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Publication numberUS3513845 A
Publication typeGrant
Publication dateMay 26, 1970
Filing dateSep 15, 1966
Priority dateSep 15, 1966
Publication numberUS 3513845 A, US 3513845A, US-A-3513845, US3513845 A, US3513845A
InventorsCallaghan Phillip B, Chesnut Merrill G, Gafvert Sven
Original AssigneeUnited Aircraft Corp
Export CitationBiBTeX, EndNote, RefMan
External Links: USPTO, USPTO Assignment, Espacenet
Bypass heart pump and oxygenator system
US 3513845 A
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Description  (OCR text may contain errors)

BYPASS HEART PUMP AND OXYGENATOR SYSTEM,

Filed Sept. 15. 1966 y 26, .3 M. G. CHESNUT ET AL GSheets-Sheet 1 SIM/75' L EVEE MANUAL PUMP OXYGEA/HTOR l VENOUS Esra/e VENOUS PssEEvo/E May 26, 1970 G. CHESNUT ETAL 3,513,845

BYPASS HEART PUMP AND OXYGENATOR SYSTEM Filed Sept. 15, 1966 e Sheets-Sheet 2 r v F .Lfl

1 1 w HP 4 4Q N i T T :B 47' A FV LV l L c I V r 15! J0 10' v I c P 2 P "7 92'- ,54 Q"- v Br F jfi'l! 1 12 w W N 1 L L May 26, 1970 M. CHESNUT ETAL BYPASS HEART PUMP AND OXYGENATOR SYSTEM Filed Sept. 15, 1966 6 Sheets-Sheet 5 May 26, 1970 M. s. CHES NUT ET AL 3,5 4

BYPASS HEART PUMP AND OXYGENATOR SYSTEM Filed Sept. 15, 1966 6 Sheets-Sheet;

6 Shets-Sheet 6 May 26, 1970 M. G. CHESNUT ET AL BYPASS HEART PUMP AND OXYGENATOR SYSTEM Filed Sept. 15, 1966 United States Patent 3,513,845 BYPASS HEART PUMP AND OXYGENATOR SYSTEM Merrill G. Chesnut, Arvatla, Phillip B. Callaghan, Westminster, and Sven Gafvert, Boulder, Colo., assignors, by

mesne assignments, to United Aircraft Corporation, a

corporation of Delaware Filed Sept. 15, 1966, Ser. No. 579,565 Int. Cl. A61m 1/03 U.S. Cl. 128214 5 Claims ABSTRACT OF THE DISCLOSURE A cardiopulmonary bypass includes a reciprocating pump to deliver blood from a reservoir fed by an oxygenator to the femoral arteries. The oxygenator receives blood from either a venous roller pump or a reciprocable pump which draws the blood from a venous reservoir that in turn is fed by catheters placed into the superior and inferior vena cavae. The pumps are controlled by computer circuitry which provides the rate, the volume and the speed of blood flow under various controls. The system includes controls to permit utilization of an arterio-arterial assist device rather than as a complete cardiopulmonary bypass device. This includes the elimination of a check valve by utilizing difierential flow crosssections. In one embodiment, a second roller pump delivers blood from a reservoir of the oxygenator to a bubble trap, the output of which is returned to the patient through a reciprocating pump.

This invention relates generally to heart pump systems for replacing the function of a patients heart, and more particularly to a heart pumping system for complete cardiopulmonary bypass.

In complete or total cardiopulmonary bypass, the total blood flow from the venous side of the patients circulatory system is withdrawn, oxygenated, and returned to the arterial side of the system under suflicient pressure to perfuse the peripheral arterial tree. In systems of this type, in contradistinction to partial veno-arterial bypass, bypass of the right heart, and bypass of the left heart, the entire oxygenating function is effected by an extra corporeal oxygenator rather than the patients lungs.

A primary limitation in prior known cardiopulmonary bypass systems is that they have not had long-term capabilities due to the development of metabolic acidosis after extended perfusions. Metabolic acidosis may be defined as the acid products released in the circulatory system resulting from a lack of oxygen in the blood causing the incomplete combustion of carbohydrates, fats and proteins. The tissue hypoxia resulting from metabolic acidosis is most prevalent in the use of steady state heart pump systems, such as those employing roller pumps, but also appears to result from the use of known pulsatile systems.

The peripheral pooling of blood during long perfusions with resulting hypoxia has been relieved somewhat using presently known techniques of hemodilution. Further, operative respiration of the lungs with helium and drainage of the left ventricle during bypass have obviated some of the alveolar collapse and interstitial hemorrhage encountered with prior known systems and techniques. However, from the circulation standpoint the high rates of flow necessary in these prior systems have produced excessive damage to the blood and overloading of the reticluoendothelial system, plugging of the renal tubules and acute renal failure. In prior systems these high rates of flow have been necessary because low flow rates lead to the development of metabolic 3,513,845 Patented May 26, 1970 ice acidosis, often times during fairly short time intervals of perfusion. However, as the length of perfusion has increased, three has been found to develop a paradoxical metabolic acidosis even when the perfusion rate matches the normal cardiac output. Many authors explain this development as due to a sl-udging of the blood in the peripheral arterial tree.

Attempts have been made by various investigators to avoid the metabolic acidosis by lowering the oxygen requirement of the tissues by hypothermia. However, hypothermia produces its own problems, e.g., at ten degrees centigrade (despite the great reduction in oxygen requirement) the temperature of the blood limits circulation profoundly.

It is therefore a primary object of the present invention to provide a new and improved total cardiopulmonary bypass system in which microcircultaion is better perfused without employing flow rate rates high enough to rupture coronary vessels, thus enabling the system to maintain artificial circulation for longer periods of time and permitting more complicated surgical procedures to be attempted on the patient.

A further object of the present invention is to provide a new and improved pulsatile cardiopulmonary bypass system which in general will eliminate the patients tendency toward metabolic acidosis even over long-term perfusions.

A further object of the present invention is to provide a new and improved pulsatile cardiopulmonary bypass system of the type described in which the arterial flow rate may be controlled by the surgeon as desired by simultaneously varying the speed and stroke length of a pulsatile pump in the system.

A more specific object of the present invention is to provide a new and improved bypass system of the type described above in which triggering signals from a variable rate pacemaker initiate each cycle of the pulsatile pump, and an automatic computer circuitry maintains coincidence between each pumping cycle and the period of the triggering signals and also varies the flow rate by automatically varying the speed of stroking of the pump in response to changes in the selected volume by the surgeon.

Another object of the present invention is to provide a new and improved cardiopulmonary bypass system of the type described above in which a roller pump is provided for withdrawing blood from the venous side of the patients circulatory system and delivering it to an oxygenator, and including a reservoir for receiving blood from the oxygenator prior to entry into a blood pumping chamber in the pulsatile pump.

A still further object of the present invention is to provide an improved cardiopulmonary bypass system, somewhat modified from that described immediately above, in which two roller pumps are provided, one for pumping blood into the oxygenator and the other for withdrawing blood from the oxygenator and delivering it to the blood pumping chamber in the pulsatile pump.

Another object of the present invention is to provide a cardiopulmonary bypass system of the type described generally above, but in somewhat modified form in that a second pusatile pump is provided for withdrawing blood from the venous side of the system and delivering it through the oxygenator in place of the roller pumps, thereby more nearly simulating the physiological function of the human heart.

Still another object of the present invention is to provide a heart pumping system which may be used as a complete cardiopulmonary bypass during an operation and as a circulation assist postoperatively.

In accordance with the present invention, a reciprocating piston pulsatile pump is provided in which each pumping cycle is initiated by triggering signals from a pacemaker in the form of an oscillator in an associated con trol circuit. A pumping cycle as defined herein is one complete reciprocation of the pumping piston Within the pulsatile pump, i.e., a push and a withdraw stroke. The reciproeating piston defines an expanding and contracting fluid chamber within the pump connected to receive oxygenated blood and deliver it in pulsatile fashion to the arterial side of the patients circulatory system. Several means are disclosed for withdrawing the blood from the patients vena cavae, oxygenating the venous blood, and delivering it to the blood pumping chamber in the pulsatile pump.

While each pumping cycle is initiated by the pump triggering signal referred to above, a computerized control circuit is provided for varying the time base of each pumping cycle so that it coincides with the period of the triggering signal by normally maintaining the length of stroke selected by the surgeon and varying the rate of travel of the pumping piston to achieve this coincidence. However in complete cardiopulmonary bypass, as there is no essential relationship between pulsatile flow initiation and the patients natural heart action, the pacemaker may be adjusted by the surgeon to the desired rate. The present computer circuitry does permit the surgeon to select the desired rate and volume and the computer circuitry automatically initially computes the proper pumping or stroke speed to eliminate dwell or overlap between the pumping cycles.

The regulation of arterial flow rate, which is one of the important aspects of the present invention, is effected in the present computerized control by the rate and volume control circuitry. The volume circuitry permits the surgeon to merely select the desired volume of blood to be pumped per cycle and the system automatically computes the correct pump speed and stroke length to achieve the selected volume without varying the time base for the pumping cycle dictated by the repetition rate of the triggering signal described above. It is this interrelationship between pumping speed, stroke length and cycle time that provides the capability of the present device of longterm perfusions and the elimination of metabolic acidosis produced by prior known systems.

The present system when used in complete bypass fashion provides adequate blood pressure profiles with low blood flow rates. These low flow rates are sufiicient to avoid metabolic acidosis but are low enough not to rupture the coronary arteries.

Other objects and advantages will be readily apparent from the following detailed description taken in connection with the accompanying drawings in which:

FIG. 1 is a diagrammatic view of the present bypass system shown with its connections into the patients circulatory system;

FIG. 2 is a sub-assembly view of a portion of the pump shown in FIG. 1;

FIG. 3 is a simplified diagrammatic view of the system shown in FIG. 1;

FIG. 4 is a diagrammatic view of a system similar to that in FIG. 1 employing two roller pumps in addition to a pressure pulse generator;

FIG. 5 is a diagrammatic view of a somewhat modified system employing two reciprocating piston pumps;

FIG. 6 is a schematic diagram of a portion of the con trol circuit including the pacemaker input circuitry and the reciprocating pump servo coils;

FIG. 7 is a schematic diagram of another portion of the control circuitry including a digital counter;

FIG. 8 is a schematic diagram of the computer portion of the control circuit which selects the volume, push and withdraw rates, and cycle time of the reciprocating pump of FIG. 1;

FIG. 9 is a schematic diagram of timing error sensors which determine the error in the pump g cycle m FIG. 10 is a schematic diagram of a reduced augmentation circuit which selects only certain heartbeats to initiate the pumping cycles; and

FIG. 11 is a partial schematic diagram of the power supply for the control circuits.

While this invention is susceptible of embodiment in many difierent forms, there is shown in the drawings and will herein be described in detail embodiments of the in vention with the understanding 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 embodiments illustrated. The scope of the invention will be pointed out in the appended claims.

Referring to the system shown in FIG. 1 of the drawings, a pressure pulse generator 10 is provided generally similar to the pump shown in our copending application, Ser. No. 406,722 filed Oct. 27, 1964, assigned to the assignee of the present invention. Reference should be made to this application for a more detailed description of the pump 10. It will only briefly be described herein with particular reference to a combined on and off check valve 11 shown in FIG. 2 which permits both complete bypass operation and arterio-aiterial circulation assist.

The pump 10 includes a multi-member housing 13 having a cylinder formed therein which slidably receives a reciprocating piston 15 defining in the housing 13 a pumping chamber 16 and an actuating chamber 17.

A controller 20 described in more detail below delivers fluid in two directions through the fluid coupling 22 to the actuating chamber 17 thereby reciprocating the piston 15 in forward and reverse strokes, sometimes referred to herein as push and withdraw phases, respectively.

As shown in FIGS. 1 and 3, the pump 10 is connected in complete cardiopulmonary bypass so that blood flow from the heart and through the lungs is terminated and circulation and oxygenation is eifected extracorporeally by the artificial system. Toward this end venous blood is drained by catheter assembly 27 from the superior and inferior vena cavae which normally drain into the right atrim 25. This blood is drained by gravity into a venous collecting chamber 28. The chamber 28 should be placed below the level of the operating table for this gravity feed.

Venous blood in the reservoir 28 is pumped by a roller pump 30 through an oxygenator 32. The roller pump 30 may be any one of a number of commercially available units of this type. The oxygenator 32 may take several forms including the bubble type, disc, screen, or even membrane type.

Blood in the oxygenator drains by gravity into a reservoir 36. No bubble trap is required in this embodiment as any bubbles in the blood will rise to the top of the reservoir.

The reservoir 36 is connected to fitting 40 shown in FIG. 2 communicating through check valve 11 with the pumping chamber 16in the pump 10.

The valve 11 may be manually operated to either block flow from the reservoir 36 into the chamber 16, or to permit flow from the reservoir but prevent back flow from the chamber 16 into the reservoir. For this purpose a manually rotatable valve member 42 is provided with a large diameter through passage 43 therein. When valve member 42 is rotated so that passage 43 is aligned with bore 44 in the fitting, fluid may flow from the reservoir 36 into the chamber 16. This is the valve open position, and is so placed when the system is used in complete bypass. During postoperative circulation assist the valve 42 may be rotated degrees from its valve open position closing bore 44. The function of this is described in more detail below.

A check valve assembly 47 is mounted within a passage 43 and permits flow from the reservoir 36 into the chamber 16, but prevents flow from the chamber into the reservoir. In this manner during the push stroke of the piston 15 blood will flow from chamber 16 into the patients arterial system rather than back into the reservoir 36.

The pumping chamber 16 is connected into the circulatory system through a catheter assembly 49 connected to fitting 50 (FIG. 2). The catheter assembly may be similar to that shown in our copending application, which in cludes branched catheters adapted for insertion into the femoral arteries up into the patients descending aorta.

It should be noted that there are no check valves in the catheter assembly 49 preventing flow from the arterial side of the patients circulatory system into the pumping chamber 16 as the piston 15 withdraws. To prevent any significant amount of blood from being removed from the arterial tree during this withdrawal action, the diameters of the bore 44 and passage 43 are large compared with the smallest diameter of the catheter assembly 49 so that the catheters offer much greater resistance to flow during the withdrawal of piston 15 than the passages connected with the reservoir 36.

As the piston 15 moves in its push phase forcing blood into the arterial tree, the check valve 47 prevents reverse flow into the reservoir 36.

As the controller and actuator 20 delivers fluid through the coupling 22, piston 15 reciprocates drawing blood into chamber 16 through check valve 47 (with valve member 42 open), and in the opposite phase forcing blood from the chamber 16 through catheter assembly 49 into the arterial tree. This action provides pulsatile perfusion of the blood in the arterial tree.

In FIG. 4 a cardiopulmonary bypass system is shown generally similar to that shown in FIGS. 1 to 3 except that the reservoir 36 is replaced by a combination of a second roller pump 52 and a bubble trap 54. With this arrangement blood is forced into the pumping chamber 16' rather than being fed by gravity from a reservoir such as noted with respect to FIGS. 1 to 3. When piston 15' withdraws during the withdrawal phase, both the pressure from roller pump 52 and the Withdrawal action of the piston 15 cooperate in filling the chamber 16'. There is only minimal leakage through the valve 47' during the push phase of the pump 10' as the pressure in chamber 16' is above the output pressure from pump 52, although the valve 47 closes somewhat slower in this embodiment than in the embodiment of FIGS. 1 to 3.

Since the system shown in FIG. 4 is essentially closed from the oxygenator to the arterial side of the patients circulatory system, the bubble trap 54 is required to removed gases entrained in the blood.

The minute flow in the system shown in FIG. 4 is somewhat higher than that shown in FIG. 3 due to the additional filling assist of chamber 16' by the pump 52.

In the FIG. 5 embodiment, a cardiopulmonary bypass system is disclosed generally similar to those described above with reference to FIGS. 1 to 4 except that an additional pulsatile pump 56 is provided for Withdrawing blood through valve 58 from the reservoir 28". Valve 58 is a check valve and may be similar in construction and operation to the valve 11 shown in FIG. 2. Pump 56 during its withdraw phase draws fluid from reservoir 28", and during its push phase pumps blood into oxygenator 32.". The withdrawal of blood from the oxygenator 32" is effected by the withdrawal action of piston 15" in pump A bubble trap 54' is also necessary as the system is a closed one from the oxygenator 32". Suitable circuitry is provided (not shown) to balance the output of the pumps 56 and 10" so that they may operate under diflerent load situations preventing the oxygenator 32" from being either flooded with too much blood or being drained dry.

In all the above systems the intermittent infusion of blood into the arterial tree produces a pulsatile arterial pressure wave. The height or amplitude of this wave may be increased by increasing the travel of the piston and for this purpose a suitable volume dial is provided on the controller in FIG. 1 which permits the surgeon to select the desired volume per cycle of the pump. In addition, the relative time duration of the positive pressure wave in comparison to the total cycle time of the pump may be altered by changing the relative time duration of the push and withdraw phases of the piston 15. Furthermore, the length of cycle is also governed from the controller 20 by changing the rate of pacemaker which provides triggering signals for each cycle of the pump. During complete cardiopulmonary bypass the phase relationships between the patients natural heart action and the synthesized wave produced by the pump 10, important in assisted circulation, are of no importance.

The control circuitry described below, contained in the controller 20, exercises control over the pumping parameters of stroke length, speed of stroking during both push and withdraw phases, and pump cycle duration. It should be understood at this point that only portions of the described circuitry are operative during complete cardiopulmonary bpyass, the other portions being operative during arterio-arterial assist as a postoperative circulation aid.

Triggering and driving circuit The triggering and driving circuit disclosed in FIG. 6 is generally adapted to develop triggering signals for initiating each cycle of the reciprocating pump 10. For complete cardiopulmonary bypass, a pacemaker 132, which may take the form of an astable multivibrator, provides triggering pulses for the pumping cycles. A suitable control 132a is provided for varying the repetition rate of the triggering signals from the pacemaker so that the surgeon may select the cyclical rate he desires. For arterio-arterial assist, which is employed as a postoperative measure in the present device, an EKG trigger level selector is provided for initiating the pumping cycles in timed relationship with the patients EKG waveform. It should be understood that the pacemaker 132 and the trigger level selector 125 are used selectively.

In the arterio-arterial mode, the triggering circuit is effective to develop triggering pulses and delay them from a selected trigger level portion of the QRS segment of the EKG wave at a predetermined trigger level, and derives and delays a triggering pulse therefrom. The triggering pulse derived initiates the push phase of the pumping cycle and effects delivery of fluid into the patients aorta increasing intra-aorta pressure at a time when the workload of the heart is the lowest. As noted above, during postoperative arterio-arterial assist the valve 11 would be closed. The delay time for the triggering pulse determines the phase relationship of the pumping cycle to the arterial pulse wave, and is determined physiologically as the phase relationship, which reduces the intraventricular pressure to a minimum for a given volume pumped and increases the postsystolic arterial pressure to an extent which returns the intra-aortic pressure to its pre-pump level or better so that coronary and peripheral circulation may be assisted. The withdrawal phase begins immediately upon completion of the pumping phase and continues during aortic valve opening thereby aspirating the left ventricle into the aorta. The withdrawal phase continues until the peak of the next intraventricular waveform at which time another pumping or push phase begins in response to another triggering signal. This phase relationship is not of any significance in the complete cardiopulmonary bypass as there is no load on the patients heart.

As shown in FIG. 6, an EKG trigger level selector 125 is connected to receive the patients EKG waveform from a conventional electrocardiogram through line 126. Line 126 connected to line 127 is adapted to drive an oscilloscope so that the patients EKG waveform may be viewed during the use of the heart pumping system in the arterio-arterial mode by the surgeon or technician. Reference should be made to the copending application Chesnut et al. for the details of construction of the trigger level selector 125.

A selector switch 131, which may be located on a convenient control panel, permits the alternate initiation of 7 the pumping cycles by the pacemaker 132, by the EKG trigger level selector 125, or by the manual stroke initiate switch 133. As noted above, during complete cardiopulmonary bypass, with which the present invention is mainly concerned, the switch 131 is placed in its uppermost position connecting the pacemager 132 to a trigger delay circuit 136 through line 135. With switch 131 in its central or middle position, it connects the trigger level selector 125 to the trigger delay circuit 136 through line 135. And when switch 131 is in its lowermost position, the manual position, the upper half of switch 131 is ineffective but the lower half of the switch is connected to a power supply which places the pump cycle under manual control.

The trigger delay circuit 136 delays the actual triggering pulse behind the selected portion of the patients EKG waveform during arterio-arterial assist. It may c nsist of a standard dual PNP transistor one shot multivibrator with base triggering. Its RC time constant is manually adjustable with a potentiometer 138 which may be mounted on the control panel. The potentiometer 138 provides an adjustment of the delay time. The output triggering pulse from the trigger delay circuit 136 triggers a refractory gate 140 which produces an 80 msec. pulse when triggered. The refractory gate 140 may be a standard dual PNP transistor one shot multivibrator with base triggering. The delay provided by the delay circuit 136 is unnecessary during complete cardiopulmonary bypass.

The pulse from the refractory gate 140 comprises one of two inputs to AND gate 141. The AND gate 141 prevents any erroneous triggering signal from the refractory gate 140 from initiating the pumping cycle. If a triggering pulse from the refractory gate 140 is conducted to the AND gate 141 prior to the receipt of a signal from the other input to the AND gate 141, i.e., line 143, which indicates a partial completion of the withdrawal stroke, the 80 msec. pulse will be held by the AND gate for 80 msec. and if line 143 is not energized by that time, the triggering pulse will be dropped. If a triggering pulse from the refractory gate 141 and a signal from line 143 are received by the AND gate 141 simultaneously, blanking gate 144 passes a signal. The blanking gate 144 consists of a dual PNP transistor switch normally deactivated by the lock-out gate 450. A signal in line 145 to the blanking gate 144 blanks out selected triggering pulses in response to the reduced augmentation circuitry described in more detail below. A pulse derived from the blanking gate 144 is one of the two required enabling signals to the AND gate 146. The AND gate 146 will not trigger the pumping cycle, however, until a signal is received in line 151 which is the other input to the AND gate 146 indicating that the pump has completed its withdrawal stroke. This circuitry eliminates out-of-phase pumping during arterio-arterial assist by combining electronic control features with the physiological phenomenon known as the refractory period.

A pulse in line 150 from the AND gate 146 initiates the push phase of the pumping cycle by turning on a regenerative switch 148.

The regenerative switch 148 may consist of two transistor stages connected so that the output is fed back into the input. In addition, another NPN transistor switch shunts the first stage of the regenerative switch. A high voltage at the first stage turns on the regenerative switch, whereas the high voltage at the shunting NPN transistor turns it off regeneratively. The presence of a high voltage in line 150 turns the first stage of the regenerative switch on and the second stage off thereby energizing line 151 and deenergizing line 152. Line 151 energizes the pump servo coil 147 through the volume and rate computer circuitry shown in FIG. 8, while the energization of line 152 energizes the withdrawal servo coil 154 through the same volume and rate computer circuit.

A pump phase signal, modulated by the volume and rate computer circuitry described in more detail below with reference to FIG. 8, energizes line in FIG. 6 which drives the pump phase driver 161. The magnitude of the current in line 160 determines the bias on the pump driver 163 and thereby the magnitude of excitation of the pump servo coil 147. The rate of travel of the piston 15 is directly proportional to the magnitude of the current in the pump coil 147 and the withdraw coil 154.

The servo coils 147 and 154 drive a servo valve (not shown) which ports fiuid to a reciprocating hydraulic actuator. The hydraulic actuator drives fluid in two directions through coupling 22 shown in FIG. 1 which in turn reciprocates the piston 15. Such a system is shown in our above mentioned copending application. Alternatively, the coils 147 and 154 may be arranged to drive an electromagnetic actuator eliminating the need for a hy- (lraulic servo valve and the associated hydraulic circuitry. Connected to the actuator is a linear feedback potentiometer 118 for indicating the actual position of the pumping piston 15.

As the piston 17 begins its push stroke driving the potentiometer 118, a feedback displacement signal from the potentiometer drives an inverting amplifier 170, the output of which is used to determine the error between the actual pump and withdrawal displacement and the reference pump and withdrawal displacement signals determined by the computer circuitry shown in FIG. 8. Any error is fed back from the computer circuitry to line 172 in FIG. 6 whereby it is amplified by an AC amplifier 173 and integrated by a pump integrator 174 which excites a pump error driver 177 for varying the bias on the pump driver 163.

An identical closed loop feedback circuit is provided as shown for the withdrawal stroke and consists of line 189, AC amplifier 190, withdraw integrator 191, a withdraw error driver 192, a withdraw phase driver 188, and a withdraw driver 194. The error drivers are two-stage DC amplifiers which are collector coupled to the base of the NPN servo drivers for isolation purposes. The phase drivers 161 and 188 are PNP transistors which are collector coupled in parallel with the error drivers. The sevo drivers 163 and 194 may each be an NPN transistor which is connected in push-pull fashion to the associated servo coil. The servo drivers 163 and 194 are driven from an isolated source and are collector coupled to the servo valve coils 147 and 154 for isolation purposes.

As the feedback potentiometer 118 and the actuator shaft (connected with the actuator described above) reach the end of the push stroke and a suitable mechanical stop (not shown), an end of pump stroke sensor 186 provides a high voltage to the regenerative switch 148 to turn the switch off, thereby energizing line 152 and deenergizing line 151. The end of pump stroke sensor 186 may consist of an inverting PNP amplifier which drives the PNP comparator. The PNP amplifier emitter modulates the PNP comparator. The base of the comparator is biased by a potentiometer offset amplifier 187. The output of the end stroke sensor 186 immediately initiates the withdrawal stroke at the end of the pump stroke. The potentiometer and the offset amplifier 187 which provides a Vernier adjustment for the length of stroke may also be used to bias an end of withdrawal stroke sensor 200 to eliminate offset errors due to mechanical alignment tolerances between the piston 15 and the feedback potentiometer 118.

The energization of line 152 by the regenerative switch 148 enables the withdrawal phase driver 188. The withdrawal rate errors are impressed in the form of an error signal on line 189 and through the AC amplifier 190, the withdrawal integrator 191 and the withdraw error driver 192 providing a closed loop feedback on the withdrawal rate in the same manner as the corresponding components in the pump error driver circuit described above. A withdraw driver 194 energizes the withdraw servo coil 154 to drive the pump piston 15 in its withdraw stroke thereby drawing blood from the reservoir 36 through valve 11 into the fluid chamber 16.

As the pump proceeds in its withdrawal stroke, the feedback potentiometer 118 is driven as shown in FIG. 6. An isolated oscilloscope jack 197 is provided so that the inverted pump displacement wave may be viewed by the surgeon during heart augmentation.

As the feedback potentiometer 118 is driven toward ground, it drives the end of withdraw stroke sensor 200. This sensor, modulated by the volume comparator 201, indicates when the pump has reached the end of the desired withdraw stroke. The length of stroke is determined by a volume limit generator described in more detail below with respect to FIG. 8. Sufiice it to state at this point that a signal representative of the desired volume is carried in line 203. The end of withdraw stroke sensor 200 and the volume comparator 201 may consist of an inverting PNP amplifier which drives one-half of an NPN differential amplifier. The other side of this comparator is driven by the signal representing the desired volume in line 203 so that a suitable switch (not shown) is turned on whenever the displacement wave amplitude from the potentiometer 118 exceeds the volume amplitude in line 203. This switch drives an end of withdrawal oscilloscope output 205 as well as the pump trigger AND gate 146 and an end of withdraw one shot multivibrator 207 for purposes described in more detail below. Line 151 is then energized and awaits a new pump triggering signal from the blanking gate 144 to initiate another pumping cycle as described above. In this manner the volume of blood pumped by the reciprocating pump may be accurately controlled and selected by the surgeon as desired.

During the withdraw stroke, the feedback potentiometer 118 also drives a blanking threshold sensor 210 which provides one of two inputs to AND gate 141 preventing triggering pulses occurring earlier than a predetermined time before the completion of the withdraw stroke. This portion of the circuit is usually only operative during arterio-arterial assist when the EKG trigger level selector 125 is operative. The blanking threshold sensor 210 may be identical to the end of withdraw sensor 200 except that the input volume signal in line 203 is divided by a ratio of .9 in a volume comparator to 211. The output of the volume comparator to 211 which occurs a predetermined time before the end of the Withdrawal stroke enables AND gate 141. Upon receipt of the refractory gate output signal AND gate 141 enables in turn the AND gate 146 through the normally inactive blanking gate 144.

Rate and volume computer circuitry The purpose of the rate and volume computer circuitry shown in FIGS. 7, 8 and 9 is, generally speaking, to vary the pump and withdraw rates in response to changes in the triggering signal rate from the pacemaker 132 selected by the surgeon, and to vary the pump and withdraw rates in response to the desired volume per cycle selected by the surgeon. Thus, during complete cardiopulmonary bypass, the surgeon need only select the desired repetition rate for the pumping cycles by adjusting control 132a on the pacemaker and the computer circuitry will automatically vary the speed of the pumping piston so that the duration of each pumping cycle coincides with the period of the signal from pacemaker 132. Further, in response to the selection of a certain volume per cycle by the surgeon, the computer circuitry varies the length of stroke of piston 15 and automatically corrects the speed or rate of stroke of the piston so that the pump cycle duration remains constant with that determined by the period of the signal from pacemaker 132.

When used in arterio-arterial assist, the rate and volume computer circuitry is used to vary the pump and withdraw rates and the volume pumped in response to varying heart cycle times and also in response to dilferent desired volumes selected. Further, the output volume pumped decreases as a function of the heart rate as the maximum withdrawal and pumping rates are reached.

These maximum rates are determined by blood hemolysis and catheter size. The push and withdrawal rates for each cycle are determined from the heart rate computed on the previous cycle. To eliminate any stacking errors, a closed loop error detection circuit is provided at the end of the withdraw stroke.

The computer logic described briefly above is based upon the selected rate of pacemaker 132 (or the cardiac cycle time during postoperative assist) and the selected volume per stroke as variable input parameters. From these the pump and withdraw rates are computed and the pump controlled.

During arterio-aiterial assist it is sometimes more desirable to determine the cardiac cycle time from a pres sure Wave which is sensed from the patient. False triggering of the counting circuitry can be eliminated to a great extent by counting the time interval between the patients pressure waveforms rather than his EKG wave.

Referring to FIG. 7 wherein the digital counting circuit is shown in schematic form, a pacemaker input signal is received in line 266 from pacemaker 132. With switch 255 connecting line 266 to one shot multivibrator 271 the digital counting circuitry counts each period of the pacemaker pulses and provides an output voltage to the modulation circuitry shown in FIG. 8 which in turn computes the proper speed of the pumping piston 15. Thus, during complete cardiopulmonary bypass, the digital counting circuitry shown in FIG. 7 determines the period or rate of the pulses from pacemaker 132.

Alternatively, during arterio-arterial assist the digital counting circuitry in FIG. 7 may be employed to determine the patients cardiac cycle time as represented either by the patiens EKG wave or a pressure signal. Toward this end, if switch 131 is placed in its intermediate position connecting trigger level selector to line 266 the digital counter in FIG. 7 will determine the period between the patients R wave. Alternatively, if switch 255 connects the one shot multivibrator 271 with line 250a the counting circuit will determine the cardiac cycle time from a pressure input circuit driven by a suitable transducer attached to the patient and impressed on input line 250.

Regardless of the input to the counting circuit, the interval between successive waveforms is counted by one of the digital counting banks 251, channel A, or 252, channel B, and the resulting signal proportional to the interval between waves or pulses is delivered to the modulation circuitry in FIG. -8 through output line 250a. The digital counting circuitry produces a linear output voltage with time as shown in the graph in FIG. 7.

When the pressure wave is used as an input, it is attenuated by variable attenuator 266a which includes a voltage divider adjustable from a suitable control panel. This waveform drives an AGC circuit 256a which consists of dual cascaded field effect transistors. This waveform is impressed on a pressure trigger level selector 268 which consists of an integrator, an AC amplifier, an AC signal detector, and a Schmitt trigger. The integrator 267 is provided to prevent the digital counter from viewing the portion of the synthesized pressure waveform produced by the pump as a separate triggering wave. The output of the AC amplifier (not shown) is integrated by the integrator 267 and biases the AGC circuitry 256a to maintain a constant amplitude at the AC amplifier output. Potentiometer 269 in the pressure trigger level detector 268 provides an adjustable bias for the detector so that a selected portion of the patients pressure wave may be employed to initiate the counting circuitry. With this circuitry, false signals may be filtered and prevented from initiating the digital counting circuit. The Schmitt trigger (not shown) in the pressure trigger level de tector 268 triggers the one shot multivibrator 271 through selector switch 255. The trigger level of the pressure wave may be viewed on the oscilloscope through an isolation oscilloscope driver 272a to output 272. The one shot 11 multivibrator 271 generates a 100 microsecond wide pulse.

Of course, the one shot multivibrator 271 may also be triggered by the EKG waveform. It should be remembered that during cardiopulmonary bypass the one shot multivibrator 271 is triggered only by pulses from the pacemaker 132.

The one shot multivibrator 271 triggers a master flipflop 27212 and an inverter 273. The master flip-flop 272b selects either. channel A or channel B for counting on alternatively successive cardiac cycles. The master flipflop may be a standard dual PNP transistor flip-flop with base steering. The inverter 273 resets the digital counting banks 251 and 252 to erase the count held by the counting banks at the beginning of every other cycle by providing a reset pulse to each of the AND gates 274 and 275 upon receipt of each triggering pulse. The master flip-flop 27217 provides an enabling pulse to each of the AND gates 274 and 275 on alternate triggering pulses so that after a first triggering pulse channel A resets and after a second triggering pulse channel B resets. The reset circuits 276 and 277 are conventional transistor switches which reset each appropriate channel by simultaneously base triggering its associated flip-flops.

The master flip-flop 2721) also enable the AND gates 280 and 281 selectively on alternate cardiac cycles, so that the first signal from the master flip-flop through line T enables AND gate 280 allowing counting bank 251 to count, While the second signal from the master flip-flop 27212 disables the AND gate 280 and enables the AND gate 281 through line P so that the flip-flop bank 252 is activated and ready to count.

The digital pulses for the counting channels A and B are provided by an astable clock 285 which may be a multivibrator or a standard unijunction transistor oscillator which triggers a driver flip-flop. When the trigger pulse from the patients pressure wave energizes the master flip-flop 27212 which activates one of channels A or B, clock 285 delivers digital pulses to it.

Each of the counting banks 251 and 252 consists of a plurality of series connected flip-flops which count pulses from the astable clock 285 in binary coded fashion. Flipflops 290 to 297 in bank 251 and flip-flops 300 to 307 in bank 252 are standard dual PNP transistor flip-flops with base steering. Each flip-flop divides the frequency of its input signal by 2, successively. Each bank is capable of counting 256 pulses. A three input AND gate 309 is provided between the flip-flops 304 and 305 so that counter circuitry output may not indicate less than a count of 224 if it has been inadvertently allowed to exceed a count of 256. The inputs to the AND gate 309 are driven from the false side of the flip-flops 305 and 306 and 307 so that the trigger pulse to flip-flop 305 from 304 is inhibited after each of these three flip-flops has been triggered once. A similar circuit is provided for bank 251 (channel A).

The false side of each of the counter flip-flops 290 to 297 and 300 to 307 may be diode coupled to constant current switches 310 to 317 and 320 to 327. Each of the constant current switches is composed of a single PNP transistor stage which delivers a current flow proportional to its corresponding count when its associated flipfiop is triggered. For example, constant current switch 321 will produce a current proportional to a count of two, while constant current switch 325 will produce a constant current proportional to a count of 32 so that the sum total of the currents in the constant current switches associated with the triggered fiipfiops represents the time duration of the cardiac cycle. A composite analog readout of the constant current switches is obtained by summing the currents through a precision resistor 330 associated with channel A or a precision resistor 331 asociated with channel B. Channel switches 335 and 336 are single transistor stages which alternatively shunt the precision 12 output resistors 330 and 331. The channel which is counting is always shunted by the channel switches which are driven by master flip-flop 27212 through either line F or T. An added 337 receives a constant DC voltage at its input from one or the other of channel switches 335 or 336.

As described in more detail below, a timing error sensor circuit is provided for correcting any error between the actual displacement cycle time and the period of the input signal from the pacemaker, or the actual cardiac cycle time (in arterio-arterial assist). This error, if it is negative, is delivered to the isolation circuitry output 337a through line 340 where it subtracts from the counted input signal period to provide a new effective period. If it is positive it is also delivered to the isolation circuitry output 337a where it adds to the counted period. As the period of the input signal delivered to the one shot multivibrator 271 is the controlling parameter for the computer modulating circuitry it is possible to correct errors in the displacement cycle time by varying the effective period signal to the modulation circuitry through line 252a.

It is apparent that each of the channels 251 and 252 count on alternate input signals so that upon completion of any one cycle channel A will read out to the modulating circuitry in response to a change in state of master flip-flop 272b, while at the same time channel B counts. Upon the initiation of the counting circuitry by the next input signal pulse which changes the state of master flip-flop 2721: again, channel B reads out to the modulating circuitry and channel A counts.

Referring to FIG. 8 wherein the modulating circuitry for determining the push and withdrawal rates is schematically shown, a saw toothed oscillator 350 produces a saw tooth waveform as shown at 351 with a rising ramp voltage and a very steep retrace. The saw tooth oscillator 350 may consist of a unijunction relaxation oscillator driven by a constant current source. The output may be capacitively coupled to a variable voltage divider which is internally adjusted to compensate for the voltage gain tolerance of the unijunction oscillator.

A duty cycle modulator 352 is provided for relating the desired or limited blood displacement volume to a time base. The output of the duty cycle modulator 352 is a pulse having a width modulated proportionally to the desired or limited volume of blood to be displaced.

A clipping amplifier 353 is provided for setting the maximum volume, and hence the maximum pump and withdraw rates, for given heart cycle period. This circuit is only operative during arterio-arterial assist when the system responds to the patients EKG wave. To achieve this, the linear voltage which represents the heart period is impressed on line 355 and constitutes an input to the clipping amplifier 353. It will be recalled that there is a maximum volume which may be pumped or should be pumped for every heart cycle period. In this regard, the clipping amplifier 353 prevents excessive withdrawal and pump rates caused by an extremely fast heart beat rate counted by the digital counter in FIG. 7. The clipping amplifier 353 may consist of a two stage DC amplifier having an NPN transistor driving a PNP transistor, with gain quiescent bias adjustment provided by the patients computed cardiac cycle time.

A catheter size selector switch 356 is provided for varying the maximum volume limitation determined by the clipping amplifier 353 for dilferent catheter sizes.

The relationship of heart rate or pacemaker rate to the maximum permissible volume may be drawn as exponential curves based upon the transfer characteristics of each individual catheter size. Therefore, the maximum permissible volume for any catheter size is a complex function of the flow characteristics of the particular catheter used. The catheter size selector switch 356 in response to the selection of any of the catheter sizes biases the clipping amplifier 353 to vary the gain thereof and hence the maximum permissible volume in accordance with calculated data. An emitter resistor for the PNP stage in the clipping amplifier 353 is selected from one of eight resistors which correspond to eight catheter sizes. Hence, the gain of the clipping amplifier 353 is selected by the selector switch 356. In addition an output resistive voltage divider is connected to the collector of the PNP transistor and one of eight resistors is simultaneously selected to adjust the output quiescent level of the DC amplifier. By this circuitry the clipping amplifier 353 produces output voltages varying with the heart period, or pacemaker, for each catheter size.

It should be noted that the clipping amplifier 353 merely sets the maximum limit of the volume pumped during any cycle and that for any lesser volume selected by the surgeon will normally determine the push and withdrawal rates.

A volume limit generator 358 biases the duty cycle modulator 352 to achieve a pulse output from the duty cycle modulator having a width proportional to either the maximum permissible volume or the manual volume selected, if it is less than the maximum permissible volume which is determined by the clipping amplifier 353.

The duty cycle modulator 352 may consist of a twotransistor stage DC amplifier, followed by a regenerative switch and an NPN reset transistor. The sawtooth oscillator output is amplified and provides a forward bias to the first stage of the regenerative switch. Simultaneously, a back bias is provided by the volume generator circuitry. When the sawtooth amplitude exceeds the volume amplitude, the regenerative switch turns on. When the sawtooth waveform retraces to start another ramp, the NPN reset transistor turns on, resetting the regenerative switch to its normally off condition. The DC amplifier gain is manually adjustable to accurately match the volume generator output.

A manual volume adjust 360 is provided which normally controls the level of the bias of the volume limit generator on the duty cycle modulator 352. Therefore, the width of the pulse from the duty cycle modulator 352 is normally proportional to the manually adjusted volume unless the maximum volume limit for a particular heart rate is exceeded. The output of the clipping amplifier 353, which is discussed above, is a signal proportional to the maximum permissible volume and hence the maximum permissible pump and withdraw rates. It drives an emitter follower in the volume limit generator 358' which limits the output of the volume adjust potentiometer in the manual volume adjust 360 to a value proportional to the maximum permissible volume. The emitter of the emitter follower is diode coupled to the wiper of the volume adjust potentiometer and hence the low driving point impedance of the emitter follower effectively modulates the relatively high impedance of the circuit associated with the volume potentiometer. The volume adjust potentiometer is connected in series with a voltage divider wherein its wiper is the output of the volume limit generator 358. The potentiometer in the manual volume adjust 360 may be on the control panel so that the surgeon may select the desired volume.

A catheter select switch 363 is provided which selects either one or two catheters. The catheter select switch 363 selects the range of the voltage divider in the volume limit generator 358. Switch 363 divides the bias to the PNP limiting emitter follower from 2:1. Thus, catheter assembly 49 shown in FIG. 1 may consist of either one or two catheters. When one catheter is used instead of two, the catheter select switch 363 approximately halves the maximum permissible volume limit, which controls or limits the output of the volume adjust potentiometer 360. When two catheters are used, the two catheter switch in the catheter select switch 363 approximately doubles the maximum voltage the voltage adjust potentiometer 360 may produce over the single catheter select switch position. It should be noted again that the catheter select switch 363,

the catheter size selector switch 356 and the clipping amplifier 353 limit only the maximum volume which may be pumped without exceeding the maximum withdrawal and pump rates. They do not affect the width of the output pulse from the duty cycle modulator 352 unless the potentiometer in the manual volume adjust circuit 360 attempts to produce a voltage representing a volume which would produce excessive rates.

The volume limit generator 358 drives an output line 203 for delivering a signal representing the desired volume to: (l) a volume indicator (not shown) for visual representation, (2) to the withdrawal volume comparator 201 which enables the pump trigger AND gate 146 as described above with respect to FIG. 6, and (3) to the blanking volume comparator 211 which enables the AND gate 141 which transfers the delayed pump triggering signal to I blanking gate 144.

The above circuitry produces a pulse width modulated pulse train from the duty cycle modulator 352 proportional to the desired volume of blood to be pumped. As the heart rate varies or the pacemaker rate is varied by the surgeon, it is apparent that the withdrawal or the pump rate must change to maintain the same volume of blood delivered and to achieve a time coincidence between the displacement cycle and the period of the pacemaker signal (or the cardiac cycle). For this purpose, an amplitude modulator 370 and a gain break amplifier 371 are provided for amplitude modulating the pulse train from the duty cycle modulator 352. The input to the gain break amplifier 37-1 (line 355) is a voltage proportional to the period of the pacemaker signal (or the heart cycle) determined by the digital counter in FIG. 7. The voltage in line 355 is linear with respect to time. Since the Withdrawal and pump rates must be proportional to the frequency of the pacemaker signal or the patients heart, linear voltage at 355 must be converted to an exponential voltage representing frequency. Hence, the patients heart period or pacemaker period which is on a time base must be converted to a frequency base so that the amplitude modulator 370 produces a pulse train having an amplitude proportional to the frequency of the heart signal or pacemaker signal.

For this purpose the gain break amplifier 371 produces an exponential output voltage as shown in graph 375 in response to a linear voltage input from line 355. The gain break amplifier 371 consists of an emitter follower which drives a voltage divider. The impedance of the voltage divider is a function of the divider output voltage. As the voltage increases diodes in the circuit become for ward biased, coupling successive resistive loads to the output. There are six of these gain breaks which generate an exponential output curve from the linear input voltage at line 355. The gain break output is amplified by a single NPN stage which modulates the amplitude modulator. Since heart frequency is an exponential reciprocal function of the heart period or pacemaker period, the gain break amplifier 371 effectively provides an exponential voltage output proportional to the frequency of the pacemaker signal or cardiac cycle.

The amplitude modulator 370 may consist of a single PNP transistor inverting amplifier driven by the pulse width modulator pulse train from the duty cycle modulator 352. The emitter of this PNP inverter amplifier in the amplitude modulator 370 is back biased by the output of the gain break amplifier 371. The net collector current flow in the amplitude modulator 370 is therefore directly proportional to the differential of these two biases. In this manner the pulse train output from the amplitude modulator 370 has an amplitude proportional to the frequency of the pacemaker signal or cardiac cycle.

An integrating amplifier 377 is provided for integrating the pulse train which is modulated both with volume and with pacemaker frequency. The integrating amplifier 377 is a conventional RC integrator which drives two NPN DC amplifiers in parallel. The integrating amplifier 377 produces a DC analog voltage equal to the duty cycle times the peak output voltage which is actually proportional to the desired volume times the pacemaker frequency. Therefore, the DC voltage output has a level proportional to the desired pump and withdraw rates. It is apparent that as the frequency of the pacemaker increases, as computed by the gain break amplifier 371, the pulse amplitude from the amplitude modulator 370 will increase as will the DC level from the integrating amplifier 377. Therefore, as the pacemaker frequency increases, so do the push and withdraw rates to achieve coincidence between the pacemaker signal period and the displacement cycle. Similarly, as the desired volume increases, represented by the output voltage from the volume limit generator the pulses from the duty cycle modulator 352 will increase also producing an increase in the DC level voltage output from the integrating amplifier 377 resulting in an increase in pump and withdrawal rates.

The DC voltage output from the integrating amplifier 377 biases a charge gate 378 and a discharge gate 379. The gates 378 and 379 may each consist of an NPN transistor, the collector of which modulates a pair of PNP transistor drivers. The PNP transistors are forward biased, in parallel, by the integrating amplifier. If the NPN gating transistor is turned on, both PNP transistors are back biased. It will be recalled that during the withdraw stroke, line 152 is energized by the regenerative switch 148 and during the pump phase line 151 is energized by the regenerative switch. During the withdrawal phase, line 151 turns on the NPN gating transistor in the discharge gate 379 back biasing the PNP transistors in the discharge gate 379. The absence of a signal on line 152 produces an output to charge driver 381 and line 382 which energizes the withdraw phase driver 188 as shown in FIG. 6. The withdraw phase then proceeds at a rate corresponding to the magnitude of the current in line 382.

When the regenerative switch 148 deenergizes line 151 and energizes line 152, the charge gate 378 turns on ending the withdrawal stroke and the discharge gate 379 turns oif initiating the pump stroke by energizing the push phase driver through line 160. The pump stroke rate is determined by the integrating amplifier 377 similar to the withdraw stroke. A charge driver 381 is associated with the charge gate 378 in the same manner as the discharge driver 383 is driven by the discharge gate 379.

A closed loop error feedback circuit is provided for achieving coincidence between the reference waveform produced by the modulating circuitry in FIG. 8 and the actual displacement waveform produced by the feedback potentiometer 118. Referring to FIG. 8, the charge driver 381 and the discharge driver 383 drive a reference ramp generator 385. The ramp generator 385 may be a low leakage wet tantalytic capacitor with a minimum voltage clamping circuit. The clamping circuit may be a resistor voltage divider which is diode coupled to the reference capacitor. The reference ramp generator 385 produces a triangular waveform proportional to the desired displacement of the reciprocating heart pump piston with respect to time. Thus, the reference ramp generator produces a reference signal which biases one side of a comparator 386. The sense of the reference signal is determined by the regenerative switch 148.

The comparator 386 may be simple NPN transistor comparator forward biased by the reference ramp generator 385. The back bias for this transistor is provided by the feedback voltage from the potentiometer 118 which is inverted by the inverting amplifier 170 shown in FIG. 6. It is amplified by a two stage, non-inverting DC amplifier 388, shown in FIG. 8. As long as the reference ramp waveform from the reference generator 385 coincides with the inverted displacement waveform from the amplifier 388, a fixed quiescent voltage drives both the low error amplifier 389 and high error amplifier 390. The low error amplifier 389 may consist of a dual PNP inverter which drives a resistive voltage divider, and amplifies the negative error voltage between the reference and displacement wave voltages. The high error amplifier 390 may consist of a PNP transistor stage which drives a resistor voltage divider, and amplifies the positive error between the reference and displacement wave voltages. The outputs of the low error amplifier 389 and the high error amplifier 390 are connected to their associated AC amplifiers, 173 and 190, respectively, through AND gates 395 and 396.

The AND gates 395 and 396 along with an associated chopper flip-flop 398 provide continuous repetitive error correction during the withdraw and pump phases. The chopper flip-flop 398 is driven by the sawtooth oscillator 350 through line 400 and provides alternate enabling pulses to AND gates 395 and 396 through'switches 402 and 403, respectievly. It is apparent that if the sawtooth oscillator provides sawtooth pulses on the order of microseconds that AND gates 395 and 396 will be enabled many times during each pumping cycle and therefore provide a virtually continuous error correction of the pump and withdraw rates. The pump and withdraw rates determine the volume delivered.

As discussed above, the push and withdraw rates for each cycle are determined by the modulating circuitry of FIG. 8 from the pacemaker rate counted in the previous period of the pacemaker signal by the digital counter described in FIG. 7.

Referring now to FIG. 9, a timing error sensor circuit generally designated by the numeral 415 is provided for computing the time difference between the completion of the withdrawal stroke and the receipt of a pump cycle trigger signal. The resulting error signal is used to vary the pump and withdrawal rates during the next cycle in closed looped fashion. This closed loop approach to error correction compensates for actuator gain variations and counting errors. The timing error sensors 415 are essentially two channels which integrate the error defined by the on time of a regenerative switch which is a pulse proportional to the time error between the end of withdraw stroke and the pump triggering signal. Channel 416 determines the error when the pumping cycle is completed before the next triggering pulse is received by the blank gate 144, and channel 417 determines the error when the pump triggering signal from the blanking rate is received before the end of the cycle. Therefore, channel 416 determines the error when the pump and withdraw rates are too fast for a particular pacemaker period and channel 417 determines the error when the pump and withdrawal rates are too slow for a particular pacemaker period (or cardiac cycle).

Each of the timing error sensor channels 416 and 417 are substantially identical except for having cross connected inputs. Differentiators 418 and 419 receive input pulses respectively from the end of withdraw stroke one shot multivibrator 207 and the blanking gate 144. The difierentiators 418 and 419 may be typical RC type. The differentiated input signal from the ditferentiators 418 and 419 turn on regenerative switches 421 and 422 The regenerative switches 421 and 422 may each be standard two PNP transistor cascaded switches with the output fed back to the input. Reset stages 424 and 425 are provided for turning off their respective regenerative switches 421 and 422. The input to the reset stage 424 is the blanking gate output v144, while the input to the reset stage 425 is the end of withdraw multivibrator 207. The reset stages 424 and 425 may each be diode coupled resistor logic which biases the regenerative switch off upon receipt of the appropriate logic command. The time duration that each of the regenerative switches 421 and 422 are on represents the positive or negative error between the pumping cycle and the period of the pacemaker signal (or patients cardiac cycle). The regenerative switches 421 and 422 drive integrating amplifiers 428 and 431 which produce a DC voltage proportional to the pulse time duration of their associated regenerative switch outputs. The integrating amplifier 428 and 431 may each consist of a two stage transistor switch, followed by an RC integrator and a single NPN transistor DC amplifier. The output pulse of integrating amplifier 428 drives a readout amplifier 429. The readout amplifier 529 may consist of a PNP transistor DC amplifier. Channel 417 has a similar NPN transistor readout amplifier 432.

If the displacement cycle is too fast compared with the pacemaker period or cardiac cycle, the end of withdraw stroke multivibrator 207 will turn on regenerative switch 421 before an output pulse from the blanking gate 144 initiates the reset stage 424 producing a DC error output from readout amplifier 429. The error signal from the readout amplifer 429 increases the current in the isolation circuit summing resistor in the digital counting circuit shown in FIG. 7. Since this current represents the time period of the pacemaker 'wave or cardiac cycle, the readout amplifier 429 effectively lengthens the counted pacemaker cycle or cardiac cycle time, producing a new, effective cycle time. Since the cycle time has been effectively increased, the modulating circuitry, shown in FIG. 8, views an effectively longer cycle time than it would view without the addition of the error signal. Since an increase in the pacemaker or cardiac cycle count produces a reduction in the withdrawal and push rates determined by the modulating circuitry in FIG. 8, the displacement wave slows down on the next pump cycle to correct for this error. Conversely, if the pump cycle is too slow compared with the pacemaker or cardiac cycle, a triggering pulse from the blanking gate .144 will turn the regenerative switch 422 on before the end of withdraw stroke multivibrator 207 resets the regenerative switch 422 so that readout amplifier 432 produces an output voltage proportional to the time error. This error signal decreases the current in the isolation circuit summing resistor in the'counting circuitry shown in FIG. 7, to effectively reduce the counted pacemaker or cardiac cycle time computed by the digital counter circuit. The modulating circuitry shown in FIG. 8 will View a reduced cycle time and will increase the pump and withdrawal rates on the next pump cycle to correct for this error.

During arterior-arterial assist, the pump triggering pulse occurs very early in the cardiac cycle, i.e., a predetermined time before the end of the withdrawal stroke as determined by the blanking threshold sensor 210 and the volume comparator 211, the AND gate 141 will drop the triggering signal so that there is no output from the blanking gate 144 and therefore no error output from the channel 417 timing error sensor which normally indicates that the pumping cycle is too slow. In this case, the pump and withdrawal rates are determined by the cardiac cycle time counted by the digital counting circuit on the previous cardiac cycle without the modulation of any error fromthe timing error sensors.

Single stroke initiation The present device is provided with means for initiating a single pump stroke. The purpose of providing this capability is to enable the system to be used for injecting radio-opaque dies in the cardio vascular system and/or to obtain coronary angiograms. Referring to FIG. 6, when the select switch 131 is placed in its manual position a suitable circuit is provided for withdrawing the actuating shaft to the end of the withdraw stroke. In this mode the triggering pulses from the EKG trigger selector 125 and the pacemaker 132 are disconnected from the control system so that repetitive pumping action is prevented. Upon the closure of the spring loaded manual stroke initiate switch 133 lockout gate 450 disables the blanking gate 144. This prevents the initiation of a triggering pulse from AND gate 141. At the same time, a withdraw disable gate 452 is activtaed which disables the withdraw driver 194 and prevents any initiation of the withdraw stroke. Actuation of switch 133 also energizes line 453 charging capacitor 454 which initiates a pump triggering signal to line 150 turning the regenerative switch 148 on energizing line 151 and thereby initiating a single pump stroke of the reciprocat- Reduced augmentation circuit Referring to FIG. 10, a reduced augmentation circuit 460 is provided for skipping one or three cardiac cycles. In this manner the heart pump system will operate only on selected pacemaker pulses or selected beats of the patients heart, if desired. A disable gate 461 is provided for disabling the blanking gate 144 except when selected triggering pulses are received by blanking gate 144. A switch 462 is provided for selecting one hundred percent augmentation, fifty percent augmentation and twenty-five percent augmentation as shown. With the switch in the one hundred percent augmentation position, the disable gate 461 completely disables the blanking gate 144 and permits all triggering pulses from the AND gate 141 to be coupled to AND gate 146. In this event all normal triggering pulses from refractory gate initiate a pump cycle on each pacemaker pulse or cardiac cycle. When switch 462 is in the fifty percent augmentation position, a flip-flop 463 enables the disable gate 461 to blank alternate triggering signals from AND gate 141 and effect a pump cycle only on alternate pacemaker pulses or cardiac cycles. Flip-flop 463 is tripped to either state by refractory gate 140. This is important because if reduced augmentation were responsive to a time type blanking gate, an increase in the heart rate may produce the blanking gate, an increase in the heart rate may produce the blanking of more than one triggering pulse even if fifty percent augmentation was desired. Another flip-flop 464 cascaded with flip-flop 463, and an AND gate 465 are provided to drive the twenty-five percent augmentation. Now on every fourth beat of the patients heart (or pulse from the pacemaker 132) the binary flip-flop 464 which is the fourth binary stage activates AND gate 465 and the disabled gate 461 permitting a pulse to pass through the blanking gate 144. In this manner three cardiac cycles or pacemaker pulses may be skipped regardless of variations in the repetition rate of the input signal to the system.

Power supply Referring to FIG. 11, a two contact switch 490 is provided connected to a 30 volt DC power supply. When switch 490 is in its upper position, a release coil 491 is energized deactivating the single catheter select switch 492 and a dual catheter select switch 493. This assures that each of the catheter select switches will be placed in their olf positions when the system is turned on assuring the correct catheter selection on the next use of the pumping system. When switch 490 is placxed in the lower position, i.e., the on position, release coil 491 is deenergized permitting the manual selection of either the single or dual catheter switches 492 and 493. At this time the coil 495 is energized connecting the hydraulic power supply to a source of electric power. At the same time, the electric control circuitry is energized with the exception of the pump and withdraw servo coils 147 and 154 respectively.

Before pump and Withdraw servo coils 1-47 and 154 may be energized, it is necessary that the surgeon perform the following functions; (1) select either the single or the dual catheter switches 492 or 493, (2) choose a catheter size by actuating the catheter size selection switch which energizes a holding coil 497, (3) and set the manual volume adjust potentiometer 360 to zero volume which by suitable circuitry closes switch 498. After the surgeon has made all these selections, relay 499 may be energized by closing switch 503 so that contacts 2 and 3 are made, placing the circuit in the operate position indicated by the energization of an operate lamp 501 on the control panel. If the surgeon has failed to make one of the parameter selections, no current will fiow through the coil 499 and even if the surgeon actuates the start switch 503, power will not be supplied to the pump and withdraw servo coils 147 and 154. In this case, coil 499 will not be energized and contacts 4 and 1 of relay 500 will be closed lighting a standby lamp 504 on the control panel indicating to the surgeon that he has failed to make a parameter selection. A lamp 506 is also provided on the control panel for indicating when the catheter size select switch 363 has been closed.

The operation With the components and catheters connected and arranged according to FIGS. 1 and 2, or alternatively according to FIGS. 4 or 5, the power supply circuit shown in FIG. 11 is placed in the standby position. The surgeon then makes the various catheter selections and adjusts the volume potentiometer to place the power supply circuit in the operate position indicated by the operate lamp 501. The piston 15 at this time is at the end of the pump stroke. The various oscilloscope outlets are connected so that the surgeon views the various parameters and the start switch 503 is tripped preparing the control circuitry for the receipt of a pump triggering signal. Switch 131 is placed in the pacemaker position and switch 255 is moved to a position connecting line 266 with the one shot multivibrator 271. The surgeon then adjusts the pacemaker 132 to the desired repetition rate of the pumping cycles.

The delay one shot 136 may be adjusted so that the output pulses to the refractory gate are in phase with the pulses from the pacemaker 132 as no delay time is necessary in complete cardiopulmonary bypass.

Pulses from the refractory gate 140 trigger the regenerative switch 148 and the pump proceeds through its pump stroke forcing blood into the patients arterial tree through catheter assembly 49. Then the piston 15 proceeds through a stroke dictated by the volume potentiometer 360 associated with the volume limit generator 358, the regenerative switch 148 is turned off reversing the piston 15 causing blood to be drawn from reservoir 36 into pump chamber 16. This cycle is repeated w1th each triggering pulse from the pacemaker 132.

If the surgeon desires to increase the arterial flow rate, the frequency or repetition rate of the pacemaker 132 may be increased. In response to an increase in the rate of pulses from pacemaker 132, the modulating circuitry shown in FIG. 8 will increase the speed of stroking of piston 15 so that the pump cycle corresponds with the newly selected period of the pacemaker signal. Further, the arterial flow rate may also be increased by increasing the selected volume on the manual volume adjust 360. In response to this change the volume comparator 201 (controlled by the volume limit generator 358) increases the length of stroke of the piston 15, and the duty cycle modulator 352 at the same time increases the speed of stroke of the piston 15 so that the newly selected volume per cycle will be achieved without variation in the time duration of the pumping cycle.

The proper operation of the present system in complete cardiopulmonary bypass depends upon the proper balance of the manually controlled parameters of volume per stroke and pacemaker rate. If the pacemaker rate is too high, the speed of stroking of the piston 15 will produce blood damage. If the pacemaker rate is too slow, the selected volume per stroke must be increased and this may rupture the small vessels. Thus, the volume adjust ment and rate control must be varied to obtain optimum physiological results between these two limits.

As noted above, the present system may be used postoperatively as an arterio-arterial circulation assist after complete cardiopulmonary bypass. T 0 place the system in this mode it is merely necessity to close valve 11 and connected the various inputs to the computing and triggering circuitry so that the patients EKG wave triggers the pumping cycles. The phasing in this mode is described in detail in our above mentioned copending application.

Thus far, tests on animals have indicated no tendency toward metabolic acidosis with the present complete cardipulmonary bypass system even after extended perfusion. With one exemplary test animal complete cardiopulmonary bypass proceeded as follows. The dog was not allowed to cool, but to warm slowly from 37 degrees centigrade to 39.5 degrees centigrade for one hour, then allowed to cool to 34 degrees centigrade over 45 minutes, and then warm to a normal 37 degrees over the next 30 minutes, the latter temperature being maintained for a further 30 minutes.

At the beginning of the perfusion the peak flow rate was 1,720 ccs. per minute or 62 ccs./kg./min. i.e., peak flow per pump cycle. The flow rates used herein are the peak flow rates attained at the top of the flow curves, the actual flow per cycle being much lower. A pulsed-field electromagnetic flow probe was placed around the common iliac artery, from which the adventitia had been stripped to avoid interference with the readings, for the purpose of sensing the blood flow rate in the animal. It was placed far enough away from the pump catheter so that variations in its diameter with the withdrawal and pump stroke would not affect the electrical fit for the flow probe.

The pressures attained at the initial flow rate were 150/100 which appeared adequate. After 45 minutes the flow rate was readjusted to 1,380 ccs. per minute peak flow, or 49.5 ccs./kg./min. peak flow per cycle. Pressure was then an acceptable /70. Only 350 ccs. of saline were added during the first hour and this was less than the loss. The perfusion continued for 3 hrs. The animal was acidotic (respiratory) at the outset of the perfusion with arterial pH being 7.18 and the venous pH being 7.15. After two hours of pulsatile perfusion the arterial pH had risen. Both the arterial and venous pHs increased toward normal during the third hour of perfusion. There was no tendency toward metabolic acidosis. The arterial and venous carbon dioxide contents and the pCO levels all declined during the third hour of perfusion. The lactate pyruvate levels rose, but the lactate pyruvate ratios remained stationary or declined slightly during the perfusion. Thus, the perfusion of the arterial tree seemed adequate enough to prevent pooling of blood with resultant tissue hypoxia. Similar biochemical results were found in the other animals.

We claim:

1. A heart pump system for complete cardio-pulmonary bypass, comprising: a pulsatile pump having a pumping chamber, a piston reciprocable within said pumping chamber having push and withdraw phases defining a pumping cycle, catheter means for receiving blood from the venous side of a patients circulatory system, oxygenator means connected to receive and oxygenate blood from said catheter means, means for conveying blood from said oxygenator means to said pumping chamber, second catheter means connecting said pumping chamber with the arterial side of the patients circulating system, means for reciprocating said piston so that oxygenated blood flows into the patients circulatory system in pulsatile fashion; and control means for said reciprocating means including means for deriving pump triggering signals corresponding with the desired repetition rate of the pumping cycle, means for initiating the pumping cycles in timed relationship with said pump triggering signals, manually controlled means for varying the volume of blood pumped on each pumping cycle, and means responsive to said manually controlled means for automatically increasing the blood flow rate through said second catheter means as the selected volume of blood is increased, said manually controlled means including means for varying the length of stroke of said piston, said means for automatically increasing the blood flow rate including means for increasing the rate of movement of said pumping piston.

2. A heart pump system for complete cardio-pulmonary bypass, comprising: a pulsatile pump having a pumping chamber, a piston reciprocable within said pumping chamber having push and withdraw phases defining a pumping cycle, catheter means for receiving blood from the venous side of a patients circulatory system, oxygenator means connected to receive and oxygenate blood from said catheter means, means for conveying blood from said oxygenator means to said pumping chamber, second catheter means connecting said pumping chamber with the arterial side of the patients circulating system, means for reciprocating said piston so that oxygenated blood flows into the patients circulatory system in pulsatile fashion; and control means for said reciprocating means including means for deriving pump triggering signals corresponding with the desired repetition rate of the pumping cycle, means for initiating the pumping cycles in timed relationship with said pump triggering signals, manually controlled means for varying the volume of blood pumped on each pumping cycle, and means responsive to said manually controlled means for automatically increasing the blood flow rate through said second catheter means as the selected volume of blood is increased; and means for preselecting the repetition rate of said triggering signals as desired, said manually controlled means including computer means responsive to the period of said triggering signals for providing substantial coincidence between said triggering signal period and said pumping cycle, said computer means being reponsive to a change in the selected volume for varying the speed of said piston to achieve said coincidence, whereby the speed of said piston will increase and decrease with the selected blood volume.

3. A heart pump system for complete cardio-pulmonary bypass, comprising: a pulsatile pump having a pumping chamber, a piston reciprocable within said pumping chamber having push and Withdraw phases defining a pumping cycle, catheter means for receiving blood from the venous side of a patients circulatory system, oxygenator means connected to receive and oxygenate blood from said catheter means, means for conveying blood from said oxygenator means to said pumping chamber, second catheter means connecting said pumping chamber with the arterial side of the patients circulating system, wherein said second catheter means permitting blood flow in either direction, said second catheter means and said conveying means having a relative resistance to flow such that during the withdraw stroke of said piston a substantial portion of the blood will be withdrawn from said conveying means rather than from said second catheter means, means for reciprocating said piston so that oxygenated blood flows into the patients circulatory system in pulsatile fashion; and control means for said reciprocating means including means for deriving pump triggering signals corresponding with the desired repetition rate of the pumping cycle, means for initiating the pumping cycles in timed relationship with said pump triggering signals, manually controlled means for varying the volume of blood pumped on each pumping cycle, and means responsive to said manually controlled means for automatically increasing the blood flow rate through said second catheter means as the selected volume of blood is increased.

4. A heart pump system for cardiopulmonary bypass and postoperative arterio-arterial assist, comprising: a pulsatile pump having a reciprocating piston the push and withdraw strokes of which define a pumping cycle, said piston defining a fluid chamber in said pump, means for receiving and modifying a signal representing one of the patients physiological parameters and producing pump cycle triggering signals for circulatory assist, pacemaker means for deriving pump cycle triggering signals for bypass operation, control means for initiating each pumping cycle, selectively operable means for connecting the output of said producing means or said deriving means to said control means, catheter means connecting said fluid chamber into the patients circulatory system so that fluid is delivered thereto during the push stroke of said piston, means for conveying fluid into said fluid chamber during bypass operation, said conveying means having a check valve for preventing the flow of fluid from the fluid chamber into the conveying'means, said catheter means being unrestricted so that fluid may flow in either direction therethrough during arterio-arterial assist, and valve means for closing said conveying means during arterio-arterial assist.

5. A heart pump system as defined in claim 4, wherein said check valve is mounted within said valve means.

References Cited UNITED STATES PATENTS 2,847,008 8/ 1958 Taylor et al 23258.5 2,927,582 3/1960 Berkman et al. 23258.5 3,099,260 7/1963 Birtwell 1281 3,183,908 5/1965 Collins et a1. 23258.5 3,266,487 8/ 196 6 Watkins et a1. 1281 OTHER REFERENCES Murphy: Trans. Amer. Soc. Artificial Inter. Organs, 1961, vol. VII, pp. 361-68.

DALTON L. TRULUCK, Primary Examiner US. Cl. X.R. 23-25 8.5

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Classifications
U.S. Classification604/6.1, 422/47, 604/6.11, 604/67
International ClassificationA61M1/10, A61M1/32, A61M1/36
Cooperative ClassificationA61M1/1037, A61M1/3621, A61M1/1086, A61M1/32
European ClassificationA61M1/10R, A61M1/36C