Jet Pump Cardiac Replacement And Assist Device And Method Of At Least Partially Replacing A Disabled Right Heart

Abstract

An implantable jet pump cardiac replacement device and method for replacing or assisting the right heart. The jet pump device is an elongated tubular structure including an upstream driving nozzle from which a driving flow of arterial blood under pressure is ejected into a suction nozzle creating a zone of reduced pressure to cause venous blood to be sucked into and admixed with the driving flow for distribution to the pulmonary circulation system. The pump may be powered by blood pumped by the left heart or an artificial replacement for the left heart.

Description

The invention described herein was made in the course of work under a grant or award from the Department of Health, Education and Welfare.

This invention relates to an implantable jet pump cardiac replacement device and more particularly to a jet pump right ventricle replacement device to work in conjunction with left ventricle support yielding total heart support.

The present death rate from acute transmural myocardial infarction in coronary care units today is primarily because of power failure of the heart. This includes patients who are in cardiogenic shock and/or congestive heart failure. Although heart transplants offer some hope, further significant lowering of the mortality rate depends largely on an artificial heart or satisfactory cardiac replacement and assist units.

In a great number of cases this necessarily means replacement of both the right and left ventricles since the coronary artery disease affects power to both ventricles when infarction occurs. Use of one pump to replace the entire heart is complicated by higher pressure requirements in the systemic circuit than in the pulmonary circuit. Too, exposure of the pulmonary vasculature to such high pressures is detrimental. This is amply demonstrated in patients who have a single ventricle or large ventricular septal defects. Theoretically, it is possible to pull blood through the pulmonary circuit by a negative pressure developed by an artificial left ventricle. However, the physical properties and the geometrical configuration of the pulmonary veins do not allow this, for they are prone to collapse at low negative pressures.

Left ventricular output cannot exceed right ventricular output. So typically, pump replacements of the right ventricle have been identical to their left ventricle counterpart. This arrangement is necessary when ventricle type pumps are utilized. The precise balancing control of these two units is difficult to achieve. As another limitation, flow and pressure dictate ventricle pump size. Yet, with that system, because the left and right heart flow rates are equal, no advantage can be realized from a right heart work load which is less than 20 percent of the total. Finally, overt hemolysis and delayed red cell damage, as well as the chance of mechanical failure, are at least doubled for two ventricle pumps.

The present invention is directed, as an alternative, to a jet pump system to replace, assist, or bypass the right ventricle function, which will couple with any left heart. In combination with a unit functioning for the left ventricle, this system constitutes a total heart replacement. Specifically, many of the problems associated with two identical pumps have been eliminated.

Although jet pumps have never been applied to the circulation, they have been investigated for over 100 years. Presently, they are used in deep well pumping, in aircraft lubrication systems, as booster pumps, and in many other fields in design configurations not markedly different from those of the present invention. Excellent papers are found in the literature that accurately predict the performance of a particular pump design from theoretical considerations.

The invention is illustrated by the accompanying drawings in which:

FIG. 1 is a schematic representation of a jet pump;

Fig. 2 is a longitudinal section of one jet pump design;

FIG. 3 is a transverse section on the line 3--3 of FIG. 2 and in the direction of the arrows;

FIG. 4 is a schematic representation of a jet pump in place in the heart as a replacement for the right heart;

FIG. 5 is a schematic representation of a jet pump in place in the heart as an assist to the right heart;

FIG. 6 is a schematic representation of the circulatory system including a jet pump as a right heart replacement;

FIG. 7 is a graphic representation of pulmonary arterial head generated as a function of driving flow rate;

FIG. 8 is a graphic representation of the effect of venous flow rate on pulmonary arterial pressure;

FIG. 9 is a graphic representation of pulmonary arterial head generated as a function of driving head for different values of venous flow rate; and

FIGS. 10, 11, 12, and 13 are modified forms of jet pump shown in longitudinal section.

As seen schematically in FIG. 1, a jet pump is basically a device by means of which one fluid is pumped by the action of mixing with another fluid. The pump, indicated generally at 10, is so simple that it entails no moving parts. The fluid which performs the pumping action is termed the driving flow, and the fluid to be pumped is termed the suction flow. The driving flow results from the high pressure of an independent source. As seen hereafter, the pump may take a variety of different specific forms.

The reason that a jet pump works is quite simple. High pressure driving fluid from a driving flow line 11 mixes with low pressure suction fluid from a suction plenum or chamber 12 in a mixing chamber 13, resulting in a total head intermediate between the driving and suction heads.

The suction plenum 12 and mixing chamber 13 are interconnected through a restricted annular orifice 14 between an inner driving nozzle 15 and an outer concentric suction nozzle 16. The driving nozzle may be held spaced within the suction nozzle as by means of radial struts 17. The pressure energy of the driving flow from flow line 11 is converted into velocity energy by means of the driving nozzle 15. At the plane of the driving nozzle exit, the static pressure is reduced to a value lower than in the suction plenum 12. Thus, the suction fluid is entrained into the suction nozzle 16. The two flows exchange momentum by turbulent mixing in the mixing chamber 13. It is customary to add a diverging tubular diffuser 18 to convert the velocity head in the mixing chamber 13 into pressure head at the diffuser exit 19.

The invention includes the application of a jet pump as either (1) a replacement [FIG. 4], or (2) an assist to the right heart [FIG. 5]. As seen in FIG. 4, when used as a replacement, the pump is placed into the venous circuit so that the suction flow is venous blood taken from the right atrium and the cavae. The sole purpose of the jet pump will be to generate the pulmonary arterial pressure and flow. The pressure source for the driving fluid of the jet pump is the left heart, or its artificial replacement. There is no limitation in regard to the type of left heart replacement used. One form of artificial replacement is a blood pump as described in copending application Ser. No. 816,952 filed Apr. 17, 1969 by Dorman et al., now U.S. Pat. No. 3,608,088. To consistently maintain the driving fluid pressure at a high level, the connection of flow line 11 is made to the aorta or one of its branches rather than to the left ventricle. Used with an artificial left heart, the addition of the jet pump constitutes a total heart replacement. Note that this means that the external power input to the body is only to one pump. As seen in FIG. 5, when used as an assist, the pump 10 encased in a flexible tubular housing 20 or the like is connected in parallel with the right heart.

A schematic representation of the jet pump placed into the circulation as a right heart replacement is shown in FIG. 6. The driving fluid is realized from the aorta or one of its branches and is mixed with the venous blood in the jet pump. The combined flow is then passed into the pulmonary circulation at the proper head. The pulmonary flow necessarily exceeds the systemic flow by an amount equal to the driving flow. This increase in pulmonary flow is not a significant factor, since it has been shown by several investigators that pulmonary flow may be increased three to four fold with little or no increase in pulmonary arterial pressure. However, the magnitude of this driving flow is of importance, and is minimized by careful design. Shunts such as shown schematically at 20 and 21 may be used to reduce the volume of flow to the pulmonary circulation. The pulmonary circulation system may include an artificial lung.

For the application shown in FIG. 4, typical characteristics of the pump are presented, based on the work of Cunningham reported in Jet Pump Theory and Performance with Fluids of High Viscosity, Trans. Amer. Soc. Mech. Eng. 79, 2, 1807, 1957. The design criteria are:

a. Aortic pressure of 100 mm Hg,

b. Venous pressure of 0 mm Hg, and

c. Venous flow rate of 6 L/min.

Once the pulmonary arterial pressure is dictated, an optimum pump may be designed, and the driving flow rate calculated.

FIG. 7 represents the pulmonary arterial head generated as a function of driving flow rate. Each location on the curve refers to different optimum pump dimensions. The right ventricle must generate on the average a head of 17 mm Hg. From the graph, it is seen that the driving flow rate must be about 4 liters/min. At flow rates not exceeding about 2.5 liters/min., the operation of the pump is impaired due to low Reynolds numbers.

The driving flow rate is dependent solely upon the driving pressure and the driving nozzle diameter and profile (shape of nozzle). In fact, the driving nozzle velocity depends only upon the driving pressure and nozzle profile, so that the flow rate is governed by the chosen nozzle. No control is necessary. At 100 mm Hg, the driving nozzle velocity is about 500 cm/sec.

One pump characteristic for a jet pump is a plot of pulmonary arterial head generated versus venous flow rate for fixed jet size. Consider a pump designed for the aforementioned criteria and also a pulmonary arterial pressure of 17 mm Hg. FIG. 8 is a graph of the effect of venous flow rate on pulmonary arterial pressure, assuming the pump is not operated at the design point. As specified, 17 mm Hg are generated at 6 liters/min. At 4 and 8 liters/min. the pulmonary pressure is about 22 and 11 mm Hg, respectively. Because the driving head is fixed at 100 mm Hg, the driving flow rate can be read from FIG. 7 at a head of 17 mm Hg, giving 4 liters/min.

A third characteristic is the effect of aortic pressure for a pump optimally designed to the above specifications. The aortic pressure is now allowed to assume values other than 100 mm Hg. FIG. 9 is a plot of pulmonary arterial head generated as a function of driving head (aortic pressure) for different values of venous flow rate. The top, middle, and bottom curves correspond to venous flow rates of 3, 6, and 9 liters/min. Because the driving nozzle diameter is fixed, the driving flow rate is dependent only upon driving pressure. The three driving flow rates, for pressures of 120, 100, and 80 mm Hg, are about 3.5, 4.0, and 4.4 liters/min., respectively. Referring to he graph, at a venous flow rate of 6 liters/min., the pulmonary arterial pressure varies from about 12 to 22 mm Hg, as the driving pressure varies from 80 to 120 mm Hg (e.g., pulsatile driving pressure).

A typical jet pump design for circulation application is illustrated in FIG. 2. It is designed for the specifications:

a. Aortic Pressure of 100 mm Hg

b. Venous pressure of 0 mm Hg

c. Pulmonary arterial pressure of 17 mm Hg

d. Venous flow rate of 6 liters/min.

The driving nozzle diameter is desirably between about 0.2 and 0.6 cm, with a mixing chamber diameter of between about 0.25 and 2.0 cm. The mixing chamber length may be up to 10 mixing chamber diameters. The suction nozzle tapers inwardly from a diameter between about 0.5 and 2.5 centimeters to between about 0.25 and 2 centimeters over a length between about 0.2 and 1 centimeter. A diffuser is optional, particularly in the case of infants. When present, it may have a taper between about 6.degree. and 30.degree.. The diffuser exits with a diameter which matches the pulmonary artery diameter, e.g., about 1 to 2.5 cm.

The total length of the pump, without tubing connections, is desirably on the order of 8 to 10 cm. The maximum outside diameter is desirably no more than 2 cm. Thus, the pump is quite small and compact. The driving nozzle is permanently attached to the mixing chamber, the mixing chamber and diffuser being of uni-construction.

In FIGS. 10 through 13, there are shown in longitudinal section modified forms of jet pump which are illustrative of the wide variety of specific jet pump design which may be utilized in the practice of the present invention. In each instance, the parts are numbered to correspond to those of FIGS. 1 and 2 with suffixes A through D added. In the embodiment of FIG. 10, the driving flow line 11A extends along the longitudinal axis through the end of the housing enclosing the suction chamber 12A. The mixing chamber 13C, restricted orifice 14A, driving nozzle 15A, suction nozzle 16A and diffuser 18A are all as previously described.

In FIG. 11, there is shown a Coanda type jet pump 10B in which the driving fluid is introduced through a flow line 11B and radial port into an annular channel 21 whose inner annular opening forms a jet driving nozzle 15B. The opening tapers radially inwardly and extends downstream. The driving fluid under high pressure exits from the driving nozzle and is diverted in a downstream direction by the Coanda effect creating the zone of lower pressure to draw fluid from the upstream suction chamber 12B through suction nozzle 16B. The mixing chamber 13B and diffuser 18B are as already described.

In FIG. 12, there is shown a somewhat similar type of pump 10C in which the driving fluid enters an annular chamber 22 in the housing through a radially extending port from driving flow line 11C. The driving fluid is ejected radially inwardly and is diverted downstream from the annular jet driving nozzle 15C to draw fluid through the suction nozzle 16C from the suction chamber 12C, which in the case of a circulation application may be the heart or a tube connecting to the heart. The mixing chamber 13C and diffuser 18C are as already described.

In FIG. 13, there is shown a further modified form of jet pump 10D. The driving flow line 11D, orifice 14D, driving nozzle 15D and suction nozzle 16D are generally as described in connection with FIGS. 1 and 2. However, this form of pump has no diffuser as such. Instead, the upstream end of mixing chamber 13D is provided with an annular cavity 23 of enlarged diameter surrounding the end of suction nozzle 16D. It is believed that eddy currents created by the flow into cavity 23 function to convert the velocity head into a pressure head at the exit from mixing chamber 13D and thus the modified structure performs a function similar to that of a diffuser. Alternatively, that portion of the pump housing downstream from the suction nozzle may be simply a straight segment of tubing, particularly where the pump is to be used to assist or replace the right heart of an infant.

The advantages of a jet pump are numerous. There are no moving parts, or mechanical drive-train, as with conventional pumps. No valves are needed. It can be lightweight, as rugged as desired, and small. Since the device is made of substantially rigid, substantially non-flexing material, there are no parts to wear out, dictating long term dependability. Of prime consideration is that the pump characteristics can be predicted successfully. The efficiencies are high due to all energy losses being confined to fluid friction and mixing losses.

The jet pump is a device that can take advantage, in regard to size, of the small work output of the right heart, whereas ventricular type pumps may not. The jet pump is small and can be used with any left ventricular pump, including the natural left heart. The right heart need not be excised. The absence of valves and moving parts is ideal. Surgical installation techniques are simple.

In combination with a replacement left heart, a total heart replacement is realized. External power is only to the left heart pump. If necessary, the right heart jet pump may be controlled, either through left heart control or by means of a valve in the driving flow. The latter involves regulating the driving flow.

Further advantages of the jet pump include: (1) minimal construction costs, (2) construction from almost any material, (3) disposability, (4) ease in sterilization, (5) can be coated with non-thrombogenic material, (6) high efficiency, (7) no service necessary, (8) simple and reproducible calibration, and (9) no heat transfer involvement.

Some disadvantages are present in a circulation application. The mixing process involves turbulent flow, and the velocities are modestly high. An independent pressure supply is needed for the driving flow, and this source is not considered as part of the jet pump.

Hemolysis may be important in two parts of the pump--the driving nozzle and the mixing chamber. At an aortic pressure of 100 mm Hg, the driving nozzle jets blood at a velocity of about 500 cm/sec, and the flow is turbulent. In the mixing chamber, the average velocity is reduced to about 200 cm/sec. prior to the diffuser, but the flow is also turbulent.

We have shown that by jetting plasma into whole blood at turbulent jet velocities not exceeding 1,700 cm/sec., no hemolysis is detectable. We similarly jetted whole blood through an orifice into whole blood, and detected no hemolysis at orifice velocities less than 1,000 cm/sec. Thus, hemolysis does not appear to be significant under the conditions that exist in a jet pump.

Further evidence that the jet pump will not be excessively hemolytic is found in work by Dorman et al who tested for hemolysis a centrifugal artificial left heart pump of the type described in the aforesaid application Ser. No. 816,952. The pump turbine produced turbulent velocities of 500 cm/sec., and exited the blood through a diffuser of dimensions about equal to a jet pump diffuser. Index of hemolysis of less than 0.01 (grams per 100 liters) was measured. The jet pump hemolysis should not exceed the rate found in this centrifugal pump.

Cavitation is often a concern in the mixing chamber of water jet pumps. In a circulation application, the minimum mixing chamber pressure is about - 20 mm Hg. At least 350 mm Hg of negative pressure is needed to induce blood cavitation. Thus, no cavitation is expected in the jet pump.

Blood clotting has been a continual problem in extracorporeal circuits. A jet pump has no dead spaces or stagnation regions, the diffuser being designed for no flow separation. Clotting is thus retarded, and the problem is of the same, or of a smaller, order of magnitude as in all other heart pumps.

The pump may be made of any of a variety of rigid and semi-rigid materials which are capable of maintaining the shape and alignment of the parts and are compatible with blood. For example, the pump may be electroformed by electroplating nickel over an expendable core and after removal of the core plated with gold or coated with a synthetic resin. It may be formed from polycarbonate. For long term support, the pump must be made of, or coated with, an anti-thrombogenic and anti-hemolytic material. The entire pump may desirably be manufactured from low energy materials, examples of which include fluorinated carbon based resins and fluorinated silicone rubbers.

It is apparent that many modifications and variations of this invention as hereinbefore set forth may be made without departing from the spirit and scope thereof. The specific embodiments described are given by way of example only and the invention is limited only by the terms of the appended claims.

Claims

The embodiments of the invention in which an exclusive property or privilege is claimed are defined as follows:

1. An implantable jet pump cardiac replacement and assist device comprising:

A. an elongated tubular housing, said housing being of a size adapted to implantation in a living being and of lightweight material compatible with body fluids,

B. an inwardly tapering suction nozzle inlet adjacent to the upstream end of said housing,

C. an inwardly tapering concentric driving nozzle within said tubular housing, the downstream end of said driving nozzle being in fluid communication with said suction nozzle and associated with said suction nozzle to create a zone of reduced pressure therein,

D. a suction plenum upstream from and in fluid communication with said suction nozzle, said plenum including means for connection to receive venous blood of the recipient living being,

E. means for connecting said driving nozzle to a source of arterial blood of the recipient living being,

F. a mixing chamber within said housing downstream from said suction and driving nozzles and in fluid communication therewith,

G. a discharge outlet downstream from said mixing chamber, and

H. means for connecting said discharge outlet to the pulmonary circulation system of the recipient living being.

2. A device according to claim 1 further characterized in that:

A. said driving nozzle is of lesser diameter than and upstream from the suction nozzle,

B. an annular orifice exists between said driving nozzle and said suction nozzle,

C. means are provided to hold said nozzles concentric.

3. A device according to claim 1 further characterized in that:

A. said driving nozzle is an annular orifice tapering radially inwardly and encircling the downstream end of said suction nozzle,

B. an annular chamber of greater diameter and greater cross-sectional area surrounds said driving nozzle, and

C. said driving nozzle is in direct fluid communication with said annular chamber and said annular chamber is in direct fluid communication with said means for connection with a source of arterial blood.

4. A device according to claim 1 further characterized in that said discharge outlet comprises an outwardly tapered diffuser.

5. a device according to claim 1 further characterized in that the blood contacting surfaces of the inside of the housing, the inside of the suction nozzle, the inside and outside of the drive nozzle, the inside of the suction plenum, the inside of the mixing chamber and the inside of said connecting means and discharge outlet are comprised of an anti-thrombogenic and anti-hemolytic material.

6. A jet pump cardiac replacement and assist device comprising:

A. an elongated tubular housing,

B. an inwardly tapering suction nozzle inlet adjacent to the upstream end of said housing, said suction nozzle tapering inwardly from a diameter between about 0.5 to 2.5 centimeters to between about 0.25 to 2 centimeters over a length between about 0.2 to 1 centimeter,

C. an inwardly tapering concentric driving nozzle within said tubular housing, the downstream end of said driving nozzle being in fluid communication with said suction nozzle and associated with said suction nozzle to create a zone of reduced pressure therein, said driving nozzle being of lesser diameter than and upstream from the suction nozzle, the diameter of said driving nozzle being between about 0.2 to 0.6 centimeter,

D. an annular orifice between said driving nozzle and said suction nozzle,

E. means to hold said nozzles concentric,

F. a suction plenum upstream from and in fluid communication with said suction nozzle,

G. means for connecting said driving nozzle to a source of arterial blood,

H. a mixing chamber within said housing downstream from said suction and driving nozzles and in fluid communication therewith, the diameter of said mixing chamber being between about 0.25 to 2 centimeters and its length being up to ten times its diameter,

I. a discharge outlet downstream from said mixing chamber, and

J. means for connecting said discharge outlet to the pulmonary circulation system.

7. A device according to claim 6 further characterized in that said discharge outlet comprises an outwardly tapered diffuser of between about 6.degree. and 30.degree. and an exit diameter between about 1 to 2.5 centimeters corresponding to the diameter of the pulmonary artery.

8. A method of at least partially replacing a disabled right heart in a living being which comprises:

A. connecting to a source of arterial blood under pressure the upstream end of the driving nozzle of a jet pump comprising:

1. an elongated tubular housing,

2. an inwardly tapering suction nozzle inlet adjacent to the upstream end of said housing,

3. an inwardly tapering concentric driving nozzle within said tubular housing, the downstream end of said driving nozzle being in fluid communication with said suction nozzle and associated with said suction nozzle to create a zone of reduced pressure therein,

4. a suction plenum upstream from and in fluid communication with said suction nozzle,

5. means for connecting said driving nozzle to a source of arterial blood,

6. a mixing chamber within said housing downstream from said suction and driving nozzles and in fluid communication therewith,

7. a discharge outlet downstream from said mixing chamber, and

8. means for connecting said discharge outlet to the pulmonary circulation system.

B. connecting the suction plenum of said jet pump to receive venous blood from the right atrium of the heart, and

C. connecting the discharge outlet from said jet pump to the pulmonary circulation system.

9. A method according to claim 8 further characterized in that the driving nozzle is connected to the aorta.

10. A method according to claim 8 further characterized in that the driving nozzle is connected to an artificial pump replacement for the left heart.

11. A method according to claim 8 further characterized in that the tricuspid and pulmonary valves of said right heart are made incompetent and said jet pump is implanted within the heart between the right ventricle and pulmonary artery as a total replacement of the right heart.

12. A method according to claim 8 further characterized in that one end of said jet pump is connected externally of the heart to the right atrium of the heart and the other end of said pump is connected externally of the heart to the pulmonary artery in parallel with the right heart as an assist to the right heart.