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Ann Thorac Surg 1996;61:457-462
© 1996 The Society of Thoracic Surgeons

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Pumps in Progress

Cleveland Clinic Rotodynamic Pump

Department of Biomedical Engineering, Research Institute, The Cleveland Clinic Foundation, Cleveland, Ohio

Abstract

Background. It is now accepted that 70% to 80% of patients with end-stage heart failure would benefit from a permanent implanted left ventricular assist device. Previously there was little consideration of the use of nonpulsatile pumps for this function.

Methods. An extensive 5-year engineering research and development program to develop a permanent implanted nonpulsatile blood pump has been undertaken.

Results. We have developed a continuous-flow blood pump of small size (207 g) and low power requirement (6.5 watts) producing 5 L/min flow with low hemolysis.

Conclusions. This pump has the potential to be the basis of an innovative ventricular assist system.

For a healthy person at rest, the normal cardiac index is about 3.0 L•min^-1•m^-2 and increases to an index of 4.5 L•min^-1•m^-2 during moderate activity. A surface area range of 1.25 m² to 2.25 m² would encompass more than 96% of the adult population and correspond to a flow requirement of 3.75 to 6.75 L/min at rest and 5.6 to 10 L/min at moderate activity.

For many years efforts to develop an implanted blood pump that could function as a permanent assist or replacement to ventricular function were based on devices that simulated ventricular function in producing an intermittent ejection of blood, ie, pulsatile flow. In an earlier National Heart, Lung and Blood Institute program, one design criterion was the capability to produce 10 L/min blood flow at rates up to 120 beats/min.

The Institute of Medicine's recent report strongly endorsed the need to continue the total artificial heart program, but noted that 70% to 80% of patients with end-stage heart failure would be served well by a permanent left ventricular assist device [1].

Three major concerns in the past have limited the consideration of nonpulsatile (continuous-flow) pumps as an alternative to implanted blood pumps: hemolysis, physiologic effects of nonpulsatile perfusion, and technological difficulty of the motor/blood separation by a seal that ultimately failed. To use a continuous-flow blood pump as the basis of an implanted ventricular assist system necessitates that these three problems be addressed and resolved appropriately.

Hemolysis

An early report on centrifugal blood pumps [2] implicated hemolysis as a major limitation to their use as ventricular assist devices, but the development of design technology specifically applicable to blood has overcome this. The Hemadyne pump (Medtronic Inc, Eden Prairie, MN) was an early system with very low hemolysis [3], and, since that time the Bio-Pump (Bio-Medicus, Eden Prairie, MN), the Sarns pump (Sarns 3M Healthcare, Ann Arbor, MI), the St. Jude pump (St. Jude Medical, St. Paul, MN), and the Nikkiso pump (Nikkiso, Tokyo, Japan) have been approved by the Food and Drug Administration for short-term human use, the Bio-Pump being used instead of a roller pump for routine cardiopulmonary bypass in 70,000 to 100,000 patients yearly in the United States. We have used two Hemadyne pumps for biventricular bypass in animals for up to 3 months, and the serum free hemoglobin values were maintained within normal limits [4]. In temporary support of clinical patients for days to weeks with Bio-Pumps, we have also shown minimal red cell damage [5, 6]. Thus the question of hemolysis problems has been resolved.

Pulsatility and Physiology

It used to be taught that a pulse was essential to maintain systemic organ function. This was based on many early animal studies and on the documented effects of cardiopulmonary bypass with nonpulsatile perfusion by several investigators showing renal dysfunction, cerebral edema, lymphedema, hormonal abnormalities, and other undesirable effects [7–21].

When the early animal studies on pulsatility were reviewed, it became apparent that all were acute studies with results indicating the inevitable effects from surgical trauma and anesthesia and by other cardiopulmonary effects as well as often abnormally low flows and pressures. The relationship of such studies to the situation for an awake animal long-term was questionable as no such data were available.

In 1976 Johnston and associates [28] reported the physiologic effects of continuous flow in alive, awake animals surviving up to 2 weeks. They used a continuous-flow pump as a left ventricular assist device and showed minimal hemolysis and, more importantly, no deleterious physiologic effects. The study, however, used high pump flows to overcome cardiac output from the active ventricle and so produced a hyperdynamic situation with flows of up to 15 L/min.

In 1980 we reported our initial results in animals chronically depulsed (to 34 days) using an animal model [4]. In two series of animal experiments we have demonstrated that animals can survive prolonged periods (to 99 days) without a pulse; any limitations were related to the pump flow rate provided and not the absence of a pulse pressure [4, 29–31]. The major findings from those studies were as follows: Mammalian physiology can adapt to a continuous-flow regimen provided that adequate flow and pressure are provided. The lower limit of flow for continued organ function in the absence of a pulse appears to be about 10% higher than that for pulsatile flow. No organ dysfunction was seen if a continuous-flow state was induced in awake animals beyond the effects of operation and anesthesia (ie, 7 to 10 days postoperatively). The limitation to the duration of these studies was pump failure (mechanical). Cerebral autoregulation was preserved. Autonomic nerve reflex control of the cardiovascular system in treadmill studies was normal. Since our studies, another group has reported on the absence of abnormal effects with a rapid switch to continuous flow in awake animals [32].

Human data on pulseless flow effects have been mainly from situations of cardiopulmonary bypass, where flows are rarely >2.4 L•min^-1•m^-2 and there are always significant effects from operation, anesthesia, and hypothermia. In more recent years the use of centrifugal pumps for temporary ventricular assist after cardiotomy has become the standard approach, but most of these situations have also had use of an intraaortic balloon pump, which provided a pulsatile component to the pump flow. Even in these situations with an abnormal pulse wave, organ function can be maintained for several days and patients can recover well. Certainly the results of this method in the registry [33] are the same as for those in which pulsatile pumps were used, although the average recovery and weaning was quicker with centrifugal pumps.

Even more relevant are the few situations where centrifugal pumps were used in ``bridging to heart-transplantation.'' Before our use of the Thermo Cardiosystems, Inc (Woburn, MA) system, a few patients had centrifugal pumps for cardiogenic shock before transplantation [5, 6]. We have had four successful transplants, and in all patients the intraaortic balloon pump was removed to minimize the risk of sepsis. Thus, patients underwent pulseless systemic perfusion for varying periods. This allowed for observation of effects on physiology. The main findings were that recovery of deteriorating organ function depended on providing adequate systemic flow and perfusion, cerebral function was normal, there was no lymphedema, and renal function recovered.

In 1 case, successful transplantation was done after 31 days of nonpulsatile perfusion. It is interesting to note that although this patient was usually nonpulsatile, on occasion a small pulse pressure was noted-often associated with activity or emotional stimulus. This type of data suggests that some recovery of ventricular function may occur with time and would then permit that ventricle to cope with some variability in flow.

Technological Limitations

The seal, or barrier, between the blood compartment and the motor or other driving mechanism has been a major, continuing limitation to long-term use of continuous-flow blood pumps. Dry running lip seals or face seals have been used but have never shown long-term durability, due to mechanical failure or deposition encouraged by the crevice and frictional heat at the stationary/rotating element interface.

Our rotodynamic pump is designed to eliminate the motor/seal problem. It consists of three main assemblies: the stator housing, containing the motor windings; the rotating assembly, consisting of the motor magnets, the ``primary'' impeller, and the ``secondary'' impeller; and the volute housing, which encloses the pump. Perhaps 98% of the pump flow passes from the left ventricle to the inlet, through the primary impeller, and out through the discharge system to the aorta. A small quantity of blood lubricates the bearing, creating the hydrodynamic film that separates the rotating and stationary surfaces, passes into the secondary impeller, and then passes over the rotating assembly to the discharge. Motor stator-rotor magnetic attraction acts as the thrust ``bearing''. This pump has been built and tested in three increasingly successful versions. Its major characteristics are as follows:

• Overall volume 62 mL
  
• Priming volume 12 mL
  
• Weight 100 g
  
• Nominal operating point
  

• Speed 3,000 rpm
  
• Flow 5 L/min
  
• Pressure rise 100 mm Hg
  
• Power Consumption 6.5 watts
  
• Efficiency
  

Some of the key innovative features are as follows:

• Continuous-flow, centrifugal pump
  

• No valves
  
• No compliance chamber
  
• Very low parts count (cost, reliability)
  
• No inherent oscillation or vibration
  

• Hydrodynamic journal bearing
  

• Fluid film separating surfaces
  
• Hydrodynamic action self-pumps wash flow
  
• Eliminates need for shaft seal, purge fluid
  

• ``Inside-out'' motor
  

• Gap at smaller, lower shear diameter
  
• Rotor/stator eccentricity provides stable load, unaf fected by orientation, shock, etc
  
• Innovative rotor/stator details provide high effi ciency in small, low-speed, large-gap motor
  
• Motor surrounded by cooling fluid
  

• Passive, magnetic thrust bearing
  

• Motor rotor/stator attraction eliminates thrust surfaces
  
• Thrust balancing secondary impeller
  

• Biocompatibility
  

• Hemocompatible, wear-resistant, economical materials
  
• Small wetted area
  
• Optimized flow patterns
  

• Maintainability
  

• ``Cartridge'' design permits functional element re placement without disturbing cannulation
  

Hydrodynamic journal bearings take advantage of fluid viscosity, geometry, and surface velocity to create surface-separating films of the order of 25 µm and can have essentially infinite service life because the surfaces do not touch in normal operation.

Figure 1 shows the basic physics of operation. When the journal begins to rotate, it initially rolls up the wall. The rotation drags the viscous fluid into the converging wedge between the rotating and stationary elements, which creates a pressure that lifts the journal and holds it in the final position shown in the figure. There is an optimum diameter, combined with the load, clearance, viscosity, and speed, to create the maximum film thickness. This finite diameter makes space for the motor winding inside the stator, with the gap at a smaller, lower shear radius than if the motor windings were conventionally located outside the pump. Fluid flow in a journal bearing is generally laminar, and the ``inside-out'' arrangement further suppresses the Taylor vortices that initiate transitional flow. A journal bearing has self-pumping action, so it can generate its own wash flow. The hydrodynamic journal bearing has the potential to meet performance requirements, integrate with the design of other pump elements, and have long life, high reliability, simplicity, and low production costs.

View larger version (23K): [in this window] [in a new window]   Fig 1. . Journal bearing operation.

Axial thrust is minimized by balancing the pressure distribution of the secondary impeller against the pressures on the primary impeller. The residual thrust is reacted to by the magnetic attraction of the motor rotor and stator. Clearances are adequate to prevent surface contact, and no additional design elements or active controls are required.

The pump is located against the inside of the chest wall, with all of the key components forming a cartridge that can be removed and replaced in the volute housing. Therefore the heart and aorta need not be disturbed if the functional pump assembly must be removed.

Pump Development

Feasibility studies were performed by numeric modeling and resulted in a design for a mixed-flow blood pump (model 2156) with a blood-lubricated journal bearing to eliminate seal problems [34, 35]. This pump was first assembled in May 1990 and the characteristics of the assembled device were: (1) volume, 70 mL; (2) weight, 170 g; (3) internal priming volume, 16 mL; (4) 12 watts power requirement at 5,000 rpm to produce 5 L/min and 100 mm Hg pressure rise; (5) index of hemolysis approximately 0.07. The major blood flow was from the inlet, through the mixed flow primary impeller, into the collector, and out the discharge port. Additionally, some blood passed from the impeller discharge, through windows in the impeller hub, to a secondary flow path. Part of this fluid flowed down through the small diameter front bearing back to the inlet, providing lubricating and washing flow to this area. Additional flow passed along wider and shallower grooves in the main bearing to a secondary impeller that raised its pressure and sent it along the rotor outer diameter to pump discharge. This blood provided main bearing lubrication motor cooling and wash flow along its path. A significant aspect of this design was the wash of all blood-contacting surfaces. There were no crevices, tiny pivots, or other clot-forming features on the center line of the pump, where velocities and centrifugal force gradients to induce flow are inherently very low. At the same time, attention was paid to clearances, velocities, and residence times, so that shear-induced cell damage was at low levels. Surface velocities and the geometry were appropriate to generate a surface-separating hydrodynamic film, and only a small amount of blood deliberately recirculated from discharge to inlet.

This prototype had early problems with hemolysis and rotor stability, but these were resolved and hemolysis improved to 0.07 g of hemoglobin/100 L of blood pumped. An epoxy dry film coating was used to prevent destructive titanium/titanium contact at startup. A brief in vivo study demonstrated significant deposition in the front journal bearing.

Based on those results, a second prototype (model 2336) was designed that had three major design changes: (1) use of a higher performance alloy for the motor laminations better used the neobdynium-iron-boron magnets and resulted in a one-third shortening of the motor; (2) the mixed flow primary impeller was changed to a radial flow design and thus eliminated the need for the front bearing, whereas the larger mean discharge radius reduced the necessary speed; and (3) the original collector-style discharge system was replaced by a more sophisticated volute/diffuser design to more efficiently convert velocity to pressure. Retained were the inverted motor, the grooved, blood-lubricated journal bearing, and the spiral secondary impeller. The stator housing decreased in length due to the motor redesign, shortening the path of blood through the journal bearing. The direction of secondary flow came from a midpoint of the primary impeller, along the bearing, and then into the secondary impeller, on its way over the outer diameter and into the discharge.

In vitro testing of that device showed good hydraulic function with 5 L flow against 100 mm Hg pressure rise at 3,000 rpm. The electric power input to produce this performance also dropped to less than 10 watts. Modification of the stator groove geometry and offset of the axis of the stator in relation to the motor axis increased the journal bearing loads and led to improved stability. Other modifications were surface treatment of the opposing members by vapor deposition or by coating with ceramic or silver. The best results were from a nitrided titanium/dry film epoxy coating pair demonstrating 14 days' continuous use, and hemolysis index values of 0.08 g/100 L blood pumped. Although good results were obtained in vitro, short-term animal implants in vivo showed significant deposition in the grooved journal bearing area. The spiral blade secondary impeller system was also found to be sensitive to its clearance, relative to the rear shroud.

A final prototype was detailed (model 2631) (Fig 2), eliminating the grooves in the stator housing and replacing the spiral impeller blades by simpler straight blades. The straight blades proved to be more effective hydraulically and were less affected by variation of clearance from the stator flange. In vitro studies have shown 5 L/min blood flow against 100 mm Hg pressure rise at 3,000 rpm with 6.5 watts electric power (Figs 3, 4). This model has undergone hydraulic testing in various orientations to show that no significant alteration in performance occurred. Additionally, in mock loop testing the inlet to the device was connected to both mock atrial and mock ventricular pressure sources and the pump functioned in both situations. In vitro hemolysis studies have demonstrated hemolysis index values as low as 0.03 (Table 1), and short-term animal implantations confirmed the in vitro hemolysis values (Fig 5) and showed only minor deposition within the pump at housing junctions and no systemic emboli over 48 hours of function (Fig 5).

View larger version (45K): [in this window] [in a new window]   Fig 2. . Cross-section of model 2631 rotodynamic pump.

View larger version (29K): [in this window] [in a new window]   Fig 3. . Model 2631 flow versus pressure (Delta P) versus revolutions per minute (RPM).

View larger version (28K): [in this window] [in a new window]   Fig 4. . Model 2631 input electric power versus flow versus revolutions per minute (RPM).

View larger version (21K): [in this window] [in a new window]   Fig 5. . Plot of daily free serum hemoglobin (Hgb) levels from in vivo study. (PO = postoperative.)

Conclusion

Continuous-flow blood pumps are showing encouraging results for the development of the recent generation of implanted blood pumps with potential advantages of small size, lower cost, and durability.

Footnotes

Presented at The Third International Conference on Circulatory Support Devices for Severe Cardiac Failure, Pittsburgh, PA, Oct 28-30, 1994.

Address reprint requests to Dr Golding, Biomedical Engineering (Wb-3), The Cleveland Clinic Foundation, 9500 Euclid Ave, Cleveland, OH 44195.

References

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35. Golding LAR, Smith WA. US patent #5,324,177 Sealless rotodynamic pump with radially offset motor, 1994.

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This Article Abstract Alert me when this article is cited Alert me if a correction is posted Citation Map Services Email this article to a friend Similar articles in this journal Similar articles in PubMed Alert me to new issues of the journal Add to Personal Folders Download to citation manager Author home page(s): Leonard A. R. Golding Permission Requests Citing Articles Citing Articles via HighWire Citing Articles via Google Scholar Google Scholar Articles by Golding, L. A. R. Articles by Smith, W. A. Search for Related Content PubMed PubMed Citation Articles by Golding, L. A. R. Articles by Smith, W. A.