JOURNAL

OF SURGICAL

RESEARCH

A Pusher

18, 597-605 (1975)

Plate Pump for Totally

Left Ventricular

Assist

Device

Implantable Systems



BENEDICT D. T. DALY, M.D., KENNETH G. HAGEN, M.S., JOHN M. FUQUA, B.S., STEPHEN R. IGO, FRED N. HUFFMAN, PH.D., AND JOHN C. NORMAN, M.D. Cardiovascular Surgical Research Laboratories of Texas Heart Institute of St. Luke’s Episcopal and Texas Children’s Hospitals, Texas Medical Center, Houston, Texas and Therm0 Electron Corporation, Waltham, Massachusetts. Submitted for publication July 17, 1974

A unique left ventricular assist device designed for totally implantable nuclear-fueled and electrical circulatory support systems has been tested in vitro and in vivo. This device is a pusher plate-actuated rolling diaphragm-type pump with inlet and outlet valves and a polyester flocked inner surface. Incorporation of a pressure volume transformer and blood-cooled heat exchanger within the pump have obviated the need for separate system components and reduced over-all system size. The purpose of this report is to describe this pump, its in vitro characteristics, and its in vivo performance. DESCRIPTION OF THE PUMP The pusher plate pump illustrated in Fig. 1 is a left ventricular assist device which accepts blood from the left ventricular apex and ejects into the descending thoracic aorta. The discoid pumping chamber has four major components: a rigid hollow base or plenum, a flexing bladder, and inlet and outlet ports. The pump displaces a volume of 750 ml and weighs 1.25 kg. It measures 7 in. in over-all length, 4$$ in. in width, and 3yz in. in height. The rigid stainless-steel plenum forms the base of the pump and functions as a bloodcooled heat exchanger in those systems requiring transfer of reject heat. The Ilexing bladder is composed of polyester-reinforced Supported in part by USPHS Contracts NOl-HV6-982, NOl-HT-4-2909, and NOl-HV-1-2065.

silicone rubber 40 mils thick. Its upper surface is bonded to a pusher plate and forms a circumferential roll at the periphery as the pusher plate is depressed. When the pusher plate forces the walls of the pump bladder to move together, the natural flow pattern between these walls is radial (Fig. 2). With the inflow valve closed, the lines of potential flow must be redirected toward the outlet, and as the pusher plate continues to close the gap between the walls, the fluid in the inflow side of the bladder must flow through the narrow space between the bladder walls outside the maximum diameter of the pusher plate. All of the blood-contacting surfaces are covered with a thin layer of polyurethane (2 mils thick) which is overcoated with polyester fibrils 10 mils long and 1 mil in diameter with a density of 150 filaments per square mm. This flock interface promotes the rapid deposition of red blood cells, platelets, and fibrin which provide a stable blood-compatible lining. The pump actuator which controls the pusher plate is situated above the pumping chamber but is an integral part of the energy system driving the pump (Table 1). The nuclear-fueled systems utilize heat engines which provide either pneumatic or hydraulic actuation by means of a small-diameter piston mechanically attached to the pump pusher plate [3, 5-71. In this way, the necessary pressure/volume transformation is achieved. The ejection stroke of the pump always terminates at the same location;

597 Copyright @1975by Academic Press, Inc. All rights of reproduction in any form reserved.

598

JOURNAL OF SURGICAL PUSHE?

RESEARCH VOL. 18, NO. 6, JUNE 1975 PLATE

PISTON COMPLIANCE

BAG

ELBOW

LILLEHEI-KASTER

VALVE

WICK CONNECT BLOOD OUTFLOW TO DESCENDING AORTA

BLOOD INFLW

FIG. 1. The pusher-plate pump with its attached inflow tube and outflow graft. A section has been cut away to illustrate the heat exchanger plenum.

however, the stroke volume is variable because the filling stroke termination point is not fixed, but is established by the volume output of the biologic left ventricle for each beat. The electrical systems employ either a brushless dc motor driving a face cam and cam follower or an electromagnetic solenoid. They provide a motivating force directly or hydraulically to the pusher plate

Lf31. BLADDER

ia)

[c)

‘\

FILLING FLOW TENDS TO FORM so~fo ROTATION VORJEX WITHIN PUMP (OUTLET VALVE CLOSED)

NET FLOW FIELD DURING (INLET VALVE CLOSED)

EJECTION

A collapsible air or liquid-filled compliance bag of polyester-reinforced silicone rubber surrounds the entire pump. Its function is to accommodate the volume changes created by movement of the pusher plate. The pusher plate pump counterpulses with the biologic heart. However, unlike pneumatically driven assist devices which are powered from external drive units synchronized with the EKG, this pump func-

8013~

fbj SUPERIMPOSED FLOW FIELD DEPRESSING PUSHER PLATE

DUE TO (VALVES

OPEN1

Id) NET FLOW FIELD NEAR END OF PUMP EJECTION STROKE

FIG. 2. Flow patterns within the pump relative to pusher-plate position. At the initiation of pumping, Row is radial. As the pusher plate is depressed, the inflow valve closes and blood is redirected toward the outlet valve. As the pusher plate reaches full stroke, the majority of blood flow is at the circumferential roll between the pusher plate and the sidesof the pumping chamber.

DALY ET AL.: LEFT VENTRICULAR

599

ASSIST DEVICE

TABLE 1 Pusher Plate Pump Circulatory Support Systems System

Type

Driving mechanism

Aerojet Liquid Rocket Company, Sacramento, CA. Andros Corporation, Berkeley, CA.

Nuclear

Donald W. Douglas Laboratories, Richland, WA. Statham Instruments, Oxnard, CA.

Nuclear

Sterling engine pneumatic actuator Electrical energy converter Electromagnetic solenoid actuator Sterling engine hydraulic actuator Electrical energy converter brushless dc motor driving face cam and cam follower Tidal regenerator engine-hydraulic actuator

Therm0 Electron Corp., Waltham, MA.

Electrical

Electrical

Nuclear

tions independently of the electrocardiogram. Control logic triggers the pump actuator from signals derived from the position and/or velocity of the pusher plate. Three types of valves have been used in this pump. Tricuspid valves of polyesterreinforced silicone rubber have been used and exhibit low flow resistance, but their mechanical reliability is not yet well developed. Silicone rubber disc valves have been used based on extensive experience in other types of blood pumps, and LilleheiKaster valves have been used because they combine low pressure drop with excellent durability. The inflow tube and the outflow graft are integral parts of this pump. Both have rotary quick connect/disconnect fittings for attachment to the pump. The inflow tube is composed of a curved, rigid stainless-steel tube which inscribes an angle of 70” and has an internal diameter of 17 mm. Its inner and outer surfaces are flocked directly with polyester fibrils. The outflow graft is a standard 22-mm woven Dacron vascular prosthesis. IN VITRO EVALUATION The pusher plate pump components have been characterized in vitro. Several factors which cannot be easily determined in vivo

Sensing mechanism

Control logic

Pneumatic

Fluid (pneumatic)

Displacement transducers

Electronic

Hydraulic

Fluid (hydraulic)

Hall effect sensor

Electronic

Hall effect sensor

Electronic

have been closely examined. These include the pressure volume ratio as a function of pusher-plate position; bladder compliance; pump flow relative to inflow pressure, outflow pressure, beat rate and valve configuration; and heat-exchanger temperature. As the pusher plate is depressed from the fully retracted position, increasing portions of the inflow and outflow transition parts of the bladder add to the effective pusher plate area. The effect of this is that the pressure/ volume ratio increases as the pump strokes toward the full eject position. The magnitude of this effect is illustrated in Fig. 3.

DISPLACEMENT,

INCH

FIG. 3. The ratio of effective pusher plate area to drive piston area as a function of pusher plate displacement. As the pusher plate is depressed, a portion of the pump bladder rolls downward with the pusher plate. This increases the effective area of the collapsing or driving portion of the pump bladder.

600

JOURNAL

OF SURGICAL

RESEARCH VOL. 18, NO. 6, JUNE 1975

PRESSURE ImmHgl

FIG. 4. Pump bladder compliance. The polyesterreinforced silicone rubber pumping bladder retains some elasticity. As pressure is exerted on the bladder wall, it tends to stretch. This creates a residual volume in the pump proportional to the pressure imparted to the blood and is approximately 10% of the pump’s stroke volume capability.

The asymmetry and elasticity of the pump bladder contribute to a volume compliance which is a function of output pressure. This compliance detracts from the efficiency of the pump and is shown in Fig. 4 as a function of pressure with the pusher plate in the full eject and fill positions. It represents about 10%of the pump’s stroke volume capability. In order to study the effects of filling pressures at varying beat rates and outflow resistances on pump flow with each of the three valves utilized with this pump, three fill pressures (5, 15, and 30 mm Hg) were used at beat rates of 40,60,80, 100, 120, and 150 beats per minute, and outflow pressures of 80, 100, 120, and 140 mm Hg in a mock circulatory loop. The loop used in these experiments consisted of a pump interposed between electromagnetic inflow and outflow gauges and inflow and outflow compliance reservoirs. Connecting tubes and a fluid reservoir completed the circuit. Pump outflow resistance was controlled with a variable constrictor distal to the outlet compliance reservoir. These studies demonstrated the expected inverse relationship between flow rate and output pressure and the dominating roll of fill pressure and valve configuration at high beat rates. For example, using LilleheiKaster valves, and pumping at 120 beats per minute and 120 mm Hg output pressure, the flow was 16 liters/min with a fill pressure of 30 mm Hg; with a fill pressure of 5 mm Hg,

-5 FILL

0 PRESSURE

5 -mm Hi

Ib

15

2b

25

30

3s

FIG. 5. Families of curves demonstrating the effectiveness of three valve configurations in terms of flow through the pusher plate pump at varying inflow pressures with an outflow pressure of 100 mm Hg. The superiority of the Lillehei-Kaster valves is apparent.

the flow was only 10 liters/min. With tricuspid valves, the comparative flows were 12 liters/min. and 2 liters/min. Below a rate of 60 beats per minute, the range of inflow pressures examined had no significant effect on the flow rate. Families of curves at an outflow pressure of 100 mm Hg demonstrate the comparative effects of the three valve configurations with regard to inflow pressure (Fig. 5). The hemodynamic superiority of Lillehei-Kaster valves is apparent. Since pump filling depends primarily on left ventricular pressure, and since energy requirements are of paramount importance in totally implantable systems, the resistance to filling and pressure losses resulting from the utilization of any valve are major concerns. Therefore, the pressure drop across these types of inlet and outlet valves was determined as a function of steady flow with the pusher plate fully depressed and fully retracted. Figures 6 and 7 illustrate the results. Significantly greater pressure losses across the inflow valve compared to those across the outflow valve can be attributed to the adverse effects of the transition in geometry in the inflow port rather than the valve itself. However, Lillehei-Kaster valves were clearly superior at higher flows. Although the silicone disc valves functioned well in terms of pump output at varying inflow and outflow

DALY ET AL.: LEFT VENTRICULAR

LILLEHEI-KASTER

“AWES

ASSIST DEVICE

TRICUSPID

601

VALVES,

FIG. 6. The pressure loss across three inlet valve configurations at constant flow with the pusher plate fully depressed (100% displacement) and fully retracted (0% displacement). The superiority of the LilleheiKaster and tricuspid valves at high flow rates is apparent.

FIG. 7. The pressure loss across three outlet valve configurations at constant flow with the pusher plate fully depressed (100% displacement) and fully retracted (0% displacement). The superiority of the LilleheiKaster valves is apparent.

pressures and beat rates, they exerted increased pressure demands on the pump in the outflow position. They did not exert this effect in the inflow position. In those systems utilizing nuclear fuel as the power source, approximately 30-45 W of reject heat must be eliminated from the systems and hence the body. An attempt at eliminating this heat directly to the tissue surrounding the energy systems resulted in tissue necrosis. A practical solution to this problem is to transfer the heat directly to the blood utilizing a blood-cooled heat exchanger [4]. The rigid plenum has been adapted to permit heated fluid to be circulated to it from the nuclear engines. Since a pseudointimal lining develops over the plenum, heat transfer to the blood must be efficient enough to prevent significant temperature rises at the plenum-lining interface. Pseudointimal damage and perhaps thrombosis formation, hemolysis, red cell damage, and denaturation of serum proteins might result from a high temperature interface. Experiments were designed to study the effects of blood flow and heat input on temperatures across the plenum. A mock circulatory loop was utilized for these experiments. One heater maintained the circulating fluid (water or blood) at body temperature. A second electrical heater was incorporated into a copper plenum to permit

variable heat inputs and uniform heat dissipation. Thermistors were placed on the outer surface of the plenum, and within the pumping chamber. The temperature differences on both sides of the plenum at various pump flows, stroke volumes, and watts of heat energy input were determined. The result (Fig. 8) demonstrated the expected inverse relationship between the temperature differential (AT) across the plenum and flow at varying heat (watts) inputs, and the importance of large stroke volumes for adequate heat transfer. The latter is a consequence of the increased turbulence relative to the plenum at large stroke volumes which improves the convective heat

FIG. 8. The effects of heat input on the temperature gradient between the plenum and blood at varying flow and stroke volume. The higher the flow and the larger the stroke volume, the smaller the 6T across the plenum.

602

JOURNAL OF SURGICAL

RESEARCH VOL. 18, NO. 6, JUNE 1975

FIG. 9. A

cir datory

support

system impla ntation

transfer coefficient. Although pump flow and stroke volume were important for heat transfer, there was less than a 0.1” C temperature rise in the circulating water or blood between the inlet and outlet at all pump flows, stroke volumes, and heat inputs studied.

sensors within the actuator which determine pusher plate position. When the pump is unloading the left ventricle (LVP < AoP), the aortic valve does not open, and the cardiac output (pump output) is essentially pusherplate pump output or the product of pusherplate rate and mean pump-stroke volume. A representative tracing in Fig. 9 demonstrates IN VIVO EVALUATION the in vivo hemodynamic effectivenessof this The implantation procedures have been pump. When the pusher plate pump was described in detail previously [2,5]. The actuated, the aortic pressure increased 20% blood pump is positioned on the left hemi- and was phase-shiftei’ into diastole, peak left diaphragm. The electrical or nuclear energy ventricular pressure was reduced 56% left systems are positioned in the abdomen sus- ventricular end-diastolic pressure 75% and pended from the twelfth rib and a lumbar left ventricular dP/dt 43%. The pump transverse process. Anticoagulation is stroke volume averaged 72 ml/stroke and the cardiac output (pump) was 7.26 liters/ accomplished with 500 ml dextran40 intravenously every 12 hr. Aortic, left ven- min. tricular, and central venous pressures and Table 2 summarizes the hemodynamic left ventricular dP/dt are monitored and and temperature data obtained from five rerecorded continuously. Pusher-plate pump cent nuclear system implants. The output output is determined or estimated from from the pusher plate pump and the degree

DALY ET AL.: LEFT VENTRICULAR

ASSIST DEVICE

603

TABLE 2 Hemodynamic Effects of Pusher Plate Pump in Nuclear Energy Systems

AoPs Calf Condition (mm Hg) H-l&6(N) Pumpoff 100 f 5 Pumpon 110 + 10 H-187(N) Pump on 133 f 8 H-191(N) Pumpoff 134 f 4 Pump on 137 + 9 H-197(N) Pumpon llO* 12 H-199(N) Pumpon 148 + 7

LVP (mm Hg) 100 + 5 65 f 20 71510 134 i: 4 66 f 13 95 + 15 125 f 8

LVEDP (mm Hg) 13*2 5+4 3+1.6 15 * 5.1 3.4* 2.6 6*4 10 f 3.5

LVdP/dt (mm Hg/sec) 2200t 510 1200 2 200 560555 1600 + 524 1067 i 175 1080 f 432 2100 t 380

of counterpulsation are functions of several factors: inlet tube position, system efficiency, and metabolic demands of the experimental animals. In none of these experiments was the pump functioning at maximum output. For example, inlet tube obstruction was demonstrated at autopsy in calf H-197. This was a result of the inlet tube orifice angulating toward the ventricular septum presumably due to a shift in its position postoperatively. Several new inlet tube configurations have been designed to obviate this problem. Although peak left ventricular pressure was only modestly reduced, the system pumped over 8 liters per minute and the left ventricular end-diastolic pressure was lowered. Heat transfer to the

Max. Pump stroke rectal Pump volume output Pump rate temp. (liters/min) (O’C) (B/min) (ml/B) 62 -t 4 96 f 6 5.96 k .33 39 732 15 862 11 6.17 f .71 39 -~ 110~9 72+ 8 7.99 ?I 1.13 39 85 f 20 98 f 5 8.33 f 1.85 39 103 r 13 81 +6 8.31 + 1.69 37

pump paralleled pump output in several energy-system configurations. This was related to the thermodynamics of specific engine-to-pump heat-transfer mechanisms. However, during all of the in vivo implantations, calf rectal temperature did not correlate with pump output, This finding is consistent with the minimal temperature rise observed in the circulating blood between the inlet and outlet of the pump in the in vitro tests with energy inputs almost twice those obtained in vivo. Recent studies suggest that these animals compensate for the increased heat loads by an increase in their respiratory rate with elimination of the heat via the lungs [91.

A summary of the pusher-plate pump sup-

TABLE 3 Pusher Plate Pump Circulatory Support System Implantations Calf

Engine system

Implant period (hr)

Reason for experiment termination

B-028 B-030 B-003 B-031 H-117 H-118 H-l 22 H-l 24 H-l 25 H-l 19 H-128 H-186 H-185 H-187 H-191 H-197 H-199

Tidal regenerator Sterling (hydraulic) Tidal regenerator Sterling (hydraulic) Sterling (hydraulic) Sterling (pneumatic) Tidal regenerator Brushless dc motor Sterling (pneumatic) Sterling (hydraulic) Sterling (hydraulic) Sterling (pneumatic) Sterling (pneumatic) Sterling (hydraulic) Sterling (hydraulic) Tidal regenerator Sterling (hydraulic)

8 1 7 7 12 9 112 43 73 29 5 74 175 22 54 73 176

Kinked inflow tube Overflow valve malfunction Thermal instability (engine) Overflow valve malfunction Entrained gas in hydraulic system Ventricular fibrillation Pump drive bellows leak Pump bladder tear Foil insulation vacuum seal broken Pneumonitis Ventricular fibrillation Actuator bellows leak Actuator bellows leak External gas line broken Thermal instability (engine) Respiratory distress syndrome Respiratory distress syndrome

604

JOURNAL

OF SURGICAL

RESEARCH

port systems implantations performed in our laboratories is shown in Table 3. These systems have operated in viva for as long as 176 hr after systems implantation. Experiments have been terminated because of power system failure (including problems with heat dissipation) in 13 animals, pulmonary insufficiency in three, and arrhythmia in two. In two of these experiments there were significant problems directly related to the pump. In these animals, rupture of the bladder occurred at points of stress adjacent to the pusher plate. Steps have been initiated to correct this problem. At the time of autopsy, all pumps were carefully examined for valve wear, thrombus formation, and pseudointima development. In every instance, the valves were normal and the entire surface of the bladder was covered with a smooth, evenly deposited pseudointima. Histologic and electronmicroscopic analyses of this pseudointima have shown red blood cells, white cells, platelets, and fibrin entrapped in the polyester flock. Investigations utilizing electron microscopy and biochemical techniques are being undertaken to further characterize this lining and determine its viability. Biochemical, hematologic, and coagulation studies remained within normal limits with the following exceptions. Prothrombin times and partial thromboplastin times were mildly prolonged throughout the experiments; bilirubin levels approached 1.5% (normal 0.5 mg/lOO ml) in two calves terminally. However, the serum lactic acid dehydrogenase was not elevated and red cell fragility studies were normal. No consistent trends were seen in plasma hemoglobin levels. Periodic elevations 15-45 mg/lOO ml (normal 5.9~4 mg/lOO ml) may have been related to transfusion. DISCUSSION The pusher-plate pump is a left ventricular assist device designed for coupling either to nuclear-fueled thermal engines or electrical systems. Its primary design characteristics include efficient valves with low pressure gradients and regurgitation; an

VOL.

18, NO. 6, JUNE

1975

integrated pressure volume transformer which significantly reduces system size and weight; a blood-cooled heat exchanger for dissipating reject heat; a polyester flocked blood interface and precise control over the positioning of the blood-contacting surfaces. The pump is designed for totally implantable systems which require compactness and the efficient use of energy. The flat pumping chamber and pusher plate permit relatively small pressure losses with pump actuation to near full stroke. The utilization of full flow orifice Lillehei-Kaster valves has reduced pressure losses across the valves at all beat rates, stroke volumes, and outflow resistances. Although a bladder pump can approach the versatility of a pusher plate pump if connected with an appropriate actuator via an intermediate hydraulic or pneumatic loop,the volume of such a system is considerably greater. Systems utilizing the pusher plate pump have the best potential for utilizing high-power density energizers. With a bladder-type pump, the actuating fluid pressure and volume flow are essentially equivalent to the blood pressure and pumped volume. In the pusher-plate pump, high pressure is utilized for actuation and the volume of fluid (gas or liquid) utilized to accomplish this is relatively small compared to the volume of blood moved in and out of the pump. A pressure volume transformer within the pump makes this possible. As the pusher plate is actuated, the void volume is filled by air or liquid from the compliance bag surrounding the pump. The incorporation of a blood-cooled heat exchanger in the pump also obviates the need for a separate system and eliminates an additional prosthetic interface with blood. The temperature rise of the blood is small because the heat exchanger is placed in the region of greatest flow and mixing. Its large area reduces the heat flux density into the blood minimizing the boundary layer blood temperature rise. In addition, the pumping bladder is not subjected to temperature rises which could degrade its physical properties. Although EKG triggering has been customary in other devices which counterpulse

DALY ET AL.: LEFT VENTRICULAR

with the biologic heart, the systems utilizing the pusher plate pump accomplish synchronization in response to the timing and volumetric demands of the natural heart on a beat by beat basis. While the mechanisms of sensing and actuation are different, each system senses pump filling relative to pusher-plate position and/or velocity and activates the output pump stroke as soon as the biologic heart has discharged its stroke volume into the pump. Because of the low pressure required for pump filling, these systems have provided satisfactory outputs even in the presence of ventricular fibrillation . Although the versatility of the pump has permitted its incorporation into several totally implantable circulatory support systems, the pump actuator for each particular energy system is different. Thus the in vivo performance of the pusher plate pump is also related to the particular system in which it is utilized. These systems have demonstrated hemodynamic effectiveness both in terms of circulatory support and ventricular unloading. Although the longterm effects of a pumping prosthesis on the blood and other organs remain to be established, the results thus far are encouraging. Parallel studies of the long-term effects of intracorporeal heat indicate feasibility [4]. A major problem, however, is the development of respiratory failure which has resulted in the termination of many of these experiments. Since implantation of these systems requires a thoracotomy and laparotomy and, in the case of the nuclear systems, a diaphragmatic incision, pulmonary dysfunction is not unexpected. Although the size of the system components is likely to contribute to this problem, mock pump and engine implants in animals of similar size have not resulted in respiratory failure. Miniaturization of components is in progress and future systems should provide adequate stroke volumes with reduced power density in significantly smaller configurations with improved inlet hemodynamics. The pusher plate pump has dem-

605

ASSlST DEVICE

onstrated its adaptability to several different energy systems. Its design, which incorporates sensing logic, a pressure volume transformer, heat exchanger, and blood-compatible interface has proved to be advantageous for systems intended for extended or permanent implantation. REFERENCES I. Aldrich, J. G., Chambers, J. A., Daly, B. D. T., Migliore, J. J., Newgard, P., Huffman, F. N., and Norman, J. C. An electrically activated left ventricular assist device: In viva testing. Trans. Amer. Sot. Artif Intern. Organs. 1974(in press). 2. Daly, B. D. T., Robinson, W. J., Migliore, J. J., Dove, G. B., Edmonds, C. H., Fuqua, J. M., Huffman, F. N., and Norman, J. C. lmplantable nuclear-fueled circulatory support systems IV: Respiratory management and avoidance of bovine respiratory distress syndrome. Proceedings of 26th Annual Conference on Engineering in Medicine and Biology 15:357,1973. 3. Hagen, K. Cl., Ruggles, A. E., Huffman, F. N., Daly,

B. D. T., Migliore, J. J., and Norman, J. C. Nuclearfueled circulatory support systems XIII: Augmented performance of the tidal regenerator engine. Proceedings of the 9th Intersociety Energy Conversion Engineering Conference, San Francisco, August 1974, p. 805. 4 Huffman, F. N., Bernhard, W. F., and Norman,

J. C. Thrombogenic experience with intravascular heat exchangers. Chest 58590, 1970. 5 Hughes, D., Feigenbutz, L., Daly, B., Faeser, R., Igo, S., Migliore J., Robinson, W. J., Huffman, F., and Norman, J. Nuclear-fueled circulatory support systems: current status. Trans. Amer. Sot. Artif fntern. Organs, 1974(in press). 6 Martini, W. R., Emigh, S. G., White, M. A., Griffith, W. R., Hinderman, J. D., Johnston, R. P., Perrone, R. E., and Feigenbutz, L. V. Unconventional Stirling engines for the artificial heart application. 9th Intersociety

Energy

Conversion

Engineering

Con-

jerence. San Francisco, August 1974. p.79 I. 7. Moise, J. D., Faeser, R. J., and Rudnicki, M. 1. Status of a thermocompressor-powered implantable artificial heart system. 9th Intersociety Energy Conversion

Engineering

Conference.

August 1974,

p. 799. 8. Portner, P. M., Dong, E., Jassawala, J. S., and LaForge, D. H. Performance of an implantable controlled solenoid circulatory assist system. Trans. Amer. Sot. Artif Intern. Organs 19~235, 1973. 9. Whalen, R. L., Jeffrey, D. L., and Asimacopoulos,

P. J. Chronic intracorporeal heat studies in calves. Intern. Organs 1974 (in press). Trans. Amer. Sot. Artif

A pusher plate pump for tatally implantable left ventricular assist device systems.

JOURNAL OF SURGICAL RESEARCH A Pusher 18, 597-605 (1975) Plate Pump for Totally Left Ventricular Assist Device Implantable Systems ’ BENEDI...
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