|Publication number||US3421497 A|
|Publication date||Jan 14, 1969|
|Filing date||Feb 26, 1964|
|Priority date||Feb 26, 1964|
|Also published as||US3426743|
|Publication number||US 3421497 A, US 3421497A, US-A-3421497, US3421497 A, US3421497A|
|Inventors||Ball Richard H, Chesnut Merrill G|
|Original Assignee||United Aircraft Corp|
|Export Citation||BiBTeX, EndNote, RefMan|
|Patent Citations (3), Referenced by (14), Classifications (12)|
|External Links: USPTO, USPTO Assignment, Espacenet|
Jan. 14, 1969 M. G. CHESNUT ET AL 3,421,497
HEART PUMP SYSTEM Sheet 2 of1O Filed Feb. 26, 1964 Jan. 14, 1969 M. G. CHESNUT ET AL 3,421,497
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HEART PUMP SYSTEM Sheet Filed Feb. 26, 1964 TIME IN MILL! SECONDS M mm m m m m w w w m i M United States Patent 3,421,497 HEART PUMP SYSTEM Merrill G. Chesnut, Arvada, and Richard H. Ball, Boulder,
Colo., assignors, by mesne assignments, to United Aircraft Corporation, a corporation of Delaware Filed Feb. 26, 1964, Ser. No. 347,500
US. Cl. 128-1 24 Claims Int. Cl. A61m 1/00 ABSTRACT OF THE DISCLOSURE A reciprocating pump exchanges fluid with the aorta of a patient in timed relationship with natural heart action through a catheter, the reciprocating pump being driven by an actuator pump which in turn is controlled by servo valve means for delivering hydraulic fluid to the actuator pump; electronic control means timed to the EKG waveform provides control over the length and speed of pumping and withdrawal strokes, as well as the ratio between the pump and withdrawal stroke speed; the electronic control means also provides control over the relationship for the pumping time to the EKG waveform; the pump is driven in a feedback servoloop fashion for accurate control. Also included are pump cycle reducing means including means banking selected portions of the EKG waveform.
This invention relates generally to a device for assisting the natural action of a defective heart. More particularly, it concerns an apparatus for assisting insufiicient natural heart action by synthesizing the parameters of the patients physical heart waveform.
The present invention relates to an improvement over the system shown in the copending application of Watkins and Klink, Ser. No. 285,413, filed June 4, 1963. This copending application relates to a heart-assisting system in which a closed cannula is inserted into the descending aorta and expanded and contracted in timed relation to the patients natural heart beat as derived from the patients EKG waveform. In the copending application, a control system is provided for energizing an on and off type valve which controls the flow of compressed air to a reciprocating piston pump.
Heart pumps fall into two general categories; those which augment the natural heart action by operating in series with the natural heart, and those which either partially or totally bypass the heart. The present system is of the augmentation type and in contradistinction tothe bypass type system operates in series with the heart while the heart is actually beating and supplying blood to the arterial tree. The basic objectives of the present invention are to reduce the work load of the heart by lowering the pressure head against which the ventricle must eject its contained blood, and to aid the coronary circulation by increasing the blood flow during the correct phase of the natural heart cycle. This phase corresponds to the period of time when the resistance in the coronary circulation is at a minimum and has been found to be in, what is described as, the post-systolic period. Consequently, the present system is one in which the flow and pressure developments are very accurately phased in timed relationships to the normal beating action of the heart, in contradistinction to the parallel or bypass systems which have been used in the past for bypassing blood from the veins to the arteries when the heart may be arrested. Attempts to use such total bypass systems when the heart is beating have fallen into two groups; one in which the heart is pushing against the ventricle during the systolic phase, which has been shown to be detrimental to the heart itself because it increases the heart Work load, produces "ice abnormal electrocardiograms, and fails to augment the coronary circulation at a time when its resistance 1s highest; and one in which a total bypass system is used across a veno-arterial shunt, so that its input into the vascular system does not need to be timed correctly. However, the very precise timing that is required for such a maneuver must still be maintained and difficulty is encountered in sustaining such a system because drainage from the venous side of the circulation is less dependable because of the lower pressure existing in the veins.
It is therefore, a primary object of the present invention to provide a heart pump system particularly useful for assisting vascular circulation in which the pumping of blood on the systole and diastole phases is accurately synchronized with the natural heart cycle and in which the entire arterial blood pressure waveform may be synthesized as required by the particular patients requirements. More specifically, the present invention is directed to a pumping system which may be readily coupled to the patient suffering from cardiac insufficiency, together with the electronic circuitry and manual adjustments so that a surgeon can tailor the resultant arterial blood pressure waveform to suit the specific needs of the patient. The present heart pump system when used as an augmentation device forces and withdraws blood into and out of the aortic artery through an inserted catheter with the systolic and diastolic response rates and stroke volume synthesized to the body requirements.
The use of a total bypass system with cardioplegia, as has been used in the past, has the distinct disadvantage that when the system suddenly fails an emergency is created in that cardiac standstill must be overcome, or some other means of maintaining adequate circulation must be sought within a very short period of time; whereas, with the augmentation circulation system of the present invention, should the system suddenly fail, the heart action continues at least with as good a performance as it had before such assistance was rendered, or even better performance because under the stimulus of forced flow produced by the augmentation system the diameters of the collateral vessels may be enlarged and flow through the entire coronary system increased markedly thereby providing a more permanent pathway for the normal flow of blood after discontinuance of the augmentation system; and also the myocardium has gained or increased strength of contraction. I
Another object of the present invention is to provide a heart pump system which utilizes a hydraulic system to drive the main heart pump. In the systems which utilize a reciprocating pump for supplying blood to the circulatory system some means must be provided to drive the pump. In the past, pneumatic pump systems have been provided, but these systems lack the short-response time necessary if accurate and complete control over the blood pressure waveform is to be obtained. However, with the: high efficiency insured-response time of the present hydraulic system, wave shaping is easily and effectively accomplished. For a given volume of the pump, the resulting pressure waves which develop in the aorta of the human body can be shaped accurately by metering the flow of hydraulic fluid to the system to control the main reciprocating pump, which delivers and withdraws blood from the patients aorta. In the present device, the blood pressure waveform is shaped by varying the length of stroke of the reciprocating pump, which controls the amount of blood pumped per pumping cycle, and in varying the speed of pumping in both the systole and diastole phases. The use of a hydraulically-driven pump in the present heart system because of its inherent rapid response assures that each function will terminate accurately in accordance with the predetermined demand,
thereby reducing the likelihood of fibrin, blood trauma and resulting blood cavitation.
It is a further object of the present invention to provide a new and improved heart system in which the length of stroke of the reciprocating pump decreases as the patients natural heart beat increases. In the present heart system, a synthesized pump cycle must be completed between adjacent pulses of the selected portion of the patients EKG waveform as the pump cycle is initiated by the waveform. As the repetition rate of the patients EKG signal increases, and the length of stroke of the reciprocating pump remains constant, there may be insuflicient time for the pump to complete its cycle before the appearance of a subsequent EKG pulse. If the pump stroke rate or speed were increased at this time, to compensate for the decrease in the period of the natural heart action, the excessive pumping pressures may induce undesired cavitation. To avoid this and other disadvantages, the stroke length automatically decreases thereby permitting the pump to complete its cycle in accordance with an increased repetition rate in the patients EKG signal.
A further object of the present invention is to pro- Vide a heart augmentation system in which the patients EKG signal, arterial pressure wave or other phasic related parameters initiates an artificially synthesized heart pump cycle.
Other objects and advantages will become readily apparent from the following detailed description taken in connection with the accompanying drawings, in which:
FIGURE 1 is a schematic drawing of the complete intra-corporeal augmentation system;
FIGURE 2 is a side elevation, partly in cross-section, showing the hydraulic power supply components;
FIGURE 3 is an assembly of the heart pump and servo actuator assembly shown in cross-section;
FIGURE 4 is a schematic block diagram of the hydraulic servo-actuator assembly;
FIGURE 5 is a schematic block diagram showing the input control circuit;
FIGURE 6 is a detail of the input board of the input control circuit;
FIGURE 7 is a detail of the Gate/Relay board of the input control circuit;
FIGURE 8 is a detail of the Output Board of the input control circuit;
FIGURE 9 is a schematic block diagram of the pump control circuit;
FIGURES l0 and 11 are a detail of the electric circuitry in the pump control circuit;
FIGURE 12 is an exemplary graph showing the relationship between pump delivery volume and stroking time for the maximum permissible stroke speeds; and
FIGURE 13 is a graph showing an exemplary complete heart cycle and the cycle of the present augmentation system in relation thereto.
While an illustrative embodiment of the invention is shown .in the drawings and will be described in detail herein, the invention is susceptible of embodiment in many different forms and it should be understood that the present disclosure is to be considered as an exemplification of the principles of the invention and is not intended to limit the invention to the embodiment illustrated. The scope of the invention will be pointed out in the appended claims.
GENERAL ARRANGEMENT The treatment of ailing heart patients is normally required where the patients heart action is simply not sufficient to supply the bodily needs; this is usually attributed to a lack of sufiicient muscular activity within the heart itself. Frequently, the situation is encountered that while the diastolic action of the heart will bring a volume of blood into the left ventricle of the heart sufficient to supply bodily needs, the left ventricle will not fully empty into the aorta. Or should the ventricle fill the aorta with arterial blood, the systolic action of the heart is not thereafter sufiicient in itself to completely discharge the blood 1 content of the aorta into the arterial tree, resulting in blood back-up, stagnation, particularly coronary occlusion, and the serious impairment of bodily functions. In some cases heart failure may be so complete that for all practical purposes there is no significant natural heart action.
The present system, as shown schematically in FIG- URE 1, assists the defective heart as described above by withdrawing a portion of the blood from the human aorta during the diastolic and systolic phases of the natural heart action and by delivering the withdrawn blood back to the aorta after the aortic valve has closed naturally or before closure which induces early closure thereby assisting the distribution of blood to the arterial tree. At this time the pressure requirement of the heart is at a minimum because instead of calling on the heart to supply the arterial tree, against the back pressure thereobtaining, it is merely required to supply blood to the aorta. The pressure necessary to supply the arterial tree comes from the heart augmentation system when it re-delivers the withdrawn blood to the aorta.
Referring to FIGURE 1, an EKG source, generally designated by the numeral 10, provides the basic input signal to the intra-corporeal augmentation system and may be derived from: one of the previous EKG tracings of the patient properly amplified, an amplified signal taken directly from the patients heart beat, or a suitable signal generator, such as a multivibrator. The basic purpose of employing the patients EKG signal as the input to the augmentation system is to assure the proper synchronization of the artificial heart pump with the patients natural heart action. However, in some cases, the patients heart beat may be so weak, erratic, or nonexistent that it is desirable to use either one of the patients prior normal traces or an artificial signal source, as the basic input signal rather than taking the signal directly from the patient. The EKG signal source selected is impressed upon the input control circuit, generally designated by the numeral 11.
The input control circuit 11 consists of an input circuit, a gate and delay circuit, and an output circuit. The control system is premised on the use of a parameter consisting of impulses taken in a desired selected manner from the natural heart action of the patient. The input circuit of the input control circuit shapes the EKG waveform as originally received from the patient, and a filter contained therein which is of the standard R-C type, eliminates undesirable noise potentials present in the EKG waveform, particularly preventing feedback. An emitter follower and a differential amplifier in the input circuit provide a threshold adjustment so that the particular point on the EKG wave may be selected for triggering the subsequent circuitry. The Schmitt trigger is fired at a selected threshold level on the patients R wave, QRS complex, or other phasic parameters and produces a square wave which is differentiated to achieve a sharp pulse as a function of the EKG input. This signal may be viewed on an oscilloscope by the surgeon to assure a proper synchronization in the system. The same signal also is impressed on a gate delay circuit also in the input control circuit. The gate delay circuit serves two major functions, namely, blanking undesired signals and delaying the output signal of the input control circuit with respect to the patients EKG signal. In this manner, the gate circuit serves to determine the frequency with which the pump control system operates with respect to the natural heart beat; that is, whether this pumping system works every beat of the heart, every second beat of the heart, or every third beat as selected in the adjustable gating circuit. If the patients heart beat remains strong, it may be desirable to employ the augmentation system only on selected heart beats rather than on every one. Further, the gating circuit in the input control circuit delays the output signal with respect to the EKG signal because the patients EKG signal as derived by the usual methods actually precedes the natural physical action of the heart by a short time interval ranging from .04 second to about .10 second.
Square waveform pulses from the input control circuit are impressed upon the pump control circuit generally designated by the numeral 12. The pump control circuit 12 is the component which completely synthesizes the blood pressure waveform in accordance with adjustments made by the surgeon treating the ailing cardiac patient. The surgeon inserts each of two open-ended catheters 13 into one of the femoral arteries of the patient and gently urges the end of the catheters to a distance almost to the diaphragm of the aorta or other position in the artery, depending upon the condition of the patient, and the patients response to the technique. The reciprocating piston pump 14 withdraws and delivers blood to the aorta through the catheter 13. A hydraulic actuator 15 mechanically drives the pump 14 through pump shaft 16, and is under the control of hydraulic fluid delivered to the actuator cylinder by a servo valve 17 which ports fluid to the actuator 15 from a suitable hydraulic fluid source 18. The servo valve 17 is driven by a torque motor 19 as dictated by the pump control circuit 12.
As noted above, the pump control circuit 12 completely synthesizes the blood pressure waveform by controlling the length of stroke of the pump 14, the rest period between the pump and withdrawal strokes, if any, and the speed of pumping on the variable pump and withdrawal strokes. All of these functions, or parameters, may be controlled by settings in the pump control circuit. For ex ample, if the control circuit is set for a greater pumping speed, an increased signal will be impressed upon one of the coils in the torque motor 19, thereby opening the servo valve 17 further permitting greater quantity of hydraulic fluid to flow to the actuator which results in the faster movement of the pump 14. If a change in the length of one of the pump strokes is set in the pump control circuit 12, the pump control circuit will de-energize the previously energized servo coil at the indicated time thereby cutting off flow of hydraulic fluid to the servo actuator 15. A linear variable potentiometer 20 feeds back a signal proportional to the actual position of the actuator 15 to the pump control circuit 12, so that the pump control circuit exercises its control over the actuator at the precise position thereof as dictated by the surgeons settings in the pump control circuit 12. Also, the control circuit integrates the repetition rate of the patients EKG waveform and decreases the length of stroke of the pump 14 as the rate increases. This serves to prevent cavitation when a short heart cycle is indicated by the input control circuit.
All of the components described above in addition to those described below in more detail may be contained within one console (not shown) with a control panel designed for use by a medical doctor or technician. During the systolic phase of the patients heart, the pump 14 withdraws a portion of the blood from the aorta in an amount and at a rate dictated by the surgeons settings on the pump control circuit 12, but within the pressure limits above which damage to the blood occurs through cavitation, fibrin, or trauma producing hemolysis. The pump or augmentation system thereby assists in drawing blood from the left ventricle into the descending aorta. After the heart valve closes by its natural action, or as prematurely induced, the pump control circuit 12 reverses the direction of the heart pump 14 thereby returning blood to the aorta and distributing it throughout the arterial tree. The pressure developed within the arterial tree may be controlled by adjusting the pump control circuit 12 to vary the quantity of flow of hydraulic fluid from the source 18 to the actuator 15. In this manner, the pressure of the blood within the arterial tree as well as the quantity of blood distributed may be accurately controlled by the surgeon in accordance with the specific needs of the particular cardiac or other patient. The pumping phase of the pump 14 corresponds to the period of time when the resistance in the coronary circulation is at a minimum. Consequently, the augmented system is one in which the flow and pressure developments are very accurately phased in time relationships to normal beating action of the heart, in contradistinction to parallel systems which have been used in the past for bypassing blood from the veins to the arteries when the heart is either arrested or has minimal work load.
In the event that an occlusion has occurred due to thrombosis in one of the main vessels of the coronary system, any flow that can be achieved through collateral channels will be aided by increasing the pressure head across these vessels and under the stimulus of increased flow, the diameter of the vessels may enlarge, thereby providing a more permanent pathway for the normal flow of blood even after the discontinuance of the augmentation system. For this reason, the pump control circuit 12 very precisely dictates the rate of flow of hydraulic fluid into the actuator cylinder thereby controlling the pressure waveform in the circulating system. All the components described above with respect to FIGURE 1 are described in more detail below.
HYDRAULIC POWER SUPPLY The hydraulic fluid source 18, as described above in FIGURE 1, is shown in more detail in FIGURE 2 and although the specific details of the hydraulic power supply form no part of the present invention, they will be described in order to give a more complete understanding of the over-all heart augmentation system. The basic purpose of the hydraulic fluid source 18 is to deliver hydraulic fluid, such as oil, to the servo valve 17 as demanded. In one specific embodiment of the present invention, it was found that .65 g.p.m. and with 1,000 to 1,500 psi were satisfactory for the augmentation system requirements. A gear pump 30 withdraws oil from the reservoir 31 through tube 32, compresses the oil, and delivers it to the outlet passage 33 which is connected to the servo valve 17. A coarse strainer 34 may be provided to enclose the lower end of the tube 32 to prevent the flow of any foreign matter into the gear pump. To prevent sudden surges of fluid from the gear pump 30, an accumulator 35 is provided downstream from the pump and serves to damp out pump pressure surges. Also, a relief valve 36 prevents excessive pressure developed by the gear pump from being impressed upon the outlet passage 33. A check valve 37 prevents the accumulator pressure from driving the hydraulic gear pump 30 in the reverse direction when inactive. Optionally, cooling coils 38 may be disposed in the lower portion of the reservoir 31 to cool the hydraulic fluid and insure uniform oil viscosity. Alternatively, a portion of the oil from passage 33 may be directed to a fan cooled radiator and thence returned to the reservoir.
SERVO ACTUATOR ASSEMBLY The servo actuator assembly consists of a reciprocating piston actuator 15, a hydraulic servo valve 17, and a torque motor 19, shown schematically in FIGURE 1 and in more detail in FIGURE 3. In the past, actuators for reciprocating heart pumps have usually been pneumatically operated, with the inherent lack of stability, accuracy, driving stiffness, and fast response. However, the servo valve in the present device accurately meters the flow of oil to the actuator in response to a predetermined signal impressed upon the servo valve and in this manner the hydraulic actuator assembly very accurately controls the accuracy and response-time of the reciprocating piston pump. The improved response-time and relative stability of the hydraulic actuator permits a wider range of system adjustment to meet the exact pumping characteristic desired for the particular patient. In pneumatic systems, since the compressible motive fluid has low response time and low power, the pressure differential across the catheter is very high thereby reducing the effectiveness of the pump. In a hydraulic system the pressure differential is very low.
The actuator assembly has two main functions: firstly, to control the position and speed of the reciprocating heart pump 14, and secondly, to feedback a signal to the pump control circuit (described in more detail below) proportional to the actual position of the reciprocating pump 14.
Referring now to FIGURE 3, the torque motor 19 consists of individual coils and 41 which are connected to a pump servo driver and a withdrawal servo driver, respectively, in the pump control circuit. A flapper valve 42 is connected to the torque motor 19 and extends through a nozzle 43 communicating with the hydraulic pressure input passage 33 leading from the hydraulic pressure source 18. The torque motor 19 and the flapper valve 42 constitute the first stage of the servo valve 17; while a spool valve 44 and a feedback spring 45 connected between the spool valve 44 and the flapper valve 42 constitute the second stage of the servo valve 17. The high pressure passage 33 from the hydraulic source communicates with both ends of the spool valve and the flapper valve nozzle 43 through pressure passages 46 and 47. A low pressure passage 48 communicating with the center of the spool valve and the nozzle 43 is connected to the reservoir of the hydraulic fluid source. A pump passage 49 delivers oil from the spool valve 44 to the left-hand side of an actuator piston 50, while a withdrawal passage 51 connects the spool valve 44 to the right-hand side of the pump piston which is slida'bly mounted in an actuator cylinder bore 55. A pump shaft rigidly connects the pump 14 and the actuator piston 50 so that movement of the piston 50 under the control of the servo valve will directly position the pump 14. A bracket 61 is rigidly mounted on the end of the pump shaft 60 and carries a movable potentiometer rod 62 which slides within the coil 63 of the position feed-back potentiometer 20.
Viewing FIGURES 3 and 4 together, the operation of the servo actuator assembly is as follows: when the pump control circuit impresses a signal on one of the coils 40 or 41 from one of the pump or withdrawal servo drivers, the torque motor 19 positions the flapper valve 42 in response to the magnitude of the signal to modulate pressure at the end of the spool valve 44 by restricting the nozzle 43. A differential pressure between pressure passages 46 or 47 moves the spool valve thereby connecting one of the pump inlet passage or withdrawal inlet passage 49 and 51 to the source of high pressure fluid passage 33, and the other inlet passage to the low pressure passage 48. In this manner, one surface of the actuator piston 50 is subjected to high pressure from the hydraulic source while the other surface of the piston and its associated actuator cylinder chamber is vented to the tank through the low pressure passage 48. It should be noted that the torque motor 19 is responsive to different magnitudes of the input signal from the pump control circuit to position the spool valve 44, and thereby the amount or rate of flow of high pressure fluid to the actuator is determined in accordance with the magnitude of the current impressed upon one of the coils 40 or 41. In this manner the speed of the pump shaft 60 and the pump 14 may be accurately controlled by the pump control circuit. Leaf spring 45, between the servo valve second stage spool and the flapper, feeds back the spool position to obtain accurate positioning and stable operation.
Alternatively, the position feedback potentiometer 20 maybe a linear variable differential transformer, which will increase the life of the feedback component. In this case, additional circuitry would be required such as an exciting oscillator, rectifier, and an integrator.
MAIN PUMP ASSEMBLY The reciprocating pump 14 delivers blood to the aorta during the post systolic phase of the patients natural heart action, and withdraws blood from the aorta up into the chamber of the pump during the systolic phase of the natural heart action when the heart valve is open thereby assisting the delivery of blood from the left ventricle into the aorta. This is accomplished by connecting open ended catheters to the open end of the pump 14 and inserting them into the femoral artery. In this case, the open pump 14, filled with a Dextran or other blood compatible fluid, actually draws a portion of the blood from the aorta; however, it should be noted that in contradistinction to the bypass type system only a small portion of the total circulating blood in the patient diffuses and displaces the Dextran in the open end of the main pump 14 because of the great degree of control over the pressure waveform in the circulatory system by the present augmentation system. Alternatively, a closed ended cannula may be connected to the pump 14 for insertion into the femoral arteries. In this case, a blood compatible intermediate fluid is pumped into the cannula within the aorta rather than the patients blood itself.
Referring to FIGURE 3, the main pum 14 consists of a casing which supports the actuator assembly and has an opening therein for receiving the pump shaft 61 which is driven by the actuator assembly in the manner described above. A piston 71 is rigidly fastened on the end of the pump shaft 60 and slides within a chamber 72 formed in the pump casing 70 forming a working chamber 73 and a venting chamber 74. Two rolling diap-hragms 75' and 76 are in contiguous contact with both ends of the pump piston 71 and insure reliability in the pump. An annular chamber 77 between the bellowfrarns 75 and 76 communicates with a passage 78 formed in the pum casing 70. A cylinder 79 receives an evacuating iston 89 which is threaded to a plate 81 on the end of the cylinder 79. A spring 82 urges the evacuating piston downward. Note the thumb screw in FIGURE 3 is only partially threaded. Prior to filling chamber 77, spring 82 is compressed by engaging the thumb screw 83 threads to expel air from cylinder 72. The chamber 77 is then filled with Dextran through cap 84. The thumb screw 83 is then backed out to disengage its threads permitting the spring to urge the piston downward and thus main tain a low grade vacuum in chamber 77. The low grade vacuum ap lied by the piston 80 insures proper diaphragm rolling action. The catheter (as shown in FIG- URE 1) which is inserted within the femoral artery, is connected to the pump 14 by a suitable fitting by the screw threads 89 on the pump casing 70 at one end of the working chamber 73.
INPUT CONTROL CIRCUITRY As noted with respect to FIGURE 1, the input control circuit generally designated by the numeral 11 accepts signals from an electrode preamplifier monitoring the physiological parameters of the patients heart action and provides synchronizing signals to an OscillOsCOpe and to the heart pump control circuitry described below. When the monitored signal reaches the manually selected trigger level, three distinct and simultaneous actions occur within the control panel. First, the oscilloscope sweep is triggered and the display appears on an oscilloscope. Second, the zero to four second interference blanking gate is started; and third, a zero to two second time delay is started. At the end of the zero to two second delay, a pulse is generated to trigger the self-cycling pump control circuitry. The input control circuit will not, however, provide another pump pulse until the zero to four second gate terminates and the input signal again exceeds the trigger level. In this manner interference may be blocked from the input waveform pulses by proper adjustment of the zero to four second gate and the heart pump can be triggered at the desired time by correct adjustment of the zero to two second delay. The amplitude of the pulse for delivery to the self-cyling pump control circuit can be varied or adjusted from zero to 5 volts by a suitable control. Each time a pulse is generated to trigger the pump control circuitry, a visual indication in the form of a flashing light is obtained. A second flashing light may be employed to indicate the duration of the pump phase. The heart pump pulse may be temporarily halted by a control on a suitable panel. Likewise, the pump phase may be halted or extended f r long periods by another independent control.
The input control circuit can be operated from one of three parameter monitors; a conventional EKG, in manual fashion or from an internal rate generator. The proper triggering of the input control circuit may be obtained on either the positive or negative slopes of the input signal and over a wide range of input signal levels. The EKG waveform is made up of a number of traces represent ing various phases of a heart action including the P, Q, R, S and T waves. In practice, the input control circuit signal is impressed on an oscilloscope to permit visualization of the waveform to provide the most satisfactory noise-free electrical signal. Usually only a small portion of the R wave is selected for input control circuit triggering purposes.
Referring now to FIGURE 5, the input circuit of the input control circuit is generally designated by the numeral 80 and serves to shape the EKG waveform as originally received from the patient. A filter 81, which is the standard R-C type, eliminates the undesirable noise potentials present in the EKG waveform, particularly preventing feedback into the electrode amplifiers. An emitter follower 82 and a differential amplifier 83 provide the threshold adjustment so that any particular point on the QRS complex can be selected for triggering the subsequent pump control circuitry. A manual pacemaker 84 within the input control circuit is provided to assure an input signal to the over-all heart augmentation system when the usual source of EKG waveform input does not exist. The output from the differential amplifier 83 trips a standard Schmitt trigger 84 at the selected threshold level. The threshold level may be adjusted by bias adjustment 85. The square wave Schmitt output pulse is differentiated by diiferentiator 86 and amplified by an emitter follower 89 to achieve a sharp pulse as a function of the EKG waveform input. Two triggering outputs are derived from the dual emitter follower amplifier 89, namely, for the oscilloscope sweep and for gate triggering in the following circuitry.
The gate delay circuit 90 of the input control circuit provides two major functions, namely, blanking certain pulses of the input waveform and also delaying the input pulse with respect to the EKG waveform input. This delay is provided so that the triggering pulse to the pump control circuitry which initiates the withdrawal phase or stroke of the pump will occur at a predetermined time after the EKG input pulse and also a predetermined time after the intraventricular and intra-aortic pressure rise. The reason for this is described more fully below in the description of the cycle, but briefly it assures that the main pump 14 will reciprocate, withdrawing and delivering fluid to the patients aorta in such a relation to the patients natural heart action to relieve a substantial portion of the strain on the patients heart. The output pulse from the input circuit 80 is first clipped and filtered after which it keys a blanking flip-flop 91. Following actuation of the flip-flop, a capacitor (described below) hooked to the emitter of the unijunction transistor 92 begins its charging function. The charging rate is adjustable by a potentiometer as described below with respect to the detailed circuitry. When the unijunction reaches its firing point a pulse is created resetting the flip-flop. During this charging period, all input pulses remain blanked or gated for 1 6 milliseconds to three seconds are preset by the potentiometer. Therefore, no spurious EKG signals or other stray signals can affect the systems operation.
The second portion of the Gate/Delay circuit 90 is designed similarly to the blanking stage, and consists of a delay flip-flop 94 connected to an unijunction oscillator 95. When the first flip-flop is keyed ON, the steeply rising waveform is capacitively coupled to key the flipflop 94. The unijunction oscillator 95 with its associated capacitor begins its charging, this time to delay an output pulse with respect to receipt of the keying EKG input pulse from the input circuit 80 so that the synthetic pressure waveform in the aorta controlled by the pump control circuit can be accordingly delayed. This delay ranges from 12 milliseconds to one second behind the initial rise of the blanking pulse. When the delay pulse inijunction fires, the flip-flop 94 is reset and a pulse is provided to the output circuit 100.
The output circuit of the input control circuit standardizes the delay pulse from the gate circuit 90 by the operation of a one shot multivibrator 101 which has a constant two millisecond reset time. The two millisecond pulse is transmitted to a dual emitter follower amplifier 102 to trigger the output. The output voltage provided is a pulse of approximately 1 to 5 volts amplitude and 2 milliseconds width. A proportional pulse is taken from the dual emitter follower 102 for the purpose of actuating and momentarily lighting an indicator lamp 103. Also, the output of the one shot multivibrator 101 is transmitted to an emitter follower 105 for feeding the triggering pulse to a suitable oscilloscope. Further, the signal at the balnking flip-flop 91 may be impressed upon the emitter follower 105 so that both signals may be viewed on an oscilloscope to inform the surgeon or technician of the proper cycle time delay and sequence.
For a fuller understanding of the detailed circuitry necessary in the input control circuit, the following is a detailed description taken with FIGURES 6 to 9 of the drawings:
INPUT Referring now to FIGURE 6, an input selector switch provides a means of choosing the desired mode of triggering the control panel signals. Only level two of this switch is used, and its common contact connects the selected signal to R1 of the input board. R1 and C1 form a decoupling network. T1, R2, and R3 form an emitter follower to provide a high input impedance to the control panel. T2 T3, R4, R5, R6, R7 and R8 form a differential amplifier. The operation of this circuit can best be understood by visualizing the trigger polarity switch, S2, in the positive direction and the trigger level control, R7, also in the positive direction. With no signal present, the emitter of R1 and the base of T3 will be at ground potential. However, the base of T2 will be biased to some positive potential and its emitter will also tend to rise to this level. The positive rise of the emitter of T2 will tend to place T3 in a nonconducting state since its base is at ground potential. However, when a positive signal in excess of the voltage set by R7 is obtained from the input channel, T3 will suddenly conduct and amplify the incoming signal. The sharp drop at the collector of T3 will cause the Schmitt trigger, consisting of T4 and T5, to operate. The sharp drop in the voltage at the collector of T5 is differentiated by C5 and R15. T6 serves as an emitter follower output to trigger the scope and drive T7.
The amplifier consisting of T7 and related components inverts the negative spike from the emitter of T6 and provides a trigger for the Gate/Delay circuits.
The manual circuit consists of R17, R18, C7 and S3. In the normal position of S3, C7 is allowed to charge to approximately 5 volts. When the manual button is depressed, the charge on C7 is placed on the input transistor causing a sharp rise in voltage and thereby initiating a trigger. In use of the manual circuit, the surgeon depresses the button to initiate each desired pump cycle.
GATE/ DELAY Referring now to FIGURE 7, a sharp positive spike from T7 is coupled into the base of T8 by C14 and D2. Any negative signals appearing here are bypassed by ground by D1. The transistors T8 and T9 form a flip-flop with their associated components. The positive signal at the base of T8 turns it on and T9 off. As soon as T9 is off, its collector rises to approximately 20 volts and C17 starts charging through R34, R36 and R27. When C17 reaches approximately +10 volts, the unijunction transistor, T10, triggers and discharges C17 through R32 producing a positive pulse at the junction of R31 and R32. This pulse reverts the flip-flop to its original state with T9 conducting. The time required for C17 to reach the trigger level of the unjunction is determined by the position of R36 and is adjustable from approximately 16 milliseconds to 4 seconds. R35 trims the circuit to provide exactly 4 seconds, and R34 serves to limit the charging current when R36 is at minimum value.
The circuit consisting of T11, T12, and T13 is identical in operation to the gate circuit except it is triggered by the positive rise of T 9s collector. By proper selection of R38, R39, R49, and C21 it is adjusted for a minimum time delay of 2 seconds and a minimum time of approximately 12 milliseconds. R40 is a logarithmic potentiometer to allow maximum resolution near the shorter delays. The same positive trigger that rests the flip-flop is utilized to trigger the heart pump one-shot on the output board.
OUTPUT BOARD Referring now to FIGURE 8, two transistors, T14 and T15, form a one-shot multivibrator with a pulse width of 2 milliseconds. The output of this one-shot is coupled to an emitter follower, T17, through S4, the pulse control switch. T18 serves as another emitter follower with a 50- ohm output impedance to match a coaxial cable connecting the input control circuit to a pump control circuit. If S4- is in the off position, a trigger will not appear at the pump trigger output and the pump control circuit will not be activated. The potentiometer, R72, provides an adjustable output for the pump trigger from to volts. The unijunction circuit consisting of T19 and related parts is biased so that C32 is charged to a voltage just below the firing potential of T19. When a heart pump pulse appears at the emitter of T17, it is differentiated by C33 and the positive spike developed at the emitter of T17, it is differentiated by C33 and the positive spike developed at the emitter of T19 is suflicient to trigger T19 into heavy conduction. The capacitor, C32, is then discharged through the panel lamp, L1, causing a visible flash. When the voltage across C32 has dropped to a low level, T19 ceases to conduct and C32 is recharged to the voltage determined by the voltage divider consisting of R76 and R77. The display output to the gating amplifier is obtained by utilizing an or circuit consisting of R60, R61, R62, D5, D6 and D7. The gate signal from the Gate/Delay board is normally at a positive level of approximately +18 volts, and the potential at the base of T16 will be approximately +5 volts. When the input trigger starts the gate, the level at R63 drops to near ground and T16 is biased off by virtue of the positive bias on its emitter. At the end of the delay time, the heart pump one shot is triggered and the collector of-T rises causing a positive bias on D5 and T16 producing a pulse on the output of T16. At the end of the pulse, T16, is again biased off. Normally the pump control circuits will activate with the pulse from T15.
PUMP CONTROL CIRCUITRY Referring now to FIGURE 9, wherein the pump control circuit is shown in schematic block diagram form, each square wave pulse generated at the output of the input control circuit described above is impressed upon the pump control circuit for the purpose of initiating a withdrawal stroke of the reciprocating pump 14. The general purpose of the pump control circuitry is to completely synthesize the pumping cycle of the pump 14. The only purpose of the square wave pulse from the input control circuitry is to initiate the pumping cycle at the beginning of the withdrawal stroke of the pump, i.e., movement of the pump piston to the left as shown in FIG- URE 3, at a predetermined time period after the patients EKG R wave as selected in the input control circuit. But once the withdrawal cycle is initiated, all parameters of the pumping cycle are controlled by the pump control circuit. More specifically, the volume of blood pumped by the pump 14 during each cycle, the speed of the pump during each pumping cycle, the relative speed of the pump during its pumping stroke and Withdrawal stroke, are all controlled by the pump control circuit. By adjusting the length of stroke of the pumping cycle, the amount of blood pumped to the aorta during the patients natural heart cycle may be selected by the surgeon as dictated by the requirements of the particular patient. The speed of the pump during both the pumping and withdrawal strokes thereof may be varied as desired by the surgeon to 1) reduce the intraventricle pressure within the patients heart thereby relieving the load on the heart, and (2) increase the intra-aortic pressure during a post-systolic period of the natural heart to thereby produce effective distribution of the blood through the circulatory system. After a greater portion of the systole phase of the patients natural heart action has completed and the aortic valve closes, the pumping stroke of the pump 14 begins, as delayed by the input control and forces blood into the aorta raising the aortic pressure an amount determined by the speed and length of stroke of the main pump piston 71. It is obviously very important not only to control the length of stroke of the pump but also the speed of the piston 71 during pumping as this determines the revised pressure waveform in the aorta caused by the present augmentation system. There are, of course, maximum speeds which limit the speed of the pumping cycle in both withdrawal and pumping phases, these are controlled by the pump cavitation point and breakdown, fibrin, and trauma with hemolysis of the blood. Also, a rate integrating amplifier in the pump control circuit integrates the repetition rate of the EKG signal received from the input control circuit and decreases the length of stroke of the pumping cycle as the EKG signal received from the input control circuit and decreases the length of stroke of the pumping cycle as the EKG signal rate increases. This modifies the time for the over-all pumping cycle in accordance with time changes between the input pulses of the EKG waveform. Further, the regenerative switch in the pump control circuit assures that only one of the pump driver or withdrawal driver, described above with respect to the servo actuator assembly, will be energized during any period of time thereby assuming that only one of the coils 40 or 41 of the torque motor 19, FIGURE 3, will be energized at any time.
Referring to FIGURE 9, there are two basic inputs to the pump control system: firstly, the linear variable potentiometer 20 generates a DC signal proportional to the position of the pump, see FIGURE 3, which is transmitted to the pump control circuit and, secondly, the input control circuit signal, which is a square wave impressed upon a pulse inverter in the pump control circuit. The basic output of the system is current signal delivered either from the pump driver 111 or the with drawal driver 112 to one of the pump or withdrawal coils on the servo torque motor 19, see FIGURE'3. The slow build-up of power in the DC power supply line 113 is prevented from actuating or triggering the pump cycle by the three second timer 114 during the initial cycle. An actuate switch 115 is the main switch for initiating the pumping cycle and selectively connects both pump driver and withdrawal driver to their respective coils in the torque servo motor 19. When in the off position, the switch 115 connects a biasing component 116 to the pump coil 40. In this manner when the augmentation system is not cycling, the pump coil will be energized permitting fluid to enter the left-hand side of the actuator piston 50 and move the pump 60 to the right until the bracket 61 engages the adjustable stop 64. In this manner, the cycle of operation of the 'heart augmentation system always begins with the pump piston at the end of the pumping stroke, i.e., the beginning of the withdrawal stroke.
The input delivered to the pulse inverter 110 triggers the one shot multivibrator 120, thereafter the signal is inverted by inverter 121. The inverted signal is transmitted to both a switch 122 and an integrating amplifier 123 for purposes that will appear hereinafter. The signal emitted from switch 122 initiates each cycle of operation of the pump 14 by impressing the signal on the AND gate 127. The AND gate 123 changes the state of an unlock switch 124 which turns off a regenerative switch 125. Regenerative switch 125 controls flow of current to the pump driver 111 and the withdrawal driver 112 and assures that only one of them will be energized at any period. When the unlock switch 124 turns off the regenerative switch 125, line 126 is energized thereby exciting the withdrawal coil 41 of the servo valve assembly through the withdrawal driver 112. In this fashion, the pumping cycle is initiated and movement of the pump piston 71 begins from right to left as viewed in FIGURE 3. As the piston moves from right to left, the potentiometer is driven toward ground until the stroke length amplifier 130 forwardly biases the regenerative switch, i.e., turns the regenerative switch on. The length of withdrawal stroke may be controlled by the stroke adjustment 131 on the stroke length amplifier 130. When the regenerative switch 125 is on, line 132 is energized and at the same time line 126 is de-energized, thereby turning off the withdrawal driver turning on the pump driver 111. The main pump piston 71 thereupon reverses its direction of travel and proceeds from left to right in the pumping stroke. During the movement of the pump piston from left to right, the potentiometer 20 is driven upwardly and produces an increasing feedback signal to the stroke length amplifier 130. The stroke length amplifier 130 will then turn on an end of stroke sensor 132 at a time determined by the end of stroke adjustment 133. The end of stroke sensor is calibrated, with an internal potentiometer, to insure that the mechanical pump stop 64 coincides with that position of the linear of the potentiometer 211 required to accurately indicate the end of stroke. When the end of stroke sensor is turned on, it delivers a signal to the AND gate 127. However, this will not initiate a new cycle of operation of the pump, but will merely preset the AND gate until the receipt of another of the patients EKG signal pulses from the switch 122. In this fashion, the patients EKG signal always initiates a new cycle of operation, but undesired and spurious signals cannot initiate a new cycle until the end of stroke sensor 132 indicates to the AND gate 127 that the previous pumping cycle is completed.
As noted above, the integrating amplifier 123 receives EKG signals from the inverter 121. The integrating amplifier 123 serves to emit a signal in line 140 proportional to the repetition rate of the EKG waveform pulses, and this signal is impressed upon the stroke amplifier 130 thereby biasing the amplifier and changing the level at which the amplifier will turn on the regenerative switch at the end of the withdrawal stroke of the pump '14. As the repetition rate of the EKG signal in the inverter 121 increases, the integrating amplifier 123 modulates the stroke length amplifier 130 to decrease the level at which the stroke length amplifier turns on the regenerative switch 125 to reverse the pump and begin the pumping stroke. In this manner, the time of the pumping cycle and the length of stroke is reduced so that it will complete its cycle prior to the receipt of the next EKG signal which initiates a new cycle. This also prevents blood cavitation which would occur if the pump speed were increased to compensate for an increased repetition rate of the patients EKG signal.
The time required to traverse one pumping cycle is manually adjustable in the feedback bias amplifier 141 by the speed adjustment 142. The amplifier controls the gain on the respective pump driver 111 and withdrawal driver 112 thereby controlling the current level in the pump and withdrawal coils 40 and 41. It is apparent that as the current in the torque motor coils increases, the torque motor 19 will increase the displacement of the spool valve 44 thereby permitting a greater flow of hydraulic fluid from the hydraulic fluid source 18 into the actuator piston cylinder 55. In this manner, the speed of the piston 71 may be varied as desired depending upon the needs of the particular patient. As noted above, if the speed of the pump is increased, this will increase the intra-aortic pressure during the pump stroke thereby reducing the intraventricular pressure during the systolic phase of the patients natural heart action. This is described in more detail below with respect to the operation of the system. Also, the relative bias on the pump and withdrawal drivers may be varied by a ratio adjustment 143. In this manner, the gain of one driver may be higher than the gain of the other, which results in the pump piston 17 traveling faster on one stroke (one of the withdrawal or pump strokes) than on the other during each cycle. This is extremely important because it is usually desirable for the withdrawal stroke time to be longer than the pumping stroke time to prevent pump cavitation and blood destruction. Within broad limits, each patients waveform will indicate to the surgeon the necessary relative speeds of the pumping and withdrawal strokes to minimize the workload on the patients heart while at the same time effectively distributing blood to the patients circulatory sysem.
While not disclosed in detail in this application, the various adjustable parameters in the pump control circuit may be programmed in accordance with medically estimated surgical requirements. If a program is provided, only vernier adjustments would be necessary for the individual patient.
While the operation and construction of the pump control circuit is believed obvious from the description above and as shown in FIGURE 9, a more detailed description of the specific circuitry follows for a more thorough understanding thereof.
Referring now to FIGURE 10, the powers supply is made up of a conventional transformer and full wave rectifier bridge 151. The bridge output is filtered by a conventional LC network 152 supplying 28 volts at line 113. Following closure of switch 153 and build-up of the 28 volt DC at line 113, the three second timer 114 is initiated by the charging function of capacitor of 154 through resistor 155 turning on transistor 156 and transistor 157. Transistor 157 supplies current to the torque motor through line 158.
Referring to FIGURE 11, when the switch 115 is in the inactive position, power is supplied to the torque motor 19 through line 158 limited by the biasing resistor 116, driving the pump 71 to its right-hand position.
When the actuate switch 115 is in its closed or upward position, connection is made to the pump servo driver 111 and the withdrawal servo driver 112 through lines 159 and 160, respectively. Input to the servo driver bases will be described in more detail below.
The 1 to 5 volt two millisecond pulse from the input control circuitry forwardly biases the inverter transistor 110 turning it on. The inverted decreasing pulse amplitude in line 160 fires the one shot multivibrator 120, thereby rapidly decreasing the forward bias on inverter 121 through line .161 and 162 due to the collector potential rise transmitted to line 161, divided down by resistors 164 and 166. Two outputs are taken from the collector 163 of the inverter 121; namely, one to the base of the integrating amplifier 123 through line 165, and the other to the base of switch 122 through the capacitor .167. The collector of switch 122 is connected to the junction of the AND gate 127 through line 169. Since the piston has been driven in abutment with the pump stop, and the end of stroke sensor 132 has supplied in an enabling signal from the collector of the end of stroke sensor 132 in line 170 through resistors 176 and 177, the unlock switch 124 becomes forward biased by the combination of the enabling pulse and rising collector potential of switch 122 turning the transistor 172 off thereby causing its collector voltage to rise back biasing diode 173 which permits current to be drawn through diode 174 of line 126 turning withdrawal driver 112 on due to the positive base emitter voltage. The rising collector potential on transistor 172 forces the base voltage of transistor 173 toward supply potential turning transistor 173 off thereby depriving line 132 of current which turns off the pump servo driver 111 to the lack of forward bias. Potentiometer 143 sets the withdrawal/pump ratio by inversely changing the emitter reference voltages of 111 and 112 as referenced to the voltages supplied through lines 132 and 126.
As the potentiometer is driven toward ground by the actuator piston 51 as the withdrawal stroke proceeds, the base potential of the stroke length amplifier 130 accordingly is decreased forward biasing stroke length amplifier 130 until the current level, dictated by stroke length adjustment 131, is suificient to forward bias the transistor 172 in the regenerative switch 125. In this manner the regenerative switch is turned on, drawing current through diode 173 thereby depriving the base of withdrawal servo driver 112 of sufficient drive. The forward bias on transistor 175 caused by the drop in 172 collector voltage supplies sufficient collector current through line 132 to forward bias pump servo driver 111, thereby turning the withdrawal servo driver 112 oif and the pump servo driver 111 on. As the potentiometer 20 is being driven in the pump mode, toward plus 28 v. DC, the length of stroke amplifier 130 forward bias is being decreased, thereby raising the emitter voltage of end of stroke sensor 132. This voltage is compared with the voltage set by the end of stroke adjustment 133 and is of such polarity to turn end of stroke sensor 132 on. The end of stroke adjustment 132 is set so that the end of stroke sensor 132 again provides an enabling pulse to the AND gate junction 127 at precisely the point in time which the bracket 61 strikes the adjustable stop 64. A pulse from switch 122 cannot turn unlock switch 124 on .until re-enabled as described above, thereby achieving prevention of double-beating.
The circuit is now in a position to receive another pulse from the switch 122 through line 169 at the time dictated by the input control circuit, and in this manner a new pumping cycle is initiated as described above.
The repetition rate of the signal from the input control circuitry at line 165 is integrated by the RC filter 190 forward biasing the integrating amplifier 123. The collector current of integrating amplifier 123 drawn through line 140 changes the available voltage at junction 182 which is available for bias through diode 171 to the emitter of the stroke length amplifier 130. In this manner the gain of the amplifier will vary in accordance with the frequency of the input signal from the input control circuit.
MAXIMUM CYCLE SPEEDS Since the present system is hydraulically operated, it is possible to very accurately control the speed of the pumping cycle during both the pumping and withdrawal strokes thereof, and as described above, this has the effect of changing the pressure waveform within the patients aorta. However, there are limitations upon the speed of withdrawal and the speed of the pumping, as determined by the size of the catheter, the size of the piston and other variable factors such as pump cavitation and hemolysis of the blood. On the withdrawal phase, there is obviously a limit to the maximum amount of pressure which can be exerted in an endeavor to induce blood to flow from the body into the front chamber 73 of the pump 14; such maximum pressure cannot exceed a total vacuum. Using such pressure, there is a limiting minimum time for a given volume of blood to be transferred from the aorta to the pump, thereby producing a lowering of the aortic blood pressure at the appropriate portion of the cycle. The efliciency of the pumping system must then be sufliciently great to deliver that volume back into the aorta during the balance of the cardiac cycle. For a given catheter size, there will be certain limits of time for withdrawal of blood and certain limits of time for pushing of blood into the aorta and these limits can be varied over .a very narrow range. If the withdrawal time al lotted by the electronic pump control circuit is too short, blood will be cavitated and destroyed, whereas if the withdrawal time is too long, the corresponding balance of the cycle which has been allotted for the pump phase will be so short as to produce excessive destruction of the blood by driving it through the catheter at an excessive velocity.
Referring now to the graph of FIGURE 12, where the volume of blood pumped is plotted against the time of pumping cycle for three different sized catheters as measured on the standard French scale, the base of each of the triangles formed on the graph represents the total pump cycle time, which may be equal to or less than the total natural heart action cycle. In thiseXemplary case, the pump cycle is 750 milliseconds. The height of each of the triangles formed represents the total volume of blood being transferred from the aorta to the pump and from the pump back into the aorta; the positive gradients represent the withdrawal time, and the negative gradients represent the time for the pumping stroke. The tangent of the angle indicated as divided by the bore of the piston, i e., tangent /11-R gives the limiting linear maximum speed to avoid cavitation. It should be noted that greater time must be allowed for the Withdrawal stroke than for the pumping stroke. While it is essential that the speed of the pump be below these maximum limits set forth in the graph of FIGURE 12, the necessary pump cycle requires that there not be significant deviation below the speeds set forth in order to effectively assist the defective heart or other patient and remain in synchronism therewith. The angle shown in the graph has been determined through medical research.
CYCLE AND OPERATION The heart augmentation system of the present invention is basically an intracorporeal system in that the pumping catheter is directly inserted into the patients descending aorta. In this manner it becomes possible, with the complete control over the pressure and flow of blood in the aorta by the pump control system, to precisely modify and reshape the intra-aortic pressure in a defective heart patient.
In prior systems which either partially Or totally bypass the heart, it was not possible to accurately associate the pumping cycle with the varying physiological parameters within the defective heart.
Referring to FIGURE 13, the top waveform represents the standard EKG signal derived from the patient showing the P, Q, R, S and T time relationship on the waveform. The abscissa of the figure represents the cardiac cycle time. As can be noted on the EKG waveform, the trigger level which initiates the electrical delay in the pumping system is selected on the leading edge of the R wave. It can be noted that the actual pump trigger signal designated PT is delayed from twelve milliseconds to one second behind the trigger point. The total pump delay time is designated D in FIGURE 13. The second waveform from the top, designated AP, depicts the central aortic pressure. The dotted line immediately below the central aortic pressure represents the diastolic pressure and valve opening pressure, and is designated NDP. The third waveform from the top represents the intraventricular pressure, and is designated IVP. The fourth waveform is representative of the servo valve driving current, and is designated DSC. The sawtooth waveform shown at the bottom consisting of W representing withdrawal and P representing pumping, represents the maximum limitations of volume, time and phasing to prevent cavitation of the blood in a manner pursuant to the graph shown in FIGURE 12.
From time zero, in the example of FIGURE 13, the first event noted along the time axis is the EKG trigger level taken from the ascending phase of the QR wave which is selected as the level to give one trigger per cycle. It may be taken from other points in the electrocardiac complex depending upon the waveform which is being obtained at the time. The intraventricular pressure commences to rise in the ventricle chambers until such time as the aortic valve is open at P; the central aortic pressure then follows and can rise with the ventricular pressure to reach the peak pressure shown as the first peak on both the central aortic and intraventricular. Normally, the length of the isometric phase (ICP), is .05 second; normally, the rapid ejection phase, designated EP, lasts about 0.10 second. From the peak of the pressure wave, which corresponds usually in the neighborhood of the T wave of the EKG, both the intraventricular and central aorta pressure fall until the aorta valve closes at VCI. The pressure in the central portion of the aorta is now at diastolic levels while the intraventricular pressure falls toward zero as at DP.
The operation of the device will be discussed with reference to FIGURE 13 so that the precise physiological changes within the circulatory system may be viewed with reference to the various control functions of the present intracorporeal augmentation system. The surgeon operating upon an ailing cardiac patient inserts the open ended catheters 13 into the patients femoral arteries and gently urges them to a distance almost to the diaphragm of the aorta or other position in the artery, depending upon the condition of the patient, and the patients response to the technique. The triggering level is then selected on the patients EKG wave or other related parameters, as shown in FIGURE 13, from part of the QRS segment of the wave and the proper adjustment is made on the bias adjustment 85 in the input control circuit. The triggering level may be taken from other points on the electrocardiac complex, depending upon the waveform which is being obtained at the time. The delay period, designated as D in FIGURE 13, which may vary from 12 milliseconds to one second, is set on the capacitor of the unijunction oscillator 95, and is so chosen that the beginning of the pump stroke will occur just prior to the aortic heart valve closure at the peak of the systole phase and proceed during a post-systolic period of the natural heart action. It should be noted that the exact timing of FIGURE 3 is merely exemplary, as the waveform of the particular patient and its response to augmentation may vary the delay period from that shown in FIGURE 13. The systole phase SY is merely approximated in FIGURE 13. The gain on the feedback bias amplifier 141 is adjusted so that the pumping cycle designated by the letters PC in FIGURE 13 equals the natural heart action cycle, for a maximum delivery of blood into the aorta as determined by the graph of FIGURE 12. In this regard, the stroke adjustment 31 must be coordinated with the gain adjustment on the pump driver 111 and withdrawal driver 112, so that the speed and volume of the pump are within the maximum limits set forth by the graph in FIGURE 12. After the EKG amplifiers have been connected to the filter 81 in the input control circuit, the surgeon or technician will close the actuate switch 115. It will be remembered that the heart pump at this time is in its right-hand most position with the bracket 61 engaging the adjustable stop 64. The withdrawal stroke then begins, designated at W in FIGURE 13, during the diastolic action of the natural heart a predetermined time interval subsequent to the aortic valve closure noted at VC or more specifically a predetermined time D after the EKG trigger, in FIGURE 13. The effect of the withdrawing of blood from the aorta during this portion of the natural action of the heart is to reduce the diastolic pressure within the aorta, as shown at PA in FIGURE 13. During the first portion of the withdrawal cycle, the heart is drawing blood into the left ventricle at a low pressure level as shown at DP on the intraventricle pressure waveform. Because of the reduced back pressure from the aorta on the heart valve caused by the withdrawal stroke, the isometric contraction phase NIC on the right-hand waveform in FIGURE 13 occurs at a slower rate thereby somewhat relieving the heart workload. For this reason, the aortic valve, by its natural action, will in the second waveform open (P at a lower pressure level within the left ventricle and aorta. As the systolic action of the heart proceeds, there will be a pressure rise in the ventricle and the aorta, but it will be much less than without the augmentation system in use. As the aortic and ventricular pressure approach the peak of the systole phase, the linear variable potentiometer turns the regenerative switch on and reverses the pump piston 71 thereby beginning the pump stroke of the system. After a very short lapse (a few milliseconds) the central aorta pressure is seen to rise again, after the point of valve closure. In this instance, the valve is artificially closed earlier in time due to the onset of the pump stroke of the piston 71. In other cases it may be desirable to begin the pump stroke after thenatural closure of the aortic valve. Referring to FIGURE 13, it should be noted that the intraventricular and aorta pressure are significantly lower than they were without the use of the augmentation system. Now, during this post-systolic period, the pump piston 71 forces blood through catheters 13 into the patients aorta thereby assisting the distribution of blood throughout the circulatory system. The synthesized aortic pressure caused by the pump 14 during its pumping stroke is shown at SAP in FIGURE 13, and it should be noted that while the aortic pressure is considerably above that which occurred during the unassisted natural action of the defective heart, the pressure waveform within the left ventricle, designated at SIVP, has been considerably reduced in comparison to the first waveform. As the pumping stroke nears completion, the linear variable potentiometer 20 initiates the end of stroke sensor 132 and the AND gate 123. However, a new cycle is not initiated at this time because the AND gate will not turn off the regenerative switch until 'a new EKG trigger signal is received from the input control circuit.
It should be noted that after a very short lapse after the pressure peak on the second wave, the central aorta pressure is seen to rise again after the point of valve closure. And this is because the valve is artificially closed earlier in time due to the on-set of the pumping stroke phase of the pump. It should be particularly noted that the workloads on the left ventricle from the first curve shown in the right-hand portion of FIGURE 13 before the heart augmentation system has commenced, and the right hand curves where the heart augmentation system is in operation, have radically changed. The workload on the left ventricle is roughly proportionate to the area under the WP curve, and since the area under the second wave is much less than the area under the first wave, it is evident that the cardiac workload has been diminished considerably. The pulse waveform of the pump is seen to occur in the post-systolic position in relationship to the intraventricular wave.
In summary, it is seen that the intracorporeal augmentation system reduces the amount of work required by the heart as noted by the comparison of the two areas of A1 and A2. in FIGURE 13, i.e., the first and second waveforms. In addition, the pump provides a new central aortic pressure waveform during the pump stroke phase.
The area of the second Wave represents the amount of Work the heart must do While the pumping system is operating. Since the base of the triangle (W, P) is set by the period of the heart and the angle p is fixed by cavi tation considerations, then the height of the triangle is also limited. The volume displacement of the pump is indirectly governed by the height of the triangle, being of maximum value for a given period. The apex of the triangle shown must necessarily be to the right of the apex of an isosceles triangle, due to the lower pressure availability during the withdrawal phase. The amount of time then left over for the pump phase must be sufiicient to prevent a water hammer effect which destroys blood cells. Due to these considerations, there may be no rest or dwell time available. It is to be further noted that the resultant intraventricular and central aorta waveform is displaced in time from the analogous waveform that would appear in a nonaugmented patient. The length of time of phase displacement is a complex function of pulse rate. The ratio of the withdrawal stroke time to the pump stroke time is independent of the catheter size because there is ample power available in the hydraulic system to overcome catheter pressure drop.
Assuming that the patients EKG waveform, or more specifically the R portion of the Wave, has a decreasing period, the integrating amplifier 123 will decrease the stroke length thereby maintaining the proper relationship between the pump phase and the new physiological parameters within the patients heart.
Thus, it will be seen that we have provided a system and apparatus in which the various obects hereinbefore set forth are successfuly achieved. Our system provides an effective degree of blood handling in quantities necessary for definitive results, with minimal trauma to the blood produced hemolysis, denaturization of plasma proteins, and other accidents incident to hydrodynamic stress in flowing blood. Our intracorporeal augmentation system connected into the descending aorta contributes to lowering the intraventricular pressure, and consequently is useful in the treatment of congestive heart failure, heart muscle damage, attending coronary arterial thrombosis and the several degrees of ischemic shock. Also, the present pumping system may be used in the treatment of hemorrhagic shock, in the treatment of toxic shock, as a surgical prophylactic before, during and after surgery on poor risk patients or in prolonged operations, in the treatment of heart problems where blood supply to the heart is impaired or failure is imminent such as in acute coronary occlusion, for obtaining coronary angiograms in correlation with X-ray techniques by injecting dye into the root of the aorta, as a diagnostic instrument to determine the bulk modules of the aorta.
With increased experience, other uses for the present system will present themselves.
1. In a heart pump system the combination, comprising: a pump having a pumping and a withdrawal stroke for assisting the heart during both the systole and diastole phases thereof, means for connecting the pump into the circulatory system, means for receiving the patients EKG waveform and selecting a portion of the wave thereof, means for initiating the withdrawal stroke of said pump a predetermined time after the selected portion of the wave, means for terminating the withdrawal stroke and initiating the pumping stroke approximately at the peak of the systole phase, means for automatically decreasing the length of stroke of said pump as the repetition rate of the wave increases thereby preventing cavitation in the blood.
2. In a corporeal augmentation system for assisting circulation in an ailing patient; a pump having a pumping stroke and a withdrawal stroke for assisting the heart during both the systole and diastole phases thereof, means for connecting the pump into the circulatory system, means for receiving the patients EKG waveform and selecting a portion of the wave thereof, means for initiating the withdrawal stroke of said pump a predetermined time after the selected portion of the wave, means for terminating the withdrawal stroke and initiating the pumping stroke approximately at the peak of the systole phase, means for controlling the speed of stroking of said pump thereby controlling the pressure waveform within the circulatory system, and means for automatically decreasing the length of stroke of said pump as the repetition rate of the selected wave increases thereby preventing cavitation in the blood.
3. In a corporeal augmentation system for assisting circulation in an ailing patient; a reciprocating pump having a pumping stroke and a withdrawal stroke for assisting the natural heart action during both systole and diastole phases thereof, means for connecting the pump into the circulatory system, an actuator for driving said pump, a source of hydraulic fluid under pressure, servo valve means for controlling the flow of hydraulic fluid to said actuator, means for receiving the patients EKG Waveform, means initiating the withdrawal stroke of said pump a predetermined time after a selected portion of the patients EKG waveform, means terminating said withdrawal stroke and initiating the pumping stroke approximately at the peak of the systole phase, means for controlling the speed of pump stroking thereby varying the pressure waveform within the circulatory system, and means for automatically decreasing the length of stroke of said pump as the repetition rate of the patients EKG wave increases thereby preventing cavitation in the system.
4. In a corporeal augmentation system for assisting circulation of an ailing patient; a reciprocating pump having a pump and a withdrawal stroke for assisting the heart action both systole and diastole phases, a catheter connected to said pump and adapted to be inserted into the patients aorta, an actuator for said pump, a source of hydraulic fluid under pressure, servo valve means for controlling and varying the flow of hydraulic fluid to said actuator thereby controlling movement of the pump, means for initiating the withdrawal stroke of said pump a predetermined time after the natural systole phase, means for terminating said withdrawal stroke and initiating the pumping stroke approximately at the peak of the systole phase, and means to control the speed of the pumping stroke whereby a modified pressure waveform may be produced in the aorta.
5. In a corporeal augmentation system for assisting circulation in an ailing patient; a reciprocating pump having a pumping stroke and a withdrawal stroke for assisting the natural heart action during both systole and diastole phases thereof, means for connecting the pump into the circulatory system, an actuator for said pump, a source of hydraulic fluid under pressure, servo valve means for controlling the flow of fluid to said actuator thereby controlling movement of the pump, means for receiving the patients EKG waveform, means for initiating the withdrawal stroke of said pump a predetermined time after a selected portion of the patients EKG waveform, means for terminating said withdrawal stroke and initiating said pumping stroke approximately at the peak of the systole phase, and means to control the speed of the pumping stroke whereby an increased aortic pressure waveform occurs after the systole phase of the natural heart and is accurately shaped as required by the patients needs.
6. In a heart pump system the combination, comprising: a reciprocating pump adapted to simulate the natural heart cycle during each pumping and withdrawal stroke thereof, the pumping stroke and the withdrawal stroke thereby together constituting a pumping cycle, means for initiating said pumping cycle in timed relationship to the patients natural heart action, and means responsive to changes in the repetition rate of the patients heart action
|Cited Patent||Filing date||Publication date||Applicant||Title|
|US2977765 *||Dec 23, 1959||Apr 4, 1961||Honeywell Regulator Co||Servo motor control apparatus|
|US3099260 *||Feb 9, 1960||Jul 30, 1963||Davol Rubber Co||Heart pump apparatus|
|US3266487 *||Jun 4, 1963||Aug 16, 1966||Sundstrand Corp||Heart pump augmentation system and apparatus|
|Citing Patent||Filing date||Publication date||Applicant||Title|
|US3592183 *||May 27, 1969||Jul 13, 1971||Watkins David H||Heart assist method and apparatus|
|US3712290 *||Sep 30, 1971||Jan 23, 1973||Bleifeld W||Apparatus for intra-aortal balloon pulsation|
|US3966358 *||May 17, 1974||Jun 29, 1976||Medac Gesellschaft Fur Klinische Spezialpraparate Mbh||Pump assembly|
|US4231354 *||Jul 14, 1978||Nov 4, 1980||Howmedica, Incorporated||Pulsatile blood pumping apparatus and method|
|US4284073 *||May 10, 1979||Aug 18, 1981||Krause Horst E||Method and apparatus for pumping blood within a vessel|
|US4427470||Sep 1, 1981||Jan 24, 1984||University Of Utah||Vacuum molding technique for manufacturing a ventricular assist device|
|US4473423 *||Sep 16, 1983||Sep 25, 1984||University Of Utah||Artificial heart valve made by vacuum forming technique|
|US4493697 *||Aug 14, 1981||Jan 15, 1985||Krause Horst E||Method and apparatus for pumping blood within a vessel|
|US4838889 *||Jul 23, 1986||Jun 13, 1989||University Of Utah Research Foundation||Ventricular assist device and method of manufacture|
|US4865581 *||May 29, 1987||Sep 12, 1989||Retroperfusion Systems, Inc.||Retroperfusion control apparatus, system and method|
|US5011468 *||Apr 16, 1990||Apr 30, 1991||Retroperfusion Systems, Inc.||Retroperfusion and retroinfusion control apparatus, system and method|
|US6736789 *||Oct 21, 1998||May 18, 2004||Fresenius Medical Care Deutschland Gmbh||Method and device for extracorporeal blood treatment with a means for continuous monitoring of the extracorporeal blood treatment|
|US8380281 *||Feb 19, 2013||The Johns Hopkins University||Compression device for enhancing normal/abnormal tissue contrast in MRI including devices and methods related thereto|
|US20110270079 *||Nov 3, 2011||The Johns Hopkins University||Compression device for enhancing normal/abnormal tissue contrast in mri including devices and methods related thereto|
|U.S. Classification||600/17, 604/67, 604/66, 604/6.11, 604/152|
|International Classification||F16C29/00, A61M1/10, F16C29/12|
|Cooperative Classification||A61M1/1086, A61M1/1081, A61M1/1037|