- Email: [email protected]
Long-term biventricular assist
Panel V
Long-Term Biventricular Assist Moderator: Charles-Julien Hahn, MD Panelists: William S. Pierce, MD, Don B. Olsen, DVM, Kenneth Butler, and Robert T. V. Kung, PhD
D R PIERCE: I am going to discuss an electrically powered artificial heart that we have been working on for 8 years. It is placed in the chest in the orthotopic location, and the electric motor and motion translator, which is a device called a roller screw, is positioned between the two ventricles. These are pusher plate activated ventricles. We have to have all of the component parts that most of you are familiar with. Because this is a closed nonvented system, there has to be a compliance chamber in the left thorax. The electronic control system and emergency implanted battery system is placed in the subcutaneous position. And, of course, energy is transmitted with the wireless technique of inductive coupling with a primary and secondary coil and then an external battery pack. I think most of you are familiar with the overall design principle. On cross section of the artificial heart, there is a pusher plate located at each end. There is a fixed distance between those pusher plates. In other words, they are attached but they are not attached to the blood sacs. The rotor of the motor is attached to the nut of the roller screw. As this motor rotates some four or five revolutions in one direction and then counter rotates, the pusher plate moves back and forth a distance of about 2.5 cm. The two main items of the prime mover and motion translation system are the brushless direct-current motor (Saracen) and the stainless steel roller screw (SKF transroll). These devices are widely used in the machine tool industry, and they are also a component of the wing flaps of the Concorde. Any of those of you who came to the meeting across the ocean in the Concorde, the SKF roller screws actuate the wing flaps on that plane. Thus the components include the motor and motion translator, biomer polyurethane blood sac, and the polysulfone end caps, left and right. We use Bjork-Shiley tilting-disc valves. The implanted batteries are designed to provide about 45 minutes of energy to the pump if the primary and secondary coils are separated. Because gases can diffuse in and out of the compliance chamber, we have to have some sort of a port through which we can access that compliance system and measure the gas pressure and change that pressure in accord with the needs of the device. In the laboratory, the first designed pump that we worked on instead of using a roller screw motion transPresented at the Circulatory Support 1991 Symposium of The Society of Thoracic Surgeons, San Francisco, CA, Nov 16-17, 1991.
0 1993 by The Society of Thoracic Surgeons
lator used a drum cam having three cam followers. We had pretty good luck with 1 of those animals, but we had quite a bit of problem with the mechanics and with the little tiny cam followers. We subsequently abandoned that system and began to use the roller screw system. And we are starting to have a little better luck with our animals; 1 is still going and has passed the 11-month mark. In the design of the animal experiments, we use closed systems. The pump is placed in the thorax as is the compliance chamber and battery pack, or we use an implanted battery and control system, the energy transfer system, a subcutaneous access port to get into the compliance chamber, and the external battery pack, or, of course, most of the time the animal is just kept on the house electrical supply. At 200 days after implantation, the animal weighs about 103 kg and the pump handles about 10.5 L of blood per minute. The animal has a very low venous oxygen saturation as you might expect. But he still eats vigorously and actually looks pretty good. So we are very, very encouraged about this. This particular animal did not have the energy transfer with the coil arrangement, but the ones that we are doing now do have that system. So we are looking forward to doing an additional series of animals and see if we can continue the record of having animals that will last for a long time. Our next speaker is going to be Dr Don Olsen, and we will hold our questions until the end.
D R OLSEN: The University of Utah and Institute for Biomedical Engineering has been working on an electrohydraulic artificial heart now for about 4 years. And in concept we have the following: internal batteries, internal electronics, and the TET telemetry coil system, with TET meaning transcutaneous energy transfer of power to the induction coil system. Our internal batteries are supplied by the consortium group between Ottawa Heart Institute and the University of Utah, and we are very pleased with that. The entire intrathoracic components all fit within the pericardial sac. We do not have a volume displacement or compliance chamber in our prototypes presently. We have always strived for simplicity and miniaturization. We have a titanium encasement. There is a brushless direct-current motor housed within this containment with the electrical pass through for back EMF control of the commutation. The axial flow pump reverses such that the hydraulic oil flows back and forth between the right and Ann Thorac Surg 1993;55:227-32
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left ventricle, coupling the right and left ventricle both volumetrically as well as in timed sequencing. Our test model that we have been using in a series of acute calves is the Utah 100 type artificial heart. Actually this is an 80-ml stroke volume device. Today I would like to address how we achieve balance in this coupled system without external ports and without a volume displacement or a compliance chamber. We do that by making a small atrial septal defect here between the right atrial cuff and the left atrial cuff. This, then, is a small-diameter, and we will talk about that orifice that allows a volume of blood to move back and forth. Our surgical procedure is the following. First we expose the natural heart, do our cannulation, and establish the cardiopulmonary bypass. Then we do the cardiectomy. We make sure of hemostasis then. We now are measuring through our heart-lung machine the systemic flow of the calf. But we are also able to measure the amount of bronchial flow return into the left atrium. We use this measurement in making a decision on which of the predesigned and prebuilt intraatrial septal orifices to use in these atrial cuffs. Then we go ahead and surgically implant the device and wean the calf from the cardiopulmonary bypass on to the electric powered artificial heart. The purposes of the in vivo acute experiments were the following: to establish the intraatrial shunt performance, verifying that we can indeed achieve ventricular balance and that we maintain patency of that atrial shunt; to look at the hematological effects, hemolysis, and thrombus formation within the device and all of its components and attachments; and to examine the electrohydraulic TAH 20 system’s performance both in regard to the fit anatomically and in regard to the physiological performance as far as adequate cardiac output is concerned. In a series of 6 calves the measured bronchial blood flow as a percent of the flow of the cardiopulmonary bypass machine goes from a low of 5% to a high of 17%.We have learned that in those calves that exceed 12% we have a difficult time in establishing balance. The shunt size that we have selected to implant with each particular bronchial flow has been narrowed down to two sizes; we avoid putting the electrohydraulic artificial heart in a calf that has a high left-left shunt. During the course of the experiment we have measured the left atrial pressure and subtracted from that the right atrial pressure giving the change in pressure across this intraatrial shunt. We find that the highest we had was 6 to 10 mm Hg and the lowest we had was 2 to 5 mm Hg. In the longest living calf (about 9 days), the average difference between the mean left atrial pressure and the mean right atrial pressure was about 3.5 mm Hg, greater on the left than on the right. Occasionally we would have some spikes in pressure difference. Presently our motor speed control algorithm is being addressed to see if we can minimize these spikes, but they are relatively transient. Part of it I think might be a hammer effect from the method of open port measurements that we employ. On the bench we tried to look at some of the flows using contrast media. This shows quite readily that we do
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have that shunting. There is a higher shunt from left to right than the reverse. In the calf who lived 9 days the left atrium and the orifice of the small intraatrial shunt were quite clean. The conclusions from these 20 in vivo experiments were the following: The intraatrial shunt provides adequate balance throughout a variety of bronchial blood flows. The intraatrial septal shunt has shown no evidence of thrombosis or occlusion, albeit so far our longest experiment was 9 days in these acute studies. The size constraints of the animal do exist and can affect the cardiac output and device performance even in our miniaturized system. Furthermore, we have had no evidence of intraventricular thrombosis formation except in one experiment where we had an electrical device failure and stoppage of the device without adequate heparinization. These experiments have been a great learning experience, resulting in 18 design changes. The system meets the physiological demands with its automatic control algorithm. Now we have made a totally integrated device. This device is being built now for animal testing; it has an 80-mL stroke volume and will fit totally within the pericardial sac. Doing extensive fit trials we have successfully fit this clay model into a 57-kg and a 68-kg cadaver, although the 57-kg cadaver did have cardiomegaly. But it does fit well. We also have successfully fit the model with slight modification of the length and angulation of the inflow outflow ports into a 75-kg sheep and an 80-kg calf. This device will have an 80-mL effective stroke volume. Presently we are preparing for additional animal, cadaver, and human intraoperative fit trials.
MODERATOR HAHN: Thank you so much, Dr Don Olsen. Now I will ask Mr Ken Butler to give his lecture. MR BUTLER Indeed it is a pleasure for me to be here representing my colleagues at the Cleveland Clinic and my co-workers at Nimbus and present to you a run down of our particular artificial heart device that we have been working on for the last 4 years. From the anatomical view, central is the blood pump energy converter unit that fits in the chest, replacing the natural heart. Our system is a dual pusher plate and it does have to have its compliance space managed; we do this with a flexing diaphragm chamber that is connected via a duct to the interspace. Gas is transferred back and forth between the interspace of hemodiaphragms and the flexing chamber, which faces into the lung space. We also have refill port that is accessed percutaneously. The main power is derived from primary batteries that are located on the outside, and we also use a TET system including the secondary coil, which is implanted subcutaneously over the left chest. We are using the Thermedics or Thermocardiosystems version, and it has a secondary coil that fits on the outside over that. That transmits alternating current power across the skin, which is then conducted down to what we call a power module, which
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is actually a rectifier that converts the current into directcurrent power, which then drives the energy converter. The electronics package in our system is contained within the central blood pump and energy converter. An appendage is the internal battery, which is connected into the power module. The power module also serves as a connecting point for the various cables. So you can remove the internal battery and remove the secondary coil without disturbing the energy converter. The blood pump itself is in the "human configuration" as a result of extensive anatomical studies conducted at the clinic, where every dimension was critically evaluated to make sure it would fit in a maximum number of adult patients. The blood pump energy converter has been our main focus of work over the last several years. This is an entirely new design for us. Although we have had a considerable amount of experience with left ventricular assist systems using actuation and energy converter technology, this is really a new departure for us. The energy converter is contained within a titanium superstructure. The ventricles themselves are made of a carbonimpregnated epoxy, which has saved us a considerable amount of fabrication complexity, and they certainly are lighter than our titanium castings. There is a dual pusher plate. We simply rely on an alternate ejection from the left to the right. While the actuator is engaged in ejecting on the right side, the left side is free filling. The pusher plates, when they are filling, are disengaged from the sac or the rolling diaphragm. Conversely, on the right side when it is ejecting on the left it is filling on the right. So our control system essentially is to match our actuator cycling frequency with the fill rate of the blood pumps. If we go too slow we will impede filling. If we go too fast we will actually start an ejection while one of the ventricles is filling. So all we have to do is match our motor speed to the fill rate. And we do that electronically with a rather simple system, which I do not want to get into here. The energy converter is based on an electric motor continuously driving a gear pump that puts out a flow of hydraulic fluid. That fluid cycles a piston back and forth in a container, which provides the actuation. We have a mainspring in the device that gives up force when it is ejecting on the left side and is cocked when the right side is ejecting. So it evens out our power distribution. The diaphragms are the flexon rubber that the Clinic has been using for years. And we do use the biolyzed coatings that the Clinic has developed on all the blood contacting surfaces. This is the book on our system as far as the key characteristics. I think all the systems are coming to about these weights and volumes. Our capacity is a 60-mL blood pump. It can generate up to roughly 10 L/min, 60 beats per minute. The direct-current power consumption at its maximum is 15 W. Our internal battery time, based on our latest tests, is in excess of 60 minutes, and that is at a 6 Llmin condition; that is specified for the internal battery. We have been concentrating the last few months on in
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vivo tests, and I would like to focus the rest of the presentation on that, which gives you some idea of how these systems work and possibly some of the problems that we have with our animal models as has been alluded to before. We normally are using 80-kg calves. We are focusing just on the central blood pump and energy converter. We are using external electronics and a vent for compliance. Basically, what we are trying to show that we can support one of these animals and that the system is compatible with the physiology. And of course, most important, we are using it to test our system. We have done four experiments to date and our progression has gone from a planned I-day acute model to a couple of intermediates, and our last one was 4S days. On these early animals we are using all the blood pressure instrumentation we need: inlet and outlet pressures on both sides. These animals therefore are fairly risky in that there are indwelling catheters in them. One of the things we have learned over the years is that a calf is not a human, and it is very difficult to take a human configured system and try to force it into a calf. Therefore we have changed the inlets a little bit to help u s get better anatomical fit in the calves. The big difference is that the outlet on the left side is bent so it intercepts the aorta. I would like to give some data now from a 45-day experiment. When we plot out all the data points we took, they average around 10 mm Hg in the right atrium. On the left side it i s a little more scattered, but i t would average out about 16 mm Hg. One of the things about these animals, and i n particular onc, hi, mnxcd o u r pump out the first night. He was running a t 1 0 Limin t w t ~ n t i a l l v on day 1 and that is wherr he stayed. Although t h i 5 is good healthy workout for our actuator, it does not do much for our control svstem, which indiccittbh cj prohlc~ni we have in the model. We rn'ixed out a t 160 beats t o r thc whole experiment and oiir powt'r drab\. was 15 W throughout. In some of the e,irhcr isxpcrimcnts where the calves were not as hcalthv, w c ) were ablt. to go to the lower flow rates and verifv that indeed our rontrol svstrm does function like we intend it But w c n t d to h a w a lot more experience with that. To summarim the in vivo results, w e can maintain the hemodynamics of thcse animals Thr f i t is extremely good. We have not had anv ht~molysis ( i r thrombus problems. The diaphragnis come o u t absolutely cledn after the experiment. What we are going to be doing i s continuing the in vivo studies, our sequence of involving the other system integrLitionrlements. This is due to c ~ i i n e in a couple of months. We will add that into the energy converter, add the comp1i.inc-r ' I E T s i n with the rest OT the systems, and bv the cnd of this program we will be doing a total. I think 1 speak for all of us when 1 sav that these programs are supported by the devicrs and technology of the National Institutes of Health, and as I understand the plans we have another couple vears to go on the current contracts. There probably will be a follow on for 3 or 4 years, which puts us in a time frame such that wc probably would be ready to use human implants on a trial
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basis somewhere in the year 2OOO. That is the time line we’re working on. MODERATOR HAHN Thank you so much, Mr Butler, for this exhwnely nice presentation. Now we will ask Dr Kung to come here and give his report. DR KUNG: This is a collaborative effort between Abiomed and the Texas Heart Institute on the development of the electrohydraulic total artificial heart. Because I am last on the agenda I thought the best thing for me to do is to focus on the specihc and unique features of ow system instead of going into the description of the entire system. Fxternal parts of the system are battery electronics including a TET, internal battery, blood pump, energy converter, hybrid electronics, and sensors for the control of the system. I will describe how our energy converter works, how our blood pump interfaces with the atria, the design of the inflows for the pump, and how the compliance or rather the management of the left a d right plow differences is done. I also will give some information on our in vivo results. The system is electrohydrauk, and the difference between our electrohydraulic system and that of Nimbus, which I can describe as also eledrohydraulic but eh?ctrohydraulic converted to mechanical energy, is that we use a direct electrohydraulic drive of the flexing membrane, which is what the Utah group does also. The difference between our system and the Utah system is that we use a unidirectional rotary pump. In other words, the rotary pump is not reversed on every cycle. Fluid reversal is accomplished through an electromagnetic fluid valve. The blood contacting part of the system is designed to be smooth and seamless, and it uses angioflex, which is a polyurethane proprietary to us, and also trileaflet valves made of the Same material. We use blood as the volume compensation: I believe we are the only ones taking that approach. As you heard yesterday, if you have a biventridar support system-in this case this is also classified as biventricular-you have to make sure that the left flbw and the right flow somehow match or the difference matches the bronchial shunt. In current systems that have been used clinically vou have to manage this by essentially dialing in the flow so that you do n d flood the lungs. When you flood the lungs that means the right side is flowing too much. One of the manifestations of that is that the left atrial pressure goes up. If the left atrial pressure goes up you can think of a way of using that as a means of feeding back to your system. Because for a total artificial heart you obviously cannot keep on dialing the right side flow, you have to use what physiological information is provided by the artificial heart and the living system to control or feed back to adjust the right side flow. There is a hydraulic pump with a fluid switch that switches the fluid back and forth. To manage the left and right side flow, we have put a shunt flow in the hydraiilic circuit and placed that in reference to
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the left atrial pressure, which means that you can put a flexible diaphragm in contact with the left atrial reservoir. If the left atrial pressure goes up the shunt flow increases in the hydraulic circuit. Men the shunt is krger to the right side d the hydraulic chamber, it reduces the amount of blood that is being drawn into the right-sided W d pump. So as the left atrial pressure goes up there is less stroke volume on the right side, and as the left atrial pressure goes down there will be more right-sided stroke volume. One of the advantages of this is that the flow compensationchacteristics would not vary over time. In long-term studies we are trying to demonstrate some of these characteristics. I would like to talk a little bit about the control algorithm. With it we have taken a very simple approach to control the system. In the hydraulic circuit there are pressure transducers, one pressure transducer that measures the filling pressure during the fill cycte of either the left side or the right side. That information is used to control the hydraulic pump. If you have a high enddiastolic pressure on one beat it Senses that pressure and then it tells the pump to go to pump a little bit harder on the next beat, so you are trying to bring the end-diastolic pressure to a level that you have preset, whether it is zero or some mmber that is dose to zero filling pressure. The control algorithm is also responsive to afterload because the afterload changes affect the pumping rate or the filling rate. Once you change the f i g rate you change the end-diastolic pressure, which then is sensed again, changing the pumping speed of the rotary pump. The other control is focused on trying to make every stroke a full stroke. We have an optimal sensor in the system that monitors the diaphragm position. That information is then used to determine the beat rate of the system. If there is a lot of fluid or i€ the atrial pressures are high, the beat rate will increase to maintain the fullstroke and at the same time to try to bring the end-diastolicpressure down. Now let me go to the energy converter. There are two channels that connect the hydraulic chambers to the energy converter chamber, so the flow of hydraulic fluid can go either way on each beat. It reverses the flow. As I say, there is a unidirectional rotary pump and a centrifugal one pumping hydraulic fluid. There i s a switching valve concentric to the pump that assumes two different positions, which then opens a set of ports, which then directs the flow either from one pump to another or vice versa. It is actually a very simple system. You are switching the hydraulic chambers, which pumps the blood to the two alternate pumps. This is an alternating pumping system. Now let me describe briefly the Mood pump. The valve that is used is a polyurethane trileaflet valve. Concerning the experience that we have on these bbod-contacting components, the valves have gone about 400 million cycles in vitro. The pumps we have tested up to 100 million cycles, and in vivo about 150 days. The design of the whale system is to achieve a seamless Mood contacting surface. Now in the inflow region, where you have tissue and cuff in contact and s u h w s that hold the cuff and the tissue together, because we have
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discontinuities, it is much better to have a cuff that can promote growth of tissue. But to avoid ingrowth into the inflow of the device, we have made a rough surface on the cuff and a smooth surface in the device. The cuff will be sewn onto the atrial remnants and then the rest of the device will be snapped into the atrial cuff, so this is a quick connecting approach. By this approach there are no end-to-end seams. To get around this rough and smooth configuration, we actually have a smooth surface tube project into the atrial cuff. This provides a tissue barrier to stop the growth of tissue. It is helped by the fact that you have flow going in this way and you probably have some recirculation, which prevents tissue from climbing up the smooth step. The protrusion into the atrial cuff is about 3 mm in height, and in in vivo studies it does act as a barrier quite well. It essentially tries to keep away blockage problems. The system actually can operate at more than 8 Wmin, but at 8 Wmin it takes about 20 W. In our chronic studies, we will concentrate on system integration; we will include hybrids, TETs, and in vitro study as well as the liability evaluation in this program.
MODERATOR HAHN: Thank you for this very interesting lecture. There are plenty of problems we could discuss, but first I should like the audience to ask questions, because the time is extremely limited. DR D. GLENN PENNINGTON (St. Louis, MO): I would like to ask Dr Pierce about those animals that he showed us. What is the stage of development of those animals? Is that an integrated system? In other words, are those implanted electronics? And is that an implanted compliance chamber? How complete is the system we are seeing in those animals?
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simple enough so that things are not that expensive. Our estimate is that we can do it, once the wholt. tcchnulogy is developed, for about $25,000. Ndw. I d i i t suit’ thdt 50me of the people here would think otherwise, but we have done some evaluation of that.
MR BUTLER: In our particular case, L \ C puject a cost between $35,000 and $50,000. Actually, uur total heart device is cheaper than our left ventricular ’iAst system because it is simpler. MODERATOR HAHN: Doctor Olsen, do you have comment?
DR OLSEN: 1 think when we consider the additional constraints and costs of Food and Drug Administration approval, it will be the upper limit of Mr Butler’s estimate of $50,000. DR PIERCE 1 think it is important to note that there is nothing inherently expensive about these devices. The roller screws we use in our units cost, I think, $1,000.The motors are $400 or $500,and there is one fairly complex titanium machined part, the costs of which would not be high if you were going to make a bunch of them. But by the time you put in the animal experiments and all of the careful controls that have to be done before something like this can be sold, the price escalates. 1 would estimate $5O,OOO, something like that, and I do not think these total hearts are going to be more expensive than the lett ventricular assist devices. There is very little additional complexity to them. MODERATOR HAHN: Another question?
DR PIERCE: Well there is really a progression. As I mentioned, the animal that is now over 11 months does not have energy transferred by the coil system. We do have another animal that we have subsequently operated on that has implanted batteries, a closed system, and energy transfer by the primary and secondary coil. The only thing that we are lacking now is the ability to talk to our system and get information in and out of it. Other than that, the system is pretty much complete. MODERATOR HAHN: Another question? Doctor Champsaur? DR GERARD CHAMPSAUR (Lyon, France): What is the projected price of the total artificial heart using your current techniques?
DR KUNG: All of us probably will have an answer to that. Let me just start by saying that we have in the design of the system and its development always tried to make it
DR PAUL DIDISHEIM (kthesda, MD): I would like to ask each of the panelists about thromboembolism. Because we know from past experience that an implanted device can look free of thrombus or relatively free of thrombus and still be a generator of thrombuemboli, I would like to ask each of the paiielists if they have yet had an opportunity to look at end organs, specifically the kidney, which has in the past been a useful thromboembolus trap. MODERATOR HAHN: Doctor Kung, would you answer? DR KUNG: Yes, in our experience, we have not had any long-term experience with the total artificial heart yet. But the technology that we have used in the permanent ventricular assist device is really quite similar to what we have here. And there we have had experience up to 5 months with no infarcts or emboli.
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MODERATOR HAHN: Another comment from the panelists? MR BUTLER: We just completed a 45-day experiment with the total heart. That is our longest. But it does have the biolyzed coating that we have had considerable experience with on the left ventricular assist devices. This 45-day animal was run without anticoagulants, and I do not think all the histology data are in yet, but the gross examination showed no infarct problems. DR OLSEN: Our limited experience in acute experiments with the electrohydraulic system would not lend itself to respond there. We have seen none, however, in 9 days. MODERATOR HAHN: Doctor Pierce?
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DR PIERCE If we follow the design principles that we have all learned the hard way over the past 20 or so years and then we keep our animals on warfarin sodium and their renal function has been basically normal, the devices have been clean when they have been taken out. So thromboembolism has not been a problem in our case. We have had broken wires. We have had moisture get into the devices. We have had just about everything you can think of. But thromboemboli have not been a problem. MODERATOR HAHN: I guess I must stop this session. It is a great pity, because we had so many questions to ask about these electromechanical devices that will be the future in cardiac replacement.
Long-Term Biventricular Assist Moderator: Charles-Julien Hahn, MD Panelists: William S. Pierce, MD, Don B. Olsen, DVM, Kenneth Butler, and Robert T. V. Kung, PhD
D R PIERCE: I am going to discuss an electrically powered artificial heart that we have been working on for 8 years. It is placed in the chest in the orthotopic location, and the electric motor and motion translator, which is a device called a roller screw, is positioned between the two ventricles. These are pusher plate activated ventricles. We have to have all of the component parts that most of you are familiar with. Because this is a closed nonvented system, there has to be a compliance chamber in the left thorax. The electronic control system and emergency implanted battery system is placed in the subcutaneous position. And, of course, energy is transmitted with the wireless technique of inductive coupling with a primary and secondary coil and then an external battery pack. I think most of you are familiar with the overall design principle. On cross section of the artificial heart, there is a pusher plate located at each end. There is a fixed distance between those pusher plates. In other words, they are attached but they are not attached to the blood sacs. The rotor of the motor is attached to the nut of the roller screw. As this motor rotates some four or five revolutions in one direction and then counter rotates, the pusher plate moves back and forth a distance of about 2.5 cm. The two main items of the prime mover and motion translation system are the brushless direct-current motor (Saracen) and the stainless steel roller screw (SKF transroll). These devices are widely used in the machine tool industry, and they are also a component of the wing flaps of the Concorde. Any of those of you who came to the meeting across the ocean in the Concorde, the SKF roller screws actuate the wing flaps on that plane. Thus the components include the motor and motion translator, biomer polyurethane blood sac, and the polysulfone end caps, left and right. We use Bjork-Shiley tilting-disc valves. The implanted batteries are designed to provide about 45 minutes of energy to the pump if the primary and secondary coils are separated. Because gases can diffuse in and out of the compliance chamber, we have to have some sort of a port through which we can access that compliance system and measure the gas pressure and change that pressure in accord with the needs of the device. In the laboratory, the first designed pump that we worked on instead of using a roller screw motion transPresented at the Circulatory Support 1991 Symposium of The Society of Thoracic Surgeons, San Francisco, CA, Nov 16-17, 1991.
0 1993 by The Society of Thoracic Surgeons
lator used a drum cam having three cam followers. We had pretty good luck with 1 of those animals, but we had quite a bit of problem with the mechanics and with the little tiny cam followers. We subsequently abandoned that system and began to use the roller screw system. And we are starting to have a little better luck with our animals; 1 is still going and has passed the 11-month mark. In the design of the animal experiments, we use closed systems. The pump is placed in the thorax as is the compliance chamber and battery pack, or we use an implanted battery and control system, the energy transfer system, a subcutaneous access port to get into the compliance chamber, and the external battery pack, or, of course, most of the time the animal is just kept on the house electrical supply. At 200 days after implantation, the animal weighs about 103 kg and the pump handles about 10.5 L of blood per minute. The animal has a very low venous oxygen saturation as you might expect. But he still eats vigorously and actually looks pretty good. So we are very, very encouraged about this. This particular animal did not have the energy transfer with the coil arrangement, but the ones that we are doing now do have that system. So we are looking forward to doing an additional series of animals and see if we can continue the record of having animals that will last for a long time. Our next speaker is going to be Dr Don Olsen, and we will hold our questions until the end.
D R OLSEN: The University of Utah and Institute for Biomedical Engineering has been working on an electrohydraulic artificial heart now for about 4 years. And in concept we have the following: internal batteries, internal electronics, and the TET telemetry coil system, with TET meaning transcutaneous energy transfer of power to the induction coil system. Our internal batteries are supplied by the consortium group between Ottawa Heart Institute and the University of Utah, and we are very pleased with that. The entire intrathoracic components all fit within the pericardial sac. We do not have a volume displacement or compliance chamber in our prototypes presently. We have always strived for simplicity and miniaturization. We have a titanium encasement. There is a brushless direct-current motor housed within this containment with the electrical pass through for back EMF control of the commutation. The axial flow pump reverses such that the hydraulic oil flows back and forth between the right and Ann Thorac Surg 1993;55:227-32
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left ventricle, coupling the right and left ventricle both volumetrically as well as in timed sequencing. Our test model that we have been using in a series of acute calves is the Utah 100 type artificial heart. Actually this is an 80-ml stroke volume device. Today I would like to address how we achieve balance in this coupled system without external ports and without a volume displacement or a compliance chamber. We do that by making a small atrial septal defect here between the right atrial cuff and the left atrial cuff. This, then, is a small-diameter, and we will talk about that orifice that allows a volume of blood to move back and forth. Our surgical procedure is the following. First we expose the natural heart, do our cannulation, and establish the cardiopulmonary bypass. Then we do the cardiectomy. We make sure of hemostasis then. We now are measuring through our heart-lung machine the systemic flow of the calf. But we are also able to measure the amount of bronchial flow return into the left atrium. We use this measurement in making a decision on which of the predesigned and prebuilt intraatrial septal orifices to use in these atrial cuffs. Then we go ahead and surgically implant the device and wean the calf from the cardiopulmonary bypass on to the electric powered artificial heart. The purposes of the in vivo acute experiments were the following: to establish the intraatrial shunt performance, verifying that we can indeed achieve ventricular balance and that we maintain patency of that atrial shunt; to look at the hematological effects, hemolysis, and thrombus formation within the device and all of its components and attachments; and to examine the electrohydraulic TAH 20 system’s performance both in regard to the fit anatomically and in regard to the physiological performance as far as adequate cardiac output is concerned. In a series of 6 calves the measured bronchial blood flow as a percent of the flow of the cardiopulmonary bypass machine goes from a low of 5% to a high of 17%.We have learned that in those calves that exceed 12% we have a difficult time in establishing balance. The shunt size that we have selected to implant with each particular bronchial flow has been narrowed down to two sizes; we avoid putting the electrohydraulic artificial heart in a calf that has a high left-left shunt. During the course of the experiment we have measured the left atrial pressure and subtracted from that the right atrial pressure giving the change in pressure across this intraatrial shunt. We find that the highest we had was 6 to 10 mm Hg and the lowest we had was 2 to 5 mm Hg. In the longest living calf (about 9 days), the average difference between the mean left atrial pressure and the mean right atrial pressure was about 3.5 mm Hg, greater on the left than on the right. Occasionally we would have some spikes in pressure difference. Presently our motor speed control algorithm is being addressed to see if we can minimize these spikes, but they are relatively transient. Part of it I think might be a hammer effect from the method of open port measurements that we employ. On the bench we tried to look at some of the flows using contrast media. This shows quite readily that we do
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have that shunting. There is a higher shunt from left to right than the reverse. In the calf who lived 9 days the left atrium and the orifice of the small intraatrial shunt were quite clean. The conclusions from these 20 in vivo experiments were the following: The intraatrial shunt provides adequate balance throughout a variety of bronchial blood flows. The intraatrial septal shunt has shown no evidence of thrombosis or occlusion, albeit so far our longest experiment was 9 days in these acute studies. The size constraints of the animal do exist and can affect the cardiac output and device performance even in our miniaturized system. Furthermore, we have had no evidence of intraventricular thrombosis formation except in one experiment where we had an electrical device failure and stoppage of the device without adequate heparinization. These experiments have been a great learning experience, resulting in 18 design changes. The system meets the physiological demands with its automatic control algorithm. Now we have made a totally integrated device. This device is being built now for animal testing; it has an 80-mL stroke volume and will fit totally within the pericardial sac. Doing extensive fit trials we have successfully fit this clay model into a 57-kg and a 68-kg cadaver, although the 57-kg cadaver did have cardiomegaly. But it does fit well. We also have successfully fit the model with slight modification of the length and angulation of the inflow outflow ports into a 75-kg sheep and an 80-kg calf. This device will have an 80-mL effective stroke volume. Presently we are preparing for additional animal, cadaver, and human intraoperative fit trials.
MODERATOR HAHN: Thank you so much, Dr Don Olsen. Now I will ask Mr Ken Butler to give his lecture. MR BUTLER Indeed it is a pleasure for me to be here representing my colleagues at the Cleveland Clinic and my co-workers at Nimbus and present to you a run down of our particular artificial heart device that we have been working on for the last 4 years. From the anatomical view, central is the blood pump energy converter unit that fits in the chest, replacing the natural heart. Our system is a dual pusher plate and it does have to have its compliance space managed; we do this with a flexing diaphragm chamber that is connected via a duct to the interspace. Gas is transferred back and forth between the interspace of hemodiaphragms and the flexing chamber, which faces into the lung space. We also have refill port that is accessed percutaneously. The main power is derived from primary batteries that are located on the outside, and we also use a TET system including the secondary coil, which is implanted subcutaneously over the left chest. We are using the Thermedics or Thermocardiosystems version, and it has a secondary coil that fits on the outside over that. That transmits alternating current power across the skin, which is then conducted down to what we call a power module, which
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is actually a rectifier that converts the current into directcurrent power, which then drives the energy converter. The electronics package in our system is contained within the central blood pump and energy converter. An appendage is the internal battery, which is connected into the power module. The power module also serves as a connecting point for the various cables. So you can remove the internal battery and remove the secondary coil without disturbing the energy converter. The blood pump itself is in the "human configuration" as a result of extensive anatomical studies conducted at the clinic, where every dimension was critically evaluated to make sure it would fit in a maximum number of adult patients. The blood pump energy converter has been our main focus of work over the last several years. This is an entirely new design for us. Although we have had a considerable amount of experience with left ventricular assist systems using actuation and energy converter technology, this is really a new departure for us. The energy converter is contained within a titanium superstructure. The ventricles themselves are made of a carbonimpregnated epoxy, which has saved us a considerable amount of fabrication complexity, and they certainly are lighter than our titanium castings. There is a dual pusher plate. We simply rely on an alternate ejection from the left to the right. While the actuator is engaged in ejecting on the right side, the left side is free filling. The pusher plates, when they are filling, are disengaged from the sac or the rolling diaphragm. Conversely, on the right side when it is ejecting on the left it is filling on the right. So our control system essentially is to match our actuator cycling frequency with the fill rate of the blood pumps. If we go too slow we will impede filling. If we go too fast we will actually start an ejection while one of the ventricles is filling. So all we have to do is match our motor speed to the fill rate. And we do that electronically with a rather simple system, which I do not want to get into here. The energy converter is based on an electric motor continuously driving a gear pump that puts out a flow of hydraulic fluid. That fluid cycles a piston back and forth in a container, which provides the actuation. We have a mainspring in the device that gives up force when it is ejecting on the left side and is cocked when the right side is ejecting. So it evens out our power distribution. The diaphragms are the flexon rubber that the Clinic has been using for years. And we do use the biolyzed coatings that the Clinic has developed on all the blood contacting surfaces. This is the book on our system as far as the key characteristics. I think all the systems are coming to about these weights and volumes. Our capacity is a 60-mL blood pump. It can generate up to roughly 10 L/min, 60 beats per minute. The direct-current power consumption at its maximum is 15 W. Our internal battery time, based on our latest tests, is in excess of 60 minutes, and that is at a 6 Llmin condition; that is specified for the internal battery. We have been concentrating the last few months on in
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vivo tests, and I would like to focus the rest of the presentation on that, which gives you some idea of how these systems work and possibly some of the problems that we have with our animal models as has been alluded to before. We normally are using 80-kg calves. We are focusing just on the central blood pump and energy converter. We are using external electronics and a vent for compliance. Basically, what we are trying to show that we can support one of these animals and that the system is compatible with the physiology. And of course, most important, we are using it to test our system. We have done four experiments to date and our progression has gone from a planned I-day acute model to a couple of intermediates, and our last one was 4S days. On these early animals we are using all the blood pressure instrumentation we need: inlet and outlet pressures on both sides. These animals therefore are fairly risky in that there are indwelling catheters in them. One of the things we have learned over the years is that a calf is not a human, and it is very difficult to take a human configured system and try to force it into a calf. Therefore we have changed the inlets a little bit to help u s get better anatomical fit in the calves. The big difference is that the outlet on the left side is bent so it intercepts the aorta. I would like to give some data now from a 45-day experiment. When we plot out all the data points we took, they average around 10 mm Hg in the right atrium. On the left side it i s a little more scattered, but i t would average out about 16 mm Hg. One of the things about these animals, and i n particular onc, hi, mnxcd o u r pump out the first night. He was running a t 1 0 Limin t w t ~ n t i a l l v on day 1 and that is wherr he stayed. Although t h i 5 is good healthy workout for our actuator, it does not do much for our control svstem, which indiccittbh cj prohlc~ni we have in the model. We rn'ixed out a t 160 beats t o r thc whole experiment and oiir powt'r drab\. was 15 W throughout. In some of the e,irhcr isxpcrimcnts where the calves were not as hcalthv, w c ) were ablt. to go to the lower flow rates and verifv that indeed our rontrol svstrm does function like we intend it But w c n t d to h a w a lot more experience with that. To summarim the in vivo results, w e can maintain the hemodynamics of thcse animals Thr f i t is extremely good. We have not had anv ht~molysis ( i r thrombus problems. The diaphragnis come o u t absolutely cledn after the experiment. What we are going to be doing i s continuing the in vivo studies, our sequence of involving the other system integrLitionrlements. This is due to c ~ i i n e in a couple of months. We will add that into the energy converter, add the comp1i.inc-r ' I E T s i n with the rest OT the systems, and bv the cnd of this program we will be doing a total. I think 1 speak for all of us when 1 sav that these programs are supported by the devicrs and technology of the National Institutes of Health, and as I understand the plans we have another couple vears to go on the current contracts. There probably will be a follow on for 3 or 4 years, which puts us in a time frame such that wc probably would be ready to use human implants on a trial
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basis somewhere in the year 2OOO. That is the time line we’re working on. MODERATOR HAHN Thank you so much, Mr Butler, for this exhwnely nice presentation. Now we will ask Dr Kung to come here and give his report. DR KUNG: This is a collaborative effort between Abiomed and the Texas Heart Institute on the development of the electrohydraulic total artificial heart. Because I am last on the agenda I thought the best thing for me to do is to focus on the specihc and unique features of ow system instead of going into the description of the entire system. Fxternal parts of the system are battery electronics including a TET, internal battery, blood pump, energy converter, hybrid electronics, and sensors for the control of the system. I will describe how our energy converter works, how our blood pump interfaces with the atria, the design of the inflows for the pump, and how the compliance or rather the management of the left a d right plow differences is done. I also will give some information on our in vivo results. The system is electrohydrauk, and the difference between our electrohydraulic system and that of Nimbus, which I can describe as also eledrohydraulic but eh?ctrohydraulic converted to mechanical energy, is that we use a direct electrohydraulic drive of the flexing membrane, which is what the Utah group does also. The difference between our system and the Utah system is that we use a unidirectional rotary pump. In other words, the rotary pump is not reversed on every cycle. Fluid reversal is accomplished through an electromagnetic fluid valve. The blood contacting part of the system is designed to be smooth and seamless, and it uses angioflex, which is a polyurethane proprietary to us, and also trileaflet valves made of the Same material. We use blood as the volume compensation: I believe we are the only ones taking that approach. As you heard yesterday, if you have a biventridar support system-in this case this is also classified as biventricular-you have to make sure that the left flbw and the right flow somehow match or the difference matches the bronchial shunt. In current systems that have been used clinically vou have to manage this by essentially dialing in the flow so that you do n d flood the lungs. When you flood the lungs that means the right side is flowing too much. One of the manifestations of that is that the left atrial pressure goes up. If the left atrial pressure goes up you can think of a way of using that as a means of feeding back to your system. Because for a total artificial heart you obviously cannot keep on dialing the right side flow, you have to use what physiological information is provided by the artificial heart and the living system to control or feed back to adjust the right side flow. There is a hydraulic pump with a fluid switch that switches the fluid back and forth. To manage the left and right side flow, we have put a shunt flow in the hydraiilic circuit and placed that in reference to
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the left atrial pressure, which means that you can put a flexible diaphragm in contact with the left atrial reservoir. If the left atrial pressure goes up the shunt flow increases in the hydraulic circuit. Men the shunt is krger to the right side d the hydraulic chamber, it reduces the amount of blood that is being drawn into the right-sided W d pump. So as the left atrial pressure goes up there is less stroke volume on the right side, and as the left atrial pressure goes down there will be more right-sided stroke volume. One of the advantages of this is that the flow compensationchacteristics would not vary over time. In long-term studies we are trying to demonstrate some of these characteristics. I would like to talk a little bit about the control algorithm. With it we have taken a very simple approach to control the system. In the hydraulic circuit there are pressure transducers, one pressure transducer that measures the filling pressure during the fill cycte of either the left side or the right side. That information is used to control the hydraulic pump. If you have a high enddiastolic pressure on one beat it Senses that pressure and then it tells the pump to go to pump a little bit harder on the next beat, so you are trying to bring the end-diastolic pressure to a level that you have preset, whether it is zero or some mmber that is dose to zero filling pressure. The control algorithm is also responsive to afterload because the afterload changes affect the pumping rate or the filling rate. Once you change the f i g rate you change the end-diastolic pressure, which then is sensed again, changing the pumping speed of the rotary pump. The other control is focused on trying to make every stroke a full stroke. We have an optimal sensor in the system that monitors the diaphragm position. That information is then used to determine the beat rate of the system. If there is a lot of fluid or i€ the atrial pressures are high, the beat rate will increase to maintain the fullstroke and at the same time to try to bring the end-diastolicpressure down. Now let me go to the energy converter. There are two channels that connect the hydraulic chambers to the energy converter chamber, so the flow of hydraulic fluid can go either way on each beat. It reverses the flow. As I say, there is a unidirectional rotary pump and a centrifugal one pumping hydraulic fluid. There i s a switching valve concentric to the pump that assumes two different positions, which then opens a set of ports, which then directs the flow either from one pump to another or vice versa. It is actually a very simple system. You are switching the hydraulic chambers, which pumps the blood to the two alternate pumps. This is an alternating pumping system. Now let me describe briefly the Mood pump. The valve that is used is a polyurethane trileaflet valve. Concerning the experience that we have on these bbod-contacting components, the valves have gone about 400 million cycles in vitro. The pumps we have tested up to 100 million cycles, and in vivo about 150 days. The design of the whale system is to achieve a seamless Mood contacting surface. Now in the inflow region, where you have tissue and cuff in contact and s u h w s that hold the cuff and the tissue together, because we have
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discontinuities, it is much better to have a cuff that can promote growth of tissue. But to avoid ingrowth into the inflow of the device, we have made a rough surface on the cuff and a smooth surface in the device. The cuff will be sewn onto the atrial remnants and then the rest of the device will be snapped into the atrial cuff, so this is a quick connecting approach. By this approach there are no end-to-end seams. To get around this rough and smooth configuration, we actually have a smooth surface tube project into the atrial cuff. This provides a tissue barrier to stop the growth of tissue. It is helped by the fact that you have flow going in this way and you probably have some recirculation, which prevents tissue from climbing up the smooth step. The protrusion into the atrial cuff is about 3 mm in height, and in in vivo studies it does act as a barrier quite well. It essentially tries to keep away blockage problems. The system actually can operate at more than 8 Wmin, but at 8 Wmin it takes about 20 W. In our chronic studies, we will concentrate on system integration; we will include hybrids, TETs, and in vitro study as well as the liability evaluation in this program.
MODERATOR HAHN: Thank you for this very interesting lecture. There are plenty of problems we could discuss, but first I should like the audience to ask questions, because the time is extremely limited. DR D. GLENN PENNINGTON (St. Louis, MO): I would like to ask Dr Pierce about those animals that he showed us. What is the stage of development of those animals? Is that an integrated system? In other words, are those implanted electronics? And is that an implanted compliance chamber? How complete is the system we are seeing in those animals?
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simple enough so that things are not that expensive. Our estimate is that we can do it, once the wholt. tcchnulogy is developed, for about $25,000. Ndw. I d i i t suit’ thdt 50me of the people here would think otherwise, but we have done some evaluation of that.
MR BUTLER: In our particular case, L \ C puject a cost between $35,000 and $50,000. Actually, uur total heart device is cheaper than our left ventricular ’iAst system because it is simpler. MODERATOR HAHN: Doctor Olsen, do you have comment?
DR OLSEN: 1 think when we consider the additional constraints and costs of Food and Drug Administration approval, it will be the upper limit of Mr Butler’s estimate of $50,000. DR PIERCE 1 think it is important to note that there is nothing inherently expensive about these devices. The roller screws we use in our units cost, I think, $1,000.The motors are $400 or $500,and there is one fairly complex titanium machined part, the costs of which would not be high if you were going to make a bunch of them. But by the time you put in the animal experiments and all of the careful controls that have to be done before something like this can be sold, the price escalates. 1 would estimate $5O,OOO, something like that, and I do not think these total hearts are going to be more expensive than the lett ventricular assist devices. There is very little additional complexity to them. MODERATOR HAHN: Another question?
DR PIERCE: Well there is really a progression. As I mentioned, the animal that is now over 11 months does not have energy transferred by the coil system. We do have another animal that we have subsequently operated on that has implanted batteries, a closed system, and energy transfer by the primary and secondary coil. The only thing that we are lacking now is the ability to talk to our system and get information in and out of it. Other than that, the system is pretty much complete. MODERATOR HAHN: Another question? Doctor Champsaur? DR GERARD CHAMPSAUR (Lyon, France): What is the projected price of the total artificial heart using your current techniques?
DR KUNG: All of us probably will have an answer to that. Let me just start by saying that we have in the design of the system and its development always tried to make it
DR PAUL DIDISHEIM (kthesda, MD): I would like to ask each of the panelists about thromboembolism. Because we know from past experience that an implanted device can look free of thrombus or relatively free of thrombus and still be a generator of thrombuemboli, I would like to ask each of the paiielists if they have yet had an opportunity to look at end organs, specifically the kidney, which has in the past been a useful thromboembolus trap. MODERATOR HAHN: Doctor Kung, would you answer? DR KUNG: Yes, in our experience, we have not had any long-term experience with the total artificial heart yet. But the technology that we have used in the permanent ventricular assist device is really quite similar to what we have here. And there we have had experience up to 5 months with no infarcts or emboli.
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MODERATOR HAHN: Another comment from the panelists? MR BUTLER: We just completed a 45-day experiment with the total heart. That is our longest. But it does have the biolyzed coating that we have had considerable experience with on the left ventricular assist devices. This 45-day animal was run without anticoagulants, and I do not think all the histology data are in yet, but the gross examination showed no infarct problems. DR OLSEN: Our limited experience in acute experiments with the electrohydraulic system would not lend itself to respond there. We have seen none, however, in 9 days. MODERATOR HAHN: Doctor Pierce?
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DR PIERCE If we follow the design principles that we have all learned the hard way over the past 20 or so years and then we keep our animals on warfarin sodium and their renal function has been basically normal, the devices have been clean when they have been taken out. So thromboembolism has not been a problem in our case. We have had broken wires. We have had moisture get into the devices. We have had just about everything you can think of. But thromboemboli have not been a problem. MODERATOR HAHN: I guess I must stop this session. It is a great pity, because we had so many questions to ask about these electromechanical devices that will be the future in cardiac replacement.