Mock Circulation

Abstract

The invention disclosed herein relates to a device for simulating the hydraulic impedance of the blood circulatory system for the purpose of evaluating artificial hearts, heart valves, heart bypass methods, and control systems for artificial hearts. The evaluation of these items relates both to their overall performance and to the durability of their materials of construction.

Description

With the growing interest in an artificial heart-pumping device to replace a malfunctioning natural heart, there has arisen a need for a hydraulic impedance device to "mock" the human body's blood-handling system in such a manner as to closely resemble the actual blood circulation system of the body. In this manner, an artificial heart can be realistically evaluated or tested as though it were supplying blood to the circulatory system of the body without the necessity of actually attaching the artificial heart to an experimental animal for testing. Such an impedance device should be adjustable over the expected range of impedances that could be expected from the blood circulation system of a body and be able to handle the expected flow output of an artificial heart. The device should also be relatively inexpensive to manufacture and compact in size. Such a device is disclosed herein.

The mock circulation system of this invention comprises a series of enclosed chambers through which the pumping fluid passes serially and into which a compressible gas can be individually added or withdrawn to alter the hydraulic impedance of the fluid as it passes through these enclosed chambers. The chambers can be either in the form of concentric cylindrical chambers or as chambers aligned linearly. In either modification, the fluid passes from the lower section of one chamber into the next chamber where it must be lifted over a separating baffle such that it can then pass into the next chamber in the series where it then exits from the lower portion of that chamber into another chamber. The chambers as described above can be repeated along the path of flow for as many chambers as it takes to supply the necessary impedance to the fluid.

Into each chamber a compressible gas can be either introduced or withdrawn so that the volume of the entrapped gas will suitably alter the impedance of the fluid as it passes through the chambers. The gas entrapped above the fluid in the chamber acts as a conventional surge-dampening device, or as an energy storage device, in that a pulse or surge of entering fluid causes the partial compression of the entrapped gas which compression tends to continue to force the fluid from the chamber through the outlet into the next chamber when the incoming or pressure pulse from the pump has ceased. Increasing the quantity of gas in the chamber will increase the resilience of the system which in turn will tend to decrease the maximum or peak of the incoming pressure pulse and smooth the resulting pressure waveform of an artificial heart or pulsatile pump.

The internal and frictional resistance to flow or viscosity of the fluid, the volume of the entrapped gas, the mass of the fluid, and the height that the fluid must be raised in each chamber all contribute to the impedance of the fluid in the mock circulation device when subjected to the pumping action of an artificial heart. All the above characteristics can be predetermined within certain limits by the number and size of the chambers and the diameter of the various openings between chambers through which the fluid must pass. Changes in the quantities of entrapped gas within the chambers will also alter the impedance characteristics within predetermined limits.

In most instances in the development of an artificial heart, the pulsatile pumping action of the natural heart will be duplicated. The waveform of the incoming pressure pulse or systolic pressure and the drop in pressure to its lowest point or diastolic pressure just before the next systolic pressure pulse will be duplicated by suitably adjusting the gas volume in each chamber of the mock circulation unit and in this manner, the artificial heart can be more effectively evaluated.

The quantity of compressible gas entrapped above the fluid in the chambers determines the fluid column height or head of fluid in each chamber. It is this head which determines the diastolic pressure presented by the entire mock circulation system. For example, a decrease in the volume of compressible gas above the fluid in each chamber causes a smaller difference in fluid levels between successive adjoining chambers and this decreased fluid level differential causes more rapid drop off in pressure from the peak systolic pressure to the diastolic pressure since it requires less time for the system to reach equilibrium between pulses of the pump.

A curve or graph may be plotted for an artificial heart attached to a mock circulation unit which curve represents the heart output per beat or systolic pressure as being directly proportional to the diastolic or filling pressure of the inlet reservoir to the pumping side of the artificial heart. Such a curve is called a Starling's curve and data for the curve is generated by allowing the fluid to drain away from the mock circulation unit rather than returning it to the inlet reservoir.

The relative size of the openings between the chambers should be sufficiently large to allow the free passage of fluid without imparting excessive resistance to the flow of the fluid unless of course it is desired to impart additional resistance to the flow of fluid through the mock circulation unit. Generally, to accomplish the relatively free passage of fluid, the openings between successive chambers will be larger than the inlet into the first chamber from the artificial heart. However, additional resistance can be created by a reduction in the size of the openings between successive chambers.

A venturi element in one of the chambers or in series therewith is calibrated and then utilized in conjunction with a pressure differential indicator to determine the rate of fluid flow through the mock circulation unit.

Since the heart functions essentially as two separate pumping cavities, or ventricles, each of which in turn has a separate receiving cavity, or atrium, it is envisioned that there could be a separate mock circulation unit affixed to each ventricle or pumping side of the artificial heart. In order to mock the total circulatory system of the body, the discharge conduit from each mock circulation unit will discharge into the receiving reservoir or atrium for the pumping cavity that supplies the other mock circulation unit.

This interconnection would serve to test the artificial heart for control over systemic and pulmonary pooling and to ascertain that each side responds to Starling's Law. For example, if the left side overpumps, the right atrial reservoir level will rise and if the artificial heart being tested responds correctly this will cause an increased right heart output which in turn will cause left atrial reservoir level to lower and the left heart output to decrease. In this manner the mock circulation will ascertain whether or not the right and left atrial filling pressures will assist a particular artificial heart in providing equal flows on the right and left sides. The only significant difference between the two systems is that the volumes of the chambers of the mock circulations device must be increased in a ratio corresponding to the difference between the aortic and pulmonary pressures. Thus, to obtain the proper impedance in the right or pulmonary system, the gas-containing chambers must be approximately three times as large as those in the left or aortic system.

A further advantage of this configuration is that it will be possible to monitor the response times of artificial heart control systems. System behavior can be determined by changing the reservoir levels (hence atrial filling pressure) and then observing the changes in the pool volumes and the flow rates.

As presently designed, the unit is relatively compact by reason of the serially disposed chambers which can be arranged in a variety of ways. Expense of manufacture can be held within tolerable limits since one embodiment of the mock circulation unit can be vacuum molded from sheets of commercially available plastic.

In view of the foregoing, it it an object of this invention to simulate the complete circulatory system of a body as such circulatory system would appear to an artificial heart.

Another object of this invention is to provide separate circulatory systems that simulate the fluid impedances encountered by the two pumping sides of an artificial heart.

A further object of this invention is to provide a means for varying the pumping impedance within each circulatory system.

A still further object is to provide a relatively inexpensive unit which serves as an effective means for evaluating the overall performance and durability of an artificial heart.

Another object is to provide body circulatory system simulator that is relatively compact in size.

These and other objects will become obvious when viewed in light of the accompanying drawing and description.

The drawing is a cross section of the channels of one embodiment of the mock circulation unit.

The mock circulation unit as depicted can serve as the impedance simulator for either the pulmonary or aortic circulatory system of the body, the only differences being the relative sizes of the chambers as has been previously discussed. For reasons of simplicity, the unit shown in this figure will be called the aortic circulatory impedance simulator system although the ensuing discussion would be equally applicable to the pulmonary system.

Fluid entering the mock circulation system from the artificial heart enters through inlet 10 into the first chamber 11 where it is subjected to the pressure of the entrapped gas of space 12 which gas can be either introduced or withdrawn through valve 13 to alter the gas volume of chamber 11.

Fluid departs chamber 11 through an exit located in the lower half of chamber 11 and enters the next serially disposed chamber 14 through an opening located near the bottom of that chamber. The fluid must then be elevated a distance which represents the difference between the level the fluid would rise to in chamber 14 resulting from the various gas pressures in all the chambers and the top of the barrier 15 separating the fluid of chamber 14 from the fluid of chamber 16. The volume of entrapped air above the fluid in chambers 14 and 16 can be suitably altered by means of valve 17 to impart the desired impedance characteristics to the flow of fluids through these two chambers. The fluid departs chamber 16 into the next serially disposed chamber in a manner similar to that previously described for chamber 11.

Double-chambers similar to chambers 14 and 16 repeat the functions of these two chambers in imparting the total desired impedance to the fluid as it passes through the chambers. The total number of chambers in the system will depend upon the amount of impedance it is desired to impart to the fluid. In the presently preferred embodiment there are three double-chambers similar to chambers 14 and 16 and one inlet chamber 11. The final chamber is occupied by a venturi flow meter 18 which when used in conjunction with a pressure differential indicator 24, is used for the measurement of the volume of flow of fluids through the mock circulation system.

Upon exit from the venturi flow meter 18, the fluid enters a reservoir 19 which communicates directly with a heart pump inlet 20 which is located near the upper terminus of an inverted enclosure 21. Enclosure 21 is so designed that it allows fluid communication between reservoir 19 and pump inlet 20 only below the lower extremity of enclosure 21 which lower extremity is below the pump inlet 20 to prevent the aspiration of air through pump inlet 20 when the level of reservoir 19 drops below that of the pump inlet 20.

An area for the excavation of a tunnel through the left or aortic mock circulation unit for the pulmonary circulation or right atrial connection is shown by dashed lines at 22. In the embodiment where both the left and right mock circulation units are used in conjunction to evaluate the total performance of an artificial heart, the area indicated by 22 is a tunnel through the left or aortic mock circulation unit for the passage of a connecting conduit from the reservoir of the pulmonary or right mock circulation unit to the right atrium of the artificial heart. Another area for the excavation of a tunnel 23 similar to the area for tunnel 22 is for a conduit that connects the inlet to the right or pulmonary mock circulation unit with the right side of the artificial heart. Both tunnels serve as passages through the left or aortic mock circulation unit and in no way communicate with the interior of the aortic mock circulation unit. The tunnels also allow the shortest possible connections between the artificial heart and the mock circulation units to reduce the inherent impedances in the connecting conduits.

The particular locations shown for the various connection to the artificial heart also reduce the length of the connecting conduits since they represent the relative positions of blood vessels that attach to the natural heart. In addition, the layout of the conduit connections on the artificial heart duplicate the natural heart since an artificial heart is constructed for placement in the same location in the body as the natural heart.

In operation, fluid from the pump enters the first chamber wherein the volume of gas entrapped above the fluid in the chamber is partially compressed by the surge of incoming fluid and in this manner the gas of this chamber acts as a conventional surge-dampening device. The partially compressed gas also tends to continue to force the fluid from the first chamber into the next chamber after the input pulse from the pump has ceased. The entrapped gas above the fluid in all chambers acts in a manner similar to that previously described. Altering the volumes of the entrapped gas in any of the chambers will alter the total impedance of the mock circulation unit. However, it has been found that if the volume of the entrapped gas has been changed, the level of the fluid in the atrial reservoir must be readjusted to the predetermined level since changes in the atrial reservoir will alter the output of the artificial heart.

Claims

We claim:

1. In an apparatus for applying a complex impedance to a fluid from a pumping device undergoing evaluation, said fluid impedance apparatus comprising a plurality of serially disposed, interconnected, and enclosed chambers having at least one vertical partition disposed in each chamber to direct said fluid in a predetermined path and wherein the volume of a compressible gas entrapped above the fluid in each chamber can be suitably altered to simulate various impedances to the flow of fluid through the apparatus, flow-rate-measuring means interposed in said fluid stream, and a fluid reservoir means for retention of said fluid upon exit from said apparatus and before said fluid is returned to said pumping device.

2. An apparatus for applying a complex pumping impedance to a fluid pumping device as described in claim 1 wherein the pumping device is an artificial heart.

3. An apparatus for applying a complex pumping impedance to a fluid pumping device as described in claim 1 wherein the pumping impedance simulates at least a portion of the blood circulatory system of a body.

4. An apparatus for applying a complex pumping impedance to a fluid pumping device as described in claim 1 wherein a separate pumping impedance device is utilized to simulate an aortic circulatory system and another separate pumping impedance device is utilized to simulate a pulmonary circulatory system.

5. An apparatus for applying a complex pumping impedance to a fluid-pumping device as described in claim 4 wherein the two separate pumping impedance devices are suitably interconnected with the artificial heart pumping device to simulate the total circulatory system of a body.