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IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 37. NO. 2. FEBRUARY 1990
A Multiple Disk Centrifugal Pump as a Blood Flow Device GERALD E. MILLER,
MEMBER, IEEE,
BRADLEY D. ETTER,
Abstract-A multiple disk, shear force, valveless centrifugal pump was studied to determine its suitability as a blood flow device. A pulsatile version of the Tesla viscous flow turbine was designed by modifying the original steady flow pump concept to produce physiological pressures and flows with the aid of controlling circuitry. Pressures and flows from this pump were compared to a Harvard Apparatus pulsatile piston pump. Both pumps were connected to an artificial circulatory system. Frequency and systolic duration were varied over a range of physiological conditions for both pumps. The results indicated that the Tesla pump, operating in a pulsatile mode, is capable of producing physiologic pressures and flows similar to the Harvard pump and other pulsatile blood pumps.
INTRODUCTION
H
EART assist and total heart replacement systems have been studied for decades. During this time span, there have been various types and modes of operation of these systems tested. These have included pneumaticallydriven sac pumps [ l ] , [2], roller pumps [3], centrifugal impeller pumps [4], [5], pusher plate diaphragm pumps [6]-[8], and a variety of others. Of such systems, the pneumatically-driven sac pump and the electrically-powered pusher plate pump have received the most attention as either a cardiac assist device or a total heart replacement. Clinical trials of such systems have proven that such devices operate successfully in humans over a limited period in both an assist mode and as a total cardiac replacement [9], [ 101. However, a variety of complications have arisen in the development of these devices which has served to delay the application of such systems as relatively risk free prosthetic devices for humans. These complications have included the formation of pannus [ 111, thrombus formation [ 121, operational difficulties [ 131, [ 141, calcification of blood contacting surfaces [15], limited cardiac output, as well as a host of others. Although much research has been conducted into the development of problem free versions of sac type and diaphragm type devices, research has also been conducted into the development of alternative modes of operation such as the roller pump, the impeller pump, and the disk pump. In particular, the rotating disk pump, once called Manuscript received November 28, 1988; revised June 15. 1989. The authors are with the Bioengineering Program, Texas A&M University, College Station, TX 77843. IEEE Log Number 8932198.
MEMBER,
IEEE, A N D
JEAN M. DORSI
the Tesla turbine, has not been widely studied as a cardiac pumping mechanism. In the past, such a device has been utilized as a valveless, steady flow pump which can produce extremely high flow rates and pressure heads and can be electrically powered. However, the effects of steady blood flow from a valveless pump have resulted in physiological problems which have served as a limiting factor in the further development of such devices [16]-[18]. The purpose of this study was to design and test a prototype version of a pulsatile version of a multiple disk centrifugal pump, the so-called Tesla Viscous Flow Turbine. This pump was then compared to a common blood flow device, such as a Harvard piston pump, to determine whether the Tesla pump can produce physiological pressures and flows similar to those of sac type or diaphragm type cardiac pumps. The development of a pulsed disk pump might thus prove to be a viable alternative as a heart assist or total heart replacement device. CENTRIFUGAL FLOWPUMPS Since centrifugal pumps have been widely used in the past, the theoretical analysis, design, and testing of such pumps has been widely published [19], [20]. The parameters which determine the maximum pressure head are the pump speed, impeller or disk diameter, and the number of disks in series (and their spacing). Similar parameters affect the output flow and maximum shearing stresses exerted on the fluid. The major advantages of such pumps include the ability to generate large flow rates and pressure heads, and the ability to easily power and control such systems electrically. Centrifugal pumps have been studied sporadically as blood flow devices for many years. In 1964, Dorman et al. studied the utilization of a multiple disk Tesla pump as a heart assist device [21]. They utilized the characteristics of shear action pumps as described by Hassinger and Kerht [22] and by Rice [23], and then modified the disk spacing, disk diameter, and rotation speed to produce quasi-physiological flows and pressure heads, albeit in steady flow. Further studies of centrifugal pumps as heart assist or substitute devices has centered around impeller type or vane type devices. The Biomedicus Bio-Pump (Biomedicus Inc., Minneapolis, Minnesota) propels fluid by means of a force-vortex pumping principle by rotation of smooth
0018-9294/90/0200-0157$01.OO 0 1990 IEEE
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IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 37, NO. 2. FEBRUARY 1990
DISK f ?ller
I
I
I
Magnet Rotor
Fig. 1. Medtronic impeller pump for use as a heart substitute device.
rotator cones. Blood flow can be controlled from 100 ml/min to 10 l/min. It has been used in either steady flow or pulsatile mode. Medtronic (Medtronic Inc., Minneapolis, MN) has developed an impeller pump which operates either in steady or pulsatile modes (see Fig. 1). Flows of up to 10 l/min have been exhibited in vitro. Descriptions of these and similar pumps have been reported by Bernstein [SI, Johnson [24], Olsen and Bramm [25], and Hinglais et al. [26]. Due to continuing operational and pathological difficulties, the impeller and vane type heart devices have seen limited clinical application. Blood damage has been a serious drawback to the further (b) development of these systems. The use of a multiple disk Fig. 2. Tesla pump design consisting of seven parallel polycarbonate disks centrifugal pump may offer the advantages of typical cendriven by a central shaft mounted at the end disk. The spiral volute serves as a housing and a diffusing chamber. The fluid enters through the side trifugal operation (large flow, easy control, electrical port opposite the central shaft and is collected via the housing as it is power), while limiting the typical drawbacks such as propelled radially from the disks. (a) Disk dimensions. (b) Pump and blood damage and quasi-steady flow operation. A disk housing perspective. pump utilizes shear forces rather than lifting surfaces to impart momentum to a fluid by means of a series of rotating disks. nected to each other by three stainless steel shafts. The Theoretical analyses and design studies of steady flow end disk is connected to a driving motor by means of a multiple disk pumps have been conducted by Dorman et larger steel shaft located outside the pump housing. This al. [21], Hassinger and Kerht [22], Rice [23], and many shaft was sealed at the interconnection with the housing others. However, little has been done with multiple disk by means of a graphite gland packing seal. The test fluid type pumps since the study of steady flow performance by (physiologic saline) enters into the rotor area through the Dorman et al. in the middle 1960’s. The use of such end opposite the driving shaft [Fig. 2(b)]. As the fluid pumps in a pulsatile mode or as an alternative artificial enters into the spaces between the disks, it is sent into a heart design has not been previously investigated. This spiral motion whose tangential and radical velocity compaper will describe a study which was conducted regard- ponents are generated by frictional forces. As the fluid ing the design and performance of a pulsatile version of spins off the peripheral surface of the disks, it is collected the multiple disk shear pump as a possible blood flow de- by a spiral volute chamber which serves as both a pump vice. housing and a diffusing chamber. The fluid is then ejected tangentially through the outlet port of the pump [Fig. METHODSAND MATERIALS 2(b)]. The equations governing the performance of mulThe Tesla pump which was constructed for this study tiple disk pumps have been presented by Hassinger and consists of seven polycarbonate disks configured in a par- Kerht [22] and by Rice [23]. allel arrangement as seen in Fig. 2(a) and (b). Plastic was The number of disks, disk spacing, and disk diameter utilized for the disk material to provide a low moment of for this application were calculated so as to minimize the inertia, thus allowing the pump to respond quickly to a shear rate exerted on the fluid while creating sufficient change in torque produced by a pulsatile driving signal pressures and flows required to mimic cardiac function. from a motor. The inner and outer disk diameters are 1.5 The geometry of the pump, in combination with a limiting and 3.0 in, respectively [Fig. 2(a)]. Each disk is 0.063 in disk rotation speed of 2250 rpm, would produce a theothick. The disk spacing is 0.016 in. The disks are con- retical limiting shear rate of 10 000 s-’ (400 dyn/s cm2
-
MILLER
EI U / . :
159
CENTRIFUGAL P U M P AS BLOOD FLOW DEVICE
2 4
1
Duratm
b
LM74121
c3 100
I '
I
I
115'4 NEUTRAL
ECG5641
115V HOT
1PO
Fig. 3 . Controlling circuitry for the Tesla pump which produces pulsatile motion. Frequency, systolic duration, and contraction-like ejection thrust are independently controlled (see text for description).
for blood with a 4 centipoise viscosity), which is below the limiting value (400 dyn/s cm2) to avoid blood cell damage [27]. This pump was driven by a centrifugal pump motor (Oberdorfer Model 142-12). This motor normally operates as a steady flow A / C device, but was modified for this study to produce sinusoidal pulsations in order to produce pulsatile flow from the pump [28]. The controlling circuitry for this purpose is shown in Fig. 3 . A Sylvania triac is used to control the speed of the motor by acting as an electronic switch driven by a series of 5 V pulses. The motor speed varies according to the frequency and duration of the pulses. An LM555 timer, acting as an astable oscillator, drives the controlling circuitry and acts as a heart rate simulator. The rate can be varied by controlling the timer resistors. Systolic duration is controlled by an LM74121 one shot multivibrator. This signal is then sent to an LM741 operational amplifier configured as a modified lossy integrator. This integrator serves to smooth out the square wave control signal into a smoother sinusoidal waveform. This smooth signal is then sent to an LM555 pulse width modulator (PWM) which serves to gate the triac in accordance with the voltage supplied by the integrator. Typical control signals from each of these components are shown in Fig. 4. Design analyses and characteristic signals for each component of the controlling circuit were studied by Etter [28]. Constraints on the pump motor are that no flow reversal takes place (no pure reversal of the disks), no undue forces are exerted on the pump which may damage the motor,
-
I
74121
555
r-7 0
:
- 1-
,
1
2
~
3
4
Time (sec) Fig. 4. Signals from the components of the pump controlling circuit. The LM555 timer controls the pump frequency. The LM74121 monostable multivibrator controls the systolic duration. The LM741 lossy integrator smooths the multivibrator output and controls the acceleration and deceleration of the pump. The LM555 pulse width modulator sends a signal to the triac which controls the pump speed. Typical pressure and flow waveforms are shown for reference.
and that physiological pressures and flows are generated. The design of this pump is valveless (thus, the constraint on disk reversal). Diastolic (zero flow) conditions are maintained with minimal forward disk motion. The pump is connected to an artificial circulatory system as seen in Fig. 5 . This system consists of a resistor, capacitor, and reservoir and is similar to that designed by Rosenberg [29]. The resistor consists of a series of parallel, flexible, small diameter tubes which can be con-
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IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 31. NO. 2. FEBRUARY 1990
CAPACITOR
FLOW PROBE
VENOUS R E S E R V O I R PRESSURE T R A N S D U C E R
Fig. 5 . Artificial circulatory system consisting of a resistance, compliance, and atrial reservoir as designed by Rosenberg [29].
stricted by means of a plate and adjusting nut. The capacitor is a leaf spring system where the spring force opposes the increase in capacitor pressure resulting from an inrush of fluid during systole. Pressure and flow rate are measured at the pump outlet by means of a Motorola MPXlOOA pressure transducer and a Carolina Medical EP670 electromagnetic flow probe. The pressure signal is amplified by means of an Analog Devices AD521 instrumentation amplifier, while the flow signal is processed by means of a Carolina Medical FM501 Flowmeter. One inch tygon tubing is used to connect the circulatory system components. The output of the Tesla pump was compared to that produced by a Harvard Apparatus pulsatile piston pump (model 1423). This pump contains ball valves at the inlet and outlet ports. The systolic duration and pump frequency can be varied independently by separate controls on the pump console. The systolic duration control of the Harvard pump affects the forward excursion of the piston, such that stroke volume of this pump can be indirectly manipulated within a limited range. Outlet pressure and flow were recorded using a Texas Instruments Portable Professional computer with a Metrabyte DASH-8 Analog to Digital converter. The data were sampled at 100 Hz in order to comply with Nyquist’s criteria. Sampling was done in sets of 1024 data points which resulted in 10.24 s of pressure and flow data per set. The resistance and compliance settings of the circulatory system were adjusted to produce physiological pressures and flow values while the Harvard pump was connected to the circulatory system. The settings of the circulatory system were to remain unchanged following the initial setting. The shear pump has no specific stroke volume, as there are no valves. However, the timer, pulse width modulator and lossy integrator can be adjusted to produce an effective stroke volume which approximates that of the Harvard pump. The LM555 timer adjustment affects the pump frequency; the LM74 12 1 monostable multivibrator adjustment affects the width of the systolic duration; and the LM741 lossy integrator adjustment affects the rise and decay rate of the flow and pressure waveforms, which affect the acceleration and deceleration of the disk pump. At
present, these adjustments are made qualitatively. However, present research is ongoing to develop a quantifiable relationship between control circuit component settings and pump output. Once the settings on the controlling circuitry were completed to produce this “stroke volume,” these were also to remain unchanged throughout a given experiment. Data was recorded for each pump at heart rates of 60, 70, 80 and 90 beat/min and at systolic durations of 0.3, 0.35, and 0.4 s for each heart rate. This provided a total of 12 operating conditions for each pump. Frequency and systolic duration of the disk pump were adjusted before each experiment to approximate that of the Harvard pump. Conductance of the saline was checked periodically as was fluid temperature and the output of the pressure and flow transducers in order to verify that the calibration of these devices was still valid.
RESULTSAND DISCUSSION Typical pressure and flow waveforms for both the Tesla pump and the Harvard pump are shown in Fig. 6(a)-(d). A list of peak and mean pressures and flows for both pumps at all operating conditions is given in Tables I and 11. Inspection of the time varying waveforms seen in Fig. 6(a)-(d) show good agreement between the two pumps. The pressure values for the Harvard pump were slightly lower than typical physiologic norms, although the waveform shapes are typical of physiologic aortic pressure waveforms. The lower pressure values are likely a result of the settings of the circulatory system since both pumps produced similar results. Mean pressures and flows exhibited from the disk pump were slightly higher than that of the Harvard pump. This is due to inaccurate settings of the controlling circuitry of the disk pump, which were manually adjusted to approximate the systolic duration of the Harvard pump. Current research efforts include the development of a computer controlled automatic selection of parameter settings of the control circuit of the disk pump. Once completed, such a system can significantly reduce the problems associated with inaccurate matches
MILLER
cf
ol.: CENTRIFUGAL PUMP AS BLOOD FLOW DEVICE
__
Pump Pressures
Pump Pressures
H.R.=60 S.D.=.30
H.R.=90 S.D.=.40
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-Shear
Fig. 6 . Time varying output pressure and flow rate waveforms from both the Harvard pump and the Tesla pump. HR signifies heart rate in beats per minute. SD signifies systolic duration in seconds. The oscillations in the Harvard pump data are due to bouncing of the ball valve. The disk pump is not completely synchronized with the Harvard pump since the disk pump frequency and systolic duration are only approximately adjusted to that of the Harvard pump. In addition, the timing controls of the Harvard pump are not exact.
between the disk pump and a comparable piston or sac pump. However, given the limitations of the design and testing of this early prototype of the disk pump, it appears that such a design is quite capable of producing physiological pressures and flows typical of present day artificial hearts such as the sac, diaphragm, and piston pumps. Further testing is necessary in order to analyze flow conditions inside the pump and housing. Stagnation, turbulence, and cavitation are all conditions which should be monitored in such a design in order to eliminate any flow condition which would be conducive to blood damage or thrombus formation. Future designs of a disk type pump
will also address power consumption, biocompatibility , and anatomical/physiological considerations. However, such a pump, operated in a pulsatile fashion does offer advantages over other designs which are presently used. A valveless centrifugal pump does not require a compliance chamber. A rotary action pump should provide immediate response times to changing operating conditions from a drive system or controlling circuit. A valveless pump would not suffer from calcification, pannus formation nor thrombus formation at a valve seat. The use of a disk pump may also avoid the problem of fill limited conditions of a sac pump. Such fill limited operation in a valved sac type or pusher plate pump is caused
162
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 37, NO. 2. FEBRUARY I Y Y O
TABLE I HARVARD P U M PDATA
S
Peak Diast Flow l/min
Peak Syst. Flow I/min
Peak Diast . Pres. mmHg
Peak Syst. Pres. mmHg
Mean Flow I/min
Mean Pres. mmHg
0.30 0.35 0.40 0.30 0.35 0.40 0.30 0.35 0.40 0.30 0.35 0.40
0.50 0.54 0.54 0.63 0.51 0.44 0.58 0.57 0.77 0.70 0.67 1.02
13.17 13.22 12.98 14.30 14.21 14.42 15.06 15.14 15.65 15.88 15.85 16.76
60.66 59.96 59.94 61.35 62.36 63.49 63.46 63.69 67.34 65.17 65.27 68.82
96.26 95.81 95.29 97.95 94.84 94.77 95.69 95.84 99.00 97.10 97.45 101.27
4.75 4.80 5.05 5.23 5.21 5.75 5.66 5.79 6.68 6.23 6.23 7.23
74.47 73.94 74.52 75.91 75.49 76.82 76.99 77.19 81.60 78.91 79.06 83.83
Heart Rate bPm
System Duration
60 60 60 70 70 70 80 80 80 90 90 90
TABLE I1
TESLAPUMP DATA
S
Peak Diast Flow l/min
Peak Syst. Flow I/min
Peak Diast. Pres. mmHg
Peak Syst. Pres. mmHg
Mean Flow I/min
Mean Pres. mmHg
0.30 0.35 0.40 0.30 0.35 0.40 0.30 0.35 0.40 0.30 0.35 0.40
0.47 1.45 1.29 1.76 1.32 0.79 0.14 0.85 0.58 0.00 0.14 0.09
15.78 15.91 15.78 16.22 17.67 17.97 17.48 17.92 18.12 18.09 18.91 18.85
66.51 67.53 68.82 70.73 69.79 71.28 70.14 7 1.92 71.95 70.71 73.01 73.46
104.23 1 11.47 114.52 10.82 115.71 115.44 109.51 113.01 113.08 105.71 11 1.22 112.11
5.68 6.35 6.64 6.35 6.97 7.23 6.66 7.31 7.63 6.98 7.83 7.95
79.76 83.98 86.78 84.16 87.16 88.87 85.76 88.75 89.81 85.47 89.57 90.53
Heart Rate bPm
System Duration
60 60 60 70 70 70 80 80 80 90 90 90
by either a low filling pressure or an insufficient filling period (during diastole). A valveless centrifugal pump may avoid both of these problems for the following reasons: (a) there is no inlet valve to restrict inlet flow, and (b) a centrifugal pump draws fluid into it’s inlet by a vacuum action. In fact, too great a vacuum on the inlet could result in venous collapse when such a pump is utilized in vivo. This important area of concern is presently under investigation. Simultaneous measurements of inlet pressure and flow and outlet pressure and flow for the disk pump are planned to determine the effects of inlet conditions on disk pump performance. Such results will then be compared to similar measurements in a comparable piston or sac pump to ascertain the relative effects of fill limited operation on pumping performance. The results from this study suggest that further research is warranted into the development of alternative designs of pulsatile pumps such as the disk pump for use as an artificial heart or heart assist device.
REFERENCES [ I ] L. D. Joyce, W. C . DeVries, W. L. Hastings, D. B. Olsen, R. K. Jarvik, and W. J. Kolff, “Response of the human body to the first
permanent implant of the Jarvik-7 total artificial heart,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 29; p. 81, 1983. [21 R. K. Jarvik, “The total artificial heart,” Sci. Amer., vol. 244, pp. 74-80, 1981. [3] G. P. Noon, J. E. Harrell, L. Feldman, J . Peterson, P. Kent, and M. E. DeBakey, “Development and evaluation of pulsatile roller pump and tubing for cardiac assistance,” Artif. Organs, vol. 7, pp. 49-54, 1983. [4] M. F. Lynch, D. Peterson, and V. Baker, “Centrifugal blood pumping for open heart surgery,” Minn. Med,, vol. 9 , p. 536, 1978. [ 5 ] E. F. Bernstein, “A centrifugal pump for circulatory assistance,” in Assisted Circulation, F. Unger, Ed. Berlin: Springer, 1979, pp. 231242. [6] G. Rosenberg, A. Snyder, W. Weiss, D. L. Landis, D. B. Geselowitz, and W. S . Pierce, “A cam-type electric motor-driven left ventricular assist device,” J . Biomech. Eng., vol. 104, pp. 214-220, 1982. [7] J. Moise, K. Butler, J . Payne, R. Wampler, W . Smith, L. Fujimoto, L. Golding, R. Kiraly, H. Harasaki, and Y. Nose, “Experimental evaluation of complete electrically powered ventricular assist system,” Trans. Amer. Soc. Arrif. Intern. Organs, vol. 31, pp. 202205, 1985. [8] C. W. Sherman, D. Gernes, W. Clay, V. Poirier, and W. F. Bernhard, “A permanent implantable electric blood pump,” IEEE Eng. Med. Biol., pp. 26-29, 1986. [9] W . C. DeVries, J. L. Anderson, L. D. Joyce, F . L. Anderson, E. H. Hammond, R. K. Jarvik, and W. J. Kolff, ‘‘Clinical use of the total artificial heart,” New Eng. J . M e d . , vol. 310, p. 273, 1984. [IO] E. K. Olsen, W. S . Pierce, J. H. Donachy, D. L. Landis, G . Rosenberg, W. M. Phillips, G. A. Allen, and M. J. O’Neill, “A two and
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MILLER et al.: CENTRIFUGAL PUMP AS BLOOD FLOW DEVICE
one-half year clinical experience with a mechanical left ventricular assist pump in the treatment of profound post-operative heart failure,” Int. J . Artif. Organs, vol. 2, p. 197, 1979. [ I l l R. K. Jarvik, T. R. Kessler, L. D. McGill, D. B. Olsen, W. C. DeVries, J. T. Blaylock, and W. J. Kolff, “Determination of pannus formation in long-surviving artificial heart calves and its prevention,” Trans. Amer. Soc. Artif. Intern. Organs, pp. 90-95, 1981. [I21 T. R. Kessler, A. B. Pons, R. K. Jarvik, J. H. Lawson, K. J . Razzeca, and W. J. Kolff, “Elimination of predilection sites for thrombus formation in the total artificial heart-before and after,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 24, p. 532, 1978. 1131 S . Lee, G. Rosenberg, J. H. Donachy, C . B. Wisman, and W. S . Pierce, “The compliance problem: A major obstacle in the development of implantable blood pumps,’’ Art$ Organs, vol. 8, pp. 8290, 1984. [ 141 E. Hennig, A. Mohnhaupt, and E. S . Bucheri, “The influence of the filling pressure on the output of pneumatically driven blood pumps,” Proc. Europ. Soc. Artif. Organs, vol. 3, p. 19, 1976. [15] D. L. Coleman, D. Lim, T. R. Kessler, and J. D. Andrade, “Calcification of non-textured implantable blood pumps,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 27, p. 97, 1981. [16] G . Mavroudis, “To pulse or not to pulse,” Ann. Thorac. Surg., vol. 25, no. 3 , pp. 258-271, 1978. [I71 Y. Nose, “The need for a nonpulsatile pumping system,” Artif. Org a n s , v o l . 12, no. 2 , p p . 113-115, 1988. 1181 S . Takatani, K. Ozawa, L. Golding, G. Jacobs, T. Murakamia, F. Valdes, H. Harasaki, R. Kiraly, and Y. Nose, “Comparative evaluation of nonpulsatile and pulsatile cardiac prosthesis,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 26, pp. 438-443, 1980. [I91 A. Kovats, Design and Performance of Centrifugal and Axial Flow Pumps and Compressors. New York: MacMillan, 1964, pp. 91275. [20] N. Cheremisinoff, Fluid Flow: Pumps, Pipes, and Channels. Ann Arbor, MI: Ann Arbor Science, 1981, pp. 265-337. [21] F. D. Dorman, T. E. Murphy, and P. L. Blackshear, “An application of the tesla viscous flow turbine to pumping blood,” in Mechanical Devices to Assist the Failing Heart. Nat. Res. Council, Nat. Acad. Sci., 1966, pp. 119-128. [22] S . H. Hassinger and L. G. Kehrt, “Investigation of a shear force pump,” Trans. J . Eng. Power, vol. 85, pp. 201-207, 1963. [23] W. Rice, “An analytical and experimental investigation of multiple disk pumps and compressors,” ASME Trans., Series A . , vol. 85, pp. 191-198, 1963. [24] G . G. Johnson, F . S . Hammill, K. H. Johnasen, U. Marzec, D. Gerard, R. B. Dilley, and E. F. Bernstein, “Prolonged pulsatile and nonpulsatile LV bypass with a centrifugal pump,” Trans. Amer. Soc. Artif. Inter. Organs, vol. 22, p. 323, 1976. [25] D. B. Olsen and G. Bramm, “Blood pump with a magnetically suspended impeller,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 31, pp. 395-401, 1985. [26] J. Hinglais, 1. L. Chevrier, G. Demoment, R. Le Guen, and F. Loyer, “Partial or total heart substitution with a double-centrifugal deviceTheoretical and physiologic studies,” in Assisted Circulation, F. Unger, Ed. Berlin: Springer, 1979, pp. 243-257. [27] P. L. Blackshear, “Mechanical hemolysis in flowing blood,” in Biomechanics: Its Foundations and Objectives, Y. C. Fung, N. Perrone, and M. Anliker, Eds. Englewood Cliffs, NJ: Prentice-Hall, 1972. [28] B. D. Etter, “Pulsatile control of a valveless artificial heart,” M.S. thesis, Texas A&M Univ., College Station, 1986. [29] G . Rosenberg, “A mock circulatory system for in vitro studies of artificial hearts,” M.S. thesis, Penn. State Univ., 1972.
Gerald E. Miller (M’88) received the B.S. degree in aerospace engineering in 1971, and the M.S. and Ph.D. degrees in bioengineering in 1975 and 1978, respectively, from the Pennsylvania State University, University Park. Since 1977, he has been a member of the Bioengineering faculty at Texas A&M University, College Station, and currently holds the rank of Professor. He is the director of the Human Systems Engineering Laboratory which coordinates multidisciolinarv biomedical research throughout the university. As of September 1989, he became the chairman of the Bioengineering Program at Texas A&M University. He worked with the Artificial Heart Laboratory at Penn State for many years and has current research interests including biomedical fluid mechanics, artificial organs, rehabilitation engineering, and man-machine interfacing. Dr. Miller is a member of IEEE EMBS, the Biomedical Engineering Society, and ASME. He is active in ASME as past chairman of the Bioengineering Fluid Mechanics Committee and as a current member of both the Design and Executive Committees within the Bioengineering Division. He is also a member of Phi Kappa Phi and Sigma Gamma Tau.
.
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 37. NO. 2. FEBRUARY 1990
A Multiple Disk Centrifugal Pump as a Blood Flow Device GERALD E. MILLER,
MEMBER, IEEE,
BRADLEY D. ETTER,
Abstract-A multiple disk, shear force, valveless centrifugal pump was studied to determine its suitability as a blood flow device. A pulsatile version of the Tesla viscous flow turbine was designed by modifying the original steady flow pump concept to produce physiological pressures and flows with the aid of controlling circuitry. Pressures and flows from this pump were compared to a Harvard Apparatus pulsatile piston pump. Both pumps were connected to an artificial circulatory system. Frequency and systolic duration were varied over a range of physiological conditions for both pumps. The results indicated that the Tesla pump, operating in a pulsatile mode, is capable of producing physiologic pressures and flows similar to the Harvard pump and other pulsatile blood pumps.
INTRODUCTION
H
EART assist and total heart replacement systems have been studied for decades. During this time span, there have been various types and modes of operation of these systems tested. These have included pneumaticallydriven sac pumps [ l ] , [2], roller pumps [3], centrifugal impeller pumps [4], [5], pusher plate diaphragm pumps [6]-[8], and a variety of others. Of such systems, the pneumatically-driven sac pump and the electrically-powered pusher plate pump have received the most attention as either a cardiac assist device or a total heart replacement. Clinical trials of such systems have proven that such devices operate successfully in humans over a limited period in both an assist mode and as a total cardiac replacement [9], [ 101. However, a variety of complications have arisen in the development of these devices which has served to delay the application of such systems as relatively risk free prosthetic devices for humans. These complications have included the formation of pannus [ 111, thrombus formation [ 121, operational difficulties [ 131, [ 141, calcification of blood contacting surfaces [15], limited cardiac output, as well as a host of others. Although much research has been conducted into the development of problem free versions of sac type and diaphragm type devices, research has also been conducted into the development of alternative modes of operation such as the roller pump, the impeller pump, and the disk pump. In particular, the rotating disk pump, once called Manuscript received November 28, 1988; revised June 15. 1989. The authors are with the Bioengineering Program, Texas A&M University, College Station, TX 77843. IEEE Log Number 8932198.
MEMBER,
IEEE, A N D
JEAN M. DORSI
the Tesla turbine, has not been widely studied as a cardiac pumping mechanism. In the past, such a device has been utilized as a valveless, steady flow pump which can produce extremely high flow rates and pressure heads and can be electrically powered. However, the effects of steady blood flow from a valveless pump have resulted in physiological problems which have served as a limiting factor in the further development of such devices [16]-[18]. The purpose of this study was to design and test a prototype version of a pulsatile version of a multiple disk centrifugal pump, the so-called Tesla Viscous Flow Turbine. This pump was then compared to a common blood flow device, such as a Harvard piston pump, to determine whether the Tesla pump can produce physiological pressures and flows similar to those of sac type or diaphragm type cardiac pumps. The development of a pulsed disk pump might thus prove to be a viable alternative as a heart assist or total heart replacement device. CENTRIFUGAL FLOWPUMPS Since centrifugal pumps have been widely used in the past, the theoretical analysis, design, and testing of such pumps has been widely published [19], [20]. The parameters which determine the maximum pressure head are the pump speed, impeller or disk diameter, and the number of disks in series (and their spacing). Similar parameters affect the output flow and maximum shearing stresses exerted on the fluid. The major advantages of such pumps include the ability to generate large flow rates and pressure heads, and the ability to easily power and control such systems electrically. Centrifugal pumps have been studied sporadically as blood flow devices for many years. In 1964, Dorman et al. studied the utilization of a multiple disk Tesla pump as a heart assist device [21]. They utilized the characteristics of shear action pumps as described by Hassinger and Kerht [22] and by Rice [23], and then modified the disk spacing, disk diameter, and rotation speed to produce quasi-physiological flows and pressure heads, albeit in steady flow. Further studies of centrifugal pumps as heart assist or substitute devices has centered around impeller type or vane type devices. The Biomedicus Bio-Pump (Biomedicus Inc., Minneapolis, Minnesota) propels fluid by means of a force-vortex pumping principle by rotation of smooth
0018-9294/90/0200-0157$01.OO 0 1990 IEEE
158
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 37, NO. 2. FEBRUARY 1990
DISK f ?ller
I
I
I
Magnet Rotor
Fig. 1. Medtronic impeller pump for use as a heart substitute device.
rotator cones. Blood flow can be controlled from 100 ml/min to 10 l/min. It has been used in either steady flow or pulsatile mode. Medtronic (Medtronic Inc., Minneapolis, MN) has developed an impeller pump which operates either in steady or pulsatile modes (see Fig. 1). Flows of up to 10 l/min have been exhibited in vitro. Descriptions of these and similar pumps have been reported by Bernstein [SI, Johnson [24], Olsen and Bramm [25], and Hinglais et al. [26]. Due to continuing operational and pathological difficulties, the impeller and vane type heart devices have seen limited clinical application. Blood damage has been a serious drawback to the further (b) development of these systems. The use of a multiple disk Fig. 2. Tesla pump design consisting of seven parallel polycarbonate disks centrifugal pump may offer the advantages of typical cendriven by a central shaft mounted at the end disk. The spiral volute serves as a housing and a diffusing chamber. The fluid enters through the side trifugal operation (large flow, easy control, electrical port opposite the central shaft and is collected via the housing as it is power), while limiting the typical drawbacks such as propelled radially from the disks. (a) Disk dimensions. (b) Pump and blood damage and quasi-steady flow operation. A disk housing perspective. pump utilizes shear forces rather than lifting surfaces to impart momentum to a fluid by means of a series of rotating disks. nected to each other by three stainless steel shafts. The Theoretical analyses and design studies of steady flow end disk is connected to a driving motor by means of a multiple disk pumps have been conducted by Dorman et larger steel shaft located outside the pump housing. This al. [21], Hassinger and Kerht [22], Rice [23], and many shaft was sealed at the interconnection with the housing others. However, little has been done with multiple disk by means of a graphite gland packing seal. The test fluid type pumps since the study of steady flow performance by (physiologic saline) enters into the rotor area through the Dorman et al. in the middle 1960’s. The use of such end opposite the driving shaft [Fig. 2(b)]. As the fluid pumps in a pulsatile mode or as an alternative artificial enters into the spaces between the disks, it is sent into a heart design has not been previously investigated. This spiral motion whose tangential and radical velocity compaper will describe a study which was conducted regard- ponents are generated by frictional forces. As the fluid ing the design and performance of a pulsatile version of spins off the peripheral surface of the disks, it is collected the multiple disk shear pump as a possible blood flow de- by a spiral volute chamber which serves as both a pump vice. housing and a diffusing chamber. The fluid is then ejected tangentially through the outlet port of the pump [Fig. METHODSAND MATERIALS 2(b)]. The equations governing the performance of mulThe Tesla pump which was constructed for this study tiple disk pumps have been presented by Hassinger and consists of seven polycarbonate disks configured in a par- Kerht [22] and by Rice [23]. allel arrangement as seen in Fig. 2(a) and (b). Plastic was The number of disks, disk spacing, and disk diameter utilized for the disk material to provide a low moment of for this application were calculated so as to minimize the inertia, thus allowing the pump to respond quickly to a shear rate exerted on the fluid while creating sufficient change in torque produced by a pulsatile driving signal pressures and flows required to mimic cardiac function. from a motor. The inner and outer disk diameters are 1.5 The geometry of the pump, in combination with a limiting and 3.0 in, respectively [Fig. 2(a)]. Each disk is 0.063 in disk rotation speed of 2250 rpm, would produce a theothick. The disk spacing is 0.016 in. The disks are con- retical limiting shear rate of 10 000 s-’ (400 dyn/s cm2
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159
CENTRIFUGAL P U M P AS BLOOD FLOW DEVICE
2 4
1
Duratm
b
LM74121
c3 100
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I
115'4 NEUTRAL
ECG5641
115V HOT
1PO
Fig. 3 . Controlling circuitry for the Tesla pump which produces pulsatile motion. Frequency, systolic duration, and contraction-like ejection thrust are independently controlled (see text for description).
for blood with a 4 centipoise viscosity), which is below the limiting value (400 dyn/s cm2) to avoid blood cell damage [27]. This pump was driven by a centrifugal pump motor (Oberdorfer Model 142-12). This motor normally operates as a steady flow A / C device, but was modified for this study to produce sinusoidal pulsations in order to produce pulsatile flow from the pump [28]. The controlling circuitry for this purpose is shown in Fig. 3 . A Sylvania triac is used to control the speed of the motor by acting as an electronic switch driven by a series of 5 V pulses. The motor speed varies according to the frequency and duration of the pulses. An LM555 timer, acting as an astable oscillator, drives the controlling circuitry and acts as a heart rate simulator. The rate can be varied by controlling the timer resistors. Systolic duration is controlled by an LM74121 one shot multivibrator. This signal is then sent to an LM741 operational amplifier configured as a modified lossy integrator. This integrator serves to smooth out the square wave control signal into a smoother sinusoidal waveform. This smooth signal is then sent to an LM555 pulse width modulator (PWM) which serves to gate the triac in accordance with the voltage supplied by the integrator. Typical control signals from each of these components are shown in Fig. 4. Design analyses and characteristic signals for each component of the controlling circuit were studied by Etter [28]. Constraints on the pump motor are that no flow reversal takes place (no pure reversal of the disks), no undue forces are exerted on the pump which may damage the motor,
-
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74121
555
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:
- 1-
,
1
2
~
3
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Time (sec) Fig. 4. Signals from the components of the pump controlling circuit. The LM555 timer controls the pump frequency. The LM74121 monostable multivibrator controls the systolic duration. The LM741 lossy integrator smooths the multivibrator output and controls the acceleration and deceleration of the pump. The LM555 pulse width modulator sends a signal to the triac which controls the pump speed. Typical pressure and flow waveforms are shown for reference.
and that physiological pressures and flows are generated. The design of this pump is valveless (thus, the constraint on disk reversal). Diastolic (zero flow) conditions are maintained with minimal forward disk motion. The pump is connected to an artificial circulatory system as seen in Fig. 5 . This system consists of a resistor, capacitor, and reservoir and is similar to that designed by Rosenberg [29]. The resistor consists of a series of parallel, flexible, small diameter tubes which can be con-
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IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 31. NO. 2. FEBRUARY 1990
CAPACITOR
FLOW PROBE
VENOUS R E S E R V O I R PRESSURE T R A N S D U C E R
Fig. 5 . Artificial circulatory system consisting of a resistance, compliance, and atrial reservoir as designed by Rosenberg [29].
stricted by means of a plate and adjusting nut. The capacitor is a leaf spring system where the spring force opposes the increase in capacitor pressure resulting from an inrush of fluid during systole. Pressure and flow rate are measured at the pump outlet by means of a Motorola MPXlOOA pressure transducer and a Carolina Medical EP670 electromagnetic flow probe. The pressure signal is amplified by means of an Analog Devices AD521 instrumentation amplifier, while the flow signal is processed by means of a Carolina Medical FM501 Flowmeter. One inch tygon tubing is used to connect the circulatory system components. The output of the Tesla pump was compared to that produced by a Harvard Apparatus pulsatile piston pump (model 1423). This pump contains ball valves at the inlet and outlet ports. The systolic duration and pump frequency can be varied independently by separate controls on the pump console. The systolic duration control of the Harvard pump affects the forward excursion of the piston, such that stroke volume of this pump can be indirectly manipulated within a limited range. Outlet pressure and flow were recorded using a Texas Instruments Portable Professional computer with a Metrabyte DASH-8 Analog to Digital converter. The data were sampled at 100 Hz in order to comply with Nyquist’s criteria. Sampling was done in sets of 1024 data points which resulted in 10.24 s of pressure and flow data per set. The resistance and compliance settings of the circulatory system were adjusted to produce physiological pressures and flow values while the Harvard pump was connected to the circulatory system. The settings of the circulatory system were to remain unchanged following the initial setting. The shear pump has no specific stroke volume, as there are no valves. However, the timer, pulse width modulator and lossy integrator can be adjusted to produce an effective stroke volume which approximates that of the Harvard pump. The LM555 timer adjustment affects the pump frequency; the LM74 12 1 monostable multivibrator adjustment affects the width of the systolic duration; and the LM741 lossy integrator adjustment affects the rise and decay rate of the flow and pressure waveforms, which affect the acceleration and deceleration of the disk pump. At
present, these adjustments are made qualitatively. However, present research is ongoing to develop a quantifiable relationship between control circuit component settings and pump output. Once the settings on the controlling circuitry were completed to produce this “stroke volume,” these were also to remain unchanged throughout a given experiment. Data was recorded for each pump at heart rates of 60, 70, 80 and 90 beat/min and at systolic durations of 0.3, 0.35, and 0.4 s for each heart rate. This provided a total of 12 operating conditions for each pump. Frequency and systolic duration of the disk pump were adjusted before each experiment to approximate that of the Harvard pump. Conductance of the saline was checked periodically as was fluid temperature and the output of the pressure and flow transducers in order to verify that the calibration of these devices was still valid.
RESULTSAND DISCUSSION Typical pressure and flow waveforms for both the Tesla pump and the Harvard pump are shown in Fig. 6(a)-(d). A list of peak and mean pressures and flows for both pumps at all operating conditions is given in Tables I and 11. Inspection of the time varying waveforms seen in Fig. 6(a)-(d) show good agreement between the two pumps. The pressure values for the Harvard pump were slightly lower than typical physiologic norms, although the waveform shapes are typical of physiologic aortic pressure waveforms. The lower pressure values are likely a result of the settings of the circulatory system since both pumps produced similar results. Mean pressures and flows exhibited from the disk pump were slightly higher than that of the Harvard pump. This is due to inaccurate settings of the controlling circuitry of the disk pump, which were manually adjusted to approximate the systolic duration of the Harvard pump. Current research efforts include the development of a computer controlled automatic selection of parameter settings of the control circuit of the disk pump. Once completed, such a system can significantly reduce the problems associated with inaccurate matches
MILLER
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ol.: CENTRIFUGAL PUMP AS BLOOD FLOW DEVICE
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Pump Pressures
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Fig. 6 . Time varying output pressure and flow rate waveforms from both the Harvard pump and the Tesla pump. HR signifies heart rate in beats per minute. SD signifies systolic duration in seconds. The oscillations in the Harvard pump data are due to bouncing of the ball valve. The disk pump is not completely synchronized with the Harvard pump since the disk pump frequency and systolic duration are only approximately adjusted to that of the Harvard pump. In addition, the timing controls of the Harvard pump are not exact.
between the disk pump and a comparable piston or sac pump. However, given the limitations of the design and testing of this early prototype of the disk pump, it appears that such a design is quite capable of producing physiological pressures and flows typical of present day artificial hearts such as the sac, diaphragm, and piston pumps. Further testing is necessary in order to analyze flow conditions inside the pump and housing. Stagnation, turbulence, and cavitation are all conditions which should be monitored in such a design in order to eliminate any flow condition which would be conducive to blood damage or thrombus formation. Future designs of a disk type pump
will also address power consumption, biocompatibility , and anatomical/physiological considerations. However, such a pump, operated in a pulsatile fashion does offer advantages over other designs which are presently used. A valveless centrifugal pump does not require a compliance chamber. A rotary action pump should provide immediate response times to changing operating conditions from a drive system or controlling circuit. A valveless pump would not suffer from calcification, pannus formation nor thrombus formation at a valve seat. The use of a disk pump may also avoid the problem of fill limited conditions of a sac pump. Such fill limited operation in a valved sac type or pusher plate pump is caused
162
IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 37, NO. 2. FEBRUARY I Y Y O
TABLE I HARVARD P U M PDATA
S
Peak Diast Flow l/min
Peak Syst. Flow I/min
Peak Diast . Pres. mmHg
Peak Syst. Pres. mmHg
Mean Flow I/min
Mean Pres. mmHg
0.30 0.35 0.40 0.30 0.35 0.40 0.30 0.35 0.40 0.30 0.35 0.40
0.50 0.54 0.54 0.63 0.51 0.44 0.58 0.57 0.77 0.70 0.67 1.02
13.17 13.22 12.98 14.30 14.21 14.42 15.06 15.14 15.65 15.88 15.85 16.76
60.66 59.96 59.94 61.35 62.36 63.49 63.46 63.69 67.34 65.17 65.27 68.82
96.26 95.81 95.29 97.95 94.84 94.77 95.69 95.84 99.00 97.10 97.45 101.27
4.75 4.80 5.05 5.23 5.21 5.75 5.66 5.79 6.68 6.23 6.23 7.23
74.47 73.94 74.52 75.91 75.49 76.82 76.99 77.19 81.60 78.91 79.06 83.83
Heart Rate bPm
System Duration
60 60 60 70 70 70 80 80 80 90 90 90
TABLE I1
TESLAPUMP DATA
S
Peak Diast Flow l/min
Peak Syst. Flow I/min
Peak Diast. Pres. mmHg
Peak Syst. Pres. mmHg
Mean Flow I/min
Mean Pres. mmHg
0.30 0.35 0.40 0.30 0.35 0.40 0.30 0.35 0.40 0.30 0.35 0.40
0.47 1.45 1.29 1.76 1.32 0.79 0.14 0.85 0.58 0.00 0.14 0.09
15.78 15.91 15.78 16.22 17.67 17.97 17.48 17.92 18.12 18.09 18.91 18.85
66.51 67.53 68.82 70.73 69.79 71.28 70.14 7 1.92 71.95 70.71 73.01 73.46
104.23 1 11.47 114.52 10.82 115.71 115.44 109.51 113.01 113.08 105.71 11 1.22 112.11
5.68 6.35 6.64 6.35 6.97 7.23 6.66 7.31 7.63 6.98 7.83 7.95
79.76 83.98 86.78 84.16 87.16 88.87 85.76 88.75 89.81 85.47 89.57 90.53
Heart Rate bPm
System Duration
60 60 60 70 70 70 80 80 80 90 90 90
by either a low filling pressure or an insufficient filling period (during diastole). A valveless centrifugal pump may avoid both of these problems for the following reasons: (a) there is no inlet valve to restrict inlet flow, and (b) a centrifugal pump draws fluid into it’s inlet by a vacuum action. In fact, too great a vacuum on the inlet could result in venous collapse when such a pump is utilized in vivo. This important area of concern is presently under investigation. Simultaneous measurements of inlet pressure and flow and outlet pressure and flow for the disk pump are planned to determine the effects of inlet conditions on disk pump performance. Such results will then be compared to similar measurements in a comparable piston or sac pump to ascertain the relative effects of fill limited operation on pumping performance. The results from this study suggest that further research is warranted into the development of alternative designs of pulsatile pumps such as the disk pump for use as an artificial heart or heart assist device.
REFERENCES [ I ] L. D. Joyce, W. C . DeVries, W. L. Hastings, D. B. Olsen, R. K. Jarvik, and W. J. Kolff, “Response of the human body to the first
permanent implant of the Jarvik-7 total artificial heart,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 29; p. 81, 1983. [21 R. K. Jarvik, “The total artificial heart,” Sci. Amer., vol. 244, pp. 74-80, 1981. [3] G. P. Noon, J. E. Harrell, L. Feldman, J . Peterson, P. Kent, and M. E. DeBakey, “Development and evaluation of pulsatile roller pump and tubing for cardiac assistance,” Artif. Organs, vol. 7, pp. 49-54, 1983. [4] M. F. Lynch, D. Peterson, and V. Baker, “Centrifugal blood pumping for open heart surgery,” Minn. Med,, vol. 9 , p. 536, 1978. [ 5 ] E. F. Bernstein, “A centrifugal pump for circulatory assistance,” in Assisted Circulation, F. Unger, Ed. Berlin: Springer, 1979, pp. 231242. [6] G. Rosenberg, A. Snyder, W. Weiss, D. L. Landis, D. B. Geselowitz, and W. S . Pierce, “A cam-type electric motor-driven left ventricular assist device,” J . Biomech. Eng., vol. 104, pp. 214-220, 1982. [7] J. Moise, K. Butler, J . Payne, R. Wampler, W . Smith, L. Fujimoto, L. Golding, R. Kiraly, H. Harasaki, and Y. Nose, “Experimental evaluation of complete electrically powered ventricular assist system,” Trans. Amer. Soc. Arrif. Intern. Organs, vol. 31, pp. 202205, 1985. [8] C. W. Sherman, D. Gernes, W. Clay, V. Poirier, and W. F. Bernhard, “A permanent implantable electric blood pump,” IEEE Eng. Med. Biol., pp. 26-29, 1986. [9] W . C. DeVries, J. L. Anderson, L. D. Joyce, F . L. Anderson, E. H. Hammond, R. K. Jarvik, and W. J. Kolff, ‘‘Clinical use of the total artificial heart,” New Eng. J . M e d . , vol. 310, p. 273, 1984. [IO] E. K. Olsen, W. S . Pierce, J. H. Donachy, D. L. Landis, G . Rosenberg, W. M. Phillips, G. A. Allen, and M. J. O’Neill, “A two and
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MILLER et al.: CENTRIFUGAL PUMP AS BLOOD FLOW DEVICE
one-half year clinical experience with a mechanical left ventricular assist pump in the treatment of profound post-operative heart failure,” Int. J . Artif. Organs, vol. 2, p. 197, 1979. [ I l l R. K. Jarvik, T. R. Kessler, L. D. McGill, D. B. Olsen, W. C. DeVries, J. T. Blaylock, and W. J. Kolff, “Determination of pannus formation in long-surviving artificial heart calves and its prevention,” Trans. Amer. Soc. Artif. Intern. Organs, pp. 90-95, 1981. [I21 T. R. Kessler, A. B. Pons, R. K. Jarvik, J. H. Lawson, K. J . Razzeca, and W. J. Kolff, “Elimination of predilection sites for thrombus formation in the total artificial heart-before and after,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 24, p. 532, 1978. 1131 S . Lee, G. Rosenberg, J. H. Donachy, C . B. Wisman, and W. S . Pierce, “The compliance problem: A major obstacle in the development of implantable blood pumps,’’ Art$ Organs, vol. 8, pp. 8290, 1984. [ 141 E. Hennig, A. Mohnhaupt, and E. S . Bucheri, “The influence of the filling pressure on the output of pneumatically driven blood pumps,” Proc. Europ. Soc. Artif. Organs, vol. 3, p. 19, 1976. [15] D. L. Coleman, D. Lim, T. R. Kessler, and J. D. Andrade, “Calcification of non-textured implantable blood pumps,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 27, p. 97, 1981. [16] G . Mavroudis, “To pulse or not to pulse,” Ann. Thorac. Surg., vol. 25, no. 3 , pp. 258-271, 1978. [I71 Y. Nose, “The need for a nonpulsatile pumping system,” Artif. Org a n s , v o l . 12, no. 2 , p p . 113-115, 1988. 1181 S . Takatani, K. Ozawa, L. Golding, G. Jacobs, T. Murakamia, F. Valdes, H. Harasaki, R. Kiraly, and Y. Nose, “Comparative evaluation of nonpulsatile and pulsatile cardiac prosthesis,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 26, pp. 438-443, 1980. [I91 A. Kovats, Design and Performance of Centrifugal and Axial Flow Pumps and Compressors. New York: MacMillan, 1964, pp. 91275. [20] N. Cheremisinoff, Fluid Flow: Pumps, Pipes, and Channels. Ann Arbor, MI: Ann Arbor Science, 1981, pp. 265-337. [21] F. D. Dorman, T. E. Murphy, and P. L. Blackshear, “An application of the tesla viscous flow turbine to pumping blood,” in Mechanical Devices to Assist the Failing Heart. Nat. Res. Council, Nat. Acad. Sci., 1966, pp. 119-128. [22] S . H. Hassinger and L. G. Kehrt, “Investigation of a shear force pump,” Trans. J . Eng. Power, vol. 85, pp. 201-207, 1963. [23] W. Rice, “An analytical and experimental investigation of multiple disk pumps and compressors,” ASME Trans., Series A . , vol. 85, pp. 191-198, 1963. [24] G . G. Johnson, F . S . Hammill, K. H. Johnasen, U. Marzec, D. Gerard, R. B. Dilley, and E. F. Bernstein, “Prolonged pulsatile and nonpulsatile LV bypass with a centrifugal pump,” Trans. Amer. Soc. Artif. Inter. Organs, vol. 22, p. 323, 1976. [25] D. B. Olsen and G. Bramm, “Blood pump with a magnetically suspended impeller,” Trans. Amer. Soc. Artif. Intern. Organs, vol. 31, pp. 395-401, 1985. [26] J. Hinglais, 1. L. Chevrier, G. Demoment, R. Le Guen, and F. Loyer, “Partial or total heart substitution with a double-centrifugal deviceTheoretical and physiologic studies,” in Assisted Circulation, F. Unger, Ed. Berlin: Springer, 1979, pp. 243-257. [27] P. L. Blackshear, “Mechanical hemolysis in flowing blood,” in Biomechanics: Its Foundations and Objectives, Y. C. Fung, N. Perrone, and M. Anliker, Eds. Englewood Cliffs, NJ: Prentice-Hall, 1972. [28] B. D. Etter, “Pulsatile control of a valveless artificial heart,” M.S. thesis, Texas A&M Univ., College Station, 1986. [29] G . Rosenberg, “A mock circulatory system for in vitro studies of artificial hearts,” M.S. thesis, Penn. State Univ., 1972.
Gerald E. Miller (M’88) received the B.S. degree in aerospace engineering in 1971, and the M.S. and Ph.D. degrees in bioengineering in 1975 and 1978, respectively, from the Pennsylvania State University, University Park. Since 1977, he has been a member of the Bioengineering faculty at Texas A&M University, College Station, and currently holds the rank of Professor. He is the director of the Human Systems Engineering Laboratory which coordinates multidisciolinarv biomedical research throughout the university. As of September 1989, he became the chairman of the Bioengineering Program at Texas A&M University. He worked with the Artificial Heart Laboratory at Penn State for many years and has current research interests including biomedical fluid mechanics, artificial organs, rehabilitation engineering, and man-machine interfacing. Dr. Miller is a member of IEEE EMBS, the Biomedical Engineering Society, and ASME. He is active in ASME as past chairman of the Bioengineering Fluid Mechanics Committee and as a current member of both the Design and Executive Committees within the Bioengineering Division. He is also a member of Phi Kappa Phi and Sigma Gamma Tau.
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