Pierce-Donachy pediatric VAD: Progress in development

PUMPS IN PROGRESS

Pierce-Donachy Pediatric VAD: Progress in Development Bill B. Daily, MD, PhD, Timothy W. Pettitt, MD, Salvatore P. Sutera, PhD, and William S. Pierce, MD Divisions of Cardiothoracic Surgery and Mechanical Engineering, Washington University, St. Louis, Missouri, and Division of Cardiothoracic Surgery, Pennsylvania State University, Hershey, Pennsylvania

Background. Despite the successful use of ventricular assist devices in adults over the past 15 years, relatively little has been done to develop ~ de, rices for pediatric patients. Conseque~ly, no such device is carrently available. A ~ of ~ data ~ that the majority of patients in need of a ped~ahlc veNxicular assist device, either for posW.ard/otomy ~ e n / c shock or as a bridge to cardiac t r a n s ~ n , are m.mm~es weighing 3 to 5 kg. Attempts to "scale down" an adult blood p u m p to make an appropriate device for these patients have been diJficult due te the lack of suf6cientiy small, commercially available valves and the tendency for thrombus to develop in these small pumps. Methods. We report on p r o s t ~ s in the development of

the Pierce-Donachy pediatric ventricular assist device, which incozporates 10-ram-diameter bfleaflet valve prototypes. Particle image velocimetry is used to quantify the velocity 6eld inside the pump. Rz~Its and Conclusions. Particle image velocimetry velocity maps demonstrate the complexity of the flow parterres in these p u m p s and suggest that improved flow patteams may result from the use of valves with improved h c ~ performance. Animal tests to determine whether ~ e d flow patterns and better "washing" of the p u m p ' s blood-contacting surfaces will reduce thrombus formation are underway.

he need for a pediatric ventricular assist device T (WAD) is increasing due to two general trends in pediatric heart surgery. First, corrective operation is now

ticoagulation. This exacerbates bleeding, which requires reoperation in approximately 50% of pediatric patients [2-4]. This has been statistically associated with increased mortality in larger adult series. More importantly, extracorporeal systems are intended only for short-term use. These patients are intubated, monitored with multiple transcutaneous catheters, and confined to bed. In this setting, the duration of support is limited by the increasing likelihood of a serious complication developing after a few days. Some potentially salvageable patients may not recover in time, ie, before a complication develops related to this means of support. At St. Louis Children's Hospital over a 40-month period from January 1989 to May 1992, 570 children less than 6 years of age underwent a corrective operation for congenital heart disease. Severe PCCS developed in 56 patients (10%) in the immediate postoperative period. Twenty-nine patients were supported with extracorporeal membrane oxygenation; 14 of these were successfully weaned and 15 died after decannulation. An additional 27 patients died of PCCS without extracorporeal membrane oxygenation. Extracorporeal membrane oxygenation was not used in these patients because of the concern for bleeding or because in the surgeon's judgment, there was no realistic hope of myocardial recovery in less than 10 days. During the preceding 18 months, there were also 3 cardiac transplant candidates who died of complications of cardiac failure while awaiting a donor heart. We expect this number to grow as pediatric cardiac transplantation grows. Based on a review of these data, we estimate that 10 patients per year would be candi-

feasible for many previously ~ forms of complex congenital heart disease, bgt often 7eq~res operating on younger, smaller, and sick_er patients than ever before. Postcardiotomy cardtogenic shock (PCCS) remains a significant cause of mortality in this group. Second, the success of pediatric heart h,amplantation has led to increases both in the n a m b e r of ~ candidates and in the waiting lime for scarce donor organs. Recently, the use of circulatory support for pediatric patients with PCCS has increased substantially [1]. Surgeons are using an extracorpoveal civcIlt fitted with a roller or centrifugal pump, with or without an oxygenator, to provide support [2, 3]. Although only a limited number of patients have been studied, the overall salvage rate is 25% to 50%, comparable with that in adtutts. F u r t ~ r m o f e , sarvivors showed significant improvement in left v e n ~ c t t l K furtction during and after left ventricular as6ist device (VAD) support. These results indicate that c i ~ support in children can be at least as eJfective as i~n scttalts. However, extracorporeal VADs have some significant disadvantages and do not meet the needs of every patient. Current extracorporeal systems require continuous anPresented at The Third International Conferenceon CirculatorySupport Devicesfor Severe Cardiac Failure, Pittsburgh, PA, Oct 28-30, 1994. Address reprint requests to Dr Daily,Divisionof Card~othoracicSurgery, Jewish Hospital,216 S Kingshighway,St. Louis, MO 63110. © 1996 by The Society of Thoracic Surgeons Published by Elsevier Science Inc

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Fig 1. The Pierce-Donachy pediatric ventricular assist device is a geometrically scaled-down version of the adult-sized ventricular assist device except for the valves.

dates for PVAD support at St. Louis Children's Hospital: 3 to 4 PCCS patients require a means of support with less anticoagulation than extracorporeal membrane oxygenation mandates, 3 to 4 PCCS patients require a means of support that offers the possibility of intermediate-term recovery or bridge to transplantation, and 2 to 3 cardiac transplant candidates require a bridging device. Further analysis of these data shows that the majority of these patients were less than 3 months of age and almost all were less than 2 years (mean age, 71 +_ 161 days; range, 1 to 758 days). Excluding I premature child who weighed 1.2 kg, the average weight of these children was 3.0 -+ 1.7 kg, with a range of 1.8 to 10.1 kg. This establishes the need for a PVAD that can deliver a range of flows from 300 to 1,200 mL/min.

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sac held a total volume of 17 mL; under standard operating conditions the pump had a stroke volume of 11 mL. Initial in vitro mock circulation studies were carried out to characterize the basic hydrodynamic performance of the PVAD. These tests demonstrated the ability of this pump to generate pulsatile flow under physiologic conditions. The system was easy to operate and control. Qualitative flow visualization studies using a threedimensional light source and aluminum particles were conducted to determine the overall flow pattern within the PVAD and to identify regions of stagnation or recirculation. During diastole the overall flow pattern within the p u m p was circular, around the pump's short axis. With the onset of systole, the inlet valve closed, the outlet valve opened, and particles were accelerated toward the outlet. Using this technique, minimal valve regurgitation was seen and no regions of persistent stagnation or recirculation were detected. However, this technique has limited spatial and temporal resolution and is unable to resolve potentially important flow details. Animal studies were undertaken to evaluate the in vivo performance of the PVAD. Over a period of years, 28 devices were implanted in baby goats and calves treated with various combinations of anticoagulants including heparin, warfarin, and aspirin. These experiments demonstrated the repeated propensity of thrombus to form on the surface of the blood sac (Fig 2A). This material was often seen in the flexing regions of the sac, but could be

The P e n n State P V A D P r o t o t y p e In 1986, a collaborative effort between the Division of Artificial Organs at Penn State and Sarns/3M, Inc (Ann Arbor, MI) was established to develop a PVAD. This device was a scaled-down version of the Pierce-Donachy adult VAD (Fig 1). It consisted of a pneumatically powered sac-type blood pump, a pneumatic power unit, and a control unit. The seamless U-shaped blood sac was made of segmented polyurethane (Hemothane). The sac was encased in a rigid housing of polycarbonate that contained a diaphragm made of polyurethane. A port on one side of the diaphragm admitted pulses of air, a port on the other side permitted deairing, and a Hall-effect switch detected when the pump was full. Specially designed inlet and outlet ball valves were precisely seated in the case and attached to custom-made cannulas. The valves and cannulas were mated; both had an inner diameter of 6 mm. The pump was activated by cyclically delivering and withdrawing air to and from the space between the diaphragm and case. This generated positive and negative pressures on one side of the blood sac, causing alternate emptying and filling of the blood sac. Unidirectional flow was ensured by the valves. The blood

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Fig 2. (A) Photograph of a pediatric blood sac explanted after 21 days in a calf. This shows a thin layer of thrombus commonly seen in explanted pumps. Although in this photograph thrombus is primarily seen on the inlet side of the blood sac, it can also be found at other locations. (B) Scanning electron micrograph of the blood sac surface showing thrombus rich in platelet and proteinaceous material.

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located anywhere on the sac's surface. Scanning electron microscopy demonstrated that this material was rich in fibrin and platelets (Fig 2B). Postmortem hlatologic analysis of the kidneys from these animals frequently demonstrated embolization, presumably from the pump. In contrast, thrombus formation and thromboembofism were infrequently seen in the adult pump. The finding that the adult-sized VAD is markedly less thrombogenic than its geometrically similar, scaleddown counterpart suggested that fluid dynamic factors played a key role in thrombus formation in the FVAD. In general, the thrombogenicity of assist devices can be attributed to (1) the coagulability of the blood, (2) the properties of the blood-contacting surface, and (3) fluid dynamic factors. Because in our experiments the co~gulability of the blood was the same (in the sense that both adult and pediatric systems were implanted in the same animal model) and the blood contacting surfaces were the same (both surfaces were made of Hemothane using similar fabrication methods), we concluded the difference in thrombogenicity between the adult and pediatric VADs must be due to differences in the fluid dynamics in these two devices. The fluid dynamic factors that contribute to thrombus formation likely come into play either as they cause damage to blood elements potentiating coagulation or as they govern the development of viscous wall boundary layers and therein the convective transport of blood elements and plasma solutes to and from the surface. For example, high shear rates in the vicinity of the valves may activate platelets as they enter the blood sac. These activated platelets are then transported to the sac's surface, and if the shear forces here are too low to detach microscopic platelet aggregates, ie, to "wash" the sac's surface, visible thrombus may form. Subsequent embolization of this material may cause organ damage. Comparison of the adult and pediatric VADs revealed two obvious differences that were likely to cause significant differences in the fluid dynamics in the two pumps. First, of clinical necessity, the PVAD was smaller and operated at lower average flow rates than the adult VAD. The fluid dynamic consequences of scaling down the pump were analyzed by comparing values of common dimensionless parameters, such as the Reynolds number, that help to characterize various aspects of the flow in the two systems. We focused on the effects of the inlet jet during pump filling because flow visvaliT~tion studies suggested that this jet initiates the diastolic vortex, which is thought to be important in washing the surfaces of the blood sac. The conclusions of this analysis were that the flow regimens in the adult amd pediatric devices were not, and could not be, dynamically similar. The Reynolds number of the flow in the PVAD will always be lower;, hence, viscous effects will have greater influence in the smaller device. Furthermore, the boundary layers in the PVAD generally will be thicker, relative to the flow passage width, than in the adult VAD. The consequences of greater viscous effects and thicker boundary layers may be that platelets and soluble coagulation factors that find their way to the blood sac's surface tend to remain

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Fig 3. (A) Photograph of the 6-mm ball valve (right) and the lO-mm bileafiet tilting-disc valve. (B) Axial view of both valves showing the greater orifice area of the bileafiet valve.

there longer and experience less shear force that washes the surface. The second difference between the adult and pediatric VADs, likely to cause significant differences in fluid dynamics, was that the adult VAD used relatively largediameter tilting-disc valves (24 mm inlet orifice, 21 mm outlet orifice), whereas the PVAD used relatively smalldiameter (6 ram) inlet and outlet ball valves (Fig 3). The hydrodynamic performance of the ball valves was assessed in detailed mock circulation studies. Pressure and flow measurements over a range of physiologic conditions demonstrated excessive pressure gradients and energy losses across these valves (Fig 4). Excessive energy losses may potentiate blood element damage in the vicinity of the valves and may also reduce the momentum of the inlet jet, which is needed to drive the diastolic vortex that helps to wash the surfaces of the pump.

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The likelihood that the suboptimal hydrodynamic performance of the ball valves contributed to the development of thrombus in the PVAD led us to develop a 10-mm-diameter bileaflet valve conduit (see Fig 3). The conduit was designed to minimize junctions between components where thrombus may form. In our experience, even small discontinuities in the flow stream, on the order of 100/xm, can cause thrombus formation. A bileaflet design with an 85-degree opex~ing angle was chosen because for a valve with this small of a diameter central flow via an open orifice was expected to give optimal hydrodynamic performance. In fact, subsequent in vitro hydrodynamic studies using the PVAD demonstrated minimal pressure gradients and energy losses across the bileaflet compared with the ball valves. The time required to fill or empty the PVAD using the same

operating conditions (eg, drive pressure, systole duration) was about 1.6-fold less with the bileaflet valves. These data supported the hypothesis that the fluid dynamics in the PVAD system could be significantly altered by changing the valves. Whether, and how, these changes might result in reduced thrombogenicity was the next question. Measurements of transvalvular pressure gradients and energy losses reflect the integrated effects of all of the infinitesimal fluid elements moving through the pump. The velocities of these elements vary continuously with location in the p u m p and time during the pumping cycle. A record of the velocities of these elements constitutes the fiowfield. The finding that the transvalvular pressure gradients and energy losses (per milliliter of blood pumped) were greater in the PVAD than in the adult VAD indicated differences in the flowfield in the two devices. To understand how differences in the flowfield might cause differences in thrombogenesis, the impact of the forces that generate the flowfield on blood elements that generate thrombus needs to be investigated. Ideally, the flowfield would be resolved on a scale that would explain the motion of individual blood elements. Thus, a characteristic spatial scale that enters this problem naturaUy is given by the diameter of the platelet (~3 /~m) and a time scale by a fraction of the period of p u m p filling or emptying ('~200 ms). Several experimental methods in fluid dynamics are available to study the flowfield inside VADs. Conventional flow visualization methods involve direct observation of the movement of reflective particles as they travel with the flow. These techniques are easy to use and have been useful qualitatively for identifying areas of stagnation or recirculation in VADs [5-7], but extracting quantitative information such as velocities or Reynolds stresses has been difficult. Recently, quantitative techniques capable of providing information about two of the three spatial components of the flowfield's velocity vectors have been used to study the flow in VADs [8-11]. Two of these techniques, laser Doppler anemometry and particle image velocimetry, have been used to study the flowfield inside the PVAD. These techniques are complimentary in that laser Doppler anemometry provides a high degree of temporal resolution whereas particle image velocimetry provides a high degree of spatial resolution. Results from these initial studies are emerging. Koehler and associates [12] at the University of Utah have used laser Doppler anemometry to estimate the velocity and Reynolds shear stress profiles in an infant VAD, similar to the PVAD, fitted with 6-mm ball valves. During ventricular filling they found axial velocities downstream of the inlet valve up to 1.1 + 0.3 m/s and high turbulent shear stresses at the inlet jet boundaries with a maximum of 65 -+ 190 Nm - 2 lasting about 0.1 second. They concluded that these Reynolds shear stresses are present for a sufficient duration to potentially cause blood damage. We have used particle image velocimetry to provide two-dimensional velocity vector maps of the flowfield

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Fig 5. Comparison of the pediatric ventricular assist device fiow~eld generated by bileaflet valves (A) versus ball valves (B) during diastole. Particle image velocimetry was used to generate these c~dor-enhanced instantaneous two-dimensional velocity vector maps of the pediatric ventricular assist device flow~ld. The tap cross-section was taken perpendicular to the pump's short axis, whereas the bottom cross-section was taken parallel to the short axis, in sagittal section. For both hypes of valves, the inlet jet enters the pump via ports that are angled relative to the image plane shown in the tap panel This jet initiates a diastolic vortex, which has an axis parallel to the pump's short axis. Comparison of (A) and (B) demonstrates that the inlet jet with the b'ileaflet valves has higher velocities, contains greater volume, and penetrates into the pump further than that with the ball valves. No~e that the inlet jet from the bileaflet valves has two high-velocity zones, reflecting the structure of the orifice of the open bileaflet valve (see Fig 3B). The inlet jet drives the diastolic vortex that washes the surface of the blood sac, helping to reduce throm~as formation.

inside the PVAD fitted with ball valves and bileaflet valves. In general, for both types of valves, these stadies demonstrated that the flow in the PVAD was highly complex and three-dimensional. During p u m p filling, the inlet jet initiated a diastolic vortex, which r o t a t ~ around the p u m p ' s short axis in the plane of largest crosssectional area (Fig 5). Recirculation consistently developed downstream of the inlet valve, along the curved p u m p wall. Regions of slow flow were identified over variable portions of the pumping cycle adjacen~ to the surface of the p u m p in areas remote from the valves and their associated jets. The onset of p u m p emptying is reflected by an abrupt change in the flowfield as the velocity vectors align toward the outlet in response to pressure from the p u m p ' s diaphragm (Fig 6). This convergence of fluid upstream of the outlet valve is generally more uniform than the divergence of fluid downatream of the inlet valve during p u m p filling, although small areas of recirculation can be identified at the boundaries of the outlet jet. The flowfield changed continuously with time, but at a given time during the pumping cycle the flowfields from different cycles ~ e r e remadutbly similar, except near the valves in the case of the ball valves. Here, bouncing of the bali in the valve housing caused nonperiodic flow variations. Other differences in the flowfield due to the use of

different valves were as follows: (1) The inlet jet with the bileaflet valves was more uniform (velocity vectors more aligned), contained greater volume, and penetrated into the p u m p further than that with the ball valves. (2) The diastolic vortex was formed sooner after the inlet valve opened and was better organized with the bileaflet valves. Perhaps associated with this observation, the thickness of the areas of slow flow adjacent to the boundaries of the p u m p was less with the bileaflet valves. (3) In early systole the volume of fluid entrained by the outlet jet appeared to be greater with the bileaflet valves. Although a precise estimate of the volume is difficult to obtain from two-dimensional data, it is apparent that velocity vectors further from the outlet are aligned toward the outlet when bileaflet valves are used. These experimental results are consistent with the hypothesis that the greater effective orifice area of the bileaflet valves preserves the m o m e n t u m of the inlet jet and fosters the development of the outlet jet. In turn, this strengthens the diastolic vortex and enhances systolic entrainment, which help to wash the surfaces of the blood sac. The next question was whether the PVAD fitted with bileaflet valves was, in fact, less thrombogenic in animals than the PVAD fitted with ball valves. Preliminary animal studies in weanling lambs are underway using the PVAD with bileaflet valves. He-

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A B Fig 6. Comparison of the pediatric ventricular assist device )qow)qeld generated by bilea)qet valves (A) versus ball valves (B) during systole. For both types of valves, )quid elements inside the pump converge on the outlet port in response to forced displacement of the diaphragm. In the case of the bilea)qet valves, decreased resistance to flow allows a greater volume of)quid in the pediatric ventricular assist device to be entrained by the outlet jet. This entrainment increases the velocities of j~uid elements near the wall and reduces the thickness of the layer of slowly moving fluid adjacent to the wall seen with ball valves. In this velocity range, the higher near-wall velocities may promote surface washing and help to reduce thrombus formation. mothane, a Biomer derivative with acceptably low thrombogenicity in the adult Pierce-Donachy device, is no longer available so a similar but nonidentical polyurethane (Biospan) has been used for blood sac fabrication. In addition, because the bileaflet valve conduits have a larger diameter than the ball valves, the cannuias were changed from a uniform 6 mm inner diameter to a custom-designed 10 mm to 7 mm inner diameter taper. Whether these changes alter the thrombogenicity PVAD system, independent of the valves, is unknown. Two PVADs have been implanted sequentially in a single lamb using left atrial to aortic cannulation and a paracorporeal pump location. The devices were run in the full-to-empty mode at a rate of approximately 90 beats/min. The animal was initially anticoagulated with aspirin (15 mg/kg) and heparin, maintaining the activated clotting time at approximately twice its baseline value. Warfarin administration was staffed the night after operation, and when the prothrombin time reached twice baseline the heparin administration was discontinued. The first pump ran for approximately 21/2days when we noted extravasation of blood between the sac and case. Because the pump was outside the body it was a relatively simple task to clamp the cannulas briefly and insert a new pump. This device ran 11 days before the same extravasation was identified. The animal was sacrificed and a complete autopsy performed. Careful inspection of the blood sacs from the two pumps demonstrated a small tear in each sac. The tears

occurred at the same location, in the flexing region near the periphery of the sac adjacent to the case. At the site of the tears, wear marks were identified on the external surface of the sac, suggesting friction between the case and sac. After reviewing the pump assembly procedure, we concluded that we had not injected an adequate volume of lubricant between the sac and case. Previous experience during development of the adult-sized device demonstrated the necessity of this lubricant to prevent such failures from blood sac wear. In addition to identifying the cause of blood sac failure, inspection of both blood sacs also demonstrated no microscopic thrombus on the surface of either sac. Thrombus was identified in the valve conduits of the first device along a junction between polysulfone components that developed during valve fabrication. No thrombus was identified in a second set of valve conduits that were used in the second device. The cannuias were also free of thrombus, and examination of the animal's kidneys and other organs yielded no evidence of embolization.

Future Directions The development of a VAD for use in pediatric patients has lagged behind the development of devices for adults. This is principally due to the smaller number of children that require a VAD and the historically better results of extracorporeal membrane oxygenation in children than adults. In addition, the lack of availability of a small-

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diameter valve and the apparently increased thrombogenicity of small devices operated at low flow rates have slowed the development of a PVAD. Recently, as developers and users of implantable adult VADs have focused on lengthening and expanding the use of these devices, it has become clear that all such currently available devices are too large to fit easily into patients with body surface areas less than 1.5 m 2. Poor anatomic fit may increase the risks of cannula obstruction, thromboembolism, or device infection. Thus, knowledge relating to the successful design of scaled-down devices has implications not only for the development of a PVAD but also for implantable adult pumps. Thrombus formation and thromboembolism have prevented the use of the original Penn State PVAD prototype in patients. The factors that determine the thrombogenicity of VADs are understood only in the most general terms, which provide little more insight than the ideas Virchow first expressed in the 18S0s. The inherent complexity of these systems (cannulas, valves, blood sacs, pneumatic drivers, whole animal models) makes it difficult to identify and investigate the mechanisms that enhance or reduce thrombogenicity. Nevertheless, an understanding of these mechanisms incorporated in a description of the fluid dynamics in these devices, with temporal and spatial resolution that correspond to the scales (beat frequency, blood element dimensions) that are relevant to thrombus formation, is an important part of the scientific base of information that is needed to make steady, long-term progress in VAD design possible. The advantage of studying thrombus formation in the PVAD is that this device is a geometrically scaled-down version of the adult Pierce-Donachy VAD, which is made of the same materials and has clinically proven low thrombogenicity. This suggests that d i k r e n c e s in fluid dynamics may be key to the differences in thrombogenicity. Knowledge of the flowfield inside of VADs is currently limited because all available information comes from two-dimensional techniques. Although three-dimensional techniques are under development, it may be some time before they are applied to these complex, unsteady flows. Because there is minimal cycle-to-cycle variability in the two-dimensional flowfield it may be possible to reconstruct the three-dimensional fiowtield by assembling multiple two-dimensional images from two orthog-

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onal planes obtained at corresponding times of the pumping cycle. Another limitation of current knowledge relating to VAD thrombogenicity is the lack of information concerning the initiation and progression of thrombus formation. Using an extracorporeal device and platelet or fibrinogen scintigraphy it may be possible to determine where on the surface of the blood sac these elements are first deposited and how these depositions change with time. In combination with three-dimensional flowfield data, this information on the "kinetics" of thrombus formation may provide new insights into the mechanisms of flow-related thrombus formation in VADs.

References 1. Pennington DG, Swartz MT. Circulatory support in infants and children. Ann Thorac Surg 1993;55:233-7. 2. Karl TR, Sano S, Horton S, Mee RBB. Centrifugal pump left heart assist in pediatric cardiac operations. J Thorac Cardiovasc Surg 1991;102:624-30. 3. Pennington DG, Reedy JE, et aL Univentricular versus biventricttlar assist device support. J Heart Lung Transplant 1991;10:258-63. 4. Pennington DG, Swartz MT. Circulatory support in children. Cardiac Surg State Art Rev 1989;3:381-91. 5. Lenker SA, Phillips WM. The dynamics and performance of prosthetic heart valves in vitro. Aerospace research report 74-4. Hershey, PA: Pennsylvania State University, June 1974. 6. Mussivand T, Navarro R, Chen J-F, et al. Flow visualization in an artificial heart using diffuse and planar lighting. ASAIO Trans 1988;34:317-21. 7. Woodard J, Shaffer F, Schaub R, Lund L, Borovetz H. Optimal management of a ventricular assist device. ASAIOJ 1992;38:M216-9. 8. Baldwin JT, Deutsch S, Geselowitz DB, Tarbell JM. Estimation of Reynolds stresses within the Penn State left ventricular assist device. ASAIO Trans 1990;36:M275-8. 9. Jin C, Clark C. Experimental investigation of unsteady flow behavior wit~lin a sac-type ventricular assist device (VAD).J Biomech 1993;26:697-707. 10. Shaffer F, Mathur M, Woodard J, et al. Fluorescent image tracking velocimetry applied to the Novacor left ventricular assist device. Proceedings of the ASME Fluids Engineering Meeting, Los Angeles, CA, June 1992. 11. Ramer ER, Shaffer FD. Automated analysis of multiple-pulse particle image velocimetry data. Appl Optics 1992;31:779-84. 12. Koehler MA, Hansen AC, Juretich JT, Olsen DB, Koppert E, Burns GL. Estimation of velocity and Reynolds shear stress profiles in an infant ventricular assist device. Ann Biomed Eng (in press).