|Publication number||USRE40669 E1|
|Application number||US 11/156,362|
|Publication date||Mar 17, 2009|
|Filing date||Jun 17, 2005|
|Priority date||Aug 13, 2001|
|Also published as||CA2459543A1, DE60230876D1, EP1416877A2, EP1416877A4, EP1416877B1, US6579223, US20030032854, WO2003015841A2, WO2003015841A3|
|Publication number||11156362, 156362, US RE40669 E1, US RE40669E1, US-E1-RE40669, USRE40669 E1, USRE40669E1|
|Original Assignee||Arthur Palmer|
|Export Citation||BiBTeX, EndNote, RefMan|
|Patent Citations (38), Non-Patent Citations (19), Referenced by (2), Classifications (9), Legal Events (2)|
|External Links: USPTO, USPTO Assignment, Espacenet|
The present invention relates to pumps and more specifically to blood pumps, ventricular assist devices, and artificial hearts.
The natural heart functions in a fashion similar to a positive displacement pump. Each of the two pumping chambers in the natural heart has two check valves (an inlet and an outlet valve). The walls of the natural heart are made of contractile muscle that provide the power to pump the blood. Each pumping cycle consists of a filling or diastolic phase of the pumping cycle and an ejection or systolic phase of the pumping cycle. During the filling phase, the muscle fibers making up the walls of the heart relax allowing the chamber they surround to fill with blood. During the ejection phase of the cycle the muscle making up the walls of the heart contracts ejecting a portion of the blood from the chamber. The check valves assure one-way flow.
Mechanical blood pumps have been developed for use as artificial hearts to replace or assist the natural heart. Present blood pumps which are available to assist or replace the heart fall into two general categories. One category uses a rotary impeller and includes centrifugal pumps and axial flow pumps. The other category is pulsatile pumps, the diaphragm type pump being the most common. Blood pumps may also be classified as internal (intracorporeal) or external (extracorporeal) to the body.
Diaphragm pumps are favored as they provide desirable pulsative flow and are reliable owing to their simplicity. Prior art diaphragm pumps comprise a housing, a flexible but not extensible diaphragm that divides the interior of the housing into two chambers, namely a pumping chamber and a driving chamber. Diaphragms are conventionally fabricated from polyurethane, a flexible but not elastic material. The pumping chamber portion of the housing has an inlet and an outlet, each of which is equipped with a one-way flow check valve. The diaphragm is driven into and out of the pumping chamber mechanically, pneumatically or hydraulically. Mechanical drives typically include a pusher plate on the drive side of the diaphragm connected to a cam, solenoid or other device to impart reciprocal motion to the pusher plate and diaphragm. Alternatively, a drive fluid, either liquid or gas, may be used to reciprocally drive the diaphragm into and out of the pumping chamber.
One of the problems associated with available mechanical blood pumps is the formation of blood clots (thrombosis) in the pump. To address this problem, the interior surfaces of the diaphragm and housing walls that define the pumping chamber are typically designed to have a very smooth surface, in an effort to retard clotting. Other attempts to reduce clotting have involved provision of a rough texture on the interior surfaces of the pumping chamber to encourage endothelial cells, normally lining the heart and blood vessels, to grow over the surfaces eventually providing a smooth surface. Both of these methods work to some degree, but clotting in the device, with clots breaking off and entering the circulatory system, remains a problem.
Another problem relates to the flow of blood through the pump. Significant turbulence occurs in the chamber during the pumping cycle. There is little that can be done to control the characteristics of blood flow through the pumping chamber. There are areas of high velocity and other areas of slow flow. These slow flow areas also contribute to clotting. Turbulence leads to energy loss and inefficiency of the pump. Excessive turbulence may also damage the blood cells.
An additional problem is rupture of the diaphragm. If the diaphragm is driven pneumatically or hydraulically, should a tear or rupture of the diaphragm occur, the driving fluid may be pumped into the bloodstream, causing a harmful and potentially fatal embolism. Even if the pump is mechanically driven, a diaphragm rupture can result in air entering the bloodstream causing an embolism.
The foregoing are long standing problems in the art that have defied solution. There is, therefore, a need in the field for an improved blood pump and ventricular assist device.
It is an object of the invention to provide a blood pump that reduces the incidence of blood clotting.
It is another object of the invention to provide a blood pump with improved flow characteristics, particularly, to reduce or eliminate stagnant and low velocity flow areas within the pumping chamber to reduce blood clot formation, and to minimize areas of high turbulence to avoid damage to blood cells.
It is also an object of the invention to prevent intrusion of foreign matter into the bloodstream, and especially to prevent embolisms of the driving fluid or other fluids as a result of a pump failure.
In attainment of these and other objects and advantages of the invention, a pump is provided that has an elastic, extensible or stretchable bladder that expands in the filling phase and contracts in the ejection phase of the pumping cycle. The pump is particularly well suited for pumping blood, as in a ventricular assist device or a total artificial heart. However, the pump of the invention will find applications in other industries and non-medical fields for pumping fluids other than blood. The summary and following detailed description is in reference to, but is not limited to, blood pumping applications.
In a most basic embodiment, the blood pump comprises a bladder, the interior surface area and volume of which is changeable, i.e., it stretches and expands during the filling phase, and elastically contracts to its normal relaxed size during the ejection phase. The bladder has a fluid inlet and a fluid outlet. A device, such as a vacuum pump, compressor, solenoid or cam, alternately expands and contracts the interior surface area and volume of the bladder. A majority of the interior surface area of the bladder expands and contracts a significant amount (more than a few percent) in each cycle. One or more check valves or other means for causing substantially one-way fluid flow through the bladder are also provided.
Looking at the normal heart, there is very little tendency for blood clots to form in the heart when it is working normally. When it is working normally, the muscle which comprises the walls of the heart contracts with each ejection changing the surface area of the lining of the heart. After a patient has sustained a myocardial infarction (heart attack) a portion of the heart muscle comprising the wall of the heart has become necrotic (dead) and a scar has formed in that area. Because that area of the heart is now a scar, rather than muscle, and can no longer contract, it does not change the surface area of the lining of the heart in this localized area. It has been discovered that in this localized area of the natural heart (the area that does not contract due to a previous heart attack) there is a significant tendency for blood clots to form. This suggests that the change in surface area of the lining of the heart, with each pump cycle, is important in preventing clot formation on the lining of the heart. In a similar fashion, the changing of the surface area of the bladder of the invention as it stretches and contracts with each pumping cycle will decrease or eliminate clot formation on the surface of the bladder.
In a preferred embodiment, the blood pump of the invention comprises a housing, an extensible bladder in the housing, and a void volume or space between the housing and the bladder adapted to be occupied by a driving fluid. The bladder has an inlet and an outlet. At least one check valve is provided at the bladder inlet and/or outlet to provide one-way flow through the bladder. A vacuum source, compressor or other means is provided for altering the pressure of the driving fluid to alternately expand and contract the interior surface area and volume of the bladder. In the preferred embodiment, the driving fluid is a gas, and the driving means alternates pressure between comparatively high and low pressures, the high pressure being at or below atmospheric pressure and the low pressure being significantly below atmospheric. The application of the low pressure causes the bladder to expand and application of the high pressure causes the bladder to contract.
The invention also encompasses a method of pumping. A preferred method comprises the steps of (a) providing an extensible bladder having an inlet and an outlet; (b) connecting the inlet and outlet of the bladder to a person's circulatory system; (c) expanding the interior surface area and volume of the bladder to draw blood into the bladder through the inlet; (d) contracting the interior surface area and volume of the bladder to pump blood out of the outlet of the bladder; and (e) rhythmically repeating steps (c) and (d).
The bladder is preferably made of an elastic material that changes surface area during the pumping cycle. It expands or stretches during the filling phase of the pumping cycle and it returns elastically to its contracted size during the discharge or ejection phase of the pumping cycle. A majority of the interior surface area of the pumping chamber expands and contracts a significant amount (more than a few percent) in each cycle. The change in the area of the bladder surface during the pumping cycle will reduce the incidence and growth of blood clots forming on the surface of the bladder.
In addition, the blood pump of the invention may include variations in the thickness of the bladder and the material comprising the bladder in different areas, segments or portions of the bladder. The thinner areas will stretch more than the thick areas during the filling phase of the pumping cycle. This will draw more blood into the region of the pumping chamber surrounded by the thinner areas of the bladder. Varying the material in different areas of the bladder can also change the amount that various portions of the bladder stretch during the filling portion of the pumping cycle, and in addition, can change the speed at which different areas return to their neutral positions during the ejection part of the pumping cycle. Blood in some areas of the pumping chamber can thus be ejected earlier than blood in other areas. Accordingly, the characteristics of flow into the pumping chamber, through the pumping chamber, and out of the pumping chamber can be controlled and directed. Areas of stagnation can be minimized, further decreasing the likelihood of blood clot formation. Turbulence can also be minimized improving the efficiency of the pump and mitigating damage to blood cells.
In addition to varying the thickness and the material of the bladder, struts of varying elasticity can be molded into the bladder. These struts will bridge from one side to another side of the bladder and aid in maintaining the geometrical shape of the bladder. These struts may also be stretched during the filling portion of the pumping cycle and will provide additional force for ejection during the ejection part of the pumping cycle.
The filling phase of the pumping cycle is advantageously driven pneumatically or by other means for exerting below atmospheric pressure in the space between the housing and the bladder. During the filling phase, blood will be drawn into the pumping chamber and elastic energy will be stored in the bladder. The ejection phase of the cycle will then occur when the negative pressure is released and the bladder returns elastically to its neutral position. Although some positive pressure may be used on the bladder during the ejection phase of the pumping cycle, preferably and ideally there will be no positive pressure exerted on the bladder and all the force for ejection will come from the elastic recoil of the bladder. In this situation, when the power for ejection comes entirely from the elasticity of the bladder and no positive fluid pressure is exerted on the bladder, should a break or tear occur in the bladder, there is very little chance that any significant amount of the driving fluid would enter the circulatory system as there is no positive pressure to drive it through the tear or break in the bladder.
For the foregoing reasons, the blood pump of the invention decreases the likelihood of blood clots forming, improves the pumping characteristics of the device, and decreases or eliminates the chance of foreign fluids passing into the blood stream should a tear or break occur in the bladder. Although the pump of the invention was initially conceived for pumping blood, it also will find utility for pumping fluids in industrial and non-medical fields. Other attributes and benefits of the present invention will become apparent from the following detailed specification when read in conjunction with the accompanying drawings.
The following is a detailed description of certain embodiments of the invention presently deemed by the inventor to be the best mode of carrying out his invention. The invention as defined by the appended claims is not limited to these embodiments, and additional embodiments of the claimed inventive concept will undoubtedly be apparent to those skilled in the art.
Referring to the drawings,
The bladder may be constructed of any elastic or extensible material, such as a natural or artificial latex. It is important that the bladder be fabricated of biocompatible material, that is durable and capable of withstanding numerous expansion-contraction cycles.
It is further contemplated that the bladder could be formed of a semisolid material, and that the driving fluid could be the same semisolid material, the driving material and bladder being a unitary body or mass. Thereby the adding or withdrawing of semisolid material from the housing would cause the surface area and volume of the pumping chamber to expand and contract.
The housing 14 is fabricated of a rigid material such as titanium or a semi-rigid material such as an elastomer.
The check valves 20 and 22 may be natural, such as pig valves, or artificial, both of which are commercially available.
The driving fluid may be a gas, liquid or gel. Preferably, it is a gas, such as carbon dioxide. Carbon dioxide is desirable as it can be rapidly absorbed in blood in the event of a bladder tear or rupture. Although the bladder may be driven with positive pressure, it is preferred that operating pressures be at or below atmospheric. More specifically, the fluid is driven cyclically between a high pressure that is approximately atmospheric and a low pressure substantially below atmospheric. The requisite differential in pressure depends on various factors, including the resiliency of the bladder and the volume change within the bladder that is desired in each pumping cycle.
Upon completion of the discharge or ejection phase, the bladder is fully contracted, the check valves are closed, and the pump is prepared to repeat its cycle of operation at such pulse rate as may be dictated by the attending physician.
Throughout the pumping cycle, the interior surface area of the bladder changes. Different portions of the interior surface of the bladder may change to a smaller or greater degree. For example, the portions of the bladder adjacent to the inlet and outlet may expand/contract to a smaller degree than the central portion of the bladder as shown in
Furthermore, it is desirable that a majority of the overall interior surface area of the pump (i.e., the bladder, inlet, outlet and valves) comprise elastic or stretchable material that will change, i.e., expand and contract during the pumping cycle. It may be necessary or desirable, however, in some applications, to form valves or other components of inextensible material. For blood pump applications, it is desirable to maximize the interior surface area of the pump that will expand/contract. In accordance with the present invention, a majority of the interior surface area of the bladder and pump will expand/contract at least a few percent during the pumping cycle. Consequently, the likelihood of blood clot formation will be substantially reduced.
Due to the smooth internal configuration or shape of the bladder 12 and the manner in which blood is “milked” into and through the bladder, as depicted in
A second embodiment of the invention is illustrated in
The pump 110 comprises a cup or bag shaped bladder 112, comprised of an expandable and contractible elastic material. At its upper end, the bladder is bifurcated to form an inlet conduit 116 containing an inlet check valve 120 and an outlet conduit 118 containing an outlet check valve 122. Inlet 116 corresponds more or less to the vena cava of the human heart and outlet 118 corresponds more or less to the aorta of the human heart.
The bladder is encased in and sealed to the open end of a cup or sack shaped housing or shell 114 that is complementary to but of larger size than the bladder 12 and that defines a void volume, chamber or space 124 between the shell and the bladder. A conduit 126 establishes fluid communication between the space 124 and a means, such as a vacuum pump system hereinafter described, for cyclically evacuating and filling the chamber 124 with a bladder actuating or driving fluid.
In the neutral position shown in
As with the embodiment of
Another means for controlling the expansion and contraction rates of the bladder is shown in a fifth embodiment in FIG. 6. Again, like numerals refer to like parts, but in the 500 sequence. A driving fluid flow regulating means 560 is provided in the space 524 between the housing and bladder. One or more annular walls 562, 564, 566, 568 are provided between the housing and bladder, thereby compartmentalizing the annular space 24. Each wall may have one or more holes (not shown) to connect adjoining compartments. The number and sizes of the holes will regulate the flow rate between compartments, so that the pressure in each compartment at any given time may be controlled. On the initial application of a vacuum or low pressure through port 526, the pressure in space 570 will drop rapidly, while the pressure in the remaining compartments will drop at slower rates. As a result, the portion of the bladder adjacent compartment 570 will expand initially to a greater degree than the bladder portions adjacent other compartments. As the low pressure through port 526 is maintain, the pressure in each adjacent chamber 572-578 will equalize over a period of time, dependent on the number an size of the holes in the walls 562,564,566,568 separating compartments. This in turn will cause the bladder to expand in a sequence similar to that shown in
Alternative means are likely known to those skilled in the art for regulating fluid flow and fluid pressure in the space between the bladder and housing. For example, each compartment 570,572,574,576,578 could be sealed from the other compartments, and the pressure in each compartment could be independently regulated by plural fluid pressure control tubes connected to each compartment, respectively.
While preferred embodiments of the present invention have been shown and described, it is to be understood that these represent the best mode of practicing the invention contemplated by the inventors at the present time, and that various modifications and changes could be made thereto without departing from the scope of the invention as defined in the appended claim
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|Citing Patent||Filing date||Publication date||Applicant||Title|
|US8372145 *||Feb 12, 2013||Hisham M. F. SHERIF||Implantable artificial ventricle having low energy requirement|
|US20100298932 *||Aug 28, 2009||Nov 25, 2010||Sherif Hisham M F||Implantable artificial ventricle having low energy requirement|
|U.S. Classification||600/16, 623/3.1|
|International Classification||A61M1/10, A61M1/12, A61F2/00|
|Cooperative Classification||A61M1/1062, A61M1/1037, A61M1/122|
|Dec 17, 2010||FPAY||Fee payment|
Year of fee payment: 8
|Dec 17, 2014||FPAY||Fee payment|
Year of fee payment: 12