Feb 2, 1990 - cardiac assist systems will be actuated electrically rather than by ... T. Mikami is with the Department of Biomedical Engineering, Graduate.
146
IEEE TRANSACTIONS O N BIOMEDICAL ENGINEERING. VOL. 37. NO. 2 . FEBRUARY 1990
Development of an Implantable Motor-Driven Assist Pump System YOSHINOR1 MITAMURA,
MEMBER,
Abstract-A motor-driven artificial pump and its transcutaneous energy transmission (TET) system have been developed. The artificial pump consists of a high-speed dc brushless motor driving a ball screw and magnetic coupling mechanism between the blood pump and ball screw. The ball screw transfers high-speed rotary motion into low-speed rectilinear motion by a single component. Magnetic coupling enables active blood filling without applying an excess negative pressure to the pump. The transcutaneous transformer is formed from a pair of concave/convex ferrite cores. This design minimizes lateral motion of the external core. Information on motor voltage is transmitted through the skin by infrared pulses. The motor voltage is regulated by controlling the duty ratio of the square pulse supplied to the primary coil. Pump flow of 5.6 I/min was obtained with a mean outlet pressure of 100 mmHg at a drive rate of 100 bpm under preload of 15 mmHg. The performance of synchronous pumping has been very satisfactory. Continuous pumping was maintained by the backup battery in the case of interruption of TET. 24 W were transmitted by TET system with 78 percent of efficiency. Temperature rise of the internal core was 0.2 C. The developed system is promising as an implantable assist pump system.
I. INTRODUCTION
P
NEUMATIC total artificial hearts and cardiac assist systems have been successfully applied clinically. However, this method involves the permanent piercing of the chest wall by energy carrying lines. These systems have drawbacks such as infection [ I ] and limitations of the patient’s activities. The next generation of artificial hearts and permanent cardiac assist systems will be actuated electrically rather than by compressed gases. Electrical actuation permits the design of systems free of tubes and wires passing through the skin. Electric systems promise dramatic improvements in the quality of life of pump-dependent patients. With recent advances in electric motors, electronics, and materials, several types of implantable electrically powered devices have been developed. These devices include a solenoid [2], [3], a linear electromagnetic actuator [4], a high-speed direct current (dc) motor with gear reManuscript received November 30, 1988; revised March 27, 1989. This work was supported in part by a Grant-in-Aid from the Ministry of Public Welfare and a Grant from the Akiyama Foundation. Y . Mitamura, E. Okamoto, and A. Hirano are with the Research Institute of Applied Electricity, Hokkaido University, Sapporo 060, Japan. T. Mikami is with the Department of Biomedical Engineering, Graduate School of Engineering, Hokkaido University, Sapporo 060, Japan. IEEE Log Number 8932197.
duction [5], [6], a high-speed dc motor driving a hydraulic pump [7], [8], a low-speed high-torque dc motor with cam [9]-[ 111, a high-speed dc motor driving a roller screw [12], and a high-speed, reversing axial flow pump 1131. Each of these methods has its merits and its problems; however, at present no system has a clear overall advantage. A simple solenoid driver would not be capable of producing a stroke of over 1 cm without being excessively large and inefficient. A more efficient approach involves utilizing a solenoid to cock a spring, which then drives the blood pump. This approach is in an advanced state of development and has been used as a temporary ventricular assist system in 14 patients. However, to energize the solenoid, this system requires high power ( 1 kW) pulse for a few milliseconds. The basic action of this device is an on-and-off of the solenoid, and therefore displacement of the pusher plate in the blood pump is not precisely controlled. In a high-speed dc motor with gear reduction, the likelihood of gear train wear may be high. In the system, rotary motion is transferred into rectilinear motion by a cam. Therefore, output pressure and flow waveforms are determined by the configuration of the cam, and output pressures and flows are not precisely regulated. Besides, the system has the drawback that it requires two components, gear and cam, to drive the blood pump by the motor. In the low-speed, high-torque dc motor with cam, the position of the pusher plate in the blood pump is not precisely controlled, either. Output pressures and flows are determined by the shape of the cam. Moreover, the system requires the special low-speed, hightorque dc motor which has high efficiency in lower speed. Generally, motor efficiency is high in higher speed. A high-torque, low-speed dc motor is larger than a lowtorque, high-speed dc motor. In all of these devices, the blood pump is decoupled from the actuator during a filling phase. Negative pressure is not applied to the blood pump as in a pneumatic drive pump. A hydraulic pump concept has the difficulty in fabrication of a miniature custom pump. In the hydraulic system, pressure-sensing transducers are required to control drive pressure. Otherwise, extremely high negative pressure is applied to the pump during a filling phase. After considerable thought, the choice was made to use a ball screw driven by a highspeed, reversing brushless dc motor [14], [15]. A ball screw transfers high-speed rotary motion into low-speed
001 8-9294/90/0200-0146$01 .OO 0 1990 IEEE
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MITAMURA et al.: IMPLANTABLE PUMP SYSTEM
rectilinear motion by a single component. High-speed rotation is required for a dc motor to have high efficiency. Low-speed rectilinear movement is necessary to match the physiological heart rate. The ball screw has the advantage of being able to produce a force in both directions. Thus, the system could be used to power a total artificial heart as well as an assist artificial heart. The ball screw has a theoretical efficiency of over 90 percent. Friction loss is relatively low in the ball screw, because balls are used between the ball nut and the screw shaft. Another advantage of this system is fine control of output pressures and flows. Since the position of pusher plate is directly controlled by the motor during emptying period, output pressure and flow waveforms can be precisely controlled. Moreover, in our system, a threaded screw shaft of the ball screw is coupled to the blood pump by magnetic force. A plastic magnet sheet is fixed to the ball screw shaft, and an iron plate to the blood pump. This linkage enables active blood filling in the pump without applying an excess negative pressure. Moreover, higher actuator efficiency would be expected since blood is actively filled by magnetic force. Electric artificial hearts and heart assist systems require significant power levels which cannot be met by implanted battery systems. Although power requirements of artificial hearts depend on the efficiency of artificial hearts and condition of exercise, the current power requirements are 22 W. Although there are a variety of possible methods for transporting energy into the body, schemes involving electromagnetic coupling are currently receiving the most attention. Coupling types essentially fall into two broad categories: 1) audio frequency transmission by means of closed or nearly closed magnetic systems, in which the coils are wound around a magnetic material to achieve coupling [ 161-[ 181; and 2) radio frequency transmission by means of pancake air-core coils in which magnetic material is not required [19]-[24]. Air core transformers have some potential disadvantages, such as large size, bulkiness, and less coupling between the external and internal coils. To solve these problems a new transformer was formed from a pair of concave and convex ferrite cores. The external device is concave at the base. The upper surface of the internal device is convex. The internal device causes a bulge in skin. This mound serves to locate primary core whose inner diameter approximates the perimeter of the mound. This prevents lateral motion of the primary core. The developed transcutaneous energy transmission system includes output voltage regulation and an implanted rechargeable backup battery. These functions keep constant power supply to a load even in the cases of dislocation of coils, drop of primary battery voltage, and so on. In the developed system, information on the motor voltage is transmitted through the skin by infrared pulses. It is the purpose of this study to demonstrate feasibility of a high-speed brushless dc motor driving a ball screw and magnetic coupling between the pump actuator and
I47
blood pump, and also a transcutaneous energy transmission system for delivering the required power of 22 W. 11. SYSTEMDESCRIPTION The system consists of a motor-driven artificial pump and transcutaneous energy transmission system.
A. Motor-Driven ArtiJicial Pump The motor-driven artificial pump consists of a brushless dc motor driving a ball screw, its mating blood pump, and a microprocessor-based control unit. The artificial assist heart, blood pump, and electric motor drive are shown in Fig. 1 . I) Brushless de Motor Driving a Ball Screw: The basic system is comprised of a ball screw and plastic magnet. A pusher plate is fixed to the threaded shaft of the ball screw. A doughnut-shaped plastic magnet sheet (Magx, NT-5, Japan) is glued to the pusher plate. An iron disk is also fastened firmly to a diaphragm of the blood pump by polyurethane. The blood pump is linked to the ball screw by magnetic force between the plastic magnet sheet fixed to the pusher plate and the iron disk fixed to the blood pump. This mechanism enables active blood filling without applying an excess negative pressure to the pump (Fig. 2). As the ball screw is rotated in one direction, the pusher plate moves forward pushing the blood pump (emptying phase). The magnet does not work as driving force. As the ball screw reverses, the pusher plate is pulled backward by the ball screw and the iron disk is pulled by magnetic force (filling phase). When inflow to the blood pump is enough, the iron disk and pusher plate move backward together. When pump filling is poor and the iron disk moves backward more slowly than the plastic magnet, the iron disk is pulled back by the magnetic force of the plastic magnet (active filling). However, if the pressure lower than -25 mmHg ( = -3.33 kPa) is applied to the blood pump, the plastic magnet is automatically detached from the iron disk and excess negative pressure is not applied to the pump. The area of the magnetic sheet (7 cm2) is determined so that the negative pressure less than -25 mmHg is not applied to the pump. The ball screw (Shinohara Manufacturing Co., FBS0803B-50R60SP10, Japan) used here has a threaded screw shaft which is 8 mm in diameter with a lead of 3 mm/revolution. The brushless dc motor turns the ball screw four revolutions in one direction to produce a 12 mm stroke of the pusher plate. The screw of the ball screw is kept from rotating by the pusher plate. The pusher plate is constrained to rectilinear axial motion by passing the high steel rods attached to the pusher plate through linear ball bearings (NSK, L B 4 Y , Japan). The ball-screw nut is embedded in a stainless steel cylinder, which is fixed firmly to the rotor. The ball-screw nut is constrained to only rotational motion by a large single bearing. The dc brushless motor (SierracWMagnedyne, 563-06A) utilized here is a three-phase, delta wound, 14 pole motor weighing 141 g. It employs samarium cobalt magnets and peak torque is 0.53 N * m.
I 148
IEEE TRANSACTIONS O N BIOMEDICAL ENGINEERING. VOL 37. NO 2. FEBRUARY 1990 MAGNET SHEET
PUSHER HALL PLATE SENSOR
LINEAR BALL BEARING
CHAMBER
U
0
5 0 mm
Fig. 1 . Brushless dc motor-driven artificial assist heart. The system uses a reversing dc motor to actuate a ball screw and pusher plate.
pusher p l a t e ball-screw
EJECTION PHASE
pressure i n blood Pump > -25mH9
pressure i n blood pump a -2krmHg
FILLING PHASE Fig. 2 . Schematic operation of magnetic coupling between the blood pump and ball screw.
Three Hall sensors (Toshiba, THS-103, Japan) fixed to the motor housing detect rotor position. A Hall IC (Matsushita, DN835, Japan) fixed to the bearing mount detects the position of the pusher plate. Magnets are fixed to the pusher plate facing the Hall sensor. 2) Blood Pump: The blood pump has been specifically designed for use with the motor drive. The pump cases are made of epoxy resin (Hysol, C8-W795 and H-W796). The blood-contacting surface of the pump housing is coated with segmented polyurethane (TM-3, Toyobo, Japan). The smooth blood contacting diaphragm within the pump cases is also fabricated of segmented polyurethane. Bjork-Shiley inlet ( # 2 5 ) and outlet ( # 2 3 ) valves are employed. The motor housing is machined of stainless steel. The maximum stroke volume of the pump is 89 ml and the stroke of pump diaphragm (also iron disk) is 12 mm. The diaphragm has the shape of a truncated cone. To give a maximum stroke volume of 89 ml, the diameters of the
pusher plate and diaphragm are designed to be 90 and 00 mm, respectively. 3) Control System: There are three functions to be performed by the electronic system associated with the motor drive-position sensing, power switching, and control. A block diagram of the control system is shown in Fig. 3. The position sensing system consists of four Hall sensors; three for electrical commutation and one for position of the pusher plate. The commutator controls power MOSFET switches that route power to the motor windings according to the rotor position and rotational direction. The driver consists of complementary saturating Nchannel and P-channel power MOSFET’s (Hitachi, K399 and J113, Japan). The power MOSFET’s are switched on and off at 20 kHz. The speed of the motor is changed by controlling the duty cycle of an oscillator whose output is the pulse-width modulation signal. Three Hall sensors on the motor stator provide motor rotor position information to the control system. The Hall IC attached to the motor housing provides pusher plate position information to the control system via A/D converter (National Semiconductor, ADC0809). The transducer must be small enough to fit in the motor case and its output signal must be of sufficient resolution and accuracy for the control functions. The current system uses a Hall sensor and three magnets of specific combination as shown in Fig. 4 . The magnets consist of a permanent magnetic disk and a pair of rectangular parallelepiped magnets placed on the magnet disk. This specific combination of magnets provides the Hall sensor with linear output to pump stroke. The relationship between magnetic flux density and pump stroke is shown in Fig. 5. The correlation coefficient is r = -0.99, and maximum error is 7 percent for the stroke of 20 mm. Control of the motor-driven artificial heart occurs at three levels: velocity control, pump output control, and synchronous pumping to allow the pump to be synchronized to the ECG signal. The control system consists of a CMOS one-chip microcomputer (Toshiba, TMP84CO15, Japan) with an A/D converter (National Semiconductor, ADC0809), and D/A converter (Analog Device, ADC723). The pump output controller attempts to drive the pusher plate between the preset pusher plate plate end points within the preset duration. Information on pusher plate position (output of the Hall sensor) and preset end-systolic and end-diastolic positions are fed into the processor through the A/D converter. The velocity control section has several interrelated functions. It ensures present systolic duration and diastolic duration. The reference velocity signal is varied in accordance with instructions from the pump output control system. Motor speed information is obtained from pulse sequences from the three Hall sensors. The output of each Hall sensor is shaped into a square wave through the Schmitt trigger circuit. Two pulses are generated by the ascending and descending edges of the Schmitt trigger output signal. These pulse sequences are fed into the F-V converter (National Semiconductor, LM-29 17). Motor speed signal from the
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M I T A M U R A et al.: I M P L A N T A B L E P U M P S Y S T E M
phoiolroisisior I r onscut oieous 1r o n s f o r m e r
infrored
LED
I
I
1 chip microcomputer
s y s t o l ic d u r a t i o n d i a s t o l i c d u r d t i o n ( d e l a y dad e n d - d i d s t o l i c and end-systolic posiiian
Hall-sensor x 3
w
Fig. 3. Block diagram of control system for brushless dc motor-driven assist pump. main-magnet
1H:constant-current
Fig. 4. Schematic diagram of a simple magnetic position sensor
; m m E
-
;2
;6
20
strokelml
c a l c u l a t e d m a g n e t i c f l u x density
Fig. 5. Characteristics of the developed magnetic position sensor.
converter is fed into the control system. The systolic duration and diastolic period are fed into the CPU via the A/D converter. For synchronization, an ECG signal is processed through the Schmitt trigger circuit, and then is fed into CPU via counter timer circuit. After delay interval, the assist pump is actuated synchronously.
B. Transcutaneous Energy Transmission System A transcutaneous energy transmission system has been developed for use with electric artificial hearts. The sys-
tem is comprised of a transcutaneous transformer, power oscillator, output power conditioning system, rechargeable back-up battery, and alarm system (Fig. 6 ) . I ) Transcutaneous Transformer: The concavelconvex core-type transformer is formed from two ferrite cores (TDK, specific permeability of 3300) (Fig. 7). A problem which has been observed in other devices of this type is that the external part moves with respect to the implanted part, resulting in reduced coupling and higher loss. To minimize this possibility, the external device is designed to be concave at the base. The internal device causes a bulge in skin. This mound serves to locate primary core whose inner diameter approximates the perimeter of the mound. This prevents lateral motion of the primary core. The implanted secondary device of the transformer is a ferrite core measuring 4.6 cm in diameter at the base, 3.0 cm in diameter at the upper surface, 0.8 cm tall, and weighing 40 g. The windings have 20 turns of Litz wire ( 4 0.12/45). The superficial primary coil contains 15 turns of the same wire wound inside a concave ferrite core with a diameter of 4.6 cm at the base, a diameter of 3.5 cm at the top surface, and weight of 54 g. 2) Theoretical Considerations: In the transcutaneous energy transmission system, both the primary and secondary coils are series-tuned at the operating frequency using capacitors (Fig. 6 ) . To find the conditions for minimizing the losses in the coils, we shall first consider the equivalent circuit of Fig. 8. Solving the loop and mode equations for operation at the resonant frequency:
+ jwMI, V, = jwMIl + R212
(2)
V2
(3)
V , = R,I,
(1)
and =
-RLI2
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I E E E TRANSACTIONS ON BIOMEDICAL ENGINEERING. VOL. 37. NO. 2. FEBRUARY 1990
UI b-!+--$
The received load power
Plead
is given by
0 AT T E R Y
GATE DRIVER
(5)
Plod = (RL 112 1 r / 2 Equations (4) and ( 5 ) yield BATTERY CHARGER
p l ~ a s = Pload(RI(RL
+
R2)2/w2M2
+
R2)/RL.
(6)
At this point, it is convenient to introduce a set of dimensionless quantities to eliminate R , and R2. These parameters are defined by
D,
-+ THE INTERIOR OF THE BODY
Fig. 6. Transcutaneous energy transmission system
=
Rl/wLI;
D2 = R 2 / w L 2 .
(7)
Using (7), the equation (6) becomes
Plosa = Pl,,d(Dl(RL + &Dd2/wk2L*
+ WL*D*)/RL
(8)
where k is the coefficient of coupling ( M = k J L l L 2 ) .By differentiating PI,,, with respect to L 2 and setting the result equal to zero, one finds that the load condition for minimum power loss under the power supply of Plead to the load RL is
Lz
+ D 2 k 2 / D l ) = N:L2s
= R ~ / ( W J D ~
(9)
where N2 = turns of the secondary coil L 2 , = self inductance for N2 = 1. When the resistance R I is small enough, (1) can be approximated by
‘e_
I VI I = uMII*( = wN1N2MsI12I
+--*
~
Fig. 7 Concaveiconvex core-type transcutaneous transformer.
(10)
where
N I = turns of the primary coil N2 = turns of the secondary coil M , = mutual inductance for N I = N, = 1. Equation ( 5 ) gives the unique current Z2 which provides the load RL with the power of Plead. Therefore, for a given VI, the equation ( 10) becomes I1
Cl
I M
c2
I2
Fig. 8. Equivalent circuit of the transformer.
where w is the angular resonant frequency M is the mutual inductance of the two coils V I is the voltage applied to the transcutaneous transformer V2 is the output voltage of the transformer I , is the current supplied by the source I2 is the current in the load R I is the loss in the primary coil R, is the loss in the secondary coil
The power loss is given by
I
U N N,M, ~ = V,/Z*
I
=
constant.
(11)
Equations (9) and (1 1) yield optimum number of terms which provides the load with the given power. Fig. 9 shows the relationship between the efficiency and the number of turns of the coils. 3) Power System and Output Voltage Regulation: The power MOSFET’s (Hitachi, 2SJ122 and 2SK428, Japan) convert the power drawn from an external 18-V dc source into a 50 kHz square wave to excite the transformer primary coil. The loaded transformer presents a low impedance to the power FET at the fundamental frequency of the square wave and relatively high impedance at the frequencies of the higher harmonics. The result is that the primary current waveform approximates the sine wave. The implanted portion of the system contains the secondary coil and series capacitor (polypropylene capacitor) followed by full-wave rectifier (Schottky barrier diodes, Toshiba 5FWJ2S41, Japan) and filter.
M I T A M U R A er a l . : I M P L A N T A B L E P U M P S Y S T E M
IS1
aration of 5 mm. The gap between two ferrite cores was filled with meat. The pump was connected to a mock circulation system, which consisted of an aortic compliance (air cushion chamber), peripheral resistance (constrictor valve), and atrial reservoir. The pump filling pressure was 60 maintained constant by returning the outflow to the atrial F = 5 0 KHZ reservoir. Pump outflow was measured by electromagRL=IO n 40 netic flow meter (Nihonkoden, MF-25, Japan) and outlet K=0.55 pressure (afterload) by pressure transducer (NEC-Sanei, TAN6=O.OI 45307, Japan). Electric currents in the drive system were monitored by inserting a resistance of 0.1 Q in the circuit and by measuring the voltage drop across the resistance. Afterload, pump outflow, displacement of the pusher 0 IO 20 30 40 50 plate, voltage of the external battery, current from the exTURNS OF SECONDARY COIL(TURN) ternal battery, voltage at the output power conditioning Fig. 9. Relationship between efficiency of the transformer and the number circuit, current through the output power conditioning cirof turns of the coils. cuit, and motor current under preload of 15 mmHg at drive rate of 70 bpm are shown in Fig. 10. Full stroke pumping The output power conditioning circuitry is responsible was obtained by the energy transmitted through the meat for holding its output voltage constant as the external bat- of 5 mm. At the beginning of emptying phase, peak curtery voltage falls, as coupling between the coils varies, rent of 2.7 A flowed to the motor to accelerate the ball and as the load current changes. This circuit is a variation screw and pusher plate. During the constant speed period, on the switching regulator. The output voltage is con- the motor current flowed depending on the load. At the verted into its proportional pulse frequency by V-F con- end of emptying phase motor current dropped toward zero. verter (National Semiconductor, LM566). The informa- At the start of filling phase, peak current of about 2.5 A tion on pulse frequency is transcutaneously transmitted by was also observed. During the constant speed period, the infrared LED and phototransistor (Sharp, GL513F and current was lower than that in the emptying phase. When PT550F). The pulse frequency is proportionally changed the motor stopped, the motor current dropped to zero into voltage by F-V converter (National Semiconductor, while the secondary current flowed through the output LM2917). This voltage is compared with the nominal conditioning circuit. The secondary voltage was kept alvalue for the output voltage, and according to the error most constant while the motor current changed widely. signal, duty cycle of the primary pulse is varied by Relationships between pump flow and preload, afterswitching regulator control circuit (Motorola MC-3420). load, or drive rate are demonstrated in Fig. 11. Pump out4) Rechargeable Backup Battery and Alarm System: A flow was measured for various preloads against mean sealed rechargeable lead storage battery (Matsushita, afterload of 70 mmHg at drive rate of 100 bpm. Without LCT-812 0.8Ah, 12V dc, Japan) measuring 61 X 25 X magnetic coupling (i.e., decoupling) pump outflow de95 mm and weighing 320 g is connected in parallel to the creased with the decrease of preload. However, with load through a relay. A constant voltage is applied from magnetic coupling pump flow did not decrease so much the dc-dc regulator (Maxim, MAX 630) to the battery ter- with the decrease of preload. The assist pump ejected outminals sufficient to maintain an approximately constant flow greater than 5 I/min even under zero preload. Restorage of charge. If transcutaneous energy transmission lationship between pump outflow and afterload was meais interrupted due to dislocation of transformers, the com- sured under preload of 15 mmHg at drive rate of 100 bpm. parator drives the relay and electric energy is supplied With the increase of afterload, pump flow decreased gradfrom the backup battery for at least 20 min. If the voltage ually. Outflow of 5.6 I/min was obtained against mean to the load becomes normal, the comparator switches the outlet pressure of 100 mmHg. Pump outflow was mearelay, and the backup battery is disconnected from the sured for various drive rate under preload of 15 mmHg load and recharged. against mean outlet pressure of 70 mmHg. Pump flow inThe alarm system always monitors the transmitted in- creased with the increase of drive rate, reached a peak at frared signal. If the amplitude decreases below a thresh- around 140 bpm, then decreased for a higher drive rate. old level due to dislocation of transformers, or pulse freTo examine the function of magnetic coupling between quency drops below a threshold level due to the decrease the blood pump and the ball screw, hemodynamics were of output voltage, the alarm system makes a warning sig- monitored in total inflow obstruction (Fig. 12). Under nal. preload as low as 3 mmHg, negative pressure of -15 to -29 mmHg was applied to the pump by the magnetic 111. EXPERIMENTS coupling, and normal level of pump inflow and outlet A . I n Vitro Pump Performance pressure was maintained. When the inflow cannula was The motor-driven artificial pump was actuated by the completely clamped, the magnet sheet was detached from transcutaneous energy transmission system at a core sep- the blood pump, because the pump diaphragm could not EFFICIENCY(%)
I IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING. VOL. 37. NO. 2. FEBRUAKY IY90
move backward. This prevented an extremely negative pump pressure. The pump pressure increased to -25 mmHg after the initial decrease in total inflow obstruction.
B. In Vivo Pump Performance Acute animal experiments were conducted to verify the performance of the motor-driven artificial heart. The dogs weighing 17.5 and 17.0 kg were anesthetized with sodium pentobarbital (25 mg/kg) and were placed on the Bird respirator. After thoracotomy through the fifth intercostal space, the motor-driven pump was implanted between the left atrium and the descending aorta for about 5 h. Arterial catheter was inserted into the aorta through the left carotid artery and left ventricular catheter into the left ventricle through the left ventricular free wall for pressure measurements. Electromagnetic flow probe was implanted on the ascending aorta and pump inflow cannula. Fig. 13 shows hemodynamics during synchronous pumping. Stable counterpulsation was obtained. Pump flow and cardiac output appeared alternately. Two peaks were observed in the aortic pressure due to the outputs from the motor-driven assist pump and the natural heart.
external b a t t e r y current secondary voltage
20 10
(Volt.)
0
motor
drive r o t e 70bpm. t o t o l input p o w e r 2 5 W
Fig. 10. Hemodynamics of the motor-driven assist pump powered by the transcutaneous energy transmission system.
'1
yagnetlc coupling
decoupling
zt/'
P
e 4I :'2
CL
I
01
100
50
150
200
drivt. r d L i . ( b p i s . ) (1
',U
11lO (I
l',U
("0
I 1i.r l 1 / ~ 1 d ( ~ t u i ~ l l q )
Fig. 11. Characteristics of the motor-driven assist pump
C. Transcutaneous Energy Transmission System Efficiency of the energy transmission system was measured when 24 W were transmitted at a core separation of 5 mm. The result showed that overall efficiency was 78 percent (Fig. 14). Four percent of the input power were consumed in the transformer, 12 percent in the inverter, rectifier, and filter, and 6 percent in the control circuit. Output voltage control for the changes in tissue gap and radial misalignment was tested in the pot core-type transformer (TDK, P36/22, 36 mm in diameter and 11 mm thick). The output voltage was kept almost constant for the changes in tissue gaps of 3-8 mm (Fig. 15). Output voltage was also maintained almost constant for the changes in radial displacement of 0-10 mm. Fig. 16 demonstrates that the motor was driven by the backup battery when energy transmission through the transformer was interrupted. When energy transmission was interrupted, the secondary voltage suddenly dropped to zero. The drop of the secondary voltage was sensed by the comparator, and at once the backup battery was connected to the motor by the relay. Almost the same pump stroke, pump outflow, and outlet pressure were maintained by the backup battery. The backup battery could keep pumping for 20 min. Temperature rise was tested in a chronic experiment by implanting the device in a dog, measuring temperature rises by implanted thermistors (National Semiconductor, LM35Z). The internal core was covered with polyester fabric. The thermistor was fixed on the outside of the internal core. The dog weighing 10 kg was anesthetized and the internal core was inserted just beneath the skin of the back. The internal core was fixed in position by means of several sutures of the fabric. The leads from the core were tunneled beneath the skin and externalized at a remote
~
MITAMURA et al. : IMPLANTABLE PUMP SYSTEM
153
+lsec+ mean motor ] current (3mD.) 2I
G Dreload
3mqg
d r i v e rate
7Gbsm
Fig. 12. Hemodynamics in total inflow obstruction.
-
20v1
-e--*-
WITH CONTROL
tc
WITHOUT CONTROL
2ovl
WITH CONTROL
-.--.-WITHOUT
CONTROL
2’ 1 10
= I P
lo
, - 10 0
2
4 6 8 TISSUE GAP
lOrnrn
?
oi7---, 0
-
10 2 4 6 8 lOmm RADIAL MISALIGNMENT
5
Fig. 15. Voltage control for the changes in tissue gap and radial misalignment when 20 W were transmitted.
mean motor current [amp.]
I
3
1
*1seC.*
~
I
I
I
I
synchronous drlve mode Wlth natural heart heart rate:drlve r a t e = l : 1 heart r a t e 100bpm. Fig 13 Hemodynamm during synchronous pumplng
lost power
a:’m
transformer
2
/.
‘
capacitor
received power 6z
control circuit
Fig. 14. Power consumption in each component of the transcutaneous energy transmission system.
location. Ten days afer operation, the external core was placed on the surface of the back in such a position as to be coaxial with the internal core. The thermistor was fixed on the outside of the external core. The leads from the coil within the body were connected to a resistance load (incandescent lamp). The temperatures were monitored delivering power of 24.5 W to the lamp load. The temperature rise in the internal core was 0.2 C , and that in the external core 3.0 C (Fig. 17). IV. DISCUSSION The developed motor-driven assist pump system consists of a dc brushless motor driving a ball screw, magnetic coupling mechanism between the blood pump and the ball screw, and a transcutaneous energy transmission system. The system has shown very promising results in testing thus far. The system pumped flows of 5.6 l/min with a mean outlet pressure of 100 mmHg at a drive rate
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tive pressure is required to obtain enough blood flow. For this purpose, the blood pump and motor should be coupled directly. However, direct linkage may cause excess negative pressure in the pump. the thin atrial wall is sucked by the excess negative pressure, and the inflow port of artificial pump is occluded by the atrial wall (atrial collapse). Excess negative pressure does not increase pump inflow. Atrial pressure should be maintained within a normal range. Linkage of the blood pump and motor by magvoltage netic force enables active blood filling without applying excess negative pressure to the pump. As shown in Fig. 11, enough pump outflow was obtained for low preload with magnetic coupling, although pump flow decreased with the decrease of preload without magnetic coupling. Since the blood pump are decoupled from the actuator secondary voltage when pump pressure decreased below -25 mmHg, ex(Volt.) cess negative pressure was not applied to the pump even secondary in total inflow obstruction (Fig. 12). current Long-term durability of the developed system has not (Amp. been intensively tested. However, during a half million cycles of the system in vitro, no failure occurred. As far motor current as durability is concerned, a low-speed system (up to sev(Amp eral hundred rpm) would be more favorable than a highspeed system (over ten thousand rpm). drlve rate 80b~rn. In this system, motor speed is less than 2000 rpm. This Dreload l5mnHg motor speed would not extremely impair the durability of Fig. 16. Continuous pumping of the motor-driven pump by the backup the system. However, further long-term test is required to battery. Transcutaneous energy transmission was interrupted, but almost investigate the durability of the system. the same pump flow was maintained. Various types of transcutaneous energy transmission systems have been developd to power the implanted electric artificial hearts 1161-[24]. However, several problems 40'c 1 RECEIVER remain to be solved. They are increase of transmission 30 TRANSMITTER efficiency, allowable temperature rise, easy fitting, regulation output voltage, and safety. f i High energy transmission efficiency is essential for a battery powered biomedical system because it prolongs IO life of the external battery and minimizes the possibility ROOM TEMPERATURE 25'C of burn due to high temperature. High transmission efficiency was realized in our system by employing the op60 120 180mm timum number of windings and also using MOSFET's and TIME Schottky barrier diodes with low on-resistance. Total eff=50kHz Pour=24.5W ficiency of 78 percent in our systems is reasonably high Fig. 17. Temperature rise of the transformer when 24.5 W was transmitted comparing with the reported values; total efficiency was in a dog. 65-80 percent for nominal loads of 12-24 W [7], and 5367 percent for 10 W supply [8]. of 100 beats/min under preload of 15 mmHg. The perThe temperature rise in the tissue when 24.5 W was formance of synchronous pumping has been very satisfac- transmitted was within 0.2 C. This temperature rise may tory during in vivo studies. Continuous pumping of the not give harmful effects on the tissue. Heat diffusion in artificial heart was maintained by the backup battery when the core and heat removal by blood flow probably detranscutaneous energy transmission was interrupted. creased the temperature rise. Similar temperature rises are In most of the motor-driven artificial pumps, the blood reported; Maximum tissue temperature of 39.7 C was obpump and motor are decoupled during diastole. Blood fill- served at the transmission of 1 kW of power [20]. ing is only performed passively. However, at times it is In a number of applications it is important to keep the desirable to suck blood actively by applying mild negative received power within narrow limits despite coupling pressure to the pump as in a pneumatic artificial pump. variations. Several effective means were included in our Especially, when blood is taken from the atrium (i.e., in study to achieve this goal. To minimize the possibility of total artificial hearts or atrial-aortic bypass), mild nega- lateral motion of the coils, the concave/convex core-type !
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transformer was made. Since the internal device causes a bulge in skin, the projections on the external portions can engage the internal portion and prevent lateral motion. Although several approaches were developed to prevent lateral motion of cores, these employed an additional device such as a shallow cup [17], and annular permanent ferrite magnets [ 181. Output voltage control circuit is another effective method for maintaining the received power constant despite coupling variations with some extent. In this study the information on the voltage to an implanted motor is transcutaneously transmitted by infrared pulses. Optical transmission has advantages over conventional radio-frequency transmission. Optical signal is not interfered with radio-frequency signals induced by the transcutaneous energy transmission system and motor. In this study duty cycle of the primary pulse is varied to maintain the secondary voltage constant. Sherman reported that 6 W was dissipated when power to the output stage was switched off by short circuiting the output of the tuned secondary transformer to keep motor voltage constant, while the losses were reduced from 6 to 0.75 W when the primary voltage was reduced by a factor of 10 [21]. Energy loss due to output voltage regulation is less when primary voltage is regulated rather than when secondary voltage is regulated. From the above results it can be concluded that the developed system is promising as an implantable electric artificial heart system.
REFERENCES [I] D. L. Holmberg, P. Dew, C. Crump, G. Bums, Y. Taenaka, and D. B. Olsen, “Percutaneous access devices in calves receiving an artificial heart,” Artif. Organs, vol. 12, pp. 34-39, 1988. [2] P. M. Portner, “The Novacor heart assist system: Development, testing and initial clinical evaluation,” in Artificial Heart-2, T. Akutsu, Ed. Tokyo: Springer, 1988, pp. 89-97. [3] P. T . Miller, G. F. Green, H. Chen, N. Ramasamy, D. H. LaForge, J. S . Jasseawalla, A. K. Ream, P. E. Oyer, and P. M. Portner, “ I n vivo evaluation of a compact, implantable left ventricular assist system (LVAS),” Trans. Amer. Soc. Artif. Intern. Organs, vol. 29, pp. 551-555, 1983. [4] H. Yamada, M. Nirei, H. Ota, K. Kawakatsu, M. Karita, T. Maruyama, M. Chimura, T. Ogasawara, N. Nishizawa, T. Takeuchi, Y . Yamatomato, and T . Akutsu, “Linear electromagnetic actuators for implantable artificial heart,” in Artificial Heart-2, T. Akutsu, Ed. Tokyo: Springer, 1988, pp. 273-279. [SI S . R. Igo, J. M. Fuqua, M. G. Mcgee, G. J . Creager, G. E. Pool, T. W. Krudwing, and 0. H. Frazier, “An implantable ventricular assist system; Chronic in vivo performance,” Trans. Amer. Soc. Arrif. Intern. Organs, vol. 30, pp. 81-85, 1984. [6] S . Fukunaga, Y. Hamanaka, H. Ishihara, T. Sueda, and Y. Matsuura, “Implantable motor-driven artificial heart,” in Artificial Heart-2, T. Akutsu, Ed. Tokyo: Springer, 1988, pp. 344-349. 171 J. C. Moise, “Energy system for chronic circulatory support,” in Artificial Herrrt-2. T. Akutsu. Ed. Tokyo: Springer. 1988, pp. 246253. [8] A. L. Blubaugh, K. C. Bulter, J. A. Schneider, J. C. Moise, L. Fujimoto, R. Kiraly. W. A. Smith, and Y. Nose, “Thermally and electrically powered left ventricular assist devices,” in Progress in Artificial Organs. K . Atsumi, M. Maekawa, and K . Ota. Eds. 1983. Cleveland: ISAO Press, 1984, pp. 91-97. 191 D. B . Gernes, W. F. Bemhard, W. C. Clay, C. W. Sherman, and D. Burke, “Development of an implantable, integrated, electrically powered ventricular assist system.” Trans. Amer. Soc. Artif. Intern. Organs. vol. 29, pp. 546-550, 1983. [ I O ] G. Rosenberg, A. J. Snyder, D. L. Landis, D. B. Geselowitz, J. H. Donachy, and W. S . Pierce. “An electric motor-driven total artificial
I55 heart: seven months survival in the calf,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 30, pp. 69-74, 1984. 1 I I ] S . Takatani. H. Takano, Y. Taenaka, T. Nakatani. H. Noda, M. Kinoshita, S . Fukuda, and T . Akutsu, “Toward a completely implantable TAH : A left-right simultaneously ejecting motor-driven TAH system.” Trans. Ainer. Soc. Artif. Intern. Organs. vol. 3 3 , pp. 235239. 1987. [ 121 G . Rosenberg, A. Snyder, W. Weiss, D. Landis, D. Geselowitz, and W. S . Pierce, “A roller screw drive for implantable blood pump,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 28, pp. 123-126, 1982. [I31 A. P. Lioi, J . L. Orth, K. R Crump, G . Diffee, P. A . Dew, S . D. Nielson, and D. B. Olsen, “In vitro development of automatic control for the actively filled electrohydraulic heart,” Artif. Organs, vol. 12, pp. 152-162, 1988. [I41 Y. Mitamura, T. Ishizuka, and T. Mikami, “Development of a motor-driven artificial heart,” The Heart, vol. 14, pp. 1060-1061, 1982. [15] Y. Mitamura, E. Okamoto, and T. Mikami, “Motor-driven artificial pump,” in Artificial Heart, T. Akutsu, Ed. Tokyo: Springer, 1986, pp. 71-75. [16] C. F. Andren, M. A. Fadall, V. L. Gott, and S . R. Topaz, “The skin tunnel transformer: A new system that permits both high efficiency transfer of power and telemetry of data through the intact skin,” IEEE Trans. Biorned. Eng., vol. BME-15, pp. 278-280, 1968. [I71 G. H. Myers, G. E. Reed, A. Thumin, S . Fascher, and L. Cortes, “A transcutaneous power transformer,” Trans. Amer. Soc. Artif. Inr. Organs, vol. 14, pp. 210-214, 1968. [I81 G. W. Sutton, L. M. Rivera, and P. T. Kirby, “A miniaturized device for electrical energy transmission through intact skin: Concepts and results of initial tests,” Artif. Organs, vol. 5 , suppl, pp. 437440, 1981. 1191 J . W. Fuller, “Apparatus for efficient power transfer through a tissue barrier,’’ IEEE Trans. Biomed. Eng., vol. BME-15, pp. 63-65, 1968. (201 J . C. Schuder, J . H. Gold, and H. E. Stephenson, Jr., “An inductively coupled RF system for the transmission of I KW of power through the skin,” IEEE Trans. Biomed. Eng., vol. BME-18, pp. 265-273, 1971. (211 C. W. Sherman, W. C. Clay, K. A. Dasse, and B. D. T. Daly, “A transcutaneous energy transmission system for high-power prosthetics,” in Proc. IEEE 7th Annu. Conf. Eng. Med. Biol. Soc., 1985, pp. 804-808. [22] Y. Abe, T. Chinzei, I. Fujimasa, K. Imachi, K. Mabuchi, K. Maeda, M. Asano, A. Kouno, T . Ono, and K. Atsumi, “Development of transcutaneous energy transmission system for totally implantable artificial heart,” in Artificial Heart-2, T. Akutsu, Ed. Tokyo: Springer, 1988, pp. 257-261. 1231 P. M. Portner, P. E. Oyer, J. S . Jassawalla, H. Chen, P. J. Miller, D. H. LaForger, G . F. Green, and N. E. Shumway, “A totally implantable ventricular assist device for end-stage heart disease,” in Assisted Circulation 2 , F. Unger, Ed. Berlin: Springer, 1984, pp. 1 15-141. [24] J. S . Brugler, D. H. LaForge, J. Lee, F. K. Beering, J . S . Jassawalla, and P. M. Portner, “Transcutaneous power transmission and electronic control of a ventricular assist system,” in Proc. IEEE 8th Ann. Conf. Eng. Med. B i d . S o c . , 1986, pp. 73-76.
Yushinori Mitamura (M’71) received the B.S. degree in instrument and control engineering from Nagoya Institute of Technology, Nagoya, Japan, in 1966, and the M.S. and Ph.D. degrees in electronic engineering from Hokkaido University. Sapporo, Japan, in 1968 and 1971, respectively. He worked as a Research Assistant with the Department of Biomedical Control, Research Institutz of Applied Electricity at Hokkaido University from 1971 to 1978. During 1974-1976 he was .... a National Institutes of Heal; Fellow at Cleveland Clinic Foundation. Cleveland. OH, where he studied artificial organs. From 1978 to 1989 he was an Associate Professor in the Department of Medical Electronics. Research Institute of Applied Electricity. Since 1989 he has been a Professor in the Department o f Electronic and Information Engineering. Hokkaido Tokai University, Sapporo, where his activities have included researches on implantable artificial hearts, control of artificial hearts, bioelectronics, and biomaterials. Dr. Mitamura is a member of the American Society for Artificial Internal Organs, the International Society for Artificial Organs, and the Japan Society of Medical Electronics and Biological Engineering.
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Eiji Okamoto (M’85) received the B.E. degree in electrical engineering, and the M.E. and Ph.D. degrees in biomedical engineering from Hokkaido University, Sapporo, Japan, in 1984, 1986, and 1989, respectively. Since 1989 he has been a research fellow in the Department of Biomedical Engineering, Faculty of Engineering, Hokkaido University. His current research interests include artificial hearts, particularly the total implantable artificial heart system. Dr. Okamoto is a member of the Japan Society of Medical Electronics and Biological Engineering, the Japanese Society for Artificial Organs, and the Institute of Electronics, Information and Communication Engineers. Atsushi Hirano received the B.E. degree in electronic engineering and the M.S. degree in biomedical engineering from Hokkaido University, Sapporo, Japan, in 1986 and 1988, respectively. Since 1988 he has been a researcher in the Engineering Research Center of Tokyo Electric Power Company.
Tomohisa Mikami received the B.S. degree in physics and the Ph.D. degree in medical science from Hokkaido University, Sapporo, Japan, in 1951 and 1959, respectively. From 1982 to 1981 he worked at the Research Institute of Applied Electricity, Hokkaido University, as an Associate Professor, and as a Professor from 1971. In 1981 he joined the Faculty of Engineering, Hokkaido University, where he is currently a Professor of Biomedical Engineering. His current activities in research include systems analysis of respiration and circulation, optical and ultrasonic measurements for the body, and biomedical information processing. Dr. Mikami is a member of the International Society for Artificial Organs, the Japan Society of Medical Electronics and Biological Engineering, the Japanese Society for Artificial Organs, and the Institute of Electronics, Information and Communication Engineers.
