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Publication numberUS3208448 A
Publication typeGrant
Publication dateSep 28, 1965
Filing dateFeb 2, 1962
Priority dateFeb 2, 1962
Publication numberUS 3208448 A, US 3208448A, US-A-3208448, US3208448 A, US3208448A
InventorsKenneth E Woodward
Original AssigneeKenneth E Woodward
Export CitationBiBTeX, EndNote, RefMan
External Links: USPTO, USPTO Assignment, Espacenet
Artificial heart pump circulation system
US 3208448 A
Images(6)
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Description  (OCR text may contain errors)

p 28, 1965 K. E. wooDwAR-D 3,208,448

ARTIFICIAL HEART PUMP CIRCULATION SYSTEM Filed Feb. 2, 1962 6 Sheets-Sheet l T l l T I 3K 4l 1 42 PULMONARY VEIN SUPERIOR VENA CAVA PULMONARY ARTERY AORTA 512 5| 2 28 52 k FLUID AMPLIFIER 2| I2 F e. 1 j

3 PRESSURE x 35 souRcE 59 ARTIFICIAL HEART l0 1 INVENTOR 4 F G. 3 KENNETH E. WOODWARD 5 W BY if 521%;

Sept. 28, 1965 Filed Feb. 2, 1962 K. E. wooDwARD ARTIFICIAL HEART PUMP CIRCULATION SYSTEM 6 Sheets-Sheet 2 FIG.6

' 53 RESISTANCE p 1965 K. E. WOODWARD 3,208,448

ARTIFICIAL HEART PUMP CIRCULATION SYSTEM Filed Feb. 2 1962 6 Sheets-Sheet 3 FIG.?

INPUT POWER Sept. 28, 1965 K. E. WOODWARD ARTIFICIAL HEART PUMP CIRCULATION SYSTEM 6 Sheets-Sheet 4 Filed Feb. 2, 1962 Sept. 28, 1965 K. E. WOODWARD ARTIFICIAL HEART PUMP CIRCULATION SYSTEM 6 Sheets-Sheet 5 Filed Feb. 2. 1962 FIG.9

MEAN BLOOD PRESSURE (mm Hg) 3,208,448 Sept. 28, 1965 K. E. WOODWARD ARTIFICIAL HEART PUMP CIRCULATION SYSTEM Filed Feb. 2, 1962 6 Sheets-Sheet 6 United States Patent The invention described herein may be manufactured and used by or for the Government for governmental purposes without the payment to me of any royalty thereon.

This invention relates to an artificial heart pump system hereinafter referred to as an artificial heart and more particularly to a fluid amplifier controlled artificial heart pump system.

The advent of open heart surgery has presented to the medical profession the opportunity of repairing damaged or diseased hearts of individuals who without such correction face premature death. Many devices are involved in this type of surgery. One important apparatus is a pump can assume the hearts role of pumping blood while the heart is emptied and possibly stopped during repair. Existing pumps are fairly complicated both in design and in the methods of control. These prior pumps included moving mechanical parts, often elaborate electronic equipment and were expensive. Also, the prior pumps failed to satisfy the main physiological and psychological requirements thereon.

Fluid amplification recently invented by a team of Army scientists offers the possibility of extreme design and control simplification for pulsable type extracorporeal heart pumps and a more accurate duplication of the hearts physiological pumping functions. Because all moving control parts and electronics can be eliminated, reliability can be measurably improved and production costs can be significantly reduced.

The requirements for any extracorporeal heart pump may be broken down into three major categories. The pump must:

(1) Duplicate the hearts essential pumping functions (2) Possess adequate reliability and life (3) Be appropriately packaged Each of these will be considered in detail:

Physiological requirements: If the pump is to function in lieu of the subjects own heart for a necessary period, the pump must possess enough of the essential functional pumping characteristics of the human heart to sustain life. Further irreversible blood damage can not be tolerated either during or after the pumping run. From a cursory examination of the literature and prior pump designs it would appear that a single set of functional requirements for extracorporeal pumps in general has not been established. In developing the fluid amplifier pump of this invention the following heart functions are considered to be important:

Output pressures and flows: Certainly the pump must be capable of adequate perfusion of the subject. Moreover it would seem desirable to have a single pump satisfy the needs of a relatively large range of subjects regardless of age or size. The pertinent cardiovascular pump parameters are pressure, flow, age, size, pulse and activity level. (The female imposes generally less severe requirements on the pump.) For surgical applications only resting conditions are of concern.

Load-matching capabilities: it is believed that the human heart recognizes the flow resistances of the cardiovascular circulations and exerts just enough myocardial force per stroke to achieve the necessary blood flows. Excess force causes unnecessary increases in hemolysis.

'To control hemolysis the pump should be designed to propel blood with minimal force application. It is believed that myocardial heart forces are minimal and tightly bounded for large variations in blood flow. Further myocardial energies are minimal for these same flows. Excess force causes unnecessary increases in hemolysis. To control hemolysis and to function in a compatible physiological manner with the heart, the pump must produce a required dynamic outflow.

Pulsatile blood flows: Disagreement exists among cardiovascular experts relative to the need for pulsatile' blood flows for extracorporeal pumps. Both pulsing and steady flow pumps are presently operating successfully for short periodsone to two hours. It may be, however, that long term perfusion requires pulsed flows. A suggestion is given later in this specification that resonant principles may operate in the cardiovascular system and may, among other things, enhance capillary perfusion. If so, the pump must then be capable of duplicating the subjects normal pulse to exploit the elastic responses of the circulatory systems.

Vasoconstrictive considerations: The heart is allowed to be a relatively constant force pump because of vessel distensibility and the Vasoconstrictive characteristics of the cardiovascular system. Flow resistances are made to vary autonomously by the nervous system to suit the physical and emotional needs of the individual. To minimize the pump forces propelling the blood, and to allow the pump to regulate its output in response to Vasoconstrictive changes in flow resistance, output flows must decrease with increased flow resistance and vice versa. To preclude packing or depletion of the circulations the output flows must also be made to increase with increased filling pressures as in the case of the human heart.

Other functional requirements: Filling of the human heart ventricle is achieved mainly by the difference between atrial and ventricular pressures during the period the difference is maximal. However atrial contraction plays a small but significant part in the filling process and serves to increase the flow rate of the heart. If very high flows are required or if cannulae resistances are significant, it might be desirable to cause the pump to suck slightly in diastole. Additionally if the pump is to be used as an augmentation device for failing heart patients, the pump should probably be capable of synchronizing its pulses with that of the heart. By pumping blood during diastole, both peak myocardial forces and the work of the heart are diminished. By pumping blood during systole, the mean blood pressure can be increased to aid in tissue and organ perfusion in those situations where the heart is not failing but cannot pump enough to meet tissue and organ needs. I

Apart from the contributions to hemolysis caused by excess propelling forces, the pump should be as low in hemolysis as the best available commercial pump.

Reliability and life considerations: High reliability is obviously necessary. Reliability should be measurable, predictable and should accommodate operating conditions. In the event of power failure, the pump should be designed to operate on stored energy and should require minimum maintenance.

Packaging: The following design attributes would be generally desirable for the pump. It should be: Uncomplicated to operate and maintain, controllable with respect to pulse and flow, and fabricated out of transparent materials to allow visual observation of performance. Impending failures can often be observed in time for proper corrections of such failures. The pump should be sterilizable preferrably by autoclave methods, functional with negligible heat liberation, fabricated out of materials acceptable to the blood, easily transportable, scalable to preclude blood contamination and air embolism, inexpensive to manufacture, and operable with low audible noise levels,

Prior extracorporeal pumps may be broadly classified with respect to the type of flows they produce, i.e., pulsatile or essentially non-pulsatile. Non-pulsatile pumps are the more uniform in design. They usually consist of a plastic or rubber tube or sleeve wrapped around or within a non-flexing cylinder. A roller driven by an electric motor squeezes the tube as it rolls around the cylinder. Blood is forced ahead of the roller. This type of pump has the following basic disadvantages: a relatively short ventricle life, a lack of compactness, the use of an electric motor and the moving mechanical parts.

Pulsatile pumps are less uniform in design. They usually consist of a plastic or rubber ventricle squeezed by electric motors, pistons, electromagnets or pressurized fluids; Valves are required to control the blood flows. Such pumps have the following basic disadvantages: Excess propelling pressures (usually), valves which are both hemolytic and short lived, relatively complicated control mechanisms as well as the disadvantage of using an electric motor and the moving mechanical parts.

It is, therefore, a feature of this invention to present an artificial heart pump circulation system or artificial heart which duplicates the human hearts essential functions.

One feature of this invention is reliability and long life for an artificial heart.

Another feature of this invention is to provide an artificial heart which is properly packaged to fit the human chest cavity.

Still another feature of this invention is to provide for the various requirements of the human heart in an artificial heart.

A further feature of this invention is to control blood damage in an artificial heart to tolerable limits.

A still further feature of this invention is to provide an artificial heart pump which is capable of adequate perfusion of blood for any human.

A feature of this invention is an artificial heart in which the pertinent cardiovascular pump parameters which are pressure, flow, age, size, pulse and activity level, are accommodated.

Another feature is to provide an artificial heart having load matching capabilities with minimal propelling pressures.

A further feature is to provide an artificial heart which is capable of producing propelling pressures and energies compatible with the minimal and bounded pressure and minimal energy requirements of the human circulatory system.

Still another feature of this invention is an artificial heart which provides pulsatile blood flows.

A further feature of this invention is to provide an artificial heart pump which accommodates for the elastic responses 'of the circulatory system.

A still further feature is to provide a heart pump which is capable of regulating its output in response to vascular changes in flow resistance such as occurs in response to physical and emotional needs of the individual.

A feature is to provide an artificial heart in which filling of the artificial ventricle is accomplished as demanded.

Another feature is to provide an artificial heart pump in which hemolysis is minimum.

Still another feature of this invention is to provide an artificial heart in which reliability is very high.

A further feature is to provide an artificial heart pump that is packed so as to render the blood flow therethrough visible.

A still further feature of this invention is to provide a heart pump which operates at low audible noise levels.

One feature of this invention is an artificial heart in which the only moving parts are the fluid operated diaphrams and valves.

Another feature of this invention is an artificial heart pump for which the propelling pressures do not exceed the useful limits.

Still another feature of this invention is an artificial heart pump which has a portable power source, such as a hand operated bellows pump.

A further feature of this invention is to provide an artificial heart pump in which the features of a fluid amplifier are incorporated. A further object is to provide an artificial heart in which a septum separates two adjacent ventricles to affect simulated heart operation.

According to the present invention, the foregoing and other objects are obtained by providing within an artificial heart the combination of a fluid amplifier and a fluid operated diaphragm type pump which simulates the operation of the human heart. The pulsing of the circulatory system is provided by the output of the fluid amplifier being used to compress a blood filled diaphragm until such diaphragm uncovers a port through which the fluid amplifier is fed back a pressure signal which causes the fluid amplifier output to be changed such that the diaphragm is no longer compressed and can then return to its original configuration.

The specific nature of the invention, as well as other objects, uses and advantages thereof, will clearly appear from the following description and from the accompanying drawings, in which:

FIG. 1 is a block diagram of the artificial heart of this invention as connected to the human being.

FIG. 2 shows a plan view of one form of the artificial heart as shown in FIG. 1.

FIG. 3 shows a cross section of the pump as viewed along line 3-3 in FIG. 2.

FIGS. 4 through 8 show various modifications of the artificial heart.

FIG. 9 is a graphic showing of the parameters considered in the design of this invention.

FIG. 10 shows still another modification of this invention which includes a septum.

FIG. 11 shows a cross section view of the modification shown in FIG. 10.

In the drawings, like reference numerals designate identical or corresponding parts throughout the several views.

. Fluid amplifiers which will operate properly in the artificial heart of this invention are disclosed in United States Patent No. 3,016,066, issued Jan. 9, 1962 by Raymond W. Warren and the United States application Serial No. 30,691, filed May 20, 1960 by Billy M. 'Horton, now Patent No. 3,024,805 issued Mar. 13, 1962. The most satisfactory fluid amplifier employed in this invention is shown in FIG. 2 of Serial No. 58,188 filed Oct. 19, 1960 by Romald E. Bowles and Raymond W. Warren for Fluid Amplifier Employing Boundary Layer Effect which is a continuation-in-part application of the two applications Serial No. 855,478, filed Nov. 25, 1959 by 'Romald E. Bowles and Raymond W. Warren entitled Multistable Fluid-Operated System, now abandoned, and Serial No. 4,830, filed Jan. 26, 1960 by the same inventors entitled Fluid Multi'stable Memory System, and now abandoned. These last two said applications are part of the disclosure of the Billy M. Horton Patent No. 3,024,805 set forth above.

FIG. 1 shows in block diagram form, the basic part of the artificial heart pump circulation system or artificial heart 10 and the possible connections to the circulatory system. Within the artificial heart 10 are: the fluid pres sure source 11 supplying fluid amplifier 12 through connector 21, fluid amplifier 12 which controls the operation of pump 13, 14 through the power connecting lines 27 and 28 through control lines 31 and 32. In the human circulatory system the pulmonary artery 15, the pulmonary vein 16, the aorta 17, and the superior vena cava 18 are shown connected to the pump 13, 1'4.

FIG. 2 shows the details of the artificial heart. The fluid amplifier 12 is powered through tube 21 from the fluid pressures source as indicated in FIG. 1. Included in the fluid amplifier 12 are the power jet 22, the control jets 23 and 24 and the splitters 25 and 26 which separate the power fluid stream to go to tube 27 or 28 or to the bleeder 47. Bleeder 47 is open to the atmosphere or to a sump not shown, for the equalization of fluid pressure within the fluid amplifier 12. The several connectors 39 are provided to connect the tubes such as 27 and 28 to the fluid amplifier 12. These connectors 39 can be of any material which provides a seal between the tubing and the amplifier body.

The pumps 13 and 14 are identical in structure. Pump 13 is shown with the lower cup shaped structure 33 and the upper cover section 37 with a bladder 36 secured between the cup shaped and the upper structure. In a typical embodiment as shown in the drawings, the pump 13 is substantially block shaped and is made up of substantially two basic structural parts 33 and 37. In the cup shaped part 33 is a cavity 34 in which the membrane 36 can extend to substantially line the edges of chamber 34. The fluid to be pumped, in this case blood, is contained within the membrane 36 in a chamber which is signified by the number 35.

The upper section 37 has holes bored therein into which are secured the connection pieces 45 and 46. Included in connection piece 45 is the valve 43 which permits the outflow of the fluid from the chamber 35 within the diaphragm 36. In connection piece 46 is a valve 44 which permits the fluid to flow into chamber 35.

FIG. 3 is a cross sectional view of the pump 13 as seen along line 3-3 in FIG. 2. The two body sections 33 and 37 are shown with membrane 36 secured therebetween. Input tubing 27 is shown entering body 33 through a channel 30 to communicate with chamber 34. The output tubing 31 is shown in communication with chamber 34 through a channel 40 which is covered by the membrane 36 until a suflicient amount of fluid has entered chamber 34 to separate the membrane from the opening of channel 40.

FIG. 4 shows a fluid amplifier 12 controlling the operation of a pump 13. In the outflow line is a capacitance 51 which regulates the frequency of operation of the pump. This capacitance 51 can be a chamber in which the control fluid flow is delayed so as to provide a delay means for the control jet of the fluid amplifier 12. FIG. 5 shows a structure similar to that of FIG. 4 where instead of a capacitance, an orifice 52 is used to give the pressure variation which results in a delay in the control for fluid amplifier 12.

FIG. 6 shows still another variation of the structure of FIG. 4 wherein a resistance 53 is provided instead of the capacitance of FIG. 4. This resistance provides the delay for the control of the fluid amplifier 12 as shown in FIG. 6. In FIGS. 4, 5 and 6, it is to be noted that the channel 40 in the wall of the pump 13 is situated at a point somewhat lower than the point at which the diaphragm is secured to the walls of pump 13 in the modification as shown in FIGS. 2 and 3 and channel 40 is always open.

FIG. 7 shows still another modification in which the fluid amplifier 12 is used, instead of controlling the pump operation and supplying the input power therefor, only to control a valving device 61 which, in turn, controls the flow of power fluid directly from the input power source 11 and the pump 13 through tubes 71 and 72. The normal power outputs of the fluid amplifier 12 are connected to the opposite ends of valve 61 through tubes 77 and 78 respectively. Fluid entering through tube 77 is admitted into chamber 62 of the valve 61. The piston is mounted within valve 61 to reciprocate in chamber 62. This piston is made up of three sections: 63A, 63B and 63C. These three sections of the piston are connected and spaced by elements 66 and 67. The piston parts and 6 the connectors are constructed such that a chamber 64 is provided within valve 61 by part 63A, connector 66 and part 63B and a second chamber 65 is provided within valve 61 by the part 63B, connector 67 and part 63C. Port 70 is diaphragm controlled.

The control lines for fluid amplifier 12 are tubes 79 and 80.

The exhaust lines from the pump 13 are through tube 73, chamber 65, tube 74, venturi 75 and exhaust line 76. The exhaust line 76 can be connected to a sump (not shown) or can be opened to the atmosphere.

The modification shown in FIG. 8 makes use of the fluid amplifier 12 as does the other modifications. The structure of the FIG. 8 diflFers from the other figures in that the diaphragm type pump of the previous modifications has been replaced by a piston type pump 85. The power supplying lines are 81, 83 and 92, 94 which are connected to the output lines of the fluid amplifier 12 and to the piston chambers 91 and 97 in which pistons 89 and 96 reciprocate. Also connected to the output pressure line 81 and 92 are the valve control lines 84 and respectively through junctions 82 and 93, respectively. These valve control lines 84 and 95 are connected to opposite ends of a cylinder which includes a piston made up of elements 86A, 86B, 86C which are separated by connectors 87 and 88 to form a three-part piston which has two intermediate chambers therein. The fluid to be pumped from cylinder 91 is exhausted through tube 102 which is the equivalent to the connection to the pul monary artery 15 in the human as shown in FIG. 1. The filling of cylinder 91 is by a fluid which enters tube 104 when the section 87 of the piston 86 is in position to permit such flow and this is equivalent to the pulmonary vein 16 as shown in FIG. 1. For cylinder 97, tube 105 is equivalent to the aorta 17 connection in FIG. 1 and tube 103 is equivalent to the superior vena cava 18 in FIG. 1. In order to show a closed mechanical system, tubes 102 and 103 are joined by a triple junction 106 which is further connected to load 101 and to a second triple junction 107 at which the tubes 104 and 105 are joined. The control lines for the fluid amplifier 12 are lines 98 and 99 which are located so that the power fluid in cylinders 91 and 97 will be delivered to such lines when the pistons 89 and 96 have traveled a predetermined distance to uncover the openings of lines 98 and 99 so that the fluid amplifier 12 will be switched in response to pressure in such control lines.

FIG. 9 is a graphic summary of the cardiovascular parameters of pressure, flow, age, size, pulse and activity for the human male from age 1 thru 80. It is readily seen that a wide range of general output requirements for the pump is present.

It is to be noted that in the several modifications of this invention, it is possible to include only one pump Y13 and exclude pump 14. This would mean, with regards to the modification as shown in FIG. 2, that the power connector 28 and the control connector 32 would be cut and it is advisable to put flow valves on these lines. A valve on line 28 would control the pulse rate and a valve on line 32 would control pulse duration. Variations of the fluid pressure source through line 21 by a valve would control the systolic amplitude of the pump.

In FIGS. 2 and 4 through 8, a bleeder 47 is shown as being the output vent for the center power output channel in the two splitter fluid amplifier 12. The several elements 39 are connections and sealers at junctions of dissimilar materials or elements.

FIG. 10 shows a modification in which an artifical septum 111 has been added and both ventricles are included into one housing 33. This view is similar to FIG. 3 with the remainder of top section 37 and the valves having been omitted in the drawing. The structure in FIG. 10 further differs from the structure of FIG. 3 in that the septum 111 divides the power fluid chamber 34 into two separate parts, each with a power input thereto.

Only one output part 40 is needed to return through tube 31 the switching signal, derived by the uncovering of part 40 when the left ventricle 36L is sufficiently compressed, to the left control jet of the two splitter fluid amplifier 12' as shown in FIG. 11. The septum 111 is flexible and functions like the septum of the human heart. Because of the septum 111, the pressures in chambers 34L and 34R can be different by a factor of large magnitude, such as 6 to l.

The illustration of the septum in relation to the ventricles in FIG. 11 is on a larger scale than FIG. 10. The power input tubes are not shown in the same wall of the structure 33 since it is only necessary that both sections of chamber 34 be powered without being covered by ventricles 36.

The ventricles can be made integral with the septum or can be individual sacks abutting the septum. The septum is stronger than the ventricles and is made of suitable plastic material.

In the fluid pressure source line 21 is provided a valve 116 to control the systolic amplitude. The right control line 32 is provided with a valve 115 to control the duration of the pumps output pulse. The right power output line 28 is provided with a valve 114 which controls pulse rate.

In power line 27, resistance means 112 and 113 are provided to divide the pressures to chambers to 34L and 34R to suit the resistive demands of the associated cardiovascular circulations.

The return signal from port 40 in FIG. 10 through line 31 is presented to the left control jet of the fluid amplifier to switch the power stream to line 28. Blood flows into the left ventricle L to subsequently close port thereby making line 31 effectively a closed chamber. Entrainment flows of the power jet cause the pressures in chamber 31 to become lower than the pressures in the right control jet 32. This unbalance of pressure causes the power stream to switch back to receiver 27.

This artificial heart can be designed to conform to the exact external configuration of the human heart. For modification using the structure of FIG. 2, this can be accomplished by placing the fluid amplifier between the two pumps and by proper dimensioning, connect the three units directly without the tubing 31, 27, 28 or 32. For a modification such as in FIGS. 10 and 11, the fluid amplifier can be designed into the wall 33 so that the proper interconnections can be made. Then, the outside configuration of either species can be changed so as to fit into the body cavity for the human heart.

In the operation of the artificial heart pump circulation system or artificial heart as in FIG. 1, the fluid pressure source 11 provides a fluid pressure which is applied to a fluid amplifier 12 through connector 21. The fluid amplifier and ventricle provide the pulsing action which is necessary for the operation of the pump. The flowing power fluid enters the fluid amplifier after passing through a valve, the systolic volume control, that regulates the amount of fluid flowing. On leaving the power jet as a high velocity stream the flowing fluid begins to entrain additional fluid particles from the stationary surroundings. Because the walls adjacent to the power jet in the interaction region, that is the region bounded by the power and control jets, are not symetrically spaced with respect to the power jet, entrainment on the side with the closer wall is impeded causing a drop in ambient pressure. The power stream is forced to deflect slightly toward the closer wall as a result of entrainment flows resulting from the unbalance in pressures. Entrainment on this closer side is further impeded with deflection of the power stream creating still lower ambient pressure in this region and greater stream deflections. Eventually the power stream is caused to lock on to the closer wall.

Lack of wall symmetry is caused by the inability to machine the amplifier precisely. It can be quite small, a thousandths of an inch or so, and the stream deflection phenomena still operates. Further the pump will start to pulse regardless of the wall the stream locks-on to initially.

If the power stream locks-on to the left wall upon admission to the interaction region it would subsequently flow through the left receiver, through tube 27 and into the power fluid chamber 34 between the housing 33 and ventricle 36. As it continues to flow the gradual increase in fluid pressure squeezes the ventricle 36 as shown in the side view in FIG. 3 until the opening 40 is uncovered. A fluid wave traveling at the speed of sound proceeds down the deflection control line 31 through left control nozzle 23 to entrain sufficiently with the power stream to switch the power stream into the right receiver and line 28. In this process of squeezing the ventricle by air pressure, a cardiac systole has been duplicated. With the power stream now in the right receiver and tube 28, the ventricle 36 expands. The fluid amplifier 12 refuses to switch the power stream back into the left receiver until the ventricle stops expanding or until it expands at a much reduced rate at which time the fluid amplifier will switch. Since entrainment provided by the reverse flows through tube 27 and the left receiver from the power fluid chamber 34 exceeds the entrainment provided through the control orifice 24 which normally switches the power stream into the left receiver, the power stream refuses to switch during such entrainment. When the ventricle stops expanding, these entraining flows from the power fluid chamber 34 drop substantially to zero. Now flows in the control orifice 24 are suflicient to switch the fluid amplifier.

This phenomena always allows the ventricle to distend to its maximum and such maximum is always related to filling pressures presented to the pumps inlet.

With the power stream in the right receiver and tube 28 of pump 14, the power fluid chamber of pump 14 is being filled and the ventricle thereof is being squeezed while the power fluid chamber 13 is being emptied and the ventricle 36 of pump 13 is being extended. So it is seen that when pump 13 has systole function, pump 14 has a diastole function. The operation of pump 14 is exactly like the operation of pump 13 except that they operate degrees out of phase with each other.

When two pumps are operated in series, that is when blood output of a first pump is connected as the blood input to a second pump, the phenomena of pump operation discussed above allows the flows to. be operated in balance within the designed flow limits established.

In the case when the second pump 14 is not used, the open bleeder 47 is eliminated and a single splitter fluid amplifier is used. A fixed bleeder is added to tube 28. A valve for pulse duration control is connected to tube 32 to control the flow at jet opening 24 and a second valve is connected to tube 28 and the right receiver to control pulse rate. The fixed bleeder is located between the second valve and the amplifier. When the power stream from power source 21 through jet nozzle 22 lockson to the left receiver and passes through tube 27 to fill the power fluid chamber 34 in pump 13, the ventricle 36 is squeezed until the pumped fluid chamber 35 is reduced in size and until the ventricle 36 no longer covers the opening 40. During this time of lock-on with the power stream in the left receiver and tube 27, the pressure in tube 31 has been minimal and the pressure through right control nozzle 24 has contributed only a small amount of fluid which has been readily entrained into the power stream. Through the right receiver 28, a flow of fluid from the ambient condition has also been entrained by the power stream to further lock the power stream into the left receiver, assuring the stability of the power stream to fulfill its function of filling the power fluid chamber 34 without interruption of filling pressure. As the ventricle 36 uncovers opening 40, a fluid pulse at the speed of sound travels through tube 31 and left control nozzle 23 to introduce an entrainment on the left side of the power stream to switch the power stream to the right receiver and tube 28. The ventricle 36 closes opening 40 as it starts to expand. The remaining fluid in power fluid chamber 34 is then sucked through line 27, the left receiver, and out the right receiver and line 28 by the entrainment of such remaining fluid into the power stream. Entrainment serves to exhaust the fluid in chamber 34 until the pressures appearing in line 27 match the pressures available in line 28 that maintain the power stream in line 28. These pressures in line 27 include the effects of the elasticity of the diaphragm 36 and the pressure of the pumped fluid filling chamber 35 within the diaphragm 36, as well as the pressure provided by the pumping fluid Within power fluid chamber 34. This means that the amount of pumping that this system provides is dependent upon the amount of blood made available to the pump and, therefore, the pump is pressure sensitive to the demands of the patient to which it is connected. When the pressure presented to tube 27 and the left receiver are sufficiently lowered to decrease the entrainment into the power stream, and the pressure from the left control nozzle 23 is sufliciently less than the pressure from the right control nozzle 24-, the power stream will then return to the left receiver and tube 27 to repeat the pumping cycle. The right control nozzle 24 with its valve adjustment, therefore, provides a variation in the entrainment on the right side of the power stream when it is in the right receiver and tube 28.

The pulse rate control in the right receiver line 28 serves to create a variable load for the amplifier 12 and as a consequence, pulse rate can be controlled. The fixed bleeder by-passing the pulse rate control precludes power stream switching on back pressures which are intolerably high. Control is achieved over the pumps systolic amplitude by metering the amount of air entering the ventricle chamber. The systolic amplitude control (not shown) performs this function. The systolic amplitude control is a valve in line 21 to control the output of the fluid pressure source 11. To compensate for increased flow resistances presented to the pump, the pulse duration control by regulating entrainment in the right control jet 24 allows receiver pressures to build to higher values before power stream integrity is lost.

Output blood flows are made to vary directly with auricular filling pressures by causing corresponding changes in the filled volume of the ventricle. The plastic ventricle has negligible stiffness and can be over extended into the surrounding void in the housing. The fixed maximum deflection of the ventricle in systole produces the desired effect. This characteristic provides a regulation of filling pressures without electronic or other level sensing devices.

Minimal propelling pressures are realized by controlling entrainment in the right control jet and by having the power stream control both stroke volume and pulse frequency. This condition is evidenced by visualizing the small amounts of excess kinetic energy in the pumps output when the pump is required to pump against a vertical column of fluid. The fact that stroke volume and pulse are controlled by the power stream olfers the possibility of exploiting with a fair degree of preciseness the elastic responses of the vascular system by duplicating stroke volume and pulse for any particular value of flow resistance and rate. The pump approximates this over a very protracted range of flow.

A trial contraction can be approximated by having the power stream create by entrainment a slightly negative pressure around the ventricle during diastole.

An unusual characteristic of the pump is that a pumping mode similar to ventricular fibrillation can be produced. For flow conditions of the power stream above certain energy levels (this level can bevaried) the power stream switches from receiver to receiver rapidly at 10 about fibrillation frequencies, probably because of exces sive receiver back-pressures. Blood flows fall to zero.

When the amount of blood entering the ventricle of the pump varies, the frequency of operation of the pump changes correspondingly. That is, with the normal flow, the pump operates in the pulse rate range from 70 to 190 or so cycles per minute. As the flow into the ventricle lessens, the frequency decreases correspondingly to compensate for the reduced flows coming to the ventricle, and vice versa. If the flows coming to the ventricle are reduced to near zero, then the pump will go into a mode of fibrillation with pulsing frequencies similar to the human heart fibrillation frequencies.

Pulse synchronization of the pumps output with the heart is achieved by closing the right control jet 24 in synchronism with the hearts pulsing with a balloon catheter.

In the operation of the modifications shown in FIGS. 4-6, the following delay means provide the pulse rate controlled in cooperation with the contraction of the diaphragm 36 and the operation of fluid amplifier 12. The delay means in FIG. 4 is a tank 51 which acts as a capacitor to the feedback flow pressure that switches the power stream in the fluid amplifier 12. In FIG. 5 the delay means is an orifice 51 which causes a delay of build-up of pressure in the feedback loop to control the delay time for the operation of fluid amplifier 12. In FIG. 6 the delay means is a resistance in the form of a coil 53 where the feedback pressure is subjected to additional resistive forces to give a controlled delay time for the fluid amplifier 12.

In FIGS. 4, 5 and 6, channel 40 can be located at any place within housing 33 in communication with chamber 34 so long as diaphragm 36 is not capable of covering the opening to such channel 40.

The modifications of FIGS. 4-6 operate in a pulsing manner because the time delay provided by the specific capacitance, orifice or resistance and the impedance to entrainment afforded the corresponding control jet control the power stream in the fluid amplifier.

In the operation of the embodiment shown in FIG. 7, input power source 11 provides flow through tube 21 a pressure to the fluid amplifier 12 as well as to tube 71. Tube 71 is selectively connected and disconnected by valve 61 to tube 72 which is connected to the cavity outside the diaphragm in the pump 13 to compress the diaphragm. Valve 61 responds to the control given by the fluid amplifier 21. When the amplifier provides fluid pressure through tubes 77, the input power is directed into channel 62 to move the piston 63 to the position shown in FIG. 7. Upon the arrival of the piston in the extreme right hand position as illustrated, the fluid pressure in chamber 62 is now available to tube 79 to enter into the control nozzle of fluid amplifier 12 to switch the power stream to enable tube 78 to carry the power stream. While piston 62 is in the extreme right position as shown in FIG. 7, the input power is available through tube 71, cavity 64, and tube 72 to operate the pump 13. With the fluid pressure stream carried by tube 78, the piston 63 will shift toward the left, the portion 63B of piston 63 will cover the openings of tubes 71 and 72 to effectively cut off the input power pressure to pump 13. When the connection portion 67 of piston 63 is in a position where the cavity 65 surrounding 67 is between the openings of tubes 73 and 74, the pressure stored in the chamber surrounding the diaphragm of pump 13 is pumped to exhaust through tube 73, cavity 65, tube 74 by venturi 75. The tube 76 is open to atmosphere or is connected to a sump. As a result of this exhaust, the diaphragm returns to its original configuration and blood is sucked into the diaphragm.

In FIG. 7, vent 70 precludes the ventricle from deflecting beyond a predetermined minimum volume in response to the flows emanating from power source line 72. It is the action of venturi 75 to forceably exhaust the power fluid contained in chamber 34 once the ventricle starts to refill with blood and after it closes exhaust 70.

The operation of the embodiment shown in FIG. 8 is controlled by the fluid amplifier 12. The fluid power source enters through tube 21 and is selectively channeled into tube 81 or 92 depending upon the lock-on characteristics of the fluid amplifier 12. When the pistons are in the poistions illustrated, the power stream has passed through tubes 92, 94 and 95 to force piston 96 to its extreme limit away from the power source to the position shown.

The power stream has also passed through tube 92, junction 93, tube 95 to force piston 86 to its extreme upward position. The opening in cylinder 97 which allows the power stream to be fed back through tube 99 has just been uncovered and the fiuid amplifier 12 has just received the pulse from tube 99 to cause the power stream to shift so as to begin to go through tubes 81, 83 and 84. With the power stream now in tube 81, piston 89 will be driven away from the power source toward the other end of cylinder 91 and piston 86 will immediately move downward when the pressure is applied through tube 81, junction 82 and tube 84 thereto. Piston 89 will force the fluid from cylinder 91 through channel 87, tube 104, junction 107, load 101, junction 106, tube 103, cavity 88, into cylinder 97 on the right side of piston 96 to force said piston toward the left. When the luid amplifier 12 has directed that the pressure from the fluid pressure source is directed through tube 92 into cylinder 97 to force piston 96 forces the residual fluid from the right side of cylinder 97 into the cavity 88 through tubes 105, junction 107, load 101, junction 106, tube 102, cavity 87 and into cylinder 91 to force piston 89 to the left as illustrated. It is seen that the passage of fiuid pressure through load 101 is always in the same direction.

It is noted that the fluid amplifier 12 has its third output channel a bleeder.

The heart by its design is able to adjust pressures and flows to the demands of the pulmonary and systemic circulations. The circulations assist in control of flows through vasoconstrictive and vasodilation operations. Regulation of flow by the heart may be as described in the following paragraphs.

In the heart the spiral and constrictor muscles surround both the cylindrical left ventricle and the spherical right ventricle. In contracting, these muscle bundles produce maximal blood pressures in the right ventricle only about /6 those in the left. But the fact that the pressures are different and are apparently compatible with the flow resistances of the associated circulations is important.

Forces in excess of those required to produce adequate flows must be dissipated as turbulence and other energy losses, not as increased average flow, because the heart is a constant stroke pump. The fact that these forces may be minimal for every level of activity is suggested by the dynamic responsiveness of the elastic circulation to the rythmically beating heart (to be discussed later) and by the fact that the heart beats less fast and hard upon lower activity of the individual.

This difference in maximum systolic pressures between the right and left ventricle using common and individual muscle bundles is apparently achieved by the shape and position of the ventricles. It should be noted that the right heart resembles a sphere and the left heart a cylinder.

Related to the pressure phenomenon is the hearts ability to control and balance the output flows between left and right heart. If the flows of the pulmonary and systemic circulations become unbalanced due to resistance changes, the heart adjusts its output flows between halves accordingly. Mr. Billy M. Horton, Technical Director of the Diamond Ordnance Fuze Laboratories suggests that the septum performs this flow balancing role by adjusting stroke volume between the two sides. When, for example, the venous pressure at the right heart exceeds that of the left, the septum during diastole will deflect less far into the right ventricle, allowing the right ventricle to acquire a larger filled volume. During systole the higher pressures within the left heart cause the septum to move further toward the right ventricle causing blood to be ejected from the right slightly ahead of the left. The sum of opposing fluid pressures and the developed tension stresses created in the septum limit its rightward advance. Consequently the septum may operate to balance the flows between the right and left heart in response to venous pressures and opposing resisting forces in the circulations.

The vasoconstrictive properties of the vascular system apparently operate by altering flow resistance to establish the general magnitude of the blood flows in the human. Because of this peak myocardial forces become less variable.

Numerous factors operate in the human to cause the resistances of the circulation to vary with time. Emotional disturbances, sickness, and shock all have an effect. These are not always such that the ratio of output pressures between the halves of the heart remain a constant. Consequently if the heart can not adjust its output pressures to suit, mismatching will occur.

It may be that the heart accomplishes this pressure adjustment simulataneously with flow balancing. For example if the flow resistance of the systemic circulation decreased, less myocardial squeezing force would be required to sustain its former flow rate. As a result its output flows would increase because of increased stroke volumes. Simultaneously filling pressures to the left heart would tend to decrease because of the increased stroke volume. A smaller left heart radius results.

Within a relatively short period the increased flows from the left heart cause a rise in the filling pressures to the right heart. As a consequence stroke volume increases because of an increase in right heart radius caused by the increased filling pressures. While the myocardial squeezing force and pulse may not have changed, the decreased radius of the left heart and the increased radius of the right may have caused the pressure ratio to shift just enough to suit the needs of the heart and circulations.

After the pressure ratio has been adjusted, perfusion can be maintained by varying the pulse rate. The Vagus nerves slow the heart down while the Sympathetic nerve fibers accelerate the heart. These nerves are part of the autonomic nervous system.

A similar sequence of events can operate for increased systemic resistances or for changes in flow resistances within the pulmonary circuit.

Thus it may be that the heart when matching its outputs to the changing demands of the circulatory systems does so maintaining minimal pressure expenditures.

The hearts ability to balance flows and adjust pressures is probably associated with the general problem known to electrical engineers as load-matching. In so doing the heart apparently chooses to minimize the energies involved by minimizing pressures.

Many other (engineering) factors must also be considered to achieve a satisfactory cardiovascular system design and all must be compatible with the hearts physical and dynamic design. For example the repetitive stresses incurred by the muscle fibers of the heart and the vessels of the vascular system must be low enough to be commensurate with the bodys ability to repair damaged tissue materials. If the stresses exceed some definable limit, the tissue materials will fail faster than the body can accomplish repair. A comparable limit in engineering materials such as steel, aluminum, etc., is known as the endurance limit and may be considerably lower than the ultimate or breaking strengths of the materials. As the magnitude of the repeated stresses become progressively smaller, tne

13 life of the material grows until a stress, the endurance limit, is reached below which the material can presumably live indefinitely without failure.

In the vascular system excess pressure and/ or a weak vessel wall can produce an aneurysm, a ballooning of the wall. Consequently blood pressures must be controlled below the limit causing aneurysms.

Similarly the forces and energies required to achieve adequate blood pressures and flows must be commensurate with the hearts ability to produce them. They obviously cannot exceed these limits.

These factor-pressures, flows, stresses, energies and forces-are individually related to a host of other variables each of which must be compatible with total system design. These factors establish the bounds to system design and cause myocardial forces in particular to become tightly bounded.

A clue to the method employed by the human cardiovascular system to satisfy such system requirement may exist in the principles governing the design of the familiar spring actuated pendulum clock. The clock is able to operate because a sul'ficiently large force stored as potential energy in a wound spring is caused to act at precisely the right moment on the swinging pendulum. The pendulum is then forced to swing a necessary distance beforea subsequent impulse can act on the pendulum. Yet the clock can operate for days or even weeks on relatively small energy investments because minute amounts of energy and corespondingly small forces are operating for each swing of the pendulum. This is made possible because the force is caused to act in tune with the natural frequency of the pendulum. Any tampering with the natural frequency of the pendulum or the periodicity of the force can, if great enough, stop the clock. It is suggested that the human cardiovascular system may employ these same resonant principles to achieve its functional character.

Now since the cardiovascular system consists of an elastic circulatory system pulsed by a beating heart, it

.might be interesting and rewarding to examine relationships which govern resonant systems and similar relationships for the cardiovascular system.

It is known that the pulse rate for a normal young adult varies from about 60 c.p.m. (resting) to 160-180 c.p.m. (heavy exercise) with rates for short exhaustive work to 240-270 c.p.m. No similar information was available for older persons or children, however it is suspected that about the same limits hold for older persons.

By using this data and examining it in the manner customarily used to describe resonant systems in general,

it is observed that the following similarities between resonant phenomena and the cardiovascular system exist.

The frequency ratio bounding the operation of the cardiovascular system corresponds to the region in simple resonant systems where maximum exciting forces are minimal for a desired fixed excursion of the mass.

For simple resonant systems the variation in force between these limits is small compared with force changes above the upper limit. The cardiovascular system demonstrates a similar force characteristic.

As the frequency ratio approaches zero, resonant systems respond less to the exciting forces. It would appear that the cardiovascular system at ratios approaching zero -also acts stitfiy. These few observations do not prove conclusively that resonant principles govern the operation of the cardiovascular system. However, an attempt was made to calculate the changes required in viscosity and some characteristic artery diameter for the range of blood pressures and flows observed in the human. Poiseulles law was used. It was hoped that these calculations would show that the large range of flows could not be achieved solely by the induced pressure drops and the changes in viscosity and artery diameter. However the assumptions rigid piping systems.

mote.

One could now suggest a few reasons for a pulsating heart and an elastic circulation system by comparing the performance of the cardiovascular system as described herein with that of other engineering pump systems.

Positive-displacement reciprocating pumps in small and medium capacities are better adapted for high pressures than centrifugal pumps. Or as a corollary, a reciprocating positive-displacement pump of small size can pump higher flows at higher pressures than centrifugal pumps of the same size. (Note: Centrifugal pumps have a non-pulsating output in contrast to the former type.) Hence, it might be suggested that flow requirements and com pactness necessitated a positive-displacement type pump.

Conversely, pulsating pumps in contrast to steady flow pumps produce large variations of stress in the associated By employing resonant principles in the cardiovascular design, it has been demonstrated that variation of peak stresses for up to a ninefold increase in flow is only about 38%. If the vascular vessels were rigid rather than elastic, a suggestion of the variation in stress might be had by examining the myocardium force curve for frequency ratios less than 0.3, the region where the vessel tissue tends to act stifliy. The slope of the curve is much steeper in this region than in the region between 0.3 to 0.8 with steepness increasing with the higher ratios. It might be surmised therefore that if the phenomena of resonance didnt operate above the 0.3 ratio, the myocardial forces required to achieve the required flows would be much greater than they are. If these forces were not limited, the alternative presented to the designer is to use higher strength vessel materials. Higher strength materials might penalize the designer in the areas of weight, energy expenditures, mobility and porosity. An added advantage occurs by having tightly bounded forces operating in the system. The vessel materials are efficiently utilized. There is no need for excess materials. This characteristic is often used by the engineer to reduce waste in his designs.

It is not feasible to continue a justification of the cardiovascular design on the basis of alternative approaches, but it would appear that it may uniquely satisfy the requirements imposed upon it.

An extension of these considerations as regards the apparent resonant character of the system might lead to the development of-a simplified empirical or mathematical model of the cardiovascular system assuming existing unknown elastic and related parameters can be determined. A model of this sort, if it did in fact describe the performance of the cardiovascular system, would serve as a powerful tool since now the contributions of individual 1 changes in elastic properties of the circulatory system or No model of any sort (mathematical or otherwise) appears to exist to describe the system. The subsequent paragraphs present relevant dynamic properties of the 1 Vascular systems as recorded in the literature.

The flow of blood is pulsatile in the arteries, and because of their elastic properties and flow resistance flow becomes continuous beyond the arterioles. This statement is probably based on present instrumentations abil ity to measure pulsed flows. It might be pointed out at this point that flow in the veins is non-pulsatile, and the vessel walls contain less elastic tissue than the arteries although, they can distend appreciably under prolonged elevated intravenous pressures.

There is one particularly important reason why the vascular system might require pulsed flows to the capillaries and this regards critical closing pressures. Smaller vessels require higher pressures to hold them open for flow than larger vessels. At low pressures, it is possible that closure of the vessel might result particularly if the tension exerted by the wall is increased appreciably by active contraction of the smooth muscle of the wall.

If, however, resonant principles operate in the vascular system, it is suggested that very small pulsed forces acting in conjunction with the elastic properties of vessels could establish flow in tubes which under steady pressures of higher magnitude might be closed or at least only partially open. This is especially true for lightly damped vessels. Hence, the small vessels such as the capillaries (which are about one millimeter in length) might sustain flows only because flows are pulsatile for pressures of the magnitude presented by the heart of the circulatory systern.

It has been observed that the metarterioles and the corresponding precapillary sphincters demonstrate periodic contractions. Briefly, blood flows generally into a metarteriole from the arteriols, and then into capillaries. The metarterioles lead directly into the channels which are main thoroughfares from the capillary bed to the venules. The true capillaries concerned with interchange between blood and tissues are inter-anastomosing side branches of the main channels through the bed. At the ostia of each capillary is a small precapillary sphincter of smooth muscle, which is controlled by nerves presumably from the sympathetic nervous system, in the same manner that these nerves control the arterioles and metarterioles. In the body, the metarterioles and their precapillary sphincters undergo periodic contractions at intervals of 15 seconds to 3 minutes. When the tissue is in a resting state, the constrictor phase of this rhythm predominates and the precapillary sphincters may be completely closed. When the tissue becomes active, the dilator phase of the metarterioles predominate and the precapillary sphincters are open. Thus, in skeletal muscle, it is believed that the increase in blood flow with exercise comes in large measure from this opening up of large numbers of additional capillaries.

The fact that the sphincter acts as an elastic valve and that the sphincter and associated capillary branch off from the metarterioles may be significant. This elastic valve may require pulsed flows for adequate perfusion of the capillary. If it did not receive pulsed flows for its proper response, it could refuse to open sufliciently and the blood could be partially shunted through the metarteriole to the venule.

Incidentally, if resonant principles hold for the cardiovascular system, probably extracorporeal pumps for long-term perfusion applications should be of the pulsatile variety with a pulse frequency adjustable to the patents normal resting pulse. Adequacy of perfusion is in part judged by the maintenance of blood pressures and flows. Inadequate perfusion of the capillary beds could be masked by the higher pressures required for perfusion with steady flow pumps. The inadequacy might not manifest itself for short-time pump runs.

In a prior study of the behavior of the dogs cardiovascular system and elastic tubes of varying degrees of distensibility to artificially generated pulses of controlled character the following conclusions were reached: that the inertial, viscous friction, and wall tension forces and their corresponding effect on the artificially generated pulses vary in a non-linear manner and each effects the other; that the vascular pulse pressure is the sum of these forces; and that pulse pressure varies considerably in magnitude and in its component forces under various circulatory states even though stroke volume remains constant.

When the artificial pulses were introduced into very distensible tubes, the effective mass and viscous friction were so reduced that no induced pressure pulse was seen.

The corollary is also true, that in rigid tubes these forces become very great. Nothing in the study relates the frequency responsiveness of the circulations of the pulsed flows, and it would seem that herein lies the real message of this work.

A comprehensive mathematical attempt to describe the dynamic responsiveness of portions of the human cardiovascular system to the rhythmic pulsing of the human heart has been made previously. A conclusion was that there is no doubt that the heart pumps the blood and that it generates a means increase in pressure. Also, muscle contractions which compress the larger veins will propel the blood through the veins toward the heart. Further, the mechanical waves which are emitted by the heart and the larger arteries give up their energy in the small arteries, the arterioles and the capillaries, and that a fraction of this energy is used to propel the blood and overcome flow losses in the arterioles and capillaries.

The mathematical attempt also produced the suggestion that the flow smoothing role played by the larger arteries may be performed further away from the heart than originally thought to be the case since the fluctuations do not decay very much in the brachial and femoral arteries. The complexity of the oscillations in the flow involving standing waves, reflected waves, self-excited waves and traveling waves are emphasized.

A process is suggested whereby axisymmetric waves can excite lateral waves further downstream as, for example, beyond a bend so that oscillating energy can be gradually transferred to shorter and shorter wavelengths of higher frequencies as the flow proceeds from the heart. Since the cardiovascular system has complex modes of vibration and oscillations, it would be expected from the point of view of bounding energies and forces that this would be the case for compatibility with the stiffer elastic properties of the smaller vessels. In other words the same bounding frequency ratios would prevail as the pulsed flows, now of higher frequency, coursed through smaller vessels with higher natural frequencies.

The mathematical attempt further produced the awareness of the possibility of a neuro-muscular coupling of the vibrations, which may be under central control as far as coupling gain is concerned. This may be the bodys method for altering the circulations responsiveness to pulsed flows under situations of stress, sickness or disease to maintain bounded conditions.

It is suggested that the human cardiovascular system may operate on principles which govern resonant systems generally. This conclusion is based on the close correlation observed between bounding pulse-arterial resonance ratios and the similarities existing between force and energy expenditures for the cardiovascular system and for simple resonant systems. This suggests that any changes to the resonant character of the system or the pulsed flows generated by the heart, if sufficiently great and uncorrected, can cause the heart to overwork and die prematurely or can cause the cardiovascular system to malfunction in other ways. It is further suggested without rigorous proof that only because resonant principles are employed, the heart and circulatory system are able to meet the design requirements and restrictions imposed on them.

The elimination of moving control parts and electronics raises reliability and lengthens life because of the following reasons:

(1) Apart from ventricles and valves there are no moving parts to wear or electronic components to fail. The passageways are too large to clog easily.

(2) The basic fluid amplifier components can be reproduced with consistent accuracy-possibly photographically.

(3) Maintenance is minimal. All parts can be autoclaved.

The ventricle has been designed for negligible stress. One ventricle has been tested at a two liter per minute 1 7 Water flow and 190 mm. Hg blood pressure equivalent for 1,500 hours without failure.

The packaging aspects of the pump are excellent. It consists of eleven basic parts exclusive of assembly screws. Apart fromthe fluid amplifier, dimensional tolerances are large. Both the housing and amplifier can be fabricated out of clear sterilizable plastics. Molding techniques can be used in the production of all pump elements. If required the ventricle and valves can be thrown away after each use because of their low cost.

The complete pump weighs 7% pounds and has overall dimensions of 7 /8" X 6% X 5". It produces a muflied audible pulsing noise when operating. Production costs for the complete pretested pump in its final design should be in the order of several hundred dollars. The pump would appear to have the following performance characteristics:

(1) It can produce blood pressures from 0 to about 500 millimeters of mercury.

(2) Blood flows from about /2 liter/minute to nine or ten liters/minute are possible depending on circuit flow resistances.

(3) Pulse rate can be varied from about 30 c.p.m. to 125 c.p.m. The upper limit depends on flow resistance. The upper limit has been raised to above 190 cycles per minute.

Hemolysis appears to be about 50% less than one tested prior pump when tested occlusively and several times less than another tested prior pump.

Using fluid amplification principles a heart pump circulation system has been designed which can approximate many of the human hearts pumping functions. It operates without moving control parts or electronics. Without such parts reliability and life are measurably increased and production costs significantly decreased. Its pumping and hemolytic characteristics are at least equal to those of most commercial blood pumps.

The fluid amplifier has an inherent audible pulse rate signal produced by the power stream switching from one receiver to the other. This audible signal is readily usable for monitoring any malfunction in any part of the system, including the extra corporeal circuit if such changes would affect the rates of flow to and from the pump.

Of particular importance is that the pump can be transported by hand, it can be completely disassembled in a matter of minutes, and except for ventricles and heart valves all parts can be autoclaved. Power can be pro vided by tanks of compressed air or eventually perhaps by exhaust gases from a vehicle.

Substantially all of the materials of the artificial heart of this invention are plastic materials. The body parts 11 and 13 and the fluid amplifier 12 can be made of Lucite or Teflon or any sterilizable material, a plastic or otherwise, that would be compatible with the body chemistry, such as polyurethane. The ventricles and valves can be made of polyurethane, polypropylene, or polyvinylchloride. The tubing is a conventional plastic which is compatible with body chemistry.

In the fluid amplifier, the control jets and the receivers can be made of metal to resist abrasion. This metal can be stainless steel, silver or gold, among others.

It will be apparent that the embodiments shown are only exemplary and that various modifications can be made in construction and arrangement within the scope of the invention as defined in the appended claims.

What is claimed is:

1. In an artificial heart pump circulation system: a fluid amplifier comprising a fluid power input channel, power orifice means connected to said input channel for providing a power jet, first and second output channels in fluid communication with said power jet, and control orifice means for providing fluid control signals adjacent to said power jet for directing said power jet into either of said output channels; pump means comprising a body having a cavity therein and a diaphragm suspended Within said cavity to separate said cavity into a pumped fluid chamber and a power fluid chamber, said pumped fluid chamber being adapted to receive blood from the venous system; first means connecting said first output channel to said power fluid chamber; and second means connecting said power fluid chamber to said control orifice means for providing a feedback signal for temporarily diverting said power jet from said first output channel to said second output channel, whereby said pumped fluid chamber is adapted to provide blood to the arterial system.

2. In an artificial heart pump circulation system: pump means comprising a body having a cavity therein and a diaphragm suspended within said cavity to separate said cavity into a pumped fluid chamber and a power fluid chamber, said pumped fluid chamber being adapted to receive blood from the venous system; a fluid amplifier comprising a fluid power input channel, a power orifice means connected to said input channel for providing a power jet, first and second output channels in fluid communication with said power jet, and control orifice means for providing fluid control signals adjacent to said power jet for directing said power jet into either of said output channels; a first conduit connecting said first output channel of said fluid amplifier to said power fluid chamber; and a second conduit connecting said power fluid chamber to said control orifice means so as to provide a fluid feedback signal for temporarily diverting said power jet from said first output channel to said second output channel whereby said pumped fluid chamber is adapted to provide blood to the arterial system.

3. In an artificial heart pump circulation system, a fluid pressure source issuing a power stream, a fluid amplifier means having an input jet connected to said fluid pressure source, said fluid amplifier having a pair of oppositely disposed control jet means adjacent to said input jet and perpendicular thereto and a pair of receivers which are spaced so as to selectively receive said power stream, pump means having a body means for inclosing a chamber, diaphragm means suspended within said chamber to separate said chamber into a power fluid chamber and a pumped fluid chamber, first fluid conducting means for connecting said power fluid chamber and one of said receivers and second fluid conducting means for connecting said power fluid chamber to one of said control jets, said second fluid conducting means having a port positioned in said power fluid chamber such that said port is covered by said diaphragm until a predetermined amount of fluid is pumped from said pumped fluid chamber.

4. In an artificial heart pump circulation system, a fluid power source, fluid amplifier means, said fluid power source connected to said fluid amplifier, pump means, diaphragm means in said pump means for providing a pumped fluid chamber and a power fluid chamber, and means for connecting said fluid amplifier and said pump means including a fluid power connector connected to said power fluid chamber and a feedback control connector which is operative to pass a feedback signal to said fluid amplifier upon preselected movement of said diaphragm whereby the frequency and stroke volume of said artificial heart pump circulation system vary in response to the same pressure stimulii as the human heart.

5. In combination: a fluid power source for providing a fluid power stream; fluid amplifier means including power input means connected to said power source, a pair of receiver means and bleeder means positioned between said pair of receiver means, and a pair of control means for directing said power stream into a selected one of said receiver means; valve means including a housing means with a plurality of ports therein and a reciprocative piston housed therein, said piston opening and closing a plurality of said ports; pump means including a diaphragm to separate said pump means into a pumping fluid chamber and a pumped fluid chamber; and connecting means including first means for connecting said fluid power source to a first one of said ports in said valve means, second means for connecting a second of said ports to said pumping fluid chamber, said first and second ports being opened simultaneously and closed simultaneously to control the application of said power stream from said source through said valve means to said pump means, third means for connecting said bleeder means to a third one of said ports, fourth means for connecting a fourth one of said ports to said pumping fluid chamber, said third and fourth ports being aligned so as to be opened and closed simultaneously by said piston to control the exhaust from the pumping fluid chamber, venturi means in said third means for connecting, said first and second ports being open when said third and fourth ports are closed and closed when said third and fourth ports are open, fifth means for connectin a first one of said power output means to a fifth one of said ports to supply said power stream to move said piston to open said first and second ports, sixth means for connecting a sixth One of said ports to a first one of said fluid amplifier control means, said control means positioned so as to direct said power stream from said first receiver to the second of said receivers, seventh means for connecting a seventh one of said ports to said second receiver to supply said power stream to move said piston to close said first and second ports and eighth means for connecting an eighth one of said ports to the second of said fluid amplifier control means, said sixth and eighth ports being positioned so as to be alternately open and to alternately conduct said power stream to said first and second control means after said piston has moved a distance determined by the desired frequency of opertion of the pump.

6. In combination: a fluid power source for providing a fluid power stream; fluid amplifier means including power input means connected to said power source, a pair of opposed control means, a bleeder means and a pair of receiver means aligned to receive said power stream from said power input means; pump means including a pump housing with two cavities therein, a pair of piston means, one piston means in each of said cavities separating said cavities into a power fluid chamber and a pumped fluid chamber, a power port and a control port through said housing into each of said pair of cavities, said control ports being positioned to determine the extent of travel of said pistons within said cavities; a pumped fluid means; valve means to maintain a uniform direction of flow of said pumped fluid including a valve housing means with a third cavity therein and, a pair of input ports and a pair of output ports in communication with said third cavity, one of said input and one of said output ports being also in communication with the pumped fluid chamber of one of said two cavities and the other of said input and the other of said output ports being also in communication with the pumped fluid chamber of the other of said two cavities, third piston means reciprocatively operative within said third cavity to alternately open one output port of one cavity simultaneous with the input port of the other of said two cavities, and a pair of valve power ports through said valve housing; and first connecting means for connecting one of said control means with one of said control ports, second connecting means for connecting the other of said control means with the other of said control ports, third connecting means for connecting one of said receivers with one of said power ports and one of said valve power ports, fourth connecting means for connecting the other of said receivers with the other of said power ports and the other of said valve power ports, and means containing said pumped fluid connected to said input and output ports for maintaining unidirectional flow of the pumped fluid contained therein.

7. In an artificial heart pump circulation system, a fluid power source; fluid amplifier means including a power nozzle issuing a power stream, a first and a second control nozzle, an interaction chamber, a first and a second receiver means and means for limiting the power stream to be in the plane determined by its deflection, said power nozzle, said control nozzles and said receivers having openings into said interaction chamber, said receivers being located downstream of said power nozzle and positioned to alternatively receive said power stream in response to a feedback signal applied to said first control nozzle and an alternately occurring predetermined loading of said first receiver; pump means including a body means having a cavity therein, diaphragm means in said cavity positioned and shaped to provide for a power fluid chamber and a pumped fluid chamber, first conductor means connected at one end to said first receiver and at the other end to said body means at a location to conduct the power stream from said first receiver to said power fluid chamber, port means located in said body means so as to be covered by said diaphragm means until said pumped fluid chamber is emptied a predetermined amount, second conductor means connected to said port means and said first control nozzle to conduct a fluid signal pulse from said power fluid chamber through said first control nozzle into the interaction chamber to switch the power stream to said second receiver, said second receiver being loaded such that in combination with the second control nozzle and the sucking of the fluid from the power fluid chamber filling the pumped fluid chamber an amount determined by the load presented thereto whereupon the power stream switches back to the first receiver.

8. In an artificial heart pump circulation system, as set forth in claim 7, including means for pulse synchronizing said pump means with means for periodically occluding said second control nozzle.

9. In an artificial heart pump circulation system; a fluid power source; fluid amplifier means including a power nozzle issuing a power stream, a first and a second control nozzle, an interaction chamber, a first and a second receiver means and means for limiting the power stream to be in the plane determined by its deflection, said power nozzle, said control means and said receivers having openings into said interaction chamber, said receivers being located downstream of said power nozzle and positioned to alternatively receive said power stream in response to a feedback signal applied to said first control nozzle and an alternately occurring predetermined loading of said first receiver; pump means including a body means having a cavity therein, diaphragm means in said cavity positioned and shaped to provide for a power fluid chamber and a pumped fluid chamber, first conductor means connected at one end to said first receiver and at the other end to said body means at a location to conduct the power stream from said first receiver to said power fluid chamber, port means located in said body means so as to be out of contact with said diaphragm means at any time, a fluid flow impedance means, second conductor means connected to said port means and said fluid flow impedance means and third conductor means connected to said impedance means to said first control nozzle to conduct a properly delayed feedback signal from said pumped fluid chamber to effect the switching of the power stream to the second receiver.

it). In an artificial heart pump circulation system; a fluid power source; a fluid amplifier having a power nozzle for producing a fluid power stream, first and second control nozzles, an interaction chamber and a first and a second receiver, said fluid power source connected to said power nozzle, said control nozzles and said receiver means having entry into said interaction chamber, said receivers positioned to alternately receive said power stream in response to control signals applied thereto, said second receiver and said second control nozzle being valved to the ambient condition; pump means including a body means having a cavity therein, diaphragm means for dividing said cavity into a pumped fluid chamber and a power fluid chamber, first conductor means connecting said first receiver to said cavity in said body means into the power fluid chamber whereby application thereto of the power stream effects the compression of the diaphragm and lessens the volume of the pumped fluid chamber, second conductor means connected to said power fluid chamber and to said first control nozzle so as to give a feedback signal indicative of the predetermined compression of said diaphragm to switch the power stream to said second receiver, said first receiver and said first conductor means serving to conduct the fluid from the power fluid chamber as a result of the sucking produced by the power fluid being in said right receiver in response to filling pressures presented to the pumped fluid chamber whereby when such filling pressures reach a minimum the artificial heart pump circulation system will go into fibrillation as does the human heart.

11. In an artificial heart pump circulation system, a fluid power source, a pump means comprising a body means having a cavity therein, diaphragm means suspended within said cavity to separate said cavity into a pumped fluid chamber and a power fluid chamber and first and second port means extending through said body means into said power fluid chamber; control means for directing the inflow and the outflow of power fluid to said power fluid chamber and for sensing the amount of fluid entering said pumped fluid chamber comprising a fluid amplifier having an interaction chamber, a power nozzle issuing a power stream, first and second control nozzles located on opposite sides of said power nozzle, said nozzles being directed into said interaction chamber, first and second receiver means positioned downstream of said power nozzle for alternately receiving said power stream in response to a feedback signal produced by said pump means when the pumped fluid chamber is emptied a predetermined amount and to a predetermined condition of pressures in said pumping fluid chamber, said first and second receivers and said interaction chamber, said condition being a function of the filling pressure in the pumped fluid chamber, first conductor means connected from said first port to said first control nozzle for carrying said feedback signal to said amplifier and switching the power stream into said second receiver, second conductor means connected from said second port to said first receiver for carrying power fluid into said power fluid chamber when said power stream is in said first receiver and for carrying said power fluid away from said power fluid chamber when said power stream is in said second receiver, and unidirectional means connected to said pumped fluid chamber to maintain unidirectional flow of the pumped fluid.

References Cited by the Examiner UNITED STATES PATENTS 1,594,217 7/26 Smith 91-290 X 2,952,217 9/60 Lindbom 103152 X 3,001,698 9/61 Warren 13781.5 3,016,066 1/62 Warren 137624.14 3,016,840 1/62 Frick 103-152 X 3,024,805 3/ 62 Horton 137-597 3,037,504 6/62 Everett 128214 3,048,165 8/62 Norton 1281 3,052,238 9/62 Broman 128214 3,122,062 2/64 Spivak 913 3,124,999 3/ 64 Woodward 913 OTHER REFERENCES Dennis et al.: Development of a Pump-Oxygenator To Replace the Heart and Lungs; an Apparatus Applicable to Human Patients, and Application to One Case Annals of Surgery, vol. 134, #4, October 1951, pages 709-721.

RICHARD A. GAUDET, Primary Examiner. JORDAN FRANKLIN, Examiner.

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Classifications
U.S. Classification600/17, 137/83, 137/835, 623/3.21, 137/624.14, 417/347, 128/DIG.300, 417/395, 91/3
International ClassificationA61M1/10
Cooperative ClassificationY10S128/03, A61M2001/1065, A61M1/106
European ClassificationA61M1/10E4H