CROSS-REFERENCE TO RELATED APPLICATIONS  
       [0001]     This application claims priority of U.S. Provisional Patent Application Ser. No. 60/778,170 filed Feb. 27, 2006, which is incorporated herein by reference 
     
    
     FIELD OF THE INVENTION  
       [0002]     The present invention relates generally to a cardiac assist device, and in particular to a left ventricular assist device (“LVAD”) intended to work in series with a patient heart to augment tissue perfusion.  
       BACKGROUND OF THE INVENTION  
       [0003]     For purposes of background, the disclosures of the following patent documents are hereby incorporated by reference in their entirety: U.S. Pat. Nos. 4,051,840; 4,630,597; 4,692,148; 4,733,652; 4,809,681; 5,169,379; 5,761,019; 5,833,619; 5,904,666; 6,042,532; 6,132,363; 6,471,633; 6,511,412; 6,735,532; and U.S. patent application Ser. Nos. 10/746,543; 10/770,269; 10/865,965; 11/178,969; and 60/709,323.  
         [0004]     The scarcity of human hearts available for transplant, as well as the logistics necessary to undertake heart transplant surgery, make an implantable cardiac assist device the only viable option for many heart patients. An aortic blood pump, for example, can be permanently surgically implanted in the wall of the aorta to augment the pumping action of the heart  
         [0005]     A known aortic blood pump includes a flexible bladder to be inflated and deflated in a predetermined synchronous pattern with respect to the diastole and systole of the patient to elevate aortic blood pressure immediately after aortic valve closure. Inflation and deflation of the bladder is accomplished by means of a supply tube connected to the bladder and to a percutaneous access device (“PAD”). The PAD is permanently surgically implanted in a patient&#39;s body to provide a through-the-skin coupling for connecting the supply tube to an extra-corporeal fluid pressure source. Electrical leads from electrodes implanted in the myocardium are likewise brought out through the skin by means of the PAD, The “R” wave of the electrocardiograph is used to control the fluid pressure source to inflate and deflate the inflatable chamber in a predetermined synchronous relationship with the heart action.  
         [0006]     The aortic blood pump acts to assist or augment the function of the left ventricle and is typically restricted to use in patients who have some functioning myocardium. The aortic blood pump does not need to be operated all the time, and in fact, can be operated periodically on a scheduled on-time, off-time regimen, or on an as-needed basis. Typically, the patient can be at least temporarily independent of the device for periods of one to four hours or more, depending on their heart function and level of activity. The general structure of known aortic blood pumps is a semi-rigid concave shell, and a flexible membrane that is integrally bonded to the outer surface of the shell, forming an inflatable and deflatable chamber. A fabric layer is then bonded over the exterior surface of the shell that projects clear of the shell forming a suture flange. These blood pumps have been tested and demonstrated to last a few million cycles. None of the known blood pumps disclose or suggest that any modification can be made to the geometry of the shell and membrane to increase the durability of the pump, much less what such modification would be.  
         [0007]     A known dynamic aortic patch has an elongate bladder having a semi-rigid shell with walls of uniform thickness and a relatively thicker peripheral edge and a flexible, relatively thin membrane defining an inflatable chamber. At least one passage extends through the shell defining an opening in the inner surface of the shell. The flexible membrane is continuously bonded to the shell adjacent the peripheral side edge to define the enclosed inflatable chamber in communication with the passage. The membrane may have a reduced waist portion, defining a membrane tension zone adjacent to the opening of the passage into the chamber to prevent occluding the opening to the pneumatic supply while deflating the chamber. An outer fabric layer can be bonded to the outer side of the shell of the aortic blood pump, and present a freely projecting peripheral edge to provide a suture flange for suturing the aortic blood pump in place within an incision in the aorta.  
         [0008]     Known aortic blood pumps use an inflatable bladder and an envelope. The envelope is sutured to the aorta and then the bladder is placed inside the envelope. Although this design successfully augments the blood pumping capacity of the heart, it has two major disadvantages. First, fluid may accumulate inside the envelope, between the envelope and the inflatable bladder. This accumulation of static fluid within the body commonly leads to infection. Second, due to the geometry of the bladder, the volume of blood displaced by the device is limited, and has been determined to be insufficient.  
         [0009]     Experience with patients has shown that it is relatively easy to construct a pump that will last a few million inflation-deflation cycles (on the order of weeks). However, it is very difficult to design, reproducibly manufacture, and implant a pump that will last for at least two years (on the order of a hundred million of inflation-deflation cycles, or more) without membrane failure.  
         [0010]     The top surface of the pump&#39;s shell can be overlaid with a non-tissue adhesive substance, such as silicone, to prevent scar tissue from adhering to the back of the pump and to allow the pump to be explanted later. But clinical experience has shown that even this improved design may last less than the two-year target in a patient.  
         [0011]     Known blood pumps have a suture ring placement that constrains the movement of the blood pump during each inflation-deflation cycle. In these designs, the suture ring is located closely adjacent to the shell bead, in a location outside of the periphery of the shell, and at approximately the same height (measured as the axial distance from the centerline of the aorta) as that of the bead. When the implantation wound heals, the suture line itself, as well as the scar tissue that grows into the suture line, constrain the movement of the shell during each inflation-deflation cycle. This occurrence results in effectively stiffening the shell near the region where it interacts with the membrane, thus forcing the membrane to absorb all of the stress during the inflation-deflation cycles.  
         [0012]     The hose barb provides the connection between the internal conduit and the blood pump. Known blood pumps have hose barbs that are glued into place to the back of the shell of the blood pump. This design can be improved to increase the strength of the hose barb&#39;s attachment to the shell.  
         [0013]     As seen in  FIG. 11 , the shells of prior art blood pumps are relatively flat across their length, other than slightly turning downwards at the longitudinal ends, and have relatively thin walls of uniform thickness with slightly thicker peripheral edges. However, despite the simplified drawings of aorta in  FIGS. 2, 3 , and  11 , the human aorta is not a straight circular cylinder. Rather, it has a complex three-dimensional shape, sometimes described as a “twisted question mark.” Accordingly, the known flat blood pumps are not well configured to fit to a typical human aorta, and there is a need in the art for a blood pump having a contour that generally matches the contour of a typical human aorta. Further, because of their general cylindrical configuration and relatively thin walls, the permanent deformation of these pumps during surgical implantation into the non-cylindrical aorta can affect their durability.  
         [0014]     Thus, although the art discloses the basic concept of an “in-series” mechanical ventricle assist device blood pump), having a semi-rigid shell, and a flexible membrane, nothing in the art teaches or suggests how to construct a device that will be durable enough to survive inflation-deflation cycles for the number of years desired. To the contrary, clinical experience has shown that the known blood pumps generally last less than the two-year target. Thus, there remains a need in the art for a blood pump design providing increased durability.  
       SUMMARY OF THE INVENTION  
       [0015]     A blood pump for placement in an incision in an aorta is provided that after placement contacts blood passing through the aorta. The pump inflates and deflates in order to provide left ventricular assistance. The pump has an elongated shell with a generally elliptical shape, an outer convex surface and an inner concave surface. A peripheral side edge located between the inner and outer surfaces terminates in a bead edge. A passage is provided through the shell to provide fluid communication between the outer surface and inner surface. A flexible airtight membrane has a membrane edge bonded to the outer shell surface adjacent to the bead edge to form an enclosed internal chamber in fluid communication with the passage. Preforming the membrane edge looped with a maximum linear span of curvature that is greater than a maximal transverse linear extent of the bead edge, membrane operational wear during inflation and deflation cycles is reduced in the region around the bead edge.  
         [0016]     A process of forming a blood pump with a membrane preform is provided that includes placing an airtight membrane around a platen having a platen edge bead with a curvature of maximal transverse linear extent and a platen footprint substantially identical to a footprint of the blood pump shell. By heat setting the membrane, a looped membrane edge is formed as complementary to the curvature of the platen edge bead to yield a membrane preform. The membrane preform is secured to the outer surface of the shell having a bead edge of maximum linear extent less than the curvature of the looped membrane edge. 
     
    
     BRIEF DESCRIPTION OF THE DRAWINGS  
       [0017]     Preferred features of the present invention are disclosed in the accompanying drawings, wherein similar reference characters denote similar elements throughout the several views, and wherein:  
         [0018]      FIG. 1  is a schematic perspective view of the major components of an LVAD system, as known in the art, implanted in a patient.  
         [0019]      FIG. 2  is a schematic view of a longitudinal cross-sectional view of a blood pump in a deflated state.  
         [0020]      FIG. 3  is a schematic view of a longitudinal cross-sectional view of a blood pump in an inflated state.  
         [0021]      FIG. 4  is a schematic view of a transverse planar cross-sectional view along line  4 - 4 ′ of the pump shown in  FIG. 2  in a deflated state.  
         [0022]      FIG. 5  is a schematic view of a transverse planar cross-sectional view of a blood pump in a deflated state, in an embodiment with a shell having a hollow bead.  
         [0023]      FIG. 6  is a schematic view of a transverse planar cross-sectional view of a blood pump in a deflated state, in an embodiment with the shell&#39;s wall region decreasing in thickess adjacent to the bead.  
         [0024]      FIG. 7  is a schematic view of a transverse cross-sectional view of a blood pump in a deflated state, in an embodiment with an air pocket between the membrane and the bead.  
         [0025]      FIG. 8  is a perspective view of a blood pump shell, including a hose barb.  
         [0026]      FIG. 9  is a longitudinal cross-sectional view of a blood pump shell, including a hose barb.  
         [0027]      FIG. 10  is a perspective view of a blood pump shell, including a hose barb and a suture ring.  
         [0028]      FIG. 11  is a longitudinal cross-sectional view of a prior art blood pump shell.  
         [0029]      FIG. 12  is a perspective view of a cylindrical blood pump shell, including a hose barb.  
         [0030]      FIG. 13  is a longitudinal cross-sectional view of a cylindrical blood pump shell, including a hose barb.  
         [0031]      FIG. 14  is a schematic view of a longitudinal cross-sectional view of a cylindrical blood pump with no vacuum/pressure applied to the membrane.  
         [0032]      FIG. 15  is a schematic view of a transverse cross-sectional view of a cylindrical blood pump with no vacuum/pressure applied to the membrane. 
     
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS  
       [0033]     The present invention has utility as an implantable cardiac assist device. Through a modification of shell structure to include an enlarged bead edge relative to the shell wall the operational stability of the pump is enhanced.  
         [0034]     For convenience, the same or equivalent elements in the various embodiments of the invention illustrated in the drawings have been identified with the same reference numerals. Further, in the description that follows, any reference to either orientation or direction is intended primarily for the convenience of description and is not intended in any way to limit the scope of the present invention thereto.  
         [0035]     A basic known LVAD system, as shown in  FIG. 1 , consists of the blood pump  2 , an inflatable bladder sutured into the wall of the aorta; an internal conduit  4  connecting the blood pump to the percutaneous access device (“PAD”); the PAD  6 , a through-the-skin port that permits power, electrical signals and fluid (typically air) to pass between the drive unit and blood pump; and the external drive unit  10 , a device powering and controlling the blood pump. The PAD  6  allows the implanted blood pump  2  to be operatively connected to or disconnected from the external drive unit  10 . To inflate the blood pump  2 , pressurized air is supplied from the drive unit compressor (not shown). The air flows from the compressor via an interconnect tube through a valve manifold in the drive unit  10  to an external drive line  8  running to the PAD  6  and then through the implanted internal conduit  4  to the blood pump  2 . Alternatively, an isolation chamber, separating the pressure (or vacuum) source from the air flowing to the pump, can be used to isolate the subcutaneous portion of the pneumatic circuit from the supercutaneous portion.  
         [0036]     As seen in the longitudinal cross-sectional views of  FIGS. 2 and 3 , the improved blood pump  20  of the present invention is implanted within the wall of the thoracic aorta  22 . The membrane  38  of the blood pump  20  is illustrated during deflation in  FIG. 2  and during inflation in  FIG. 3 . To implant the device, a surgeon makes a longitudinal incision through the wall of the aorta, usually downward from a location just below the subclavian artery, and the device is placed within the incision and sewn firmly in position by sutures (not shown) passing through a projecting suture ring flange  48 .  
         [0037]     The outer side of the blood pump  20 , as implanted, is a relatively thick, semi-rigid shell  26 , which is molded from a suitable biocompatible material, such as urethane. As best seen in  FIG. 8 , the shell  26  is of an elongate elliptical shape, with its upper or outer surface  33  being convex in both its longitudinal and transverse extents. The lower or inner surface  31  of the shell  26  is concave in its longitudinal and transverse extents, as seen in  FIGS. 2-5 .  
         [0038]     The shell  26  can be considered as having three regions: the spine region  57 , the wall region  55 , and the bead region  53 . The spine region  57  of the shell  26  is the relatively thick area of the shell that forms the housing  41  around the hose barb  40 , together with the area immediately surrounding the housing  41 . A passage  28  extends though the shell  26  to place the interior volume  30  of the blood pump  20  in fluid communication with the internal conduit  32  of the blood pump  20 .  
         [0039]     The wall region  55  of the shell  26  is preferably thinner than the spine region  57  and extends out from the spine region  57  as seen in  FIGS. 2 and 3 .  
         [0040]     The bead region  53  of the shell has a bead  34 , and is thicker than the wall region  55 . The peripheral side edge of the shell  26  is smoothly rounded around its circumference, as best seen in  FIG. 2  and  3 , and around its cross section, as best seen in  FIGS. 2-4 , such that the bead  34 , which runs along the periphery of the shell  26 , is itself smoothly rounded throughout its entire extent, circumferentially around the elliptical shell, and also around its cross section. The shape of the bead  34  lessens local flexing stress, particularly when the membrane  38  is taut around the edge during the pump deflation cycle.  
         [0041]     The membrane  38  is flexible, thin walled and bonded to the outer surface  33  of the shell  26 . The membrane  38  is preferably not adhered to the bead  34  or inner surface  31  of the shell  26 . Known solvation bonding techniques such as chlorinated solvent welding of polymers result in the membrane  38  and the shell  26  becoming what is in effect a unitary structure, but for purposes of explanation, the membrane  38  and the shell  26  are drawn in  FIGS. 2 and 3  as separate components.  
         [0042]     The outer surface  39  of the membrane  38  which, when implanted, interfaces with the blood in the aorta is preferably provided with fibrils, forming a textured surface similar to a flocking, to promote cellular adhesion in forming a pseudointima on the outer surface  39  of the membrane  38 .  
         [0043]     A piece of sheet material  43  is attached to the outer surface  33  of the shell  26 . As indicated in  FIGS. 2 and 3 , the periphery of the material  43  is preferably not attached to the outer surface  33  of the shell  26  and projects freely from the shell  26  to create a flange  48 . An additional material strip  49 , such as a pledget strip, is secured to the flange  48  by for example stitching, to provide a suture ring  44  for implanting the device in an incision in the aorta, as best shown in  FIGS. 2, 3  and  10 . The strip  49  preferably has a fibrous surface allowing body tissues to migrate into, and mechanically interweaving with the strip  49 , to augment the sealing action initially established by the surgical implantation sutures (not shown) between the flange  48  of the suture ring  44  and the wall  50  of the aorta  22 . The sheet material  43  and strip  49  are made of various appropriate materials, such as polyester, which are commercially available and have been certified for use in implanted devices.  
         [0044]     A number of modifications are optionally made to increase the durability of the blood pump by reducing the physical wear and stress on the membrane. The membrane failure of prior art designs is understood to have often been caused by microperforations due to membrane deformation, or “creasing,” during inflation-deflation cycles.  
         [0045]     A blood pump is provided that has a contour that generally matches that of a typical human aorta. As seen in  FIGS. 2, 3 , and  9 , the blood pump shell has a pronounced, continuous curve across its entire longitudinal extent. The interior curve of the shell as seen in  FIG. 9  corresponds to the curvature of the aorta at the point of blood pump implantation. This curvature allows the pump to match the contour of a typical human aorta better than prior art pumps.  
         [0046]     An inventive blood pump is provided with a large-diameter bead relative to the adjoining wall region of a pump shell to reduce the stress on the membrane during the cyclical operation of the pump. The large-diameter bead reduces stresses on the parts of the membrane coming in contact with the bead during the deflation cycle, while still being sufficiently flexible to be twisted and flexed during surgical implantation. Such properties are derived by varying the bead diameter relative to the shell thickness. As seen in the transverse planar cross-sectional view of the pump in  FIG. 4A , the bead  34  has a membrane-contacting region  69  which contacts the membrane  38  when the pump is deflated. The portions of the pump  20  not in the plane of the transverse cut are not shown for visual clarity. As seen in  FIG. 4A , the transverse wall region  65  of the shell  26  has a thickness T adjacent to the bead region  63 . A transverse spine region  67  is also defined. Preferably, the thickness of regions  55 - 65 ,  53 - 63  and  57 - 67  are substantially equivalent. In order to further reduce localized stresses on the membrane  38  during the deflation cycle, the maximal linear extent D of the bead  34  at the membrane-contacting region  69  is preferably between about 110% to about 700% of T. In other words, ≠1.1 T≦D≦≈7 T. More preferably, D is between about 200% to about 600% of T. In other words, ≈2 T≦D≦6 T. Even more preferably, D is between about 400% to about 500% of T. In other words, ≈4 T≦D≦≈5 T. It is appreciated that while the bead  34  is depicted as spherical in cross section resulting in the maximal linear extent D corresponding to the diameter of the bead and that the bead is readily formed in a variety of shapes devoid of a sharp corner corresponding to a mathematical derivative singularity. An alternate bead shape in cross section is ellipsoidal, as shown in  FIG. 4B .  
         [0047]     In absolute dimensions, mathematical modeling/simulation and laboratory testing was conducted on a pump with the following approximate dimensions: 
        Overall length=130 mm     Overall width=38 mm     Wall thickness in wall region=2 mm        
 
         [0051]     The testing indicated that on this pump, a bead diameter of approximately 7.0 mm advantageously reduced localized stress on the membrane, while still providing a shell with sufficient flexibility for purposes of implantation, and sufficient rigidity to hold its shape as required during implantation and use in vitro.  
         [0052]     An inventive shell  70  is provided with a hollow bead  72 , as seen in the transverse cross-sectional view of  FIG. 5 . The shell  70  is comparable to shell  26  with the exception of the hollow bead construction. This bead  72  has a tubular hollow region  74  running along the entire length of the bead  72 , considering the bead length running round the circumference of the side edge of the shell  70 . This hollow bead  72  reduces the rigidity of the shell  70 , at the interface with the membrane  38 , in comparison to the shell with a solid bead of the same diameter (as shown in  FIG. 4 ). By reducing the rigidity of the shell  70  at the bead  72 , where the shell  70  interacts with the membrane  38 , the shell  70  deforms more during the inflation-deflation cycle (as compared to a shell with a solid bead), thereby allowing the shell  70  to share with the membrane  38  more of the stress caused by the cycling. It will be appreciated that if the shell is very rigid relative to the membrane (for example if it has a large-diameter solid bead), the membrane  38  is forced to absorb a greater share of the stress related to the cycling of the pump.  
         [0053]     The thickness of the wall region  86  of the inventive shell  80  decreases from a maximum near the spine region  67  to a minimum near the bead region  63 , so that T 1  is greater than T 2 , as seen in the transverse cross-sectional view of  FIG. 6 . The thinner wall region T 2  decreases rigidity of the shell  80  at the bead region  63  and consequently at the bead membrane contacting region  90  with the membrane  38 , in comparison to the shell  70  with a wall region of uniform thickness (as shown in  FIG. 5 ). By reducing the rigidity of the shell  80  at the bead region  63 , where the shell  80  interacts with the membrane  38 , the shell  80  deforms more during the inflation-deflation cycle (as compared to a shell with a wall region of uniform thickness), thereby allowing the shell  80  to share more of the stress caused by the cycling with the membrane  38 .  
         [0054]     In yet another embodiment, the membrane is modified, as compared to known pump membranes, to enhance the durability of the blood pump. In particular, the membrane  38  is preferably formed such that even without an applied pressure or vacuum (that is, without any significant pressure differential from one side of the membrane to the other), membrane shape generally matches that of the curved inner surface  31  of the shell  26 , as seen in  FIG. 2 . This curved membrane is formed on a curved surface. This curved membrane experiences less creasing during the inflation-deflation cycles, as compared to flat membranes. Practical experience has shown that flat membranes crease at certain cusp points on the curved shell when they transition between inflated and deflated states. Analysis and testing indicate that the curved membrane reduces the magnitude and occurrence of this creasing problem.  
         [0055]     In another embodiment, the blood pump is provided with an air pocket between the membrane and the bead, when the blood pump is in a deflated state, to reduce the stress on the membrane during the cyclical operation of the pump. As seen in the transverse cross-sectional view of the pump in  FIG. 4A , the bead has a membrane-contacting region  69  which typically contacts the portion of the membrane when the pump is deflated. As shown in  FIG. 7 , membrane  102  does not contact the bead  104  around the entire cross-sectional diameter of the bead  104 , due to air pockets  106  which form in proximity to the bead  104  during the deflation cycle. By heat setting the membrane on a large bead shell forming platen, and then bonding the resulting membrane  102  to shell  100  with a relatively smaller bead  104 , during deflation, the diameter D 2  of the air pocket created by the membrane  102  wrapping around the bead  104  is larger than the diameter D 1  of the bead  104 . In order to further reduce localized stresses on the membrane  102  during the deflation cycle, the diameter D 2  of the air pocket  106  created by the membrane  102  is always greater than the bead diameter D 1 . In other words, D 2 &gt;D 1 . The local diameter D 1  of the bead  104  in this embodiment is preferably between about 110% to about 700% of the wall thickness T. In other words, ¢1.1 T≦D 1 ≦≈7 T. More preferably, D 1  is between about 110% to about 300% of T. In other words, ≈1.1 T≦D 1 ≦≈3 T. An air pocket  106  is formed in the gap between the membrane  102  and the bead  104 . The air pocket increases the radius of curvature of the membrane  102  near the bead  104 , thus reducing the strain on the membrane  102  during inflation-deflation cycles.  
         [0056]     As best seen in  FIGS. 12-15 , a shell  326  is relatively flat along a shell length with slightly downturned ends at the longitudinal termini, and a general cylindrical shape, with an upper or outer surface  339  that is convex in both longitudinal and transverse extents. The lower or inner surface  331  of the shell  326  is concave in its longitudinal and transverse extents.  
         [0057]     The shell  326  can be considered as having three regions: the spine region  357 , the wall region  355 , and the bead region  353 . The spine region  357  of the shell  326  is the relatively thick area of the shell that forms the housing  341  around the hose barb  340 , together with the area immediately surrounding the housing  341 . A passage  328  extends though the shell  326  to place the interior volume  330  of the blood pump  320  in fluid communication with the conduit  332  of the blood pump  20 .  
         [0058]     The wall region  355  of the shell  326  is preferably thinner than the spine region  357  and extends out from the spine region  357  as seen in  FIG. 14 .  
         [0059]     The bead region  353  of the shell  326  has a bead  334 , and is thicker than the wall region  355 . The bead  334  extends around the periphery of the shell  326 , as best seen in  FIGS. 14 and 15 . The bead  334  is smoothly rounded throughout its entire extent, circumferentially around the cylindrical, elliptical shell  326 . The bead  334  minimizes local flexing stress, particularly when the membrane  338  is taut around the edge during the pump deflation cycle.  
         [0060]     The membrane  338  is flexible, thin walled and is bonded to the outer surface  333  of the shell  326 . The membrane  338  is preferably not adhered to the bead  334  and inner surface  331  of the shell  326 . Known solvation bonding techniques such as chlorinated solvent welding of polymers result in the membrane  338  and the shell  326  becoming what is in effect a unitary structure, but for purposes of explanation, the membrane  338  and the shell  326  are drawn in  FIGS. 14 and 15  as separate components.  
         [0061]     The outer surface  339  of the membrane  338  which, when implanted, interfaces with the blood in the aorta is preferably provided with fibrils, forming a textured surface similar to a flocking, to promote cellular adhesion in forming a pseudointima on the outer surface  339  of the membrane  338 .  
         [0062]     A piece of sheet material  343  is attached to the outer surface  333  of the shell  326 . As indicated in  FIG. 14 , the periphery of the material  343  is preferably not attached to the outer surface  333  of the shell  326  and projects freely from the shell  326  to create a flange  348 . An additional material strip  349 , such as a pledget strip, is sewn to the flange  348 , to provide a suture ring  344  for implanting the device in an incision in the aorta, as best shown in  FIG. 14 . The strip  349  preferably has a fibrous surface allowing body tissues to migrate into, and mechanically interweaving with the strip  349 , to augment the sealing action (initially established by surgical implantation sutures (not shown)) between the flange  348  of the suture ring  344  and the wall  50  of the aorta  22 . The sheet material  343  and strip  349  are made of various appropriate materials, such as polyester, which are commercially available and have been certified for use in implanted devices.  
         [0063]     At least two modifications are optionally made to increase the durability of this blood pump by reducing the physical wear and stress on the membrane. The membrane failure of prior art designs is understood to have often been caused by microperforations due to membrane deformation, or “creasing,” during inflation-deflation cycles. This embodiment has a flat membrane  338  which lies flat and parallel to the plane described by the bead  334  of the shell  326 , when no pressure/vacuum is applied to the membrane  338 , as best seen in  FIGS. 14 and 15 . The shell  326  has a general cylindrical shape with an interior surface area that is generally planar if the cylinder were “unrolled.” The surface area of the flat membrane  338  is complementary to the interior surface area of the cylindrical shell, and thus, the creasing at cusp points that is known to occur with a flat membrane and a curved shell at cusp is greatly reduced in this embodiment, resulting in increased membrane and pump durability. Further, this embodiment has improvements to decrease the occurrence of deformation during implantation. The cylindrical shell  326  includes a spine region, the relatively thicker region near the hose barb housing, which adds structural strength to the pump and decreases the likelihood of deformation during implantation; such deformations resulted in membrane creasing in prior art pumps.  
         [0064]     As seen in  FIGS. 2, 3 ,  10  and  14 , the pump  26  or  326  is constructed so that the suture ring  44  and its flange  48  are located proximally closer to the spine region  57  of the shell  26 , as compared to prior art pumps. The suture ring  44  or  344  is located inward of the peripheral side edge  31  or  331  of the shell  26  or  326  viewed from above (as in  FIG. 10  for shell  26 ): when the pump is viewed from the side, as in FIGS.  2  or  14 , the suture ring  44  or  344  is seen as being located above the outer surface  33  or  333 , as well as being entirely located between the longitudinal ends of the shell  26  or  326 .  
         [0065]     In prior art designs, as seen in  FIG. 11 , the flange  250  of the suture ring  214  is located adjacent to, and “outboard” of the bead  224  of the shell, resulting in the suture ring and the eventual incision scar interfering with the flexing of the shell, effectively increasing the rigidity of the shell. The improved suture ring placement of the present invention prevents the suture ring itself, as well as scar tissue forming on the suture ring, from interfering with the flexing of the shell during the inflation-deflation cycles, a design flaw found in prior art designs, as seen in  FIG. 11 .  
         [0066]     In order to facilitate surgical explantation of the device, the outer surface  56  or  356  of the sheet material  43  or  343 , other than at the flange  48  or  348 , is optionally overlaid with a substance to which tissue does not adhere. Biocompatible substances prohibiting cellular adhesion include fluoropolymers and silicone. The overlayer  99  prevents scar tissue from adhering to the sheet material  43 , so the blood pump can be explanted if desired. Without the overlayer  99 , as the implant incision healed, scar tissue would also adhere to the sheet material  43  covering the back of the blood pump complicating explantation of the pump  20  or  320 .  
         [0067]     In certain embodiments, the conduit  32  of the pump is connected to a hose barb  40  or  340 , as seen in  FIGS. 2, 3  and  14 . The hose barb  40  or  340  is preferably molded into the outer surface  33  or  333  of the shell  26  or  326 . The hose barb is of rigid biocompatible, nonferrous, MRI-compatible material, such as titanium. The conduit  32  or  332  connects to the hose barb  40  or  340  and runs to the PAD (not shown). The passage  42  or  342  through the hose barb  40  or  340  connects to the passage  28  or  326  and places the interior volume  30  or  330  in fluid communication with the conduit  32  or  332 .  
         [0068]     While it is apparent that the illustrative embodiments of the invention herein disclosed fulfill the objectives stated above, it will be appreciated that numerous modifications and other embodiments may be devised by those skilled in the art. For example, it will be appreciated that features of the embodiments disclosed above can be used in various combinations and permutations. Therefore, it will be understood that the appended claims are intended to cover the forgoing—and all other—modifications and embodiments which come within the spirit and scope of the present invention.
A blood pump for placement in an incision in an aorta is provided that after placement contacts blood passing through the aorta. The pump inflates and deflates in order to provide left ventricular assistance. The pump has an elongated shell with a generally elliptical shape, an outer convex surface and an inner concave surface. A peripheral side edge located between the inner and outer surfaces terminates in a bead edge. A passage is provided through the shell to provide fluid communication between the outer surface and inner surface. A flexible airtight membrane has a membrane edge bonded to the outer shell surface adjacent to the bead edge to form an enclosed internal chamber in fluid communication with the passage. Preforming the membrane edge looped with a maximum linear span of curvature that is greater than a maximal transverse linear extent of the bead edge, membrane operational wear during inflation and deflation cycles is reduced in the region around the bead edge. A process of forming a blood pump with a membrane preform is provided that includes placing an airtight membrane around a platen having a platen edge bead with a curvature of maximal transverse linear extent and a platen footprint substantially identical to a footprint of the blood pump shell. By heat setting the membrane, a looped membrane edge is formed as complementary to the curvature of the platen edge bead to yield a membrane preform. The membrane preform is secured to the outer surface of the shell having a bead edge of maximum linear extent less than the curvature of the looped membrane edge.