameter in the free state.
This will help the tube to restore its shape during the refilling
at the inlet of the pump, making the pump less sensitive to inlet
pressure.
B. Actuator Specifications
The chosen tube is made of silicone elastomer and produced
by Maagtechnic. Its internal diameter is 8 mm; the wall thick-
ness is 1 mm. This industrial tube does not fulfill the medical
norms (at first hemocompatibility), but it will be used for the
pump first prototype. The final prototype is intended to be made
using the tubes made of Tygon 3350 by Saint-Gobain (which
fulfills the norms, more precisely ISO 10993 guidelines for
contact with blood), but due to nonstandard dimension, it has
a delivery time too long to wait for the first prototype.
The elastic force of the chosen tube is experimentally mea-
sured. Finally, the necessary profile of the force which one
actuator should generate is given in Fig. 3. The position values
Fig. 3. Necessary force that should be generated by one actuator.
of +2/0/ 2 mm correspond to high-end/mean/low-end posi-
tions, respectively.
The force is shown for the actuators B and C from Fig. 2,
as they should generate higher force than A and D. Indeed,
suppose that the actuator is in the low-end position and starts
its moving toward the high-end position (such as B between
steps 3 and 4). It is the low pressure (and force F
lp
)inthelower
tube which helps the movement and the high pressure (and force
F
hp
) in the higher tube which is opposed to the movement.
Therefore, the difference F = F
hp
F
lp
is shown in Fig. 3 as
the force F which should be generated by the actuator over the
stroke of 4 mm.
After the calculations which take into account the tube di-
mensions and the necessary blood flow rate, it is obtained that
each actuator should press the tube at its length of 12 mm with
the frequency of 2 Hz (so the sequence in Fig. 2 repeats twice
each second).
IV. A
CTUATOR PRELIMINARY DESIGN
A. Actuator Configuration
Using electromagnetic actuators is the best compromise
between simplicity of construction, power consumption, and
maintenance [14]. Four candidates (1, 2, 3, and 4) are taken into
account for the actuator configuration (Figs. 4 and 5). Each of
them is analyzed using a simple analytical model and a finite-
element model (FEM) commercial software package.
After a quick preliminary analysis (the results will not be
presented here), the configurations 1 and 2 are eliminated, as
their constant force profile along the stroke does not correspond
to the force requirement for the actuator (Fig. 3). Thus, in order
to generate a significant force in extreme positions, those actu-
ators should have a huge volume (typically twice the volume
required by specifications). The configurations 3 and 4 have an
advantage of having an important reluctant force at the end of
the stroke, which allows maintaining the tube closed without
current.
Finally, after the first calculations, it turned out that the
configuration 3 is not suitable, as its two windings require
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2056 IEEE TRANSACTIONS ON INDUSTRY APPLICATIONS, VOL. 49, NO. 5, SEPTEMBER/OCTOBER 2013
Fig. 4. Actuator configurations (left) 1 and (right) 2.
Fig. 5. Actuator configurations (left) 3 and (right) 4.
significant volume and mass. Hence, the configuration 4 is
chosen, which contains one winding.
B. Actuator Modeling
The actuator geometry with all its design parameters is
shown in Fig. 6. The electromagnetic analytical model is based
on the analysis of the equivalent magnetic circuit represented
in Fig. 7. The force is determined by the derivation of magnetic
energy W
mag
by using the formula F
y
= dW
mag
/dy, with y
as the mover vertical position. The iron is assumed to be linear;
this assumption is provided by the optimization process as it
will be shown later: The magnetic flux density is limited so that
the iron stays far from the saturation.
In Fig. 7, Λ
Iron 1
, Λ
Iron 2
, and Λ
Iron 3
represent the reluc-
tances of the iron; Λ
i
is the internal reluctance of the magnets;
Λ
h
and Λ
l
represent the (variable) reluctances of the higher
and the lower air gaps, respectively; Θ
m
and Θ
coil
are the
magnetic potentials (magnetomotive forces) of the magnets and
the coil, respectively; and Φ
mh
, Φ
ml
, and Φ
coil
are the resulting
magnetic fluxes flowing in the magnetic circuit branches.
All of the previous reluctances are calculated using the geo-
metrical parameters of the actuator and the magnetic properties
Fig. 6. Chosen actuator configuration with its geometrical parameters.
Fig. 7. Equivalent magnetic circuit of the actuator.
of the iron and the magnet. The flux leakage in the air gap,
represented in Fig. 8, is neglected in the presented model, and
only the reluctance enclosed by the red rectangle is taken into
account. Again, the reluctances of the air gaps highly depend
on the vertical position (y) of the mover .
Finally, the magnetic energy W
mag
is calculated as a function
of the mover position. The energy is calculated in each part
of the magnetic circuit (magnets, air gaps, and iron yokes), by
knowing that the energy volume density is given by w
mag
=
(BH )/2. It is worth noting that the main part of the energy is
contained in the air gaps and in the magnets. Finally, all those
energy components are added, and the total energy is derived to
obtain the force as already explained.
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MAR KOVI C et al.: DESIGN AND OPTIMIZATION OF A BLOOD PUMP FOR A WEARABLE ARTIFICIAL KIDNEY DEVICE 2057
Fig. 8. Simulated magnetic flux lines in the air gap.
Fig. 9. First actuator prototype.
Fig. 10. First actuator prototype with two tubes.
C. Actuator First Prototype
The first actuator prototype is not a result of optimization.
Instead, the values of its parameters are quickly chosen to
more or less satisfy the specifications, with the goals to verify
the actuator electromagnetic analytical model, to generate the
actuator thermal analytical model, and, finally, to refine the
specifications. Based on all this, the second prototype will be
the result of a detailed optimization.
The first actuator prototype is shown in Figs. 9 and 10. Its
mass is 124 g.
Fig. 11. Force generated by the actuator: Comparison between the analytical
model, FEM, and experimental results.
Fig. 12. Actuator temperature measured by a thermal camera (the ambient
temperature is 22.2
C).
D. Measurements on the Actuator First Prototype
In order to verify the electromagnetic analytical model, the
force is measured. A constant current of 1.1 A is injected in
the winding. In addition, the actuator is simulated using a FEM
commercial software which calculates the force. Fig. 11 shows
the corresponding results and confirms that the electromagnetic
analytical model is correct. The differences between the results
for the position above 1 mm are due to the saturation of the
ferromagnetic parts.
In order to be able to generate a simple thermal analytical
model, a constant power of 2.92 W is dissipated in the winding,
and the thermal steady state is established. Fig. 12 shows the
temperature measured using a thermal camera.
Obviously, the whole mover can be regarded as an isothermal
body. Using the measured temperatures and dissipated power,
the thermal resistance between the mover and surrounding air
is obtained as 23.1 K/W. By supposing that this resistance
corresponds to the natural convection over the mover surface,
the corresponding convection coefficient is 14.7 W/(m
2
· K).
This value will be taken t o calculate the thermal resistance in
the simple thermal model which will be used to optimize the
actuator.
It is important to note here that only t he winding surface
temperature will be calculated in the model. This value has to
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2058 IEEE TRANSACTIONS ON INDUSTRY APPLICATIONS, VOL. 49, NO. 5, SEPTEMBER/OCTOBER 2013
Fig. 13. Power card.
be limited to 45
C according to the specifications. The winding
insulation will not be a critical element in this case as it can
support the temperature of 80
C.
V. C
ONTROL ELECTRONIC CIRCUITS
The control system for the operation of the blood pump is
composed of a central computer, four power cards, and four
position measurement cards (one per actuator). The central
computer communicates with the power cards via the port
RS485.
The power card (Fig. 13) contains a microcontroller
STM32F103 which gives the ON/OFF commands for the power
transistors according to the pulsewidth-modulation logic. The
actuator is driven by the MC33887 integrated H-bridge. Also,
each power card receives a signal from the position measure-
ment card.
The position is determined by measuring two capacitances:
between the upper yoke and mobile yoke and between the
lower yoke and mobile yoke. In order to do this, the position
measurement card s ends two rectangular voltage signals to the
upper and lower mobile yokes and detects the voltage of the
mobile yoke. After demodulation, the resulting analog signal,
which is proportional to the position, is sent to the power card.
VI. A
CTUATOR OPTIMAL DESIGN
The final goal is to obtain the optimal actuator configuration
using the verified analytical models. The optimization will be
performed using the software package ProDesign [18] and is
inspired by [19].
A. Required Force Profile
The actuator force has to satisfy the required profile (Fig. 3),
which means that this profile will impose some constraints
for the optimization. As the optimization tool is not able to
accept this profile as one constraint, it is decided to “sample”
the profile in eight points. Those points (values of position and
TAB LE I
P
OINTS FRO M THE FORCE PRO FI L E
TAB LE II
O
PTIMAL ACTUATOR CONFIGURATION
TABLE III
O
PTIMAL ACTUATOR MAIN PARAMETERS
force) are presented in Table I. The force corresponding to the
point 8 has to be achieved without current, i.e., only by the
reluctant force.
B. Optimization
In order to perform the optimization, the software ProDesign
needs at first the actuator analytical model. Concerning the
constraints, eight points from Table I impose eight constraints.
In addition, the flux density in the iron is limited to 1.5 T; the
temperature is limited to 45
C as required by specifications. As
the specifications impose 400 g as the maximal pump mass, it
means that the maximal actuator mass is limited to 100 g. The
value of the current density is left free, in order to be adapted to
provide the necessary force for each point.
Therefore, a complete theoretical model including electro-
magnetic force calculation in eight points and a thermal model
is introduced in the optimization software; the chosen optimiza-
tion function is simply the volume of the actuator. Thus, the
optimization software varies all the parameters simultaneously
to minimize the volume accordingly to the defined constraints.
To perform the optimization, the software applies an advanced
sequential quadratic programming algorithm, which is an itera-
tive method for nonlinear optimization.
Finally, Tables II and III present the optimal actuator con-
figuration. The iron is Armco; the magnet remanence is 1.4 T.
The optimal actuator drawing is shown in Fig. 14. The set of
four actuators is presented in Fig. 15, and the whole system is
presented in Fig. 16.
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MAR KOVI C et al.: DESIGN AND OPTIMIZATION OF A BLOOD PUMP FOR A WEARABLE ARTIFICIAL KIDNEY DEVICE 2059
Fig. 14. Optimal actuator.
Fig. 15. Set of four actuators.
Fig. 16. Whole system.
VII. EXPERIMENTAL RESULTS
The very first experiments on the optimal actuator show
promising results. Fig. 17 shows the necessary actuator force
(imposed by the specifications) compared to the simulation and
experimental results and the corresponding current injected in
the coil. All the three forces match well.
Concerning the operation of the whole pump, the first ex-
periments are performed using a constant current of 1.2 A
(maximal value from Fig. 17). In that case, the pump generates
the flow rate of 150 mL/min (0.75 L is pumped in 5 min) and
the pressure drop of 38 kPa.
The required flow rate of 200 mL/min will be achieved by
increasing the frequency above 2 Hz, and the r equired pressure
of 52 kPa will be achieved by increasing the current.
Fig. 17. Optimal actuator force.
VIII. SUMMARY AND FUTURE WORK
This paper has presented the design of a blood pump for a
wearable artificial kidney device. The first experiments have
shown that the pump satisfies the desired specifications in terms
of rate of flow and pressure drop. The prototype is now ready for
the next steps: measurement of NIH in a specialized laboratory.
The project Nephron+ continues until April 2014. A first
demonstrator containing the real-time microfluidic circuit board
integrating the physical sensors, some actuators, and several
filtering systems is now available [3]. The NIH measurement
tests of the blood pump will start in February 2013.
The main drawback of the device is the noise which it
generates. In the actual state, it cannot be used, as the noise
would be annoying for the patient. In order to address this
problem, the control algorithm will be modified for the actuator
final version, so as to reduce the speed at closing of the tubes.
R
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Miroslav Markovic was born in Arandjelovac,
Serbia, in 1970. He received the B.S. degree from
the Faculty of Electrical Engineering, University of
Belgrade, Belgrade, Serbia, in 1996 and the Ph.D.
degree from the Ecole Polytechnique Federale de
Lausanne (EPFL), Lausanne, Switzerland, in 2004.
He is currently a Project Leader with the In-
tegrated Actuators Laboratory (LAI), School of
Engineering (STI), EPFL. His research interest is
optimization design of high-performance electric
drives.
Michael Rapin was born in Payerne, Switzerland,
in 1985. He received the B.S. degree from the
Haute Ecole d’Ingénierie et de Gestion du Canton
de Vaud, Yverdon-les-Bains, Switzerland, and the
M.S. degree in electrical engineering from the Ecole
Polytechnique Federale de Lausanne, Lausanne,
Switzerland.
He is currently an R&D Engineer with the Swiss
Center for Electronics and Microtechnics (CSEM),
Neuchatel.
Marc Correvon was born in Neuchatel, Switzerland,
in 1956. He received the M.S. degree from the Ecole
Polytechnique Federale de Lausanne, Lausanne,
Switzerland.
He was the Project and System Manager as well
as the Head of Electronics with ETEL Space and
Alcatel Space Switzerland. For nine years, he has
been a Professor with the University of Applied
Sciences Western Switzerland, Yverdon-les-Bains,
Switzerland. In 2008, he joined the Swiss Center for
Electronics and Microtechnics (CSEM), Neuchatel,
as the Section Head of electronics and firmware in the Systems Engineering
Division. His research activities include embedded electronics and preprocess-
ing of physiological parameters as well as medical device development.
Yves Perriard was born in Lausanne, Switzerland,
in 1965. He received the M.Sc. degree in microengi-
neering and the Ph.D. degree from the Ecole Poly-
technique Federale de Lausanne (EPFL), Lausanne,
Switzerland, in 1989 and 1992, respectively.
He was a Cofounder and Chief Executive Offi-
cer of Micro-Beam SA, which is involved in high-
precision electric drive design. He is currently the
Director of the Laboratory of Integrated Actuators
(LAI), School of Engineering (STI), EPFL, where
he was a Senior Lecturer beginning in 1998 and has
been a Professor since 2003. His research interests are in the field of new
actuator design and associated electronic devices.
Dr. Perriard has been the Vice-Director of the Microengineering Institute in
Neuchatel since 2009.
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    [Show abstract] [Hide abstract] ABSTRACT: This paper describes the principle, design, control and performance of a novel miniature linear peristaltic pump. Exhibiting dual and bi-directional feed capabilities that allow the simultaneous pumping of two distinct fluids at measured and precise flow rate, the pump consists of two resilient tubes and two directly actuated pumping units. Both of the two flexible tubes are placed in the pumping units in such a way that when a pumping element reciprocates in any direction to collapse one adjacent segment of the tubing, it will release the other adjacent segment of the same tubing at the same time. The two pumping elements reciprocate in a sequence so that fluid in the flexible tubing is continuously pumped in preset direction. For each pumping element, two reflective optical sensors are used to detect its position. Based on the output of the four position sensors, a digital logical control circuit is built for the pump. Experimental results show that this kind of miniature pump has potential applications in portable infusion/suction systems.
    Full-text · Article · Jan 2003 · IEEE/ASME Transactions on Mechatronics
  • [Show abstract] [Hide abstract] ABSTRACT: In this paper, a new high reliability linear drive device, termed as stator-permanent-magnet tubular oscillating actuator (SPM-TOA), is proposed for artificial hearts (AHs). The key is to incorporate the concept of two independent phases into this linear AH device, hence achieving high reliability operation. The fault-tolerant teeth are employed to provide the desired decoupling phases in magnetic circuit. Also, as the magnets and the coils are located in the stator, the proposed SPM-TOA takes the definite advantages of robust mover and direct-drive capability. By using the time-stepping finite element method, the electromagnetic characteristics of the proposed SPM-TOA are analyzed, including magnetic field distributions, flux linkages, back- electromotive forces (back-EMFs) self- and mutual inductances, as well as cogging and thrust forces. The results confirm that the proposed SPM-TOA meets the dimension, weight, and force requirements of the AH drive device.
    No preview · Article · Mar 2012 · Journal of Applied Physics
  • [Show abstract] [Hide abstract] ABSTRACT: This paper presented a prototype axial blood pump with a single-axis controlled magnetic bearing. The magnetic suspension system is realized by using passive magnetic bearings for radial direction balancing, and an electromagnetic actuator to stabilize the axial direction actively. A sensor assembly for axial position sensing is also presented. This sensor can measure axial position with good accuracy. The rotor has an inner impeller structure to accommodate fluid flow, and it is driven by a brushless DC motor. A prototype pump was built and tested both in air and in water. The experimental results demonstrated successful pump operations. The rotor can be stably levitated in the air gap. The motor was operated up to 4500 rpm in air, and 6000 rpm in water. The maximum radial position error was about 0.5 mm when the motor was running at speeds near the critical frequency. Axial position errors were less than 0.1 mm for all the experiments performed.
    No preview · Conference Paper · Oct 2007
  • [Show abstract] [Hide abstract] ABSTRACT: This paper presents a magnetic centrifugal pump, driven by an external rotating magnetic field, for use in artificial heart-assist blood pumps. The magnet rotor in the pump is synchronized to a rotating magnetic field, which then causes the fabricated impeller to produce a centrifugal force in the pump. Because the pump is driven by the direct application of an external rotating magnetic field, the proposed pump does not require a shaft or mechanical bearings. Thus, the proposed mechanism and external driving source give the pump the advantages of compact size; wireless, battery-free operation; heat-free operation; and minimal risk of bacterial infections. The pump weight, length, and diameter are 34 g, 35 mm, and 20 mm, respectively. The compact size and performance make the device suitable for use in pediatric cardiac assistance pumps.
    No preview · Article · May 2012 · IEEE Transactions on Magnetics
  • [Show abstract] [Hide abstract] ABSTRACT: A bi-directional linear peristaltic pump is presented that avoids use of a rotating step motor and rollers. Fluid displacement, flow and pump function, respectively, are achieved by means of three electro-magnetic actuators. They are electronically controlled and designed to periodically squeeze an elastic tube that contains the liquid. The device reveals a very simple construction, is maintenance free, has very low power consumption, thus accounting for very low manufacture and operation costs. Upon reducing the size of the actuators, it can be easily integrated with micro fluidic arrangements. Flow and pressure pulsation can be reduced by increasing the number of actuators. A finite element-based numerical model has been developed, using the fluid structure interaction (FSI) model. The flow dynamics inside the tube has been treated, using the coupling between the Navier-Stokes and the elastic displacement equations of the elastic tube. The appropriate and optimum timely driving sequence of the actuators, as well as the response characteristics of the physical system, were also numerically evaluated. Experimental results fully agree with the numerical simulation results. The dynamic range of the flow rate was determined as ranging from 7 00μ/min to 79ml/min, while the back pressure varies between 400 mmHg to 14 mmHg.
    No preview · Conference Paper · May 2011
  • Source
    [Show abstract] [Hide abstract] ABSTRACT: Research on pulsatile blood pumps for extracorporeal life support has been widely performed because of the proven advantageous effects of blood pulsation. However, studies on the use of pulsatile blood pumps for hemodialysis are limited, although available evidence demonstrates that pulsatile blood flow has a positive influence on dialysis outcome. Therefore, the authors designed a new pulsatile pump, which is characterized by minimal-occlusion of blood-containing tubing, no requirement for valves, and no blood flow regurgitation. In-vitro hemolysis tests were conducted using fresh bovine blood, and the normalized index of hemolysis was adopted to compare blood traumas induced by the devised pulsatile pump and a conventional roller pump. In addition, experimental hemodialyses with a canine renal failure model were performed using the devised pump. Normalized index of hemolysis levels obtained was much smaller for the devised pulse pump (45 +/- 21 mg/100 L) than for the roller pump (103 +/- 10 mg/100 L), and no technical problems were encountered during dialysis sessions. Blood and dialysate flow rates were maintained at predetermined values and molecular removal was satisfactory. Postdialysis urea and creatinine reduction ratios were 61.8% +/- 10.6% and 57.4% +/- 9.0%, respectively. Pulsatile flow has usually been generated using pulsatile devices containing valves, but the valves cause concern in terms of the clinical applications of these devices. However, the described pulsatile pump does not require valves, and yet no blood flow regurgitation was observed.
    Full-text · Article · Mar 2008 · ASAIO journal (American Society for Artificial Internal Organs: 1992)
  • [Show abstract] [Hide abstract] ABSTRACT: First Page of the Article
    No preview · Conference Paper · Nov 2003
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