Sept. 28, 1965’
`
`Filed Feb. 2, 1962
`
`K. E. wooowm-D
`ARTIFICIAL HEART PUMP CIRCULATION SYSTEM
`6 Sheets—Sheet 1
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`3,208,448
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`
`FIG.2
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` 3
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`PRESSURE
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`SOURCE
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`ARTlFICIAL HEART
`-. INVENTOR
`KENNETH E. WOODWARD
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` ‘
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`X; frank} alflwwfié
`<2; m M
`
`BY
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`PAGE 1 OF 17
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`4
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`WATERS TECHNOLOGIES CORPORATION
`EXHIBIT 1012
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`WATERS TECHNOLOGIES CORPORATION
`EXHIBIT 1012
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`PAGE 1 OF 17
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`

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`Sept. 28, 1965
`
`K. E. WOODWARD
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`3,208,448
`
`ARTIFICIAL HEART PUMP CIRCULATION SYSTEM
`
`Filed Feb. 2, 1962
`
`6 Sheets—Sheet 2
`
` RESISTANCE
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`PAGE 2 OF 17
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`PAGE 2 OF 17
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`' Sept 28, 1965
`
`K. E. WOODWARD
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`3,208,448
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`ARTIFICIAL HEART PUMP CIRCULATION SYSTEM
`
`,
`
`Filed Feb. 2, 1962
`
`6 Sheets—Sheet 3
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`[fly/VI
`I_II-I_-I“fill
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`PAGE 3 OF 17
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`PAGE 3 OF 17
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`Sept. 28, 1965-
`
`K. E. WOODWARD
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`3,208,448
`
`ARTIFICIAL HEART PUMP CIRCULATION SYSTEM
`
`Filed Feb. 2, 1962
`
`6 Sheets-Sheet 4
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`PAGE 4 OF 17
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`PAGE 4 OF 17
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`Sept. 28: 1965
`
`K. E. WOODWARD
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`3:208:448
`
`ARTIFICIAL HEART PUMP CIRCULATION SYSTEM
`
`Filed Feb. 2, 1962
`
`6 Sheets-Sheet 5
`
`l0
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` 200 0PM
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`MEAN(LlTERS/MIN/METERZ)BLOODFLOW
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`INCREASING
`ACTIVITY
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` 80 YEAR
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`OLD
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`FIG.9
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`40
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`60
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`80
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`I00
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`MEAN BLOOD PRESSURE (mm Hg)
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`PAGE 5 OF 17
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`PAGE 5 OF 17
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`Sept. 28, 1965
`
`K. E. WOODWARD
`
`3,208,448
`
`Filed Feb. 2, 1962
`
`ARTIFICIAL HEART PUMP CIRCULATION SYSTEM
`6 Sheets-Sheet 6
`
`IS
`
`ws
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`3N
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`§ ‘1
`I
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`37
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`I'
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`31 /
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`PAGE 6 OF 17
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`PAGE 6 OF 17
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`

`

`United States Patent Office
`
`3,208,448
`Patented Sept. 28, 1965
`
` 1
`
`2
`
`3,208,448
`ARTIFICIAL HEART PUMP CIRCULATION
`SYSTEM
`Kenneth E. Woodward, 2526 Hunting Ave., McLean, Va.
`Filed Feb. 2, 1962, Ser. No. 170,851
`11 Claims.
`(Cl. 128—1)
`(Granted under Title 35, US. Code (1952), sec. 266)
`
`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 par-
`ticularly 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 cor-
`rection face premature death. Many devices are involved
`in this type 'of surgery. One important apparatus is a
`pump can assume the heart’s 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 con-
`trol simplification for pulsable type extracorporeal heart
`pumps and a more accurate duplication of the heart’s
`physiological pumping functions. Because all moving con-
`trol 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 heart’s 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 subject’s :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 tol-
`erated either during or after the pumping run. From a
`cursory examination of the literature and prior pump de-
`signs it would appear that a single set of functional re-
`quirements for extracorporeal pumps in general has not
`been established.
`In developing the fluid amplifier pump
`of this invention the following heart functions are con-
`sidered to be important:
`Output pressures and flows: Certainly the pump must
`be capable of adequate perfusion of the subject. More-
`over 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 para-
`meters are pressure, flow, age, size, pulse and activity
`level.
`(The female imposes generally less severe require-
`ments on the pump.) For surgical applications only
`resting conditions are of concern.
`Load-matching capabilities: it is believed that the hu-
`man heart recognizes the flow resistances of the cardio-
`vascular 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
`
`PAGE 7 OF 17
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`It is be-
`propel blood with minimal force application.
`lieved that myocardial heart forces are minimal and
`tightly bounded for large variations in blood flow. Fur-
`ther 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.
`among
`Pulsatile blood flows: Disagreement exists
`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 periods—one to two hours.
`It may be, how-
`ever, 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
`subject’s 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 pres—
`sures as in the case of the human heart,
`Other functional requirements: Filling of the human
`heart ventricle is achieved mainly by the difference be-
`tween 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 sig-
`nificant, 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.
`.
`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 condi«
`tions.
`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: Uncom-
`plicated to operate and maintain, controllable with respect
`to pulse and flow, and fabricated out of transparent ma-
`terials to allow visual observation of. performance.
`Im-
`pending failures can often be observed in time for proper
`corrections of such failures. The pump should be steri-
`lizable preferrably by autoclave methods, functional with
`negligible heat liberation, fabricated out of materials ac-
`ceptable to the blood, easily transportable, scalable to
`preclude blood contamination and air embolism,
`inex-
`
`PAGE 7 OF 17
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`

`3,208,448
`
`3
`pensive 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., pul-
`satile 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 rela-
`tively 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: Ex-
`cess propelling pressures (usually), valves which are both
`hemolytic and short lived, relatively complicated control
`mechanisms as well as the disadvantage of using an elec-
`tric 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 heart’s essential func-
`tions.
`One feature of this invention is reliability and long life
`for an artificial heart.
`'
`Another feature of this invention is to provide an arti-
`ficial 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 arti-
`ficial 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 perfu-
`sion 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 ac-
`commodated.
`Another feature is to provide an artificial heart having
`load matching capabilities with minimal prOpelling pres-
`sures.
`
`C71
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`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 pro-
`vide 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 opera-
`tion 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 ampli-
`fier 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 accompany-
`ing 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 con—
`sidered in the design of this invention.
`FIG. 10 shows still another modification of this inven-
`tion which includes a septum.
`FIG. 11 shows a cross section view of the modification
`shown in FIG. 10.
`like reference numerals designate
`In the drawings,
`identical or corresponding parts throughout the several
`views.
`. Fluid amplifiers which will operate properly in the arti-
`ficial 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 Pat-
`ent No. 3,024,805 issued Mar. 13, 1962. The most sat-
`isfactory 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 in-
`ventors 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 con-
`nector 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 pul-
`monary vein 16, the aorta 17, and the superior vena cava
`18 are shown connected to the pump 13, 1‘4.
`
`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 dia-
`phrams and valves.
`PAGE 8 OF 17
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`5
`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 be-
`tween the cup shaped and the upper structure.
`In a typi-
`cal embodiment as shown in the drawings, the pump 13
`is substantially block shaped and is made up of substan-
`tially 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.
`In-
`cluded in connection piece 45 is the valve 43 which per-
`mits 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 showu with membrane 36 secured there-
`between.
`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 oper-
`ation 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 in-
`stead 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 de-
`lay 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 dia~
`phragm is secured to the walls of pump 13 in the modifi-
`cation 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
`
`60
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`75
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`PAGE 9 OF 17
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`3,208,448
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`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.
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`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 differs from the other figures in
`that the diaphragm type pump of the previous modifi-
`cations 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
`95 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 per-
`mit 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 .13 and
`exclude pump 14. This would mean, with regards to the
`modification as shown in FIG. 2, that the power con-
`nector 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 ele-
`ments 39 are connections and sealers at junctions of dis-
`similar materials or elements.
`FIG. 10 shows a modification in which an artifical
`septum 111 has been added and both ventricles are in-
`cluded 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.
`
`PAGE 9 OF 17
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`3,208,448
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`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 diflerent by a factor of large magnitude, such
`as 6 to 1.
`The illustration of the septum in relation to the ventri-
`cles 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 plas-
`tic 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 dura-
`tion of the pump’s output pulse. The right power output
`line 28 is provided with a valve 114 which controls pulse
`rate.
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`10
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`20
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`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 cardio-
`vascular 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 35L to subsequently close port 40 there-
`by making line 31 effectively a closed chamber. En-
`trainment 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 ampli—
`fier 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 inter—
`action 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
`PAGE 10 OF 17
`
`8
`phenomena still operates. Further the pump will start to
`pulse regardless of the wall the stream locks-on to ini-
`tially.
`If the power stream locks—on to the left wall upon ad-
`mission 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 pres-
`sure, 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 pump’s 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 180 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 opera-
`tion 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 locks-
`on 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 re-
`duced 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 pres-
`sure 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 en-
`trained 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
`
`30
`
`40
`
`60
`
`75
`
`PAGE 10 OF 17
`
`

`

`“3,208,448
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`10
`about fibrillation frequencies, probably because of eXCéSh
`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