Self-regulating Artificial Heart

Abstract

This invention describes a complete artificial heart system which is capable of providing complete self-regulative control of the pumping action of the heart. It comprises a right and left ventricular bladder connected to the blood system. These bladders are housed in right and left ventricular chambers, the space between the bladder and the chamber being filled with a drive fluid. Means are provided for sensing the volume of blood in the right and the left ventricular bladders. There is a power fluid system which comprises a sump maintained at atmospheric pressure, a pump, a flow control means and a switching means. The right and left sensors in the corresponding right and left ventricular chambers have outputs which are functions of the blood volume in the right and left ventricular bladders. A control means takes these sensor outputs and computes two control signals, one of which is a function of the difference between the volumes in the right and the left ventricular bladders, and the other is a function of the sum of the volumes in the right and the left ventricular bladders. The sum function goes to control the rate of flow of the power fluid and the difference function goes to control the switching. The switching means is provided so that when one bladder, for example the right ventricular bladder, becomes fully extended the left ventricular bladder becomes fully compressed, whereupon the drive fluid is switched into the right ventricular chamber to compress the bladder and push its contained blood out into the body system; at the same time, the power fluid that was filling the left chamber is released to flow back to the sump.

Description

BACKGROUND OF THE INVENTION

This invention is in the field of artificial hearts. More particularly, it is concerned with a complete system involving two ventricular bladders into which and out of which the blood flows, and including a power fluid system which is used to alternately compress the two ventricular bladders and further including sensor means to provide feedback signals to control the flow rate and switching of the power fluid so as to produce correct self-regulation of the blood flow requirements of the individual at all times.

DESCRIPTION OF THE PRIOR ART

The need for a reliable and functionally adequate artificial heart has been clearly demonstrated in recent years (see, for example, the publication: Bolie, Victor W., "Leading U. S. Artificial Heart Research Programmes," Journal of the Institution of Engineers (India), Volume 49, Number 5, Part GE 2, pp. 47-50, January 1969). However, up to the present time there has been no adequate means developed for providing safe and effective control of the blood-pumping function of an artificial heart, in terms of feedback signals from each of the two ventricles.

The normal physiological function of the biological heart may be summarized as follows: In the normal 70 kg (154 lb) adult man the left and right ventricles beat coherently with a pulse rate R which normally is within the range of 72 to 144 strokes/minute. The stroke volume outputs of the two ventricles tend to remain equal, and the right ventricle pumps blood (having low O.sub.2 and high CO.sub.2 content) into the pulmonary artery with a stroke volume Q in the range of 70 to 140 milliliters (ml)/stroke. The returning blood (having high O.sub.2 and low CO.sub.2 content) from the lungs is pumped into the aorta with a flow rate F = QR in the range of 5040 to 20,160 ml/min. Based on the pressure scale of 760 mm Hg = 14.7 psi = 1,013,200 dynes/cm.sup.2, the pressure in the pulmonary artery pulsates from zero to about 40 mm Hg above the ambient atmospheric pressure level. The aortic pressure pulsates between 80 mm Hg (diastolic) and 120 mm Hg (systolic) above the ambient atmospheric pressure level existing in the lung alveoli. In older people a blood pressure of 140/90, rather than 120/80, is typical.

A first object of the present invention is to provide an artificial heart which is capable of accommodating wide physiological ranges of blood pressures in the pulmonary artery and in the aorta, and which responds appropriately to physiological changes in the rates of venous blood flow out of the pulmonary and systemic circuits.

A second object of the present invention is to provide adequate feedback control signals from sensors which continuously monitor the instantaneous volumes of the two artificial ventricles.

A third object of this invention is to provide artificial ventricles which are functionally and structurally independent of the particular type of valves used in the blood conduits.

A fourth object of this invention is to provide a simple mechanical system for converting the instantaneous volumes of the artificial ventricles into appropriate switching and metering control signals for an artificial heart.

A fifth object of this invention is to provide a novel artificial ventricle having a smooth-flow interior, a long fatigue life, a permanently stable bladder-volume coupler, an external contour fitting the inside wall of the rib cage, and an intercostal access tube permitting simple connections for the necessary energizing and monitoring functions.

A sixth object of this invention is to provide both a mechanical way and an electrical way of converting the bladder distension signals from an artificial heart into appropriately timed and metered fluid power pulses.

A seventh object of this invention is to provide a novel mechanism for maintaining constant symmetry and centering of the pulsating bladder in an artificial ventricle while converting the instantaneous bladder volume into an equivalent shaft rotation.

An eighth object of this invention is to provide an artificial heart which gains the several advantages of an unattenuated constant flow of circulating power fluid while still being responsive to physiological demands for moderate changes in cardiac output.

SUMMARY OF THE INVENTION

One significant feature of this invention which differs from the prior art systems is that it uses a sensor system which measures the instantaneous volume of fluid in each of two chamber-enclosed ventricular bladders, by means of which feedback signals are provided to control the power fluid flow in accordance with the requirements of the blood system in the body.

The information produced by the two ventricular-bladder-volume sensors is used to produce two control signals, the first control signal which meters the rate of total flow of power fluid, is a function of the sum of the volumes of blood in the two ventricular bladders. The second control signal, which controls the valving or switching of the outflow and return conduits of the power fluid pump with respect to the two ventricular chambers, is a function of the difference of the two ventricular bladder volumes.

The basic artificial heart system thus comprised is fully self-regulating, in that it not only automatically balances the blood outflow rates of the right and left ventricular bladders into the pulmonary artery and the aorta, but also automatically responds to changing physiological requirements in a total blood flow rate.

In this invention the mechanical sensors which couple the ventricular bladders to the ventricular chambers is positive, and far more reliable than previous sensors, and have long term durability. Also this system is devoid of the long term base line drifts associated with the gradual slippage of fluid past valves and seals, and with the slow absorption of power fluid gases through the bladder and tubing walls.

BRIEF DESCRIPTION OF THE DRAWINGS

These and other objects and an understanding of the principles of the invention will be evident from the following description, taken in conjunction with the appended drawings, in which:

FIG. 1 shows in schematic form the overall generic structure of the blood-interfacing portions of the present invention.

FIGS. 2A and 2B show structural details of one embodiment of the ventricular sensor mechanisms.

FIGS. 3 and 4 show the structures of the metering and switching control devices which are responsive to the sensors and which control the power fluid.

FIGS. 5A and 5B shows the interconnected structures of an impeller pump, a flow attenuator, and a pressure referencing sump.

FIGS. 6A and 6B illustrate the structure of one embodiment of a four-port fluid switching valve.

FIG. 7 shows a graph of the performance characteristics of the pump assembly of FIGS. 5A and 5B.

FIG. 8 shows the characteristic performance curves of the overall artificial heart.

FIG. 9 shows a schematic diagram of the interconnected components of one embodiment of the artificial heart assembly.

FIGS. 10A, 10B and 10C show the structural details of an optimally shaped artificial ventricle which will permit the control and energizing mechanism to be attached from outside the thorax.

FIG. 11 shows a schematic diagram of an electrical embodiment of apparatus for controlling the fluid power pulses applied to a pair of ventricles of the type shown in FIG. 10.

FIGS. 12A, 12B, 12C, 12D, 12E and 12F show various details of an artificial ventricle which has a continuously centered and monitored interior bladder, and is particularly suited to the use of a gas for fluid power.

FIG. 13 shows the design of a physiologically responsive artificial heart which does not require a variable attentuation of the flow rate of the circulating power fluid.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

In the subsequent description, it will be assumed that when the artificial heart of this invention is used to replace the biological heart the only portions of the biological heart which remain after excision will be the superior (craniad) remnants of the right and left atrial walls, which can be gathered around and sutured to an appropriate pair of vascular conduits which can carry the returning systemic and pulmonary blood flows into the right and left artificial ventricles, respectively.

Turning now to FIG. 1, which illustrates the basic structure of the blood-interfacing portions of the artificial heart of this invention, it is seen that the fundamental structure is not significantly altered by the particular choice of valves used to insure unidirectional blood flow.

In FIG. 1, the right blood conduit 1 is comprised of an inflow orifice 2, an inflow valve 3, an outflow orifice 4, and outflow valve 5, and a right ventricular side-arm 6. Similarly, the left blood conduit 7 is comprised of an inflow orifice 8, an inflow valve 9, an outflow orifice 10, an outflow valve 11, and a left ventricular side-arm 12.

In FIG. 1, the complete right ventricle 13 is comprised of an interior right-ventricular bladder 14 enveloped by a rigid right ventricular shell 15, which has a right ventricular side-arm aperture 16 and a right power-fluid aperture 17. Similarly, the complete left ventricle 18 is comprised of an interior left ventricular bladder 19 enveloped by a rigid left ventricular shell 20, which has a left-ventricular side-arm aperture 21 and a left power-fluid aperture 22. It is assumed that the physical shapes of the right and left ventricles 13 and 18 are generally axially symmetric in overall form, except for the right and left power-fluid apertures 17 and 22 which extend out of the circumferential surfaces of their respective right and left ventricular shells 15 and 20.

Referring still to FIG. 1 it is seen that the right and left ventricular bladders 14 and 19 are novel in structure in that each has only one moving wall and that this moving wall remains generally perpendicular to the axis of the respective side-arm apertures. Further, it is seen that the right ventricular bladder 14 is provided with a fold-back shoulder 23 which permits simple and positive circumferential sealing to the right ventricular side-arm aperture 16. Similarly, the left ventricular bladder 19 is provided with a fold-back shoulder 26 which permits simple and positive circumferential sealing to the left ventricular side-arm aperture 21. Still further, the moving wall of the right ventricular bladder 14 is provided with a centrally located stiffening disk 24 to which is affixed a linkage-anchoring eyelet 25. The moving wall of the left ventricular bladder 19 is similarly provided with a centrally located stiffening disk 27 to which is affixed a linkage-anchoring eyelet 28.

Turning next to FIGS. 2A and 2B, which show the structural details of the ventricular sensing mechanisms, it is seen that the interior medial wall of the right ventricular shell 15 is fitted with a bearing-support strip 29, which supports a protruding right-ventricular front output shaft 30 and a right-ventricular rear output shaft 31. Similarly, the interior medial wall of the left ventricular shell 20 is fitted with a bearing-support strip 32, which supports a protruding left-ventricular front output shaft 33 and a left-ventricular rear output shaft 34.

In FIGS. 2A and 2B, the right-ventricular front and rear output shafts 30 and 31 are seen to be coupled to the right-ventricular eyelet 25 through the linkage segments 35 and 36, and through the linkage segments 37 and 38, respectively. Similarly, the left-ventricular front and rear output shafts 33 and 34 are seen to be coupled to the left-ventricular eyelet 28 through the linkage segments 39 and 40, and through the linkage segments 41 and 42, respectively.

Referring still to FIGS. 2A and 2B, it is seen that the ventricular sensing mechanisms are designed so that the instantaneous volumes of the right and left ventricular bladders 14 and 19 are converted into equivalent shaft rotations. The arrangement of the various linkages may be conveniently such that the angular rotation of each shaft is confined to a range of plus or minus thirty degrees. Also, it will be seen that the directions of rotation of the front and rear output shafts 30 and 31 of the right ventricle are identical, whereas the front and rear output shafts 33 and 34 of the left ventricle are arranged to rotate in opposite directions.

It will be seen from the illustrations of FIGS. 1, 2A and 2B that the structure of the right and left ventricular bladders 14 and 19 are each of a form designed for smooth interiors and long flexion life. Further, it will be seen that the bladder volume sensing mechanisms are designed so that the front and rear output shafts 30, 31, 33, and 34 of the right and left ventricles 13 and 18 rotate freely and thus do not significantly affect the freedom of motion of the respective moving medial walls of the right and left ventricular bladders 14 and 19.

Turning next to FIG. 3, which shows the structure of the metering control devices, it is seen that the metering control device is comprised of a set of three bell-cranks 43, 44 and 45 which are interconnected by an algebraic summing linkage composed of the four link segments 46, 47, 48, and 49. The first and second bell-cranks 43 and 44 are pinned to the right and left ventricular front output shafts 30 and 33, respectively. The third bell-crank 45 is pinned to an output metering-control shaft 50. The lengths of the bell-cranks and their interconnecting link segments in FIG. 3 can be made so that the output shaft 50 rotates through an angular range of plus or minus 45.degree. even though the angular rotation of each of the input shafts 30 and 33 is restricted to a range of plus or minus 30.degree..

Referring now to FIG. 4, which shows the structure of a switching control device, it is seen that the switching control device is in part comprised of a second set of three bell-cranks 51, 52, and 53 which are interconnected by a second algebraic summing linkage composed of the four link segments 54, 55, 56, and 57. The bell-cranks 51 and 52 are pinned to the right and left ventricular rear output shafts 31 and 34. The bell-crank 53 engages a rotary escapement mechanism 58, which has the function of permitting angular rotation of its output shaft 59 only in the forward direction and only in 90.degree. increments. The stepwise angular rotation of the output shaft 59 is mechanized by the use of two toothed wheels 60 and 61 which are continuously meshed. The toothed wheel 60 is assumed to be under the influence of a continuous forward-acting torque, which may be produced only by any one of a number of suitable devices, such as a small electric motor or a viscous-drag cup emersed in a fluid vortex chamber. The toothed wheel 60 is fitted with a pair of catch-lugs 62 and 63 which are located at opposite ends of a first catch-lug diameter. Similarly, the toothed wheel 61 is fitted with another pair of catch-lugs 64 and 65 which are located at opposite ends of a second catch-lug diameter. The meshed alignment of the first and second toothed wheels 60 and 61 is such that the first catch-lug diameter is horizontal when the second catch-lug diameter is vertical, and vice versa. Further, the bell-crank 53 is shaped so as to engage at one time no more than one of the four catch-lugs 62, 63, 64 and 65. Thus, a back-and-fourth rocking motion of the bell-crank 53 results in a series of 90.degree. unidirectional rotations of the output shaft 59. However, it will be seen that a given quarter-revolution of the output shaft 59 will occur only if the angular movement of the bell-crank 53 exceeds a predetermined threshold.

Thus, it will be seen that the designs of the metering and swtiching control devices illustrated in FIGS. 3 and 4 are such that the angular rotation of the output shaft 50 is approximately proportional to the sum of the instantaneous volumes of the right and left ventricular bladders 14 and 19, and that a given quarter-revolution of the output shaft 59 occurs if and only if the magnitude of the algebraic difference between the instantaneous volumes of the right and left ventricular bladders 14 and 19 exceeds a predetermined magnitude.

Attention may now be directed to FIGS. 5A and 5B, which show the interconnected structures of an impeller pump 66, a flow attenuator 71, and a pressure-referencing sump 68. The power fluid enters the impeller pump 66 through the inlet conduit 69, the pressure-referencing sump 68, and the connection conduit 67. The power fluid leaving the impeller pump 66 flows out of the exit conduit 74 after passing through the connection conduit 70 and the flow attenuator 71.

The pressure-referencing sump 68 has the function of ensuring that the pressure of the power-fluid entering the impeller pump 66 is at all times essentially equal to the ambient atmospheric pressure existing outside the complete power-fluid circuit. The structure of the pressure-referencing sump 68 may thus be simply a segment of pliable tubing of thin rubber or other suitable material.

For reasons which will become apparent later, it will be assumed that the algebraic difference between the maximum volume and the minimum volume of the distensible pressure-referencing sump 68 is at least equal to the sum of the maximum volumes of the right and left ventricular bladders 14 and 19.

As shown in FIGS. 5A and 5B, the flow attenuator 71 may be simply an appropriately housed rotatable vane 72 pinned to a flow-attenuator control shaft 73. The plane of the vane 72 may thus be inclined from the axis of the exit conduit 74 by any desired angle .alpha.. The least attenuation of the outflow of power fluid from the impeller pump 66 occurs when the angle .alpha. is zero. Conversely, the rate of flow of power fluid out of the exit conduit 74 is practically zero when the angle .alpha. is equal to ninety degrees.

In actual operation, the flow-attenuator control shaft 73 in FIG. 5A is made to be a connected extension of the output metering-control shaft 50 in FIG. 3. The direction of rotation of this interconnection is made to be such that the inclination angle .alpha. of the flow-attenuator vane 72 decreases in approximate proportion to the sum of the instantaneous volumes of the right and left ventricular bladders 14 and 19. Thus, the outflow of power fluid from the exit conduit 74 is least attenuated when the right and left ventricular bladders 14 and 19 are both maximally distended. Conversely, the outflow of power fluid from the exit conduit 74 is practically zero when the right and left ventricular bladders 14 and 19 are both maximally compressed.

Turning next to FIGS. 6A and 6B, which show the structure of a suitable four-port fluid-switching valve, it is seen that the structure is essentially that of a rotatable vane centered at the intersection of a pair of perpendicularly crossing power-fluid conduits. The two extending arms 75 and 76 of one of the conduits are fitted with suitable output ports 79 and 80. Similarly, the two extending arms 77 and 78 of the other of the two conduits are fitted with suitable input and return ports 81 and 82.

At the center of the intersection of the two power-fluid conduits in FIGS. 6A and 6B, a pair of holes 83 and 84 are made in the conduit walls to accommodate a switching-control shaft 85 to which a symmetric vane 86 is pinned. The shape of the four-port valve of FIGS. 6A and 6B is such that the vane 86 may be continuously rotated in a given forward direction. Due to the symmetry of the vane 86, it will be seen that a succession of quarter-revolutions of the switching-control shaft 85 in a given forward direction is equivalent to a movement of the vane 86 through an alternating sequence of back-and-forth oscillations of 45.degree. in each direction. Thus, the angle .theta. between the plane of the vane 86 and the axis of the extending conduit arms 75 and 76 may be said to remain within the range of from -45.degree. to +45.degree.degrees, even if the motion of the shaft 85 is a unidirectional sequence of quarter-revolutions. The geometry of the four-port valve of FIGS. 6A and 6B is seen to be such that when .theta. = +45.degree. the input port 81 is connected to the output port 80 and the return port 82 is connected to the output port 79. Conversely, when .theta. = -45.degree. the input port 81 is connected to the output port 79 and the return port 82 is connected to the output port 80.

In actual operation, the four-port valve of FIGS. 6A and 6B is fluid-coupled to the right and left ventricular shells 15 and 20 of FIG. 1 by connecting the output port 80 to right power-fluid aperture 17 and by connecting the output port 79 to the left power-fluid aperture 22. Also, the four-port valve of FIGS. 6A and 6B is fluid-coupled to the sump-impeller-attenuator assembly of FIGS. 5A and 5B by connecting the valve port 81 to the attenuator exit port 74 and by connecting the valve return port 82 to the sump inlet port 69. Further, the valve switching control shaft 85 is made to be an interconnecting extension of the output shaft 59 of the switching control device of FIG. 4, with the rotational alignment being such that a transition from .theta. = -45.degree. to .theta. = +45.degree. can only be triggered when the volume of the right-ventricular bladder 14 is substantially greater than the volume of the left-ventricular bladder 19, and vice versa.

In summary, the artificial heart of this invention has a closed power-fluid circuit around which the power fluid flows, first out of the sump and into the impeller, then out of the impeller and into the attenuator, next out of the attenuator and into the valve, then out of the valve to the ventricular chambers and back to the valve, and finally out of the valve and into the sump. By means of the valve, the circulating power fluid is caused to alternately compress and dilate the two blood-filled bladders encased by the right and left ventricular shells. The valve has only two operating states, the first of which directs the attenuator outflow into the right shell and concurrently directs the left-shell outflow into the sump, and the second of which directs the attenuator outflow into the left shell and concurrently directs the right-shell outflow into the sump. The instantaneous volumes V.sub.1 and V.sub.2 of the right and left ventricular bladders are continuously sensed by mechanical calipers, and their associated sum and difference linkages compute the volume sum V.sub.S = (V.sub.1 + V.sub.2) and the volume difference V.sub.D = (V.sub.1 - V.sub.2). The volume sum V.sub.s controls the attenuator so that greater flow of power fluid is permitted if V.sub.S is large. Through an associated escapement device, the volume difference V.sub.D controls the state of the valve so that a state transition is triggered whenever V.sub.D exceeds a predetermined positive or negative threshold level. The sump has two functions, the first of which is to ensure that the pressure of the power fluid entering the impeller is at all times equal to the ambient external atmospheric pressure, and the second of which is to permit free expansion or contraction of its fluid volume to accommodate moment-to-moment variations in the volume sum V.sub.S.

With the aid of the foregoing summary, and with the aid of the detailed listing of the descriptive nomenclature shown in TABLE I, a mathematical analysis of the performance characteristics of the artificial heart of this invention may now be developed.

TABLE I

DEFINITIONS OF MATHEMATICAL TERMS

Symbol Definition Dimension V.sub.1 Right ventricular bladder volume cm.sup.3 V.sub.2 Left ventricular bladder volume cm.sup.3 V.sub.m Bladder maximum stroke volume cm.sup.3 V.sub.D Bladder volume difference (V.sub.1 - V.sub.2) cm.sup.3 V.sub.S Bladder volume sum (V.sub.1 + V.sub.2) cm.sup.3 V.sub.S.sup.O Initial value of V.sub.S cm.sup.3 P.sub.A Ambient atmospheric pressure P.sub.R Attenuated impeller output pressure P.sub.m Maximum value of boost pressure (P.sub.R - P.sub.A) dynes/cm.sup.2 F.sub.I Attenuated impeller outflow rate cm.sup.3 /sec F.sub.m Maximum value of F.sub.I (for P.sub.R = P.sub.A) cm.sup.3 /sec R.sub.VC Mean systemic bloodflow return rate cm.sup.3 /sec R.sub.PV Mean pulmonary bloodflow return rate cm.sup.3 /sec P.sub.1 Mean pulmonary-artery backpressure P.sub.2 Mean systemic-arterial backpressure T.sub.1 Right-systole time constant sec T.sub.2 Left-systole time constant sec .tau..sub.1 Duration of right systole sec .tau..sub.2 Duration of left systole sec .theta. Angular position of four-port valve vane radians K Ratio of .vertline.V.sub.D /V.sub.m .vertline. which triggers .theta. switch V.sub.S.sup.2 Upper limit of V.sub.S (occurs when .theta..fwdarw.+45.degree.) cm.sup.3 V.sub.S.sup.1 Lower limit of V.sub.S (occurs when .theta..fwdarw.-45.degree.) cm.sup.3 V.sub.R Right-systole volume-sum asymptote cm.sup.3 V.sub.L Left-systole volume-sum asymptote cm.sup.3

A convenient starting point for the mathematical anaylsis is the interrelationship between the pressure P.sub.R and the flow rate F.sub.I of the power fluid emerging from the exit conduit 74 of the sump-impeller-attenuator assembly of FIG. 5A. The boost pressure produced by this assembly is (P.sub.R - P.sub.A) in which P.sub.A is the ambient atmospheric pressure existing in the impeller inlet conduit 67 by reason of the continuous action of the pressure-referencing sump 68. If the attenuator shaft-angle .alpha. is set at .alpha. = 0 for minimum restriction of flow, the relationship between the flow rate F.sub.R and the pressure boost (P.sub.R - P.sub.A) is given by a parabolic formula which generally characterizes all impeller pumps, i.e., (P.sub.R - P.sub.A) = P.sub.m.sup.. [1 - (F.sub.I /F.sub.m).sup.2 ], in which P.sub.m is the maximum boost pressure and F.sub.m is the maximum exit flow rate. Typical values of the parameters P.sub.m and F.sub.m, which are essentially determined by the size and speed of the impeller 66, are P.sub.m = 8 psi = 550,000 dynes/cm.sup.2 and F.sub.m = 550 gph = 580 cm.sup.3 /sec. An impeller pump having these parameters may be obtained commercially as the Model 3-MD chemical magnetic drive pump supplied by the corporation of S. Gelber & Sons of Chicago, Ill.

In the more general case, the attenuator shaft angle .alpha. of the sump-impeller-attenuator assembly of FIG. 5A is determined by the sum V.sub.S of the right and left ventricular bladder volumes V.sub.1 and V.sub.2. It is assumed here that V.sub.1 and V.sub.2 vary within the ranges of 0.ltoreq.V.sub.1 .ltoreq.V.sub.m and 0.ltoreq.V.sub.2 .ltoreq.V.sub.m, in which the upper-limit volume V.sub.m may typically have a value of V.sub.m = 150 cm.sup.3. It is further assumed that when the right and left ventricular bladders 14 and 19 are maximally compressed they will still contain small residual quantities of blood which are not counted in the determinations of V.sub.1 and V.sub.2.

Under the foregoing conditions it will be seen that when the attenuator control shaft 73 of the assembly of FIG. 5A is mechanically coupled to the output shaft 50 of the metering-control device of FIG. 3, the resultant relationship between the exit flow rate F.sub.I and the boost pressure (P.sub.R - P.sub.A) of the assembly of FIG. 5A can be approximated by the formula

(P.sub.R - P.sub.A) = P.sub.m.sup.. [1 -(2V.sub.m /V.sub.S.sup.. F.sub.I /F.sub.m) .sup.2 ]

which is illustrated graphically in FIG. 7 for two different values of V.sub.S. This formula may also be written in the form

F.sub.I = F.sub.m.sup.. V.sub.S /2V.sub.m .sup.. .sqroot.1 - (P.sub.R - P.sub.A /P.sub.m)

which will subsequently be used.

The next step in the mathematical analysis is the development of the appropriate equations for V.sub.1 and V.sub.2 , in which the dot denotes the time rate of change of the pertinent variable. The fluid connections of the four-port valve of FIGS. 6A and 6B with the right and left ventricles of FIG. 1 and with the sump-impeller-attenuator assembly of FIG. 5A and 5B, are such that right-systole and left-diastole occur simultaneously when the vane 86 has the position angle .theta. = +45.degree., and conversely such that left-systole and right-diastole occur simultaneously when .theta. = -45.degree..

During the time interval .tau..sub.1 when .theta. = +45.degree. the attenuator exit pressure P.sub.R is equal to the mean pulmonary-artery back-pressure P.sub.1, and the left ventricular dilation rate V.sub.2 is equal to the mean pulmonary bloodflow return rate R.sub.PV multiplied by the nearly 2.0 ratio of (.tau..sub.1 + .tau..sub.2)/.tau..sub.1, in which .tau..sub.1 and .tau..sub.2 represent the respective durations of right-systole and left-systole. Conversely, during the time interval .tau..sub.2 when .theta. = -45.degree. the attenuator exit pressure P.sub.R is equal to the mean aortic backpressure P.sub.2, and the right ventricular dilation rate V.sub.1 is equal to the mean systemic bloodflow return rate R.sub.VC multiplied by the nearly 2.0 ratio of (.tau..sub.1 + .tau..sub.2 )/.tau..sub.2. Consequently, with R'.sub.PV = R.sub.PV .sup.. (.tau..sub.1 + .tau..sub.2)/.tau..sub.1 and R'.sub.VC = R.sub.VC .sup.. (.tau..sub.1 + .tau..sub.2)/.tau. being the analytically compensated bloodflow return rates, the differential equations governing V.sub.1 and V.sub.2 are found to be

T.sub.1 V.sub.1 + V.sub.1 = -R'.sub.PV and V.sub.2 = R'.sub.PV for .theta. = +45.degree. T.sub.2 V.sub.2 + V.sub.2 = -R'.sub.VC and V.sub.1 = R'.sub.VC for .theta. = -45.degree.

in which T.sub.1 =(2V.sub.m /F.sub.m) .sup.. [1 - (P.sub.1 - P.sub.A )/P.sub.m ].sup.-.sup.0.5 is the right-systole time constant, and in which T.sub.2 = (2V.sub.m /F.sub.m ) .sup.. [1-(P.sub.2 -P.sub.A)/ P.sub.m ].sup.-.sup.0.5 is the left-systole time constant. In practice, T.sub.1 ` 0.5 sec and T.sub.2 .congruent. 0.6 sec.

The final step in the mathematical analysis is development of the equations governing the binary angular position .theta. of the vane 86 in the four-port valve of FIG. 6A and 6B, which by means of the switching control device of FIG. 4 is triggered by the bladder-volume difference V.sub.D = (V.sub.1 - V.sub.2). If K denotes the ratio of .vertline.V.sub.D /V.sub.m .vertline. for which triggering occurs, then the equations governing .theta. may be written as follows.

.theta. = 0 and .theta. = .+-.45 when .vertline.V.sub.D .vertline..noteq.K.sup.. V.sub.m

.theta.>0 only when V.sub.D = +K.sup.. V.sub.m and V.sub.D > 0

.theta.<0 only when V.sub.D = -K.sup.. V.sub.m and V.sub.D < 0

Thus, the switching action for .theta. is that of a symmetric hysteresis switch which responds to the cyclic V.sub.D waveform. A typical value for the dimensionless threshold constant K is K = 0.90.

The foregoing system equations may be solved by the application of standard differential-equation techniques. For example, for 0.ltoreq. t= .tau..sub.1 the valve position angle .theta. has the value .theta. = +45 , and the equations for the right and left ventricular volumes V.sub.1 and V.sub.2 are found to be

V.sub.1 = V.sub.S.sup.2 + KV.sub.m /2 - R'.sub.PV t + (T.sub.1 R'.sub.PV -V.sub.S.sup.2) [1 - Exp(-t/T.sub.1)]

V.sub.2 = V.sub.S.sup.2 - KV.sub.m /2 + R'.sub.VP t in which V.sub.S.sup.2 is the initial upper-limit value of the volume sum V.sub.S = (V.sub.1 + V.sub.2). By direct substitution, it will be found that these equations satisfy the primary relationships T.sub.1 V.sub.1 + V.sub.1 = -R'.sub.PV and V.sub.2 = R'.sub.PV. Further, it will be seen that for t = 0 the volume difference V.sub.D = (V.sub.1 - V.sub.2) has its correctly assumed initial value V.sub.D = +K.sup.. V.sub.m. These equations for V.sub.1 and V.sub.2 will apply until t = .tau..sub.1, at which time the volume difference V.sub.D = (V.sub.1 - V.sub.2) will have been decreased from its initial positive triggering value V.sub.D = +K.sup.. V.sub.m down to its opposite negative triggering value V.sub.D = -K.sup.. V.sub.m. Accordingly, the transcendental equation for the duration .tau..sub.1 of right-systole is

2R'.sub.PV .tau..sub.1 - (T.sub.1 R'.sub.PV - V.sub.S.sup.2) [1 - Exp(-.tau..sub.1 /T.sub.1)] = 2KV.sub.m

It should be noted here that the expression R'.sub.PV = R.sub.PV.sup.. (.tau..sub.1 +.tau..sub.2 )/.tau..sub.1 requires that a similar transcendental equation for .tau..sub.2 be specified.

Starting with a new origin, with .theta. = -45.degree. in the new time interval 0.ltoreq. t.ltoreq. .tau..sub.2, the equations for V.sub.1 and V.sub.2 are found to be

V.sub.1 = V.sub.S.sup.1 -KV.sub.m /2 + R'.sub.VC t

V.sub.2 = V.sub.S.sup.1 +KV.sub.m /2 - R'.sub.VC t + (T.sub.2 R'.sub.VC - V.sub.S.sup.1) [1 - Exp(-t/T.sub.2)]

in which V.sub.S.sup.1 is the new initial lower-limit value of the volume sum V.sub.S. By direct substitution, it will be found that these equations satisfy the primary relationships T.sub.2 V.sub.2 +V.sub.2 = -R'.sub.VC and V.sub.1 = R'.sub.VC. Further, it will be seen that for t = 0 the volume difference V.sub.D has its correctly assumed initial value V.sub.D = -K.sup.. V.sub.m. These new equations for V.sub.1 and V.sub.2 will apply until t = .tau..sub.2, at which time the volume difference V.sub.D = (V.sub.1 - V.sub.2) will have been increased from its initial negative triggering value V.sub.D =-K.sup.. V.sub.m up to its opposite positive triggering value V.sub.D = -K.sup.. V.sub.m. Accordingly, the transcendental equation for the duration .tau..sub.2 of left-systole is

2R'.sub.VC .tau..sub.2 - (T.sub.2 R'.sub.VC - V.sub.S.sup.1) [1 - Exp(-.tau..sub.2 /T.sub.2)] = 2KV.sub.m By use of the auxiliary relationships R'.sub.PV = R.sub.PV .sup.. (.tau..sub.1 + .tau..sub.2)/.tau..sub.1 and R'.sub.VC = R.sub.VC.sup.. (.tau..sub.1 + .tau..sub.2)/.tau..sub.2, the above two transcendental equations may be solved graphically for .tau..sub.1 and .tau..sub.2 in terms of the volume-sum limit values V.sub.S.sup.1 and V.sub.S.sup.2.

It will be noted, in the above two pairs of equations for V.sub.1 and V.sub.2, that the end-point values in the first pair of equations must be the starting-point values in the second pair of equations. Consequently, the equations resulting from the matching of the end-point and starting-point values for the volume sum V.sub.S are

V.sub.S.sup.1 = V.sub.S.sup.2 + (T.sub.1 R'.sub.PV - V.sub.S.sup.2) [1 - Exp(-.tau..sub.1 /T.sub.1)]

V.sub.S.sup.2 = V.sub.S.sup.1 + (T.sub.2 R'.sub.VC - V.sub.S.sup.1) [1 - Exp(-.tau..sub.2 /T.sub.2)] These last two equations, when used together with the above two transcendental equations, permit a complete graphical solution for the four unknowns .tau..sub.1, .tau..sub.2, V.sub.S.sup.1, and V.sub.S.sup.2 in terms of the installed system parameters R.sub.VC, R.sub.PV, T.sub.1, I.sub.2, K, and V.sub.m.

From the foregoing analysis it is seen that the volume sum V.sub.S = (V.sub.1 + V.sub.2) is represented in the respective time intervals 0.ltoreq. t= .tau..sub.1 and 0.ltoreq. t .ltoreq. .tau..sub.2 by the two equations,

V.sub.S = V.sub.S.sup.2 + (T.sub.1 R'.sub.PV - V.sub.S.sup.2) [1 - Exp(-t/T.sub.1)] for .theta. = +45.degree.

V.sub.S = V.sub.S.sup.1 + (T.sub.2 R'.sub.VC - V.sub.S.sup.1) [1 - Exp(-t/T.sub.2)] for .theta. = -45.degree.

In the first of these expressions, it is seen that V.sub.S begins with the value V.sub.S.sup.2 and decreases exponentially toward the right-systole volume-sum asymptote V.sub.R = T.sub.1.sup.. R'.sub.PV. Similarly, in the second of these expressions it is seen that V.sub.S begins with the value V.sub.S.sup.1 and increases exponentially toward the left-systole volume-sum asymptote V.sub.L = T.sub.2.sup.. R'.sub.VC. A typical V.sub.S waveform is shown in FIG. 8, together with its corresponding .theta., V.sub.D , and V.sub.D waveforms. The shape of the V.sub.D waveform approaches that of a triangular wave when P.sub.m is large compared to either P.sub.1 or P.sub.2, in which case T.sub.1 and T.sub.2 will be nearly equal. Through the included cardiovascular parameters R'.sub.PV, R'.sub.VC, P.sub.1, and P.sub.2, the above system equations are seen to be properly responsive to the pulmonary and systemic bloodflow return rates, and to the pulmonary-artery and systemic-arterial backpressures.

FIG. 9 represents, in very schematic form, a complete system made up of the parts described separately in FIGS. 1, 2A, 2B, 3, 4, 5A, 5B, 6A and 6B.

In the upper part of the figure, the right and left ventricular bladders 14, 19 are connected respectively to arteries and veins as shown in FIG. 1. The sensors 88 and 89 provide signals, which go by means 88', 89', to control means 90A, 90B, respectively. The sensor signal in 88' is a function of V1, the volume of blood in the bladder 14, and sensor signal in 89' is a function of V2 the volume of blood in bladder 19. The control means 90A takes the two sensor signals and computes a function of (V1 + V2). Control means 90B takes the two sensor signals and computes a function of (V1-V2).

In the bottom of the figure is the power fluid system to drive the ventricular bladders. It comprises a sump 68, which provides fluid through line 67 to pump 66, the output of which goes by line 70 to a flow control 71. Here the total flow is controlled by a vane 72, the angle .alpha. of which is a measure of the flow constriction. Means 91 which carries a control which is a function of (V1 + V2) controls the angle .alpha. of the vane 72.

The output of the fluid flow controller 71 goes by means 74 to a double reversing valve 95. An internal vane (indicated by 86) by its angle .theta., which varies from +45.degree. to -45.degree., switches the flow entering 81 to either output 80 which goes to chamber No. 1, or to output 79 which goes to chamber No. 2. With the vane angle as shown, the flow of fluid in line 74 goes from 81 to 80 and via line 93 to ventricular chamber No. 1, and compresses the bladder 14, driving blood out through 6 to the pulmonary artery. At the same time, blood is flowing into bladder 19 from the pulmonary veins, extending it, and driving power fluid out of chamber No. 2 through line 94, into 79 and out of 82 to the sump 68. This closes the circuit on the power fluid flow. The position of vane 86 is controlled from control means 90B through means 92 with a signal which is a function of (V1 - V2).

In certain medical situations it may be desirable to keep to an absolute minimum the size and weight of any prosthetic device to be surgically installed within the chest cavity to support the failing circulation. In other situations it may be desirable to leave a possibly recovering natural heart intact and simply use the artificial heart on a temporary or intermittent basis to bypass the right and left biological ventricles. In such cases it is important that each artificial ventricle be structured with particular regard to the anatomical constraints imposed by the shape of the rib cage, the mass of displaceable lung tissue, and the relative positions of the great vessels leading to and from the natural heart. FIGS. 10A, 10B and 10C show the design details of an artificial ventricle structure with these considerations as paramount but including the vital feature of a permanently stable bladder-volume sensor.

In FIGS. 10A and 10B it is seen that the pliable bladder 101 has the general shape of a flattened ellipsoid of revolution having two nearly circular walls 102 and 103, and having two blood ports 104 and 105. One of the circular walls, 102, remains stationary and the other wall, 103, moves back and forth in a direction parallel to its axis in response to externally applied pressure pulses. The moving wall 103 has a centrally located stiffening button 106 to which is affixed a linkage-anchoring eyelet 107.

Referring still to FIGS. 10A and 10B, in order to ensure long flexural life, the ellipsoidal bladder 101 is generally shaped so that the distance x between the stationary wall 102 and the moving wall 103 varies approximately within the range of 0.050d.ltoreq. .times. .ltoreq. 0.275d, where d is the large diameter of the generating ellipsoid. For the same reason, the diameter s of the stiffening button 106 forming the central portion of the movable wall 103 is made no greater than approximately s = 0.30d.

Finally, in FIGS. 10A and 10B, the two blood ports 104 and 105 are each designed to couple the main bladder with a blood-conduit tube having a circular-cross-section diameter w of approximately w = 0.30d. The axis of each of these two blood-conduit tubes is inclined through an angle .phi. of approximately .phi. = 33.7 degrees away from the external axis of the stationary bladder-wall 102, and the extension of each blood-conduit axis intersects the midplane of the fully dilated bladder at a distance u which is approximately u = 0.40d away from the rotational axis of the generating ellipsoid. A smooth flaring-and-flattening transition characterizes the actual junction of each blood-conduit tube with the bladder 101. The two blood-conduit axes need not intersect the midplane of the dilated bladder on the same large-diameter, as is actually shown for convenience of illustration. For the normal 70 kilogram man, the large-diameter d may have a typical value of d = 5 inches.

FIG. 10C shows the pliable bladder 101 in relation to its enveloping rigid shell 108. The shell 108 has an interior contour which closely fits around the dilated bladder 101 except over the movable wall 103, in which region the shell 108 evolves through a smooth and gradually separating transition into a relatively slender hollow stem 109 whose axis coincides with the axis of the movable wall 103. The interior diameter q of the hollow stem 109 may be approximately q = 0.10d, where d is the large diameter of the ellipsoidal bladder 101.

By means of a lock-pin 110 a linkage bar 111 is connected to the eyelet 107 so as to extend within the hollow stem 109 in a direction away from the movable bladder-wall 103. The length r of the linkage bar 111 is made to be at least equal to r = 0.3d, so that its outermost end is always constrained to move within the hollow stem 109 irrespective of the position of the movable bladder-wall 103. Foldback shoulders 112 and 113 extending around the peripheries of the blood ports 104 and 105 permit simple and positive circumferential sealing to their respective shell ports 114 and 115. Circumferentially applied adhesive sealants 116 and 117 are used to firmly retain the bladder 101 within the rigid shell 108.

The length of the hollow stem 109 is made to be sufficient to extend through a surgically opened orifice between two adjacent ribs in the neighboring chest wall. Thus, the complete ventricle comprised of the pliable bladder 101, its enveloping rigid shell 108, and its associated linkage bar 111 can be surgically affixed to the interior chest wall, with external connections to the necessary fluid-power control apparatus being readily made through or under the skin of the external chest wall. A pair of suitable uni-directional valves can be positioned in the two blood conduit tubes used to connect the major blood vessels to the artificial ventricle thus comprised.

Attention may now be directed to the artificial heart system of FIG. 11, which makes use of two artificial ventricles having the optimal shape described above, and which utilizes electrical components to generate the fluid-control functions previously summarized from a mechanical viewpoint in FIG. 9.

In FIG. 11, each of the two ventricles 118 and 119 is mechanically connected, by means of its respective linkage-bar 118', 119' to a corresponding one of the two linear-motion transducers 120 and 121, which are of conventional design. The transducers 120 and 121 thus produce the voltages E.sub.R and E.sub.L which are approximately proportional to the instantaneous right and left ventricular bladder volumes, respectively. These two voltages are used as the signal inputs to a voltage summer 122 and a voltage subtractor 123, which in turn produce the output voltages (E.sub.R + E.sub.L) and (E.sub.R - E.sub.L), respectively. By means of a motor-voltage amplifier 124 and a variable speed motor 126, the output voltage (E.sub.R + E.sub.L) of the summer 122 is used to proportionally control the output flow rate F of a conventional positive-displacement pump 128. Correspondingly, by means of a conventional bistable switch 125, and reversible torque unit 127, the output voltage (E.sub.R - E.sub.L) is used to control the state of a four-port valve 129 of the type previously shown in FIGS. 6A and 6B.

In FIG. 11 the fluid connections between the valve 129, the pump 128, and the two ventricles 118 and 119 are similar to those already described in connection with FIG. 9. Thus, in normal operation the respective bladders in the two ventricles 118 and 119 are alternately dilated by inflowing blood and alternately compressed by fluid-power pulses. Since the output flow rate F of a positive-displacement pump 128 is essentially independent of its loading backpressure, the durations of left and right systole are essentially unaffected by the aortic and pulmonary-artery backpressures. Consequently, the system of FIG. 11 has the advantage of physiologically responsive, but generally equal, right and left systolic durations, as well as the flexibility of permitting extrathoracic location of all the fluid-control components.

It is readily apparent, that by operating both of the ventricles in parallel, either of the two artificial heart systems illustrated by FIG. 9 and FIG. 11 can be used as the primary blood pump for extracorporeal circulation during open-heart surgery or for circulatory support during a temporary myocardial insufficiency. For such use on a routine basis a convenient and obvious modification is to construct each ventricle so that its inlet and exit blood ports extend from the ventricle in a generally upward direction. A further and more fundamental modification, which is particularly useful when a gas rather than a liquid is employed as the power fluid, is the inclusion of structural means to assure constant centering and symmetry of the pliable bladder within its rigid-shell housing. The ventricle illustrated in FIGS. 12A through 12F is designed to meet these modification requirements.

In FIGS. 12A and 12B, the pliable bladder 130 is seen to have the general shape of a flattened ellipsoid of revolution 131 with two upward-directed inlet and exit blood ports 132 and 133. Each of the two-flattened sides 134 and 135 of the ellipsoid 131 moves in a direction parallel to its axis during normal pumping action. Attached to the two flattened sides 134 and 135, respectively, are centrally located stiffening buttons 136 and 137 to which two linkage-anchoring eyelets 138 and 139 are respectively affixed. Around the lower periphery of the ellipsoidal bladder 131 are four equally spaced retainer tabs 140, 141, 142, and 143, which are seen to be attached to the bladder 131 in such a way as to permit unobstructed motion of the moving bladder walls 134 and 135.

FIGS. 12C and 12D show the structure of a rigid interior frame 144 which holds the bladder 130 firmly in a fixed plane of symmetry. The frame 144 is comprised, in part, by two necks 145 and 146 which encase the two blood ports 132 and 133 of the pliable bladder 130. The frame 144 is comprised in further part by a thin band 147 which extends out of a smooth transition from the neck 145, around the lower periphery of the bladder 130, and back through another smooth transition into the neck 146. Four rectangular detours 148, 149, 150, and 151 are indented in the band 147 so as to make it pass around in close proximity with the four retainer tabs 140, 141, 142, and 143 of the bladder 130. Four shaft-pins 152, 153, 154 and 155 are passed through corresponding holes in the four detours 148, 149, 150, and 151 and through the four retainer tabs 140, 141, 142, and 143, in order to firmly secure the bladder 130 to the frame 144. Affixed to the lower portion of the band 147 is an offset bearing-support strip 156, which is arranged to support the upper end of a downward extending shaft 157.

FIG. 12E shows a linkage mechanism which ensures that the angular rotation of the shaft 157 is a measure of the instantaneous volume of the bladder 130, and which further ensures that the moving walls 134 and 135 of the bladder 130 will move in opposite directions by equal amounts at all times. The shaft 157 is supported at its upper end by the previously mentioned bearing-support strip 156, and is supported at its lower end by a bearing not shown in FIG. 12E for reasons of clarity. Except for the rotational torque applied through the attached cross-arm 158, the shaft 157 is assumed to rotate freely. In operation, the centerline of the cross-arm 158 is permitted to swing back and forth through an angle of approximately .+-. 45.degree. on either side of the plane of the previously mentioned four shaft-pins 152, 153, 154, and 155 which hold the bladder 130 in place. The right end of the cross-arm 158 is coupled, through a first stub-link 159, a first swing-lever 160, a first drive-shaft 161, and a first arch-lever 162, to the previously mentioned eyelet 139. Similarly, the left end of the cross-arm 158 is coupled, through a second stub-link 163, a second swing-lever 164, a second drive-shaft 165, and a second arch-lever 166, to the previously mentioned eyelet 138. By means of these coupling arrangements the movable bladder walls 134 and 135 are always constrained to move in opposite directions by equal amounts. Further, it will be seen that the angle through which the shaft 157 is rotated is a measure of the instantaneous volume of the bladder 130.

FIG. 12F shows the various components of FIGS. 12A, 12B, 12C, 12D and 12E, assembled in their respective functional positions within an enveloping rigid shell 167. The shell 167 is large enough to permit essentially unobstructed flow of power fluid around the compressible portion of the pliable bladder 130. The inflow and exit tubes 168 and 169 are fitted closely around the frame necks 145 and 146, which in turn fit closely around the blood circuit tubes 132 and 133. The appropriate sealants and fold-back shoulders for the blood conduit tubes 132 and 133, as well as the necessary support bearings for the drive-shafts 161 and 165, are omitted from FIG. 12F for reasons of clarity. The power-fluid orifice 170 for the shell 167 is arranged coaxially around the volume-monitoring shaft 157 so as to avoid the need for a fourth shell-orifice. Naturally, the entire assembly shown in FIG. 12F can be constructed in a much more compact form in actual practice.

In re-examining the artificial-heart systems of FIGS. 9 and 11, it will be seen that both incorporate power-dissipative elements in order to control the rate of flow of power fluid through the pump. A careful study of the duty-cycle variability in the normal heart reveals that the need for variable attenuation of the power-fluid flow in an artificial heart can be eliminated in those situations in which only a moderate range of induced bloodflow rates is required. Thus, by using a constant-flow pump together with special flow control logic, a physiologically responsive artificial heart can be constructed without requiring significant power dissipation. The system of FIG. 13 is designed to meet this objective.

The system of FIG. 13 makes use of two artificial ventricles 201 and 202 of the type shown in FIG. 10. The two ventricles 201 and 202 are fitted with a corresponding pair of bladder-volume sensors 203 and 204, which respectively produce the output voltages E.sub.R and E.sub.L as in FIG. 11. A constant-flow pump 205, such as a positive-displacement pump operated at constant speed, is used to generate the fluid power. The electrically operated four-port valve 206 is substantially equivalent to the two components 127 and 129 of FIG. 11, except that the downstream portion of the fluid-switching vane is somewhat elongated so that in the absence of electrical excitation, the vane assumes a neutral bypass position. The load isolator 207, consisting of a set of four emitter followers, permits the valve 206 to be operated in the right, neutral, or left position in response to a small rightward, zero, or leftward current through the resistance R1.

In FIG. 13, the voltage applied to the resistance R1 is the output of an inhibit-gate 208, which receives the four binary input signal voltages V1, V2, V3, and V4 from the logic-module 209. The logic-module 209 receives the two analog input signal voltages E.sub.R and E.sub.L from the bladder volume sensors 203 and 204. The inhibit-gate 208 is comprised of a simple arrangement of four switching transistors which are controlled by the right and left outputs of a first bistable element 210, which in turn is controlled by the binary set and reset voltages V3 and V4. The current controlled by the four switching transistors in the inhibit-gate 208 is that associated with the applied binary voltages V1 and V2. The binary voltages V1 and V2 are respectively the left and right outputs of a second bistable element 211 contained within the logic-module 209. The second bistable element 211 is controlled by the binary set and reset voltages E.sub.R ' and E.sub.L '.

The inter-connections between the inhibit-gate 208 and the resistance R1, and between the logic-module 209 and the inhibit-gate 208, are such that right systole occurs whenever the first bistable element 210 is set into its active state by the presence of a positive voltage V3, and the second bistable element 211 is set by the presence of a positive voltage E.sub.R '. Neither right nor left systole will occur if the first bistable element 210 is reset into its inactive state by the presence of a positive voltage V4. Left systole will occur whenever the first bistable element 210 is set into its active state by the presence of a positive voltage V3, and the second bistable element 211 is reset by application of a positive voltage E.sub.L ' . A pair of threshhold limit snaps 212 and 213 in the logic-module 209 are used to convert the analog input voltages E.sub.R and E.sub.L into their corresponding quantized forms E.sub.R ' and E.sub.L ' . Also included in the logic-module 209 is a pair of sign inverters 214 and 215 which reverse the polarities of the quantized voltages E.sub.R ' and E.sub.L ' . The diodes D1, D2, D3, D4, D5, and D6 ensure the correct unidirectional signal pathways in the logic-module 209.

The inter-connections of the various logic components in FIG. 13 are seen to be such that right systole will be initiated by a fully dilated right ventricular bladder, and will be terminated when this bladder is compressed to a predetermined minimum volume. A similar condition applies for the left ventricle. When neither ventricle is fully dilated, the fluid switching valve 206 remains in its neutral position. It will be noted that without further modification, the system shown in FIG. 13 may operate in a gallop rhythm, since the duration of the "half time" extending from the beginning of right systole to the beginning of left systole may not necessarily be equal to the duration of the "half time" extending from the beginning of left systole to the beginning of right systole. Automatic self correction of this gallop rhythm to a balanced symmetric rhythm is easily provided is desired by the addition of a simple feedback pathway which differentially controls the two low-thresholds of the two limit snaps 212 and 213 with a signal derived from the algebraic difference between the two "half time" durations.

While in FIG. 13 and elsewhere in the description the valve means is indicated as .+-.45.degree. valve, it will be clear that other types (such as the conventional cylinder valve) can be used.

While the invention has been described with a certain degree of particularlity, it is manifest that many changes may be made in the details of construction and the arrangement of components. It is understood that the invention is not to be limited to the specific embodiment set forth herein, by way of exemplifying the invention, but the invention is to be limited only by the scope of the attached claim or claims, including the full range of equivalency to which each element or step thereof is entitled.

Claims

What is claimed:

1. In an artificial heart including a right ventricular bladder and a left ventricular bladder and means adapted to connect said bladders into the blood system of the body in which said heart is to be implanted and including a right ventricular chamber enclosing said right bladder and a left ventricular chamber enclosing said left bladder; the improvement comprising:

a. right sensor means to measure a function of the volume V.sub.1 of blood in said right bladder, and left sensor means to measure a function of the volume V.sub.2 of blood in said left bladder;

b. power fluid means; and

c. control means responsive to a function of (V.sub.1 - V.sub.2) for switching the flow of power fluid into one of said chambers, and simultaneously out of the other of said chambers.

2. The artificial heart as in claim 1 including:

control means responsive to a function of (V.sub.1 + V.sub.2) to control the total flow of said power fluid.

3. The artificial heart as in claim 1 in which said power fluid means includes sump means maintained at a pressure equal to the ambient atmospheric pressure, and pump means.

4. The artificial heart as in claim 3 in which said pump means is an impeller pump.

5. The artificial heart as in claim 3 in which said pump is a positive displacement pump.

6. The artificial heart as in claim 1 in which said right and left sensor means comprise displacement sensor means respectively responsive to a function of the instantaneous volumes of said right and left ventricular bladders.

7. The artificial heart as in claim 6 in which said sensor means are mechanical means.

8. The artificial heart as in claim 2 in which the control of the total flow of said power fluid is by means of a flow attenuator.

9. The artificial heart as in claim 2 in which said total flow control is by means of a variable speed motor which drives said pump.

10. The artificial heart as in claim 1 in which said ventricular bladder has only one moving wall, and said sensor means measures a function of the displacement of the bladder wall with respect to the chamber wall.

11. The artificial heart as in claim 1 in which said ventricular bladders each comprise a flattened spheroid of revolution, and including means to ensure that both walls move symmetrically as blood flows into and out of said bladders.

12. The artificial heart as in claim 3 in which said means for switching the flow of power fluid comprises a four-port valve means in which one port is connected to said pump means, one port to said sump means, one port to said right ventricular chamber and one port to said left ventricular chamber.

13. In an artificial heart including a right ventricular bladder and a left ventricular bladder and means adapted to connect said bladders into the blood system of the body in which said heart is to be implanted, and including a right ventricular chamber enclosing said right bladder and a left ventricular chamber enclosing said left bladder, the improvement comprising:

a. right mechanical sensor means to measure a function of the volume V.sub.1 of blood in said right bladder, and means to convert the indications of said right mechanical sensor to electrical signals, E.sub.R ;

b. left mechanical sensor means to measure a function of the volume V.sub.2 of blood in said left bladder, and means to convert the indications of said left mechanical sensor to electrical signals, E.sub.L ;

c. fluid power means including sump means and positive displacement pump means;

d. variable speed motor means responsive to a function of E.sub.R + E.sub.L and means to couple said motor to said pump; and

e. four-port valve means responsive to a function of E.sub.R - E.sub.L to control the flow of power fluid into one of said chambers, and simultaneously out of the other of said chambers.

14. The artificial heart as in claim 13 in which said four-port valve means comprises a two-position valve, which in a first position passes power fluid into a first chamber and out of a second chamber, and in a second position passes power fluid into the second chamber and out of the first chamber.

15. In an artificial heart including a right ventricular bladder and a left ventricular bladder and means adapted to connect said bladders into the blood system of the body in which said heart is to be implanted, and including a right ventricular chamber enclosing said right bladder and a left ventricular chamber enclosing said left bladder, the improvement comprising:

a. right mechanical sensor means to measure a function of the volume V.sub.1 of blood in said right bladder, and means to convert the indications of said right mechanical sensor to electrical signals, E.sub.R ;

b. left mechanical sensor means to measure a function of the volume V.sub.2 of blood in said left bladder, and means to convert the indications of said left mechanical sensor to electrical signals, E.sub.L ;

c. fluid power means including sump means and constant flow pump means;

d. four-port valve means to control the flow of power fluid into one of said chambers, and simultaneously out of the other of said chambers; and

e. control means responsive to a function of E.sub.R and E.sub.L to control said valve means.

16. An artificial ventricle adapted to fit the interior contour of the rib cage, said ventricle being comprised of a blood conduit and an encasement shell, said blood conduit being a pliable tube of axially varying cross-section of a size to extend around either the right or left lung from a point near the sternum to a point near the spine, the two ends of said blood conduit being valve-fitted cylindrical segments which extend in generally opposite directions from two smoothly flared transitions emerging from a pliable-flattened-discoid chamber whose axis of revolution is adapted to lie in the plane of the nearby rib arch, said encasement shell being of a rigid form which envelops the discoid-chamber and flared-transition portions of said blood conduit, said shell also fitted with a hollow stem which is aligned coaxially with said discoid-chamber axis and which is adapted to extend outwardly through a surgical opening between two adjacent ribs to couple subcutaneously with a suitable fluid-power source, the medial wall of said discoid bladder-chamber being adhesively affixed to the adjacent medial wall of said shell chamber, the lateral wall of said discoid bladder-chamber being movable and fitted with a central stiffening button to which is coupled a thin rod extending distally outward through said power-fluid stem to permit measurement of the instantaneous volume of said bladder chamber by means of a suitably attached linear-motion transducer.

17. A gas-operable volume-monitored artificial heart comprised of a pliable bladder, a bladder-centering framework, a symmetric caliper, and a rigid encasement shell, said bladder being comprised of a discoidally shaped chamber having two flat-disc walls from which a pair of valve-fitted inlet and outlet conduits of cylindrical cross section extend in parallel in the upward direction, said bladder chamber having several anchoring tabs attached around its semi-circular periphery, the two flat-disk walls of said bladder chamber each being fitted with a central stiffening button, said shell being generally contoured to envelop said bladder chamber and shaped so as to have a downward-extending power-fluid stem as well as a pair of upward-extending orifice-necks which fit around said bladder inlet and outlet conduits, said framework being attached both to said bladder anchoring tabs and to the interior of said enveloping shell so as to keep said bladder centered at all times within said shell, said caliper having its two arms pinned to eyelets attached to said bladder-wall stiffening buttons and having a mechanical structure which assures that said two bladder walls move in opposite directions by equal amounts, said caliper being coupled to a rotatable shaft aligned coaxially within said power-fluid stem, the rotation of said shaft being a measure of the instantaneous volume of said bladder chamber.