For in vitro characterization, the pump was attached to a mock physiologic laboratory flow loop. Although this loop is generally equipped to simulate vascular compliance, and can even be configured with a pulsatile pump to represent a beating ventricle, the steady-state results reported here reflect time invariant measurements recorded after changed flow conditions had settled. Tests were repeated with several working fluids, including water, a blood analog aqueous solution (40% glycerin), and whole bovine blood. Flow measurements were made with an ultrasonic flow probe (Transonic Systems, Ithaca, NY), and pressure measurements were made with solid-state, catheter-type pressure transducers (Cobe Cardiovascular, Arvada, CO). Speed and motor parameters were set and controlled using software specifically designed to develop the HeartMate III. All measurements were automatically processed and recorded using Labview (National Instruments, Austin, TX).
For in vivo testing, a young bovine model (<100 kg) was used (Figure 6). Surgeries were performed at the University of Pittsburgh McGowan Center for Artificial Organ Development in accordance with Guide for the Care and Use of Laboratory Animals. 9 The pump was implanted with routine surgical techniques via left ventricular apical cannulation and anastomosis to the descending aorta.
The electronics for power and control of the LVAD and for data acquisition resided on a shelf and alongside the cages containing the calves. The animals were continuously monitored for general welfare, and certain parameters such as arterial pressure, pump and pulmonary artery volume flow rates, pump speed, and blood chemistry (including plasma free hemoglobin) were periodically recorded. In two animals, measurement of pressure head throughout the duration of the studies were made with flocked carotid lines and fluid filled LV catheters.
The pump was prescribed to run at 3,500 rpm perioperatively, then gradually adjusted to maintain volume flow rates at 5 ± 1 L/min throughout the study.
Necropsies were performed immediately after sacrifice. The positions of the heart, pump, and instrumentation were photographed at various stages of resection. The heart, lungs, kidneys, pancreas, liver, and adrenal glands were harvested, sectioned, photographed, and sampled for histologic examination. The pump was immediately disassembled and photographed; remarkable findings were recorded on a pump map.
Three sets of parts were fabricated, from which three fully functional, implantable LVADs were assembled. All three have been implanted in calves, reused, and remain available for future implants, although a second set of three LVADs, bearing a modified exit angle for the percutaneous lead (explained below), are now being prepared for continued animal studies.
The steady-state pump pressure head versus volume flow rate (H-Q) characteristic is shown in Figure 7. Note that the design point (7 L/min at 135 mm Hg) occurs between 4,500 rpm and 5,000 rpm, which happens to be near the optimal speed for maximum motor efficiency. Furthermore, the desired maximum flow (10 L/min at 135 mm Hg) is achieved below 5,500 rpm, also at high efficiency and well below the motor’s maximum capability.
The hydraulic efficiency of the pump is shown in Figure 8. Note that these efficiencies are ideally maximum near the design point and higher. (It is less important to be efficient at low speeds, where power consumption is lower.)
Five of six bovine calves implanted with the HeartMate III survived to their elective termination dates. No mechanical or electronic failures occurred in the six implanted pumps. The implant durations for the electively terminated animals were 40, 27, 59, 42, and 49 days. (Note that the actual lengths of studies that were nominally “30-day” or “60-day” were governed by coordinating surgical schedules.) Pump volume flow rates were consistently maintained between 3 L/min and 6 L/min during sedentary periods and elevated to a range of 6 L/min to 11 L/min during treadmill exercise. The implanted pumps were tolerated very well by the calves with few complications (excepting several instances of immediately postoperative, moderate bleeding, several instances of chronic pulmonary atelectasis, and one drive line infection).
One calf was sacrificed on the 27th postimplant day (earlier than the planned day 1 week later) due to severe, untreatable pneumonia secondary to left lung atelectasis. This condition was caused by the LVAD’s percutaneous lead, which interfered with the expansion of the left lung.
Except for small, apparently stable and adherent islands of fibrous tissue in several of the explanted in vivo pumps, there was only one incidence of significant tissue deposition within the pump. This was a 4 mm thick disc of thrombus under the rotor, and a nearby, similarly sized mural thrombus in the volute, both of which appeared to be very fresh and were surmised to be a consequence of a study performed just prior to sacrifice in which flows were alternately very low and retrograde. (This was a planned exercise.) In addition, rings of white thrombus, starting at the outflow edge of the pump and extending a few millimeters toward the graft, and consequent partial outflow occlusion were observed.
In vivo plasma free hemoglobin values remained statistically indistinguishable from baseline values (4–10 mg/dl) in all six animal studies, demonstrating that whatever hemolysis was caused by the pump was of no clinical importance (Table 1).
It was a fundamental goal of the HeartMate III mechanical design to achieve optimal pumping performance and to enclose a sophisticated motor in the simplest possible way, this being the best means of achieving immediate success (i.e., early implant experience characterized by no mechanical failures). Furthermore, optimal peripheral hardware was incorporated in even the earliest versions of implantable pumps, again seeking immediate success.
The very simple main flow path is a consequence of the motor’s “J” core configuration (i.e., the stator laminations are J shaped), which places an imaginary plane near the mid-height of the pump, below which are confined the motor and levitation elements, and above which the hydrodynamic elements are permitted free reign. The burden was thus placed on the motor design to fit around the optimal hydrodynamic arrangement to minimize the risk of thrombogenesis, a choice that paid off with very low deposition of thrombus in the first six implants. It has already been mentioned that the fresh thrombus found in one pump under the rotor and in the volute was likely caused by stagnant flow induced in a brief study just prior to sacrifice. Outflow stenosis seen in explants has been addressed with a modification to the outflow joint, which will be tested in the next series of implants. Plasma free hemoglobin values that do not depart from baseline in all six in vivo studies have demonstrated that HeartMate III hemolysis is of no clinical importance.
Designing with access to the resources and infrastructure of established, FDA approved LVADs (i.e., HeartMate) has afforded the wherewithal to develop the HeartMate III pump (i.e., hydrodynamic elements) and peripheral hardware (i.e., inflow and outflow components) concurrently, borrowing from the enormous clinical experience gained with presently available HeartMate systems. Although generally regarded as secondary in importance when setting out on a new pump design (compared to the primary purpose of “pumping”), items such as cannula, grafts, articulating mechanical joints, valves (if any), connectors, and cables can account for most of the propensity for clinical troubles when implanted. This is expected to be increasingly true as the indications for LVAD use are expanded, imposing requirements for longer implant durations.
For example, clinical experience has suggested that some freedom of movement between the native heart and pump ameliorates stresses imposed on the heart by an imperfectly placed pump, subdues unnecessary pump motion and potential tissue trauma caused by cardiac motion, and accommodates a certain amount of reverse remodeling in a convalescing cardiomyopathic patient, should such occur. Furthermore, a convenient way to retrieve an implanted pump when salvaging the native heart (i.e., in a recovered patient, when a transplant is not to occur) is to have a section of flexible graft suitable for ligation available near the inflow cannula. In the HeartMate III, both of these requirements are met with a flexible “recovery section” between the pump and apical cannula (i.e., in order that the pump can be removed, leaving behind the cannula).
For another example, although the high pressure in an outflow graft tends to assist patency, in certain circumstances (e.g., a severe bend in the graft), there is a risk of graft collapse or kinking. To preclude that, a reinforced e-PTFE tube has been attached to surround the graft, providing bend relief and maintaining patency.
The pumps used in the first set of animal studies had percutaneous leads that exited the pump directly beneath the outflow (visible in Figure 2), an ideal orientation for the human configuration. However, in the calf model, to exteriorize at the dorsum, the lead was forced into a half-loop in the thorax that subsumed a volume normally occupied by the left lung. This method resulted in some degree of atelectasis in all six animals, and ultimately to the demise of one of them. The orientation of the percutaneous lead has been rotated by 75° (visible in Figure 1) to avert this problem in the next series of animal implants.
Other than that change to be implemented for future bovine studies, the only difference between the pump used for in vivo studies and the eventual human configuration is the presence of the elbow in the inflow segment, included to accommodate the bovine anatomy and presently regarded as unnecessary in humans.
The progress of HeartMate III from concept, through design, and into animal studies has been rapid, and the successful performance has been very encouraging. Several enhancements to HeartMate III are presently being implemented, including incorporation of motor drive and control electronics within the pump, and integration of a TETS (transcutaneous energy transmission system) to eliminate the percutaneous drive line. These enhancements will continue the progression of this LVAD’s development as a compact, reliable device to treat CHF, and a viable commercial endeavor.
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Copyright © 2001 by the American Society for Artificial Internal Organs
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