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Permanent Circulatory Support Systems at The Pennsylvania State University
Abstract-Permanent circulatory support systems are required for patients in whom myocardial damage is irreversible and cardiac transplantation is not possible. Two systems are described which provide long term circulatory support: the left ventricular assist system and the total artificial heart. These systems are based on the design of a pusher plate actuated blood pump, driven by a small brushless dc electric motor and rollerscrew driver. An implantable motor controller maintains suitable physiologic Row rates for both systems and controls left-right balance in the total artificial heart. Other parts of the system include an intra-thoracic compliance chamber, transcutaneous energy and data transmission system, and internal and external batteries.
I . INTRODUCTION HE development of implantable circulatory support systems intended for permanent use has been a major focus of the research effort at the Pennsylvania State University. Pneumatically-powered devices developed at our institution and others are being used successfully as ventricular assist devices (VAD) or total artificial hearts (TAH), primarily as a bridge to transplantation [l], [2]. For patients requiring long term support, due to difficulty in obtaining suitable tissue matching, renal complications, or simply the growing shortage of donor organs, an acceptable permanent circulatory support system is desirable. The permanent VAD system is intended for patients with primarily left ventricular failure but with satisfactory right ventricular function. The permanent TAH system is intended as a complete cardiac replacement.
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11. THE VENTRICULAR ASSISTSYSTEM The left ventricular assist system (Fig. 1) [ 3 ] consists of implantable and external components. Blood flows into the assist pump from the apex of the weakened left ventricle during pump diastole and is returned to the patient's aorta during pump systole. The assist pump acts in parManuscript received December 1, 1988; revised May 1. 1989. This work was supported in part by Public Health Service Grants SROI HL 20356 and ROI HL 13426: Public Health Service Contract NO1 HV88105; The Robert J . Kleberg. Jr. and Helen C. Kleberg Foundation, and The H. G . Barsumian. M . D. Memorial Fund. The authors are with the Department of Surgery. Division of Cardiothoracic Surgery, The Milton S . Hershey Medical Center. The Pennsylvania State University, Hershey. PA 17033. IEEE Log Number 8932196.
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Fig. I . The left ventricular assist system consists of the electric motor driven assist pump, inlet and outlet cannulae, the intrathoracic compliance chamber with infusion port, the electronic controller and backup battery pack, and the energy transmission internal coil. The external components are the energy transmission external coil and the battery and electronics pack.
allel with the natural left ventricle, and can completely replace the left heart function. The two major components of the motor driven assist pump (Fig. 2) are the blood pump and the motor drive assembly. The blood pump consists of a flexible, seamless, segmented polyurethane blood sac with a smooth blood contacting surface. The outer case is machined polysulfone. Bjork-Shiley convexo-concave Delrin disc valves are fitted into the inlet and outlet ports. The blood sac is compressed by a pusher plate during pump systole, while during diastole the plate moves away from the sac, allowing the sac to fill passively. Two sizes of the assist pump have been developed. The clinical size device of Fig. 1 has a stroke volume of ap-
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Velocity
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Fig. 3. Functional block diagram of the VAD and TAH control system
The magnitude of the applied voltage is provided by the velocity control section according to a motor voltage versus position table, which is determined as follows. A desired velocity profile is computed for both the systolic and diastolic strokes which will minimize the inertial forces and is characterized by a constant acceleration during the first half of the stroke and constant deceleration during the second half. The systolic and diastolic portions of this profile are scaled according to the desired systolic and diastolic durations, as provided by the cardiac output control. The difference between the actual and the desired velocity, or velocity error, is calculated through each pump cycle and is used to modify the motor voltage table for the next cycle. This technique is referred to as adaptive feedforward control. Velocity errors within a given proximately 70 cc, a maximum output of 8 l/min and stroke are generally small since, as motor velocity deuses 25 mm outlet and 27 mm inlet valves. A larger size creases, the reduced back emf allows motor current to inintended for experimental implantation in calves has a 100 crease, acting as an inherent negative feedback mechacc stroke volume, 10 1/min maximum output, and 25 mm nism for velocity control. Velocity feedback is used on outlet and 29 mm inlet valves. the current cycle only if the velocity falls below half the A brushless dc motor provides the driving torque which desired velocity. is converted to the linear pusher plate motion by a rollerUsing the motor voltage table and the actual velocity, screw actuator. The three phase, 12 V motor utilizes neo- the load at the pump outlet is estimated by subtracting the dymium iron magnets. The motor rotor is attached to the estimated load contributions of the frictional components, inertia of the blood and of the rotating mechanism, outlet outer nut of the rollerscrew, which is internally threaded and houses a number of threaded rollers which move in a valve pressure drop, and the electromechanical properties planetary motion about the central threaded shaft. This of the motor. The resulting estimated load versus position arrangement eliminates sliding friction and thereby pro- is made available to the cardiac output control. vides high efficiency and long life. The rotor and rollerIn the VAD controller, pump output is varied in rescrew nut rotate on two concentric bearings mounted on sponse to the filling rate. The degree of filling of the preeither side of the stator in the titanium outer case. A guide vious pump cycle is found by noting the motor position rod and bushing prevents rotation of the pusher plate. A at which the load increases, indicating the point of contact magnet ring mounted to the rotor is positioned above 3 with the blood sac. Inadequate filling is identified when Hall effect sensors which are used to sense the rotary po- the blood sac contact occurs some time after the onset of sition. The weight of the VAD motor drive assembly and systole. The output control attempts to maintain a pump blood pump is approximately 600 g. rate as high as possible while maintaining adequate filling The operation of the microprocessor-based controller by adjusting the scale factors used to calculate the desired may be described by a number of functional levels [4], velocity profiles. The effect is to pump whatever blood is summarized in Fig. 3. At the lowest level, the Hall sensor available to the assist pump. An arterial pressure estimate is also derived by averoutputs are decoded and the position of the pusher plate is determined by keeping track of the changes in rotary aging the estimated load during systole. The beat rate is position. Motor reversal is initiated at predetermined lim- reduced if the arterial pressure estimate is greater than a its of pusher plate position. Voltage is applied to the mo- preset limit, as may occur in a hypertensive episode. tor windings by pulse width modulation of semiconductor As blood flows into and out of the assist pump, pressure power switches, in a sequence determined by the rotary changes in the air space surrounding the electric motor position and desired direction. must be minimized in order to allow adequate pump fill-
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ing and ejection. This is accomplished by allowing volume changes to occur in the intrathoracic compliance chamber, which is connected by a tube to the motor housing. The compliance chamber is a flexible segmented polyurethane sac approximately 16 cm in diameter and 0.65 cm in cross section, with a fully inflated volume of 250 cc, which is velour covered to promote a stable, flexible tissue overgrowth. Diffusion of gases through the walls of the compliance chamber and blood sac leads to a gradual loss in the compliance chamber volume. The gas volume must be replenished through a subcutaneous infusion port, typically once every three weeks. 111. THE TOTALARTIFICIAL HEARTSYSTEM
The implantable total artificial heart system (Fig. 4) [5] is functionally similar to the left ventricular assist system. The major difference is that the TAH is located in the pericardial space and completely replaces the natural ventricles. The left and right pump inlets are connected to the left and right atrial remnants, respectively. The left pump outlet connects to the aorta and the right pump outlet to the pulmonary artery. The TAH device (Figs. 5 , 6) also employs a 12 V brushless dc motor and rollerscrew. Two pusher plates are required; one on either end of the rollerscrew shaft. The left and right blood sacs are therefore compressed alternately so that left systole corresponds to right diastole, and vice versa, in contrast to the simultaneous left and right ejection of the natural heart. The segmented polyurethane blood sacs are similar to the assist pump sac. Both 70 cc clinical size and 100 cc calf size devices are being developed. Control system requirements for the TAH differ significantly from those of the VAD. Of primary importance is the maintenance of left atrial pressure within a normal physiologic range. The left atrial pressure is primarily a result of the difference in fluid volume between the systemic and pulmonary circulations. In the natural heart, ventricular output (flow) increases with increasing atrial pressure (the Starling mechanism). This mechanism results in left atrial pressures which, for a healthy left ventricle, always remain in a limited range. The artificial heart control system must provide an equivalent mechanism. In the artificial heart, valve backflow increases with increasing outlet pressure, resulting in a lower net stroke volume for the left pump than for the right pump. In addition, the left pump alone carries a portion of the bronchial circulation, resulting in a slightly higher output requirement for the left pump. The TAH cardiac output control system estimates the completeness of left pump filling by analyzing the pusher plate load, as in the VAD system. When inadequate filling is determined, left systolic speed is decreased. Since left systole corresponds to right diastole, a longer filling time is allowed for the right pump, which increases the right output and the left atrial pressure. Conversely, left sys-
primary coil secondary coil
Fig. 4. The total artificial heart system. The system components are similar to those of the left ventricular assist system, with the exception of the blood pump and motor drive configuration.
Fig. 5 . The total artificial heart device showing the left and right blood pumps and the centrally located motor and rollerscrew.
tolic speed is increased when complete left filling is achieved, thereby reducing right diastolic time and limiting right pump filling. The control system attempts to maintain a stable operating point at which the left pump is just adequately filling.
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Fig. 6. The assembled 100 cc stroke volume total artificial heart, with the compliance chamber, infusion port, and percutaneous cable for animal testing.
For the case of rising right atrial pressure, the increase in right pump stroke volume due to more complete filling, and the resultant increase in left atrial pressure, leads the controller to respond with a more rapid left systole. If the left diastolic speed is unchanged, the pump rate, and therefore the overall cardiac output, is increased. Thus, the textbook Starling response [6] to right atrial pressure is duplicated. By altering how the controller changes left diastolic speed when it changes left systolic speed, we can either enhance or suppress this response. In the natural system, this effect is modulated by neural and endocrine controls, and the Starling mechanism is solely responsible for maintenance of left-right balance. In our experience in calves, right atrial pressure has been an unreliable measure of demand for cardiac output. We therefore suppress sensitivity of cardiac output to right atrial pressure, and use the control described above only to maintain balance. We can then, through manipulation of the range of left diastolic speeds available to the balance control, vary cardiac output according to peripheral resistance, as reflected in systemic arterial pressure. Arterial pressure is estimated from the magnitude of the motor load during ejection of the left pump 171. IV. POWERDELIVERY A N D MONITORING The complete electronics system for both the VAD and TAH (Fig. 7) consists of motor control, power delivery, and monitoring functions. A significant advantage of the electrically driven systems over their pneumatically-powered predecessors is the use of a transcutaneous energy transmission system (TETS) to transfer electrical energy across the intact skin [8], [9]. A three-turn external, or primary, coil is placed in close proximity (typically 1.5-2.5 cm) to a 16-turn implanted, or secondary, coil. Both coils are tuned to 160 kHz by means of low loss series capacitors. An external power oscillator drives the primary coil and induces a cur-
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rent in the secondary coil, which is rectified and filtered by the implanted electronics. The voltage supplied to the motor controller is maintained within a 13.0-14.5 V band by the action of a voltage regulator located in the internal electronics package. The mean power requirements for the VAD and TAH are typically below 13 W depending upon flow rate and aortic pressure, but short duration, peak power demand at the beginning of left systole may be much higher. As a result, the energy transmission system is designed to provide up to 70 W if necessary. The efficiency of the energy transmission system from dc input voltage to dc output ranges from 55 to 75 percent. The external power source is either a 12 V power supply connected to a 115 VAC source, or a 12 V battery pack to allow for patient mobility. The total external pack capacity is approximately 10 amp-hours (Ah), divided between two separate battery packs to provide power source redundancy. The maximum operating time between recharges is approximately 10 h, based on a mean motor power of 9 W. At this time, Nickel-Cadmium batteries are being evaluated and Lithium technology is being pursued. The internal battery pack is intended only as a backup power source during separation of the TETS coils. The internal pack is the major determinant for the size and weight of the implanted electronics package. A discharge time of 45 min has been set as a minimum specification, and nickel-cadmium batteries in the 0.45-1 .O Ah range are being evaluated. In the future, the increased energy density of rechargeable lithium batteries is expected to allow further reductions in the size of the implanted package. A number of charging profiles are also being considered in an attempt to minimize the permissible time between discharges while retaining acceptable cycle life and reliability. A telemetry system for bidirectional communication with the implanted electronics is being developed which utilizes the energy transmission coils and power carrier [lo]. Primarily, the system will be used to monitor the performance of the motor driven TAH or VAD. This information includes the motor rate and the pusher plate load as determined by the motor controller, as well as the status of the internal battery pack. This internal to external data transmission employs impedance modulation on the secondary side and demodulation of the external coil current on the primary side. Secondarily, changes may be made to the implanted controller parameters via external to internal data transmission, utilizing frequency modulation of the power carrier. External to internal communication is not required for the normal operation of the implanted device, but this capability is extremely useful during the operative and immediate post-operative periods while the hemodynamics are being stabilized, and as a means of compensating for gradual changes in the pump inflow and outflow characteristics, if necessary. Finally, the external electronics will monitor the external power sources, indicate to the patient when the external
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Fig. 7. The complete electronics system for the TAH and VAD performs the motor control, power delivery, telemetry, and monitoring functions.
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Fig. 8. The prototype implanted electronics package contains the internal batteries, motor control board, and energy transmission circuit.
battery packs require recharging, and interface to a diagnostics terminal when necessary. The motor control, power conditioning electronics, and the backup batteries are located in a titanium can implanted abdominally. Fig. 8 shows a prototype version of the implanted electronics which contain the implanted batteries, energy transmission circuit, and the motor control board. The present system uses O-ring seals for the can lid and the cable feed-throughs. Future designs will employ welded hermetic seals. IV. In Vitro PERFORMANCE The mean power requirements for the complete 100 cc TAH system as tested on the mock circulatory loop are shown in Fig. 9. Similar characteristics exist for the VAD device. The system power includes the control electronics
7 a Flow (litershin)
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Fig. 9. Results of in vitro performance testing of the 100 cc TAH on the mock circulatory loop. The system power is power input to the complete TAH system, including the energy transmission system and control electronics.
and the energy transmission system. The left atrial pressure was maintained within the 8.5-13.0 mmHg range by the control system while the right atrial pressure varied from 5.5 to 15.0 mmHg. In vitro testing of the 70 cc TAH device is beginning. V. In Vivo TESTING The motor driven rollerscrew left ventricular assist device has been implanted in 17 calves (12 using 100 cc devices and 5 with 70 cc devices) [I 13 with the longest surviving 174 days and an average survival of 52.4 days. All of these animal studies have utilized the compliance chamber rather than externally venting the motor chamber. The electronics have remained external while the internal electronics are under development. The eight most recent animals were powered via the energy transmission
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(c) (d) Fig. IO. The results of in vivo studies in a total of 24 calves are shown for the rollerscrew VAD and TAH devices. Each bar represents the mean and standard deviation for one animal, from measurements recorded once every 8 h. Estimated flow is calculated as the product of the beat rate and the estimated stroke volume, as determined by the motor controller. The TAH nominal stroke volume is 100 cc. Flow rates are higher (through higher beat rates) for the TAH animals in order to provide complete circulatory support, while in the VAD animals, the intact natural heart contributes to the total blood flow. System power is the input electrical power provided to the motor controller or to the TETS, which was utilized on 8 VAD animals for approximately 75 percent of the time. Variations in system power among similar devices at similar flow rates are mainly due to variations in outlet graft resistance and arterial pressure, differences in mechanical and commutation efficiency among devices, and the number of days using TETS power in the VAD animals.
system, with the transmitted power returned percutaneously to the external electronics. The 100 cc rollerscrew TAH has been implanted in seven calves [ 121, with a longest survival of 134 days and an average of 52.1 days. These studies have also employed a compliance chamber and external electronics but have not utilized the energy transmission system. Fig. 10 presents a summary of in vivo results for all rollerscrew devices implanted to date. The reasons for termination of the animal studies are varied, including a broken cable, a cracked pump case, moisture leakage to the Hall sensors, and rust due to improper rollerscrew material. Physiologic complications have included infection, pneumonia, thromboembolism, and occlusion of inlet or outlet cannulae. VI. DISCUSSION Development of permanent circulatory S U P P O ~systems at OUT institution 1s progressing towards the integration of
implantable VAD and TAH systems. Valuable experience has been gained from the pneumatically-powered devices, particularly in the areas of blood pump fabrication, biocompatibility, implantation techniques, patient management and interaction, and control methods. Many of the technical challenges in designing an implantable, electrically-powered blood pump, a miniature controller and successful control method, and a transcutaneous energy and data transmission system have been met. Future efforts will emphasize the long term viability and reliability of these systems.
REFERENCES [ I ] W. E. Pae, C. B . Wisman, W. S . Pierce, J . L. Myers. D. B. Campbell. and J . A. Waldhausen. “Staged cardiac transplantation: Total artificial heart or ventricular assist pump?.” Circ. Curdiov. Surg.. to be published. [2] D. J . Farrar. J . D. Hill. L. A. Gray, D. G. Pennington, L. R. McBridc, W. s. pierce, W. E. Pae, B. Glenville, D. Ross. T. A . Galbraith, and G. L. Zumbro, “Heterotopic prosthetic ventricles as
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a bridge to cardiac transplantation. A multicenter study in 29 patients,” New Eng. J . Med., vol. 318, no. 6 , pp. 333-340, 1988. 13) W . E. Richenbacher, W . E. Pae, J . A. McGovern, G . Rosenberg, A. J . Snyder, and W. S . Pierce, “A rollerscrew electric motor ventricular assist device.” Trans. Amer. Soc. A r f q . Intern. Organs, vol. 32, pp. 46-48, 1986. 141 A. J. Snyder, W. J. Weiss, W. S . Pierce, and R. A. Nazarian. “Microcomputer control of permanently implanted blood pumps,” in Proc. 1989 S w p . Eng. Corny. Based Med. S f s t . , June 1989. IS] G . Rosenberg, T. J. Cleary, A. J . Snyder, D. L. Landis, D. B. Geselowitz, and W. S . Pierce, “A totally implantable artificial heart design,’’ Trans. Amer. Soc. Mech. Eng., vol. 85-WA/DE-l I , 1985. 161 A. J. Snyder, G. Rosenberg, D. L. Landis, W . J . Weiss, and W . S. Pierce, “Control.” in Proc. Second World Symp. Artif. H e a r f . , E. S . Bucherl, Ed. Friedr. Vieweg & Sohn, Braunschweig, Wiesbaden, 1986, pp. 167-190. [7] A. J . Snyder, G. Rosenberg, and D. L. Landis, “Indirect estimation of circulatory pressures for control of an electric motor driven total artificial heart,” Trans. Amer. Soc. Mech. E n g . , vol. 85-WA/DE-I I , 1985. [8] C. Sherman, B. Daly, K . Dasse, W. Clay, M. Szycher, J. Handrahan, J. Schuder, M. Lewis, M. Worthington, R. Hopkins, and V. Poirier, “Research and development: Systems for transmitting energy through intact skin,” Final Tech. Rep. N01-HV-0-2903-3 for Devices and Technol. Branch, DHVD, National Heart, Lung and Blood Inst., Nat. Inst. Health, July, 1983. 191 J. C. Schuder, J. H. Gold, and H. E. Stephenson, Jr., “An inductively coupled RF system for the transmission of 1 kW of power through the skin.” IEEE Trans. Biomedical Eng., vol. BME-18, July 1971. [ l o ] W . 3 . Weiss. P. J . Gibbons, R. P . Gaumond. A. J . Snyder, G . Rosenberg, and W. S . Pierce. “A telemetry system for the implanted total artificial heart and ventricular assist device,” in Proc. Ninrlr Annu. Con$, IEEE. Eng. Med. Biol. Soc., vol. I . 1987, pp. 186187. 1111 G . Rosenberg, A. I . Snyder, W . J . Weiss, T . J . Cleary, and W . S. Pierce, “A permanent left ventricular assist device: In vivo testing,” in Proc. Tenth Annu. Con$. , IEEE Eng. Med. Biol. Soc., 1988, pp. 65-67. [I21 G . Rosenberg, W. S. Pierce, A. J. Snyder, W. J . Weiss, D. L. Landis, W. E. Pae, and J. A. McGovern, “ I n vivo testing of a rollerscrew type electric total artificial heart,” in Assisfed Circulation 3, Felix Unger, Ed. New York: Springer, 1987.
William J. Weiss (M’88) received the B.S degree in electrical engineering in 1985 and the M.S. degree in bioengineering in 1988 from Penn State University, University Park, PA He is currently rerearch assistant in the Department of Surgery, Division of Artificial Organs, at the Milton S Hershey Medical Center of The Pennsylvania State University, Hershey, PA, where he has been involved with the artificial heart program since 1980 He is currently working on power delivery and data communication methods for use with the implanted total artificial heart and ventricular assist device, while pursuing the Ph.D. degree in bioengineering. Mr Weiss is a member of the American Society of Artificial Internal Organs and the IEEE Engineering in Medicine and Biology Society.
@erson Rosenberg (M’88) was born in Philadelphia, PA, on August 20, 1944 He received the B S . degree and the Ph.D degree in mechanical engineering from The Pennsylvania State University, University Park, PA in 1970 and 1975, respect iv ely He has been associated with The Pennsylvania State University since starting his graduate work and IS presently a Professor i n Bioengineering and Research Professor in the College of Medicine. His main research efforts have been concerned
with mechanical circulatory assistance, implantable blood pumps and their design and the electric artificial heart, for which he has been awarded a five-year contract by the National Institutes of Health for his research directed towards developing an implantable cardiac biventricular assist and replacement device. Dr. Rosenberg is a member of Sigma Xi-The Scientific Research Society of North America and a recipient of the Outstanding Alumnus Award of the Ogontz Campus of The Pennsylvania State University (1982).
Alan J. Snyder (S’77-M’87) was born in Philadelphia, PA, in 1956 He received the B S degree in engineering science in 1978 and the Ph D degree in bioengineering in 1987 from the Pennsylvania State University, University Park, PA He is currently a Research Associate in Surgery in the College of Medicine and an Assistant Professor of Bioengineering at The Pennsylvania State University His research interests include implantable blood pumps, design of implantable devices i n general, and control systems Dr Snyder is a member of The American Society for Artificial Internal Organs, and an Associate Member of Sigma Xi, The Scientific Research Society
Thomas J. Cleary (M’86) received the B S degree in electrical engineering technology in 1977 from The Pennsylvania State University, Harrisburg, PA. He is presently a Research Assistant in the Department of Surgery, Division of Artificial Organs, at The Pennsylvania State University, Hershey, PA. Mr Cleary is a member of the Engineering in Medicine and Biology Society.
Roger P. Gaumond (M’70-SM’86) received the B S . degree in electrical engineering from Massachusetts Institute of Technology, Cambridge, i n 1968, the M.Eng degree from California Polytechnic State University, San Luis Obispo. in 1974, and the D.Sc degree in electrical engineering from Washington University, St. Louis, MO, in 1980. From 1968 to 1973, he served in the U S Air Force as Electronics Officer, responsible for the development of telemetry and data processing systems. From 1974 to 1980 he was a National Institutes of Health Trainee in Engineering Biophysics and a Research Assistant at the Computer Systems Laboratory, Washington University. He joined the Bioengineering Program, Pennsylvania State University, University Park, in 1980 where he was named Associate Professor in 1986. His research interests are in instrumentation and signal processing applications to electrophysiological monitoring and to the development of implantable prostheses Dr Gaumond is a member of Eta Kappa Nu, the IEEE Engineering in Medicine and Biology Society, the Biomedical Engineering Society, and serves on the steering committee for the Northeast Bioengineering Conference
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PERMANENT CIRCULATORY SUPPORT SYSTEMS
David B. Geselowitz ( S ' 5 l-A'54-M'6ILSM'62F'78) was born in Philadelphia, PA. He received the B.S.E.E., M.S.E.E., and Ph.D. dcgrces from the University of Pennsylvania, Philadelphia, in 1951, 1954, and 1958. respectively. He served on the faculty of the University of Pennsylvania from 1951-1971. and since then has been on the faculty of the Pennsylvania State University, University Park, where he is now Distinguished Alumni Professor of Bioengineering and Professor of Medicine. He has served as viiiting professor at M.I.T.. Duke University, and the University of Oklahoma Health Sciences Center. He is a former Guggenheim Fellow. He is a former ON BIOMEDICAL ENGINEERING. editor of the IEEE TRANSACTIONS Dr. Geselowitz is a fellow of the American College of Cardiology. He is a founding member of the Biomedical Engineering Society and the International Society of Computerized Electrocardiology. and a member o f the American Society of Artificial Internal Organs. He received the IEEEi EMBS Career Achievement Award in 1985. He has been elected to the National Academy of Engineering.
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William S. Pierce was born in Wilkes Bdrre. PA. on January 12, 1937. He received the B S degree in chemical engineering i n 1958 from Lehigh University, Bethlehem, PA, dnd the M D degree from the University of Pennsylvania, Philadelphia, in 1962 He hd\ been associated with The Penn\ylvania State University \ince 1970 and is presently d cardiothoracic surgeon in the Department of Surgery, College of Medicine He has organized an interdisciplinary group of physicians. engineers, vcterinarians, and technicians to develop a ventricular assist device and total artificial heart He has developed the total artificial heart tor which he ha5 received FDA approval for clinical implantation He I \ Chief of the Division of Artificial Organs in the Department of Surgery, College of Medicine, The Pennsylvania State University He has had eight U S patent\ awarded to him, as well a5 patents from Canada. Greece, and Australia for the right ventricular assist device Dr Pierce is an Evan Pugh Professor of Surgery, the highest academic distinction awarded by Penn State University He also is the recipient of the Jane A. Fetter Professorship of Surgery. Among his awards and honors. he has most recently been awarded an Honorary Doctor ot Science from Lehigh University
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING. VOL. 37. NO. 2 . FEBRUARY 1990
Permanent Circulatory Support Systems at The Pennsylvania State University
Abstract-Permanent circulatory support systems are required for patients in whom myocardial damage is irreversible and cardiac transplantation is not possible. Two systems are described which provide long term circulatory support: the left ventricular assist system and the total artificial heart. These systems are based on the design of a pusher plate actuated blood pump, driven by a small brushless dc electric motor and rollerscrew driver. An implantable motor controller maintains suitable physiologic Row rates for both systems and controls left-right balance in the total artificial heart. Other parts of the system include an intra-thoracic compliance chamber, transcutaneous energy and data transmission system, and internal and external batteries.
I . INTRODUCTION HE development of implantable circulatory support systems intended for permanent use has been a major focus of the research effort at the Pennsylvania State University. Pneumatically-powered devices developed at our institution and others are being used successfully as ventricular assist devices (VAD) or total artificial hearts (TAH), primarily as a bridge to transplantation [l], [2]. For patients requiring long term support, due to difficulty in obtaining suitable tissue matching, renal complications, or simply the growing shortage of donor organs, an acceptable permanent circulatory support system is desirable. The permanent VAD system is intended for patients with primarily left ventricular failure but with satisfactory right ventricular function. The permanent TAH system is intended as a complete cardiac replacement.
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11. THE VENTRICULAR ASSISTSYSTEM The left ventricular assist system (Fig. 1) [ 3 ] consists of implantable and external components. Blood flows into the assist pump from the apex of the weakened left ventricle during pump diastole and is returned to the patient's aorta during pump systole. The assist pump acts in parManuscript received December 1, 1988; revised May 1. 1989. This work was supported in part by Public Health Service Grants SROI HL 20356 and ROI HL 13426: Public Health Service Contract NO1 HV88105; The Robert J . Kleberg. Jr. and Helen C. Kleberg Foundation, and The H. G . Barsumian. M . D. Memorial Fund. The authors are with the Department of Surgery. Division of Cardiothoracic Surgery, The Milton S . Hershey Medical Center. The Pennsylvania State University, Hershey. PA 17033. IEEE Log Number 8932196.
Battery pack
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Electr;"os
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Fig. I . The left ventricular assist system consists of the electric motor driven assist pump, inlet and outlet cannulae, the intrathoracic compliance chamber with infusion port, the electronic controller and backup battery pack, and the energy transmission internal coil. The external components are the energy transmission external coil and the battery and electronics pack.
allel with the natural left ventricle, and can completely replace the left heart function. The two major components of the motor driven assist pump (Fig. 2) are the blood pump and the motor drive assembly. The blood pump consists of a flexible, seamless, segmented polyurethane blood sac with a smooth blood contacting surface. The outer case is machined polysulfone. Bjork-Shiley convexo-concave Delrin disc valves are fitted into the inlet and outlet ports. The blood sac is compressed by a pusher plate during pump systole, while during diastole the plate moves away from the sac, allowing the sac to fill passively. Two sizes of the assist pump have been developed. The clinical size device of Fig. 1 has a stroke volume of ap-
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Velocity
Estimate Cardiac Output
Fig. 3. Functional block diagram of the VAD and TAH control system
The magnitude of the applied voltage is provided by the velocity control section according to a motor voltage versus position table, which is determined as follows. A desired velocity profile is computed for both the systolic and diastolic strokes which will minimize the inertial forces and is characterized by a constant acceleration during the first half of the stroke and constant deceleration during the second half. The systolic and diastolic portions of this profile are scaled according to the desired systolic and diastolic durations, as provided by the cardiac output control. The difference between the actual and the desired velocity, or velocity error, is calculated through each pump cycle and is used to modify the motor voltage table for the next cycle. This technique is referred to as adaptive feedforward control. Velocity errors within a given proximately 70 cc, a maximum output of 8 l/min and stroke are generally small since, as motor velocity deuses 25 mm outlet and 27 mm inlet valves. A larger size creases, the reduced back emf allows motor current to inintended for experimental implantation in calves has a 100 crease, acting as an inherent negative feedback mechacc stroke volume, 10 1/min maximum output, and 25 mm nism for velocity control. Velocity feedback is used on outlet and 29 mm inlet valves. the current cycle only if the velocity falls below half the A brushless dc motor provides the driving torque which desired velocity. is converted to the linear pusher plate motion by a rollerUsing the motor voltage table and the actual velocity, screw actuator. The three phase, 12 V motor utilizes neo- the load at the pump outlet is estimated by subtracting the dymium iron magnets. The motor rotor is attached to the estimated load contributions of the frictional components, inertia of the blood and of the rotating mechanism, outlet outer nut of the rollerscrew, which is internally threaded and houses a number of threaded rollers which move in a valve pressure drop, and the electromechanical properties planetary motion about the central threaded shaft. This of the motor. The resulting estimated load versus position arrangement eliminates sliding friction and thereby pro- is made available to the cardiac output control. vides high efficiency and long life. The rotor and rollerIn the VAD controller, pump output is varied in rescrew nut rotate on two concentric bearings mounted on sponse to the filling rate. The degree of filling of the preeither side of the stator in the titanium outer case. A guide vious pump cycle is found by noting the motor position rod and bushing prevents rotation of the pusher plate. A at which the load increases, indicating the point of contact magnet ring mounted to the rotor is positioned above 3 with the blood sac. Inadequate filling is identified when Hall effect sensors which are used to sense the rotary po- the blood sac contact occurs some time after the onset of sition. The weight of the VAD motor drive assembly and systole. The output control attempts to maintain a pump blood pump is approximately 600 g. rate as high as possible while maintaining adequate filling The operation of the microprocessor-based controller by adjusting the scale factors used to calculate the desired may be described by a number of functional levels [4], velocity profiles. The effect is to pump whatever blood is summarized in Fig. 3. At the lowest level, the Hall sensor available to the assist pump. An arterial pressure estimate is also derived by averoutputs are decoded and the position of the pusher plate is determined by keeping track of the changes in rotary aging the estimated load during systole. The beat rate is position. Motor reversal is initiated at predetermined lim- reduced if the arterial pressure estimate is greater than a its of pusher plate position. Voltage is applied to the mo- preset limit, as may occur in a hypertensive episode. tor windings by pulse width modulation of semiconductor As blood flows into and out of the assist pump, pressure power switches, in a sequence determined by the rotary changes in the air space surrounding the electric motor position and desired direction. must be minimized in order to allow adequate pump fill-
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ing and ejection. This is accomplished by allowing volume changes to occur in the intrathoracic compliance chamber, which is connected by a tube to the motor housing. The compliance chamber is a flexible segmented polyurethane sac approximately 16 cm in diameter and 0.65 cm in cross section, with a fully inflated volume of 250 cc, which is velour covered to promote a stable, flexible tissue overgrowth. Diffusion of gases through the walls of the compliance chamber and blood sac leads to a gradual loss in the compliance chamber volume. The gas volume must be replenished through a subcutaneous infusion port, typically once every three weeks. 111. THE TOTALARTIFICIAL HEARTSYSTEM
The implantable total artificial heart system (Fig. 4) [5] is functionally similar to the left ventricular assist system. The major difference is that the TAH is located in the pericardial space and completely replaces the natural ventricles. The left and right pump inlets are connected to the left and right atrial remnants, respectively. The left pump outlet connects to the aorta and the right pump outlet to the pulmonary artery. The TAH device (Figs. 5 , 6) also employs a 12 V brushless dc motor and rollerscrew. Two pusher plates are required; one on either end of the rollerscrew shaft. The left and right blood sacs are therefore compressed alternately so that left systole corresponds to right diastole, and vice versa, in contrast to the simultaneous left and right ejection of the natural heart. The segmented polyurethane blood sacs are similar to the assist pump sac. Both 70 cc clinical size and 100 cc calf size devices are being developed. Control system requirements for the TAH differ significantly from those of the VAD. Of primary importance is the maintenance of left atrial pressure within a normal physiologic range. The left atrial pressure is primarily a result of the difference in fluid volume between the systemic and pulmonary circulations. In the natural heart, ventricular output (flow) increases with increasing atrial pressure (the Starling mechanism). This mechanism results in left atrial pressures which, for a healthy left ventricle, always remain in a limited range. The artificial heart control system must provide an equivalent mechanism. In the artificial heart, valve backflow increases with increasing outlet pressure, resulting in a lower net stroke volume for the left pump than for the right pump. In addition, the left pump alone carries a portion of the bronchial circulation, resulting in a slightly higher output requirement for the left pump. The TAH cardiac output control system estimates the completeness of left pump filling by analyzing the pusher plate load, as in the VAD system. When inadequate filling is determined, left systolic speed is decreased. Since left systole corresponds to right diastole, a longer filling time is allowed for the right pump, which increases the right output and the left atrial pressure. Conversely, left sys-
primary coil secondary coil
Fig. 4. The total artificial heart system. The system components are similar to those of the left ventricular assist system, with the exception of the blood pump and motor drive configuration.
Fig. 5 . The total artificial heart device showing the left and right blood pumps and the centrally located motor and rollerscrew.
tolic speed is increased when complete left filling is achieved, thereby reducing right diastolic time and limiting right pump filling. The control system attempts to maintain a stable operating point at which the left pump is just adequately filling.
WEISS er al.: PERMANENT CIRCULATORY SUPPORT SYSTEMS
Fig. 6. The assembled 100 cc stroke volume total artificial heart, with the compliance chamber, infusion port, and percutaneous cable for animal testing.
For the case of rising right atrial pressure, the increase in right pump stroke volume due to more complete filling, and the resultant increase in left atrial pressure, leads the controller to respond with a more rapid left systole. If the left diastolic speed is unchanged, the pump rate, and therefore the overall cardiac output, is increased. Thus, the textbook Starling response [6] to right atrial pressure is duplicated. By altering how the controller changes left diastolic speed when it changes left systolic speed, we can either enhance or suppress this response. In the natural system, this effect is modulated by neural and endocrine controls, and the Starling mechanism is solely responsible for maintenance of left-right balance. In our experience in calves, right atrial pressure has been an unreliable measure of demand for cardiac output. We therefore suppress sensitivity of cardiac output to right atrial pressure, and use the control described above only to maintain balance. We can then, through manipulation of the range of left diastolic speeds available to the balance control, vary cardiac output according to peripheral resistance, as reflected in systemic arterial pressure. Arterial pressure is estimated from the magnitude of the motor load during ejection of the left pump 171. IV. POWERDELIVERY A N D MONITORING The complete electronics system for both the VAD and TAH (Fig. 7) consists of motor control, power delivery, and monitoring functions. A significant advantage of the electrically driven systems over their pneumatically-powered predecessors is the use of a transcutaneous energy transmission system (TETS) to transfer electrical energy across the intact skin [8], [9]. A three-turn external, or primary, coil is placed in close proximity (typically 1.5-2.5 cm) to a 16-turn implanted, or secondary, coil. Both coils are tuned to 160 kHz by means of low loss series capacitors. An external power oscillator drives the primary coil and induces a cur-
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rent in the secondary coil, which is rectified and filtered by the implanted electronics. The voltage supplied to the motor controller is maintained within a 13.0-14.5 V band by the action of a voltage regulator located in the internal electronics package. The mean power requirements for the VAD and TAH are typically below 13 W depending upon flow rate and aortic pressure, but short duration, peak power demand at the beginning of left systole may be much higher. As a result, the energy transmission system is designed to provide up to 70 W if necessary. The efficiency of the energy transmission system from dc input voltage to dc output ranges from 55 to 75 percent. The external power source is either a 12 V power supply connected to a 115 VAC source, or a 12 V battery pack to allow for patient mobility. The total external pack capacity is approximately 10 amp-hours (Ah), divided between two separate battery packs to provide power source redundancy. The maximum operating time between recharges is approximately 10 h, based on a mean motor power of 9 W. At this time, Nickel-Cadmium batteries are being evaluated and Lithium technology is being pursued. The internal battery pack is intended only as a backup power source during separation of the TETS coils. The internal pack is the major determinant for the size and weight of the implanted electronics package. A discharge time of 45 min has been set as a minimum specification, and nickel-cadmium batteries in the 0.45-1 .O Ah range are being evaluated. In the future, the increased energy density of rechargeable lithium batteries is expected to allow further reductions in the size of the implanted package. A number of charging profiles are also being considered in an attempt to minimize the permissible time between discharges while retaining acceptable cycle life and reliability. A telemetry system for bidirectional communication with the implanted electronics is being developed which utilizes the energy transmission coils and power carrier [lo]. Primarily, the system will be used to monitor the performance of the motor driven TAH or VAD. This information includes the motor rate and the pusher plate load as determined by the motor controller, as well as the status of the internal battery pack. This internal to external data transmission employs impedance modulation on the secondary side and demodulation of the external coil current on the primary side. Secondarily, changes may be made to the implanted controller parameters via external to internal data transmission, utilizing frequency modulation of the power carrier. External to internal communication is not required for the normal operation of the implanted device, but this capability is extremely useful during the operative and immediate post-operative periods while the hemodynamics are being stabilized, and as a means of compensating for gradual changes in the pump inflow and outflow characteristics, if necessary. Finally, the external electronics will monitor the external power sources, indicate to the patient when the external
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IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 37. NO. 2. FEBRUARY 1990 I
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Fig. 8. The prototype implanted electronics package contains the internal batteries, motor control board, and energy transmission circuit.
battery packs require recharging, and interface to a diagnostics terminal when necessary. The motor control, power conditioning electronics, and the backup batteries are located in a titanium can implanted abdominally. Fig. 8 shows a prototype version of the implanted electronics which contain the implanted batteries, energy transmission circuit, and the motor control board. The present system uses O-ring seals for the can lid and the cable feed-throughs. Future designs will employ welded hermetic seals. IV. In Vitro PERFORMANCE The mean power requirements for the complete 100 cc TAH system as tested on the mock circulatory loop are shown in Fig. 9. Similar characteristics exist for the VAD device. The system power includes the control electronics
7 a Flow (litershin)
9
10
Fig. 9. Results of in vitro performance testing of the 100 cc TAH on the mock circulatory loop. The system power is power input to the complete TAH system, including the energy transmission system and control electronics.
and the energy transmission system. The left atrial pressure was maintained within the 8.5-13.0 mmHg range by the control system while the right atrial pressure varied from 5.5 to 15.0 mmHg. In vitro testing of the 70 cc TAH device is beginning. V. In Vivo TESTING The motor driven rollerscrew left ventricular assist device has been implanted in 17 calves (12 using 100 cc devices and 5 with 70 cc devices) [I 13 with the longest surviving 174 days and an average survival of 52.4 days. All of these animal studies have utilized the compliance chamber rather than externally venting the motor chamber. The electronics have remained external while the internal electronics are under development. The eight most recent animals were powered via the energy transmission
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a l . : PERMANENT CIRCULATORY SUPPORT SYSTEMS 200,
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(c) (d) Fig. IO. The results of in vivo studies in a total of 24 calves are shown for the rollerscrew VAD and TAH devices. Each bar represents the mean and standard deviation for one animal, from measurements recorded once every 8 h. Estimated flow is calculated as the product of the beat rate and the estimated stroke volume, as determined by the motor controller. The TAH nominal stroke volume is 100 cc. Flow rates are higher (through higher beat rates) for the TAH animals in order to provide complete circulatory support, while in the VAD animals, the intact natural heart contributes to the total blood flow. System power is the input electrical power provided to the motor controller or to the TETS, which was utilized on 8 VAD animals for approximately 75 percent of the time. Variations in system power among similar devices at similar flow rates are mainly due to variations in outlet graft resistance and arterial pressure, differences in mechanical and commutation efficiency among devices, and the number of days using TETS power in the VAD animals.
system, with the transmitted power returned percutaneously to the external electronics. The 100 cc rollerscrew TAH has been implanted in seven calves [ 121, with a longest survival of 134 days and an average of 52.1 days. These studies have also employed a compliance chamber and external electronics but have not utilized the energy transmission system. Fig. 10 presents a summary of in vivo results for all rollerscrew devices implanted to date. The reasons for termination of the animal studies are varied, including a broken cable, a cracked pump case, moisture leakage to the Hall sensors, and rust due to improper rollerscrew material. Physiologic complications have included infection, pneumonia, thromboembolism, and occlusion of inlet or outlet cannulae. VI. DISCUSSION Development of permanent circulatory S U P P O ~systems at OUT institution 1s progressing towards the integration of
implantable VAD and TAH systems. Valuable experience has been gained from the pneumatically-powered devices, particularly in the areas of blood pump fabrication, biocompatibility, implantation techniques, patient management and interaction, and control methods. Many of the technical challenges in designing an implantable, electrically-powered blood pump, a miniature controller and successful control method, and a transcutaneous energy and data transmission system have been met. Future efforts will emphasize the long term viability and reliability of these systems.
REFERENCES [ I ] W. E. Pae, C. B . Wisman, W. S . Pierce, J . L. Myers. D. B. Campbell. and J . A. Waldhausen. “Staged cardiac transplantation: Total artificial heart or ventricular assist pump?.” Circ. Curdiov. Surg.. to be published. [2] D. J . Farrar. J . D. Hill. L. A. Gray, D. G. Pennington, L. R. McBridc, W. s. pierce, W. E. Pae, B. Glenville, D. Ross. T. A . Galbraith, and G. L. Zumbro, “Heterotopic prosthetic ventricles as
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a bridge to cardiac transplantation. A multicenter study in 29 patients,” New Eng. J . Med., vol. 318, no. 6 , pp. 333-340, 1988. 13) W . E. Richenbacher, W . E. Pae, J . A. McGovern, G . Rosenberg, A. J . Snyder, and W. S . Pierce, “A rollerscrew electric motor ventricular assist device.” Trans. Amer. Soc. A r f q . Intern. Organs, vol. 32, pp. 46-48, 1986. 141 A. J. Snyder, W. J. Weiss, W. S . Pierce, and R. A. Nazarian. “Microcomputer control of permanently implanted blood pumps,” in Proc. 1989 S w p . Eng. Corny. Based Med. S f s t . , June 1989. IS] G . Rosenberg, T. J. Cleary, A. J . Snyder, D. L. Landis, D. B. Geselowitz, and W. S . Pierce, “A totally implantable artificial heart design,’’ Trans. Amer. Soc. Mech. Eng., vol. 85-WA/DE-l I , 1985. 161 A. J. Snyder, G. Rosenberg, D. L. Landis, W . J . Weiss, and W . S. Pierce, “Control.” in Proc. Second World Symp. Artif. H e a r f . , E. S . Bucherl, Ed. Friedr. Vieweg & Sohn, Braunschweig, Wiesbaden, 1986, pp. 167-190. [7] A. J . Snyder, G. Rosenberg, and D. L. Landis, “Indirect estimation of circulatory pressures for control of an electric motor driven total artificial heart,” Trans. Amer. Soc. Mech. E n g . , vol. 85-WA/DE-I I , 1985. [8] C. Sherman, B. Daly, K . Dasse, W. Clay, M. Szycher, J. Handrahan, J. Schuder, M. Lewis, M. Worthington, R. Hopkins, and V. Poirier, “Research and development: Systems for transmitting energy through intact skin,” Final Tech. Rep. N01-HV-0-2903-3 for Devices and Technol. Branch, DHVD, National Heart, Lung and Blood Inst., Nat. Inst. Health, July, 1983. 191 J. C. Schuder, J. H. Gold, and H. E. Stephenson, Jr., “An inductively coupled RF system for the transmission of 1 kW of power through the skin.” IEEE Trans. Biomedical Eng., vol. BME-18, July 1971. [ l o ] W . 3 . Weiss. P. J . Gibbons, R. P . Gaumond. A. J . Snyder, G . Rosenberg, and W. S . Pierce. “A telemetry system for the implanted total artificial heart and ventricular assist device,” in Proc. Ninrlr Annu. Con$, IEEE. Eng. Med. Biol. Soc., vol. I . 1987, pp. 186187. 1111 G . Rosenberg, A. I . Snyder, W . J . Weiss, T . J . Cleary, and W . S. Pierce, “A permanent left ventricular assist device: In vivo testing,” in Proc. Tenth Annu. Con$. , IEEE Eng. Med. Biol. Soc., 1988, pp. 65-67. [I21 G . Rosenberg, W. S. Pierce, A. J. Snyder, W. J . Weiss, D. L. Landis, W. E. Pae, and J. A. McGovern, “ I n vivo testing of a rollerscrew type electric total artificial heart,” in Assisfed Circulation 3, Felix Unger, Ed. New York: Springer, 1987.
William J. Weiss (M’88) received the B.S degree in electrical engineering in 1985 and the M.S. degree in bioengineering in 1988 from Penn State University, University Park, PA He is currently rerearch assistant in the Department of Surgery, Division of Artificial Organs, at the Milton S Hershey Medical Center of The Pennsylvania State University, Hershey, PA, where he has been involved with the artificial heart program since 1980 He is currently working on power delivery and data communication methods for use with the implanted total artificial heart and ventricular assist device, while pursuing the Ph.D. degree in bioengineering. Mr Weiss is a member of the American Society of Artificial Internal Organs and the IEEE Engineering in Medicine and Biology Society.
@erson Rosenberg (M’88) was born in Philadelphia, PA, on August 20, 1944 He received the B S . degree and the Ph.D degree in mechanical engineering from The Pennsylvania State University, University Park, PA in 1970 and 1975, respect iv ely He has been associated with The Pennsylvania State University since starting his graduate work and IS presently a Professor i n Bioengineering and Research Professor in the College of Medicine. His main research efforts have been concerned
with mechanical circulatory assistance, implantable blood pumps and their design and the electric artificial heart, for which he has been awarded a five-year contract by the National Institutes of Health for his research directed towards developing an implantable cardiac biventricular assist and replacement device. Dr. Rosenberg is a member of Sigma Xi-The Scientific Research Society of North America and a recipient of the Outstanding Alumnus Award of the Ogontz Campus of The Pennsylvania State University (1982).
Alan J. Snyder (S’77-M’87) was born in Philadelphia, PA, in 1956 He received the B S degree in engineering science in 1978 and the Ph D degree in bioengineering in 1987 from the Pennsylvania State University, University Park, PA He is currently a Research Associate in Surgery in the College of Medicine and an Assistant Professor of Bioengineering at The Pennsylvania State University His research interests include implantable blood pumps, design of implantable devices i n general, and control systems Dr Snyder is a member of The American Society for Artificial Internal Organs, and an Associate Member of Sigma Xi, The Scientific Research Society
Thomas J. Cleary (M’86) received the B S degree in electrical engineering technology in 1977 from The Pennsylvania State University, Harrisburg, PA. He is presently a Research Assistant in the Department of Surgery, Division of Artificial Organs, at The Pennsylvania State University, Hershey, PA. Mr Cleary is a member of the Engineering in Medicine and Biology Society.
Roger P. Gaumond (M’70-SM’86) received the B S . degree in electrical engineering from Massachusetts Institute of Technology, Cambridge, i n 1968, the M.Eng degree from California Polytechnic State University, San Luis Obispo. in 1974, and the D.Sc degree in electrical engineering from Washington University, St. Louis, MO, in 1980. From 1968 to 1973, he served in the U S Air Force as Electronics Officer, responsible for the development of telemetry and data processing systems. From 1974 to 1980 he was a National Institutes of Health Trainee in Engineering Biophysics and a Research Assistant at the Computer Systems Laboratory, Washington University. He joined the Bioengineering Program, Pennsylvania State University, University Park, in 1980 where he was named Associate Professor in 1986. His research interests are in instrumentation and signal processing applications to electrophysiological monitoring and to the development of implantable prostheses Dr Gaumond is a member of Eta Kappa Nu, the IEEE Engineering in Medicine and Biology Society, the Biomedical Engineering Society, and serves on the steering committee for the Northeast Bioengineering Conference
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PERMANENT CIRCULATORY SUPPORT SYSTEMS
David B. Geselowitz ( S ' 5 l-A'54-M'6ILSM'62F'78) was born in Philadelphia, PA. He received the B.S.E.E., M.S.E.E., and Ph.D. dcgrces from the University of Pennsylvania, Philadelphia, in 1951, 1954, and 1958. respectively. He served on the faculty of the University of Pennsylvania from 1951-1971. and since then has been on the faculty of the Pennsylvania State University, University Park, where he is now Distinguished Alumni Professor of Bioengineering and Professor of Medicine. He has served as viiiting professor at M.I.T.. Duke University, and the University of Oklahoma Health Sciences Center. He is a former Guggenheim Fellow. He is a former ON BIOMEDICAL ENGINEERING. editor of the IEEE TRANSACTIONS Dr. Geselowitz is a fellow of the American College of Cardiology. He is a founding member of the Biomedical Engineering Society and the International Society of Computerized Electrocardiology. and a member o f the American Society of Artificial Internal Organs. He received the IEEEi EMBS Career Achievement Award in 1985. He has been elected to the National Academy of Engineering.
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William S. Pierce was born in Wilkes Bdrre. PA. on January 12, 1937. He received the B S degree in chemical engineering i n 1958 from Lehigh University, Bethlehem, PA, dnd the M D degree from the University of Pennsylvania, Philadelphia, in 1962 He hd\ been associated with The Penn\ylvania State University \ince 1970 and is presently d cardiothoracic surgeon in the Department of Surgery, College of Medicine He has organized an interdisciplinary group of physicians. engineers, vcterinarians, and technicians to develop a ventricular assist device and total artificial heart He has developed the total artificial heart tor which he ha5 received FDA approval for clinical implantation He I \ Chief of the Division of Artificial Organs in the Department of Surgery, College of Medicine, The Pennsylvania State University He has had eight U S patent\ awarded to him, as well a5 patents from Canada. Greece, and Australia for the right ventricular assist device Dr Pierce is an Evan Pugh Professor of Surgery, the highest academic distinction awarded by Penn State University He also is the recipient of the Jane A. Fetter Professorship of Surgery. Among his awards and honors. he has most recently been awarded an Honorary Doctor ot Science from Lehigh University
