Blood pump device with a no-leak aortic connector assembly and method of device implantation

Abstract

A blood pump device with a leak-proof aortic adapter assembly includes a T-shaped flow connector comprising a catheter insertion portion, a collar portion, and a truss. The catheter insertion portion is connected to the collar portion and the truss is disposed in the catheter insertion portion. The inner wall of the catheter insertion portion is tapered at both ends of the catheter insertion portion and is compliantly matched to the implant site artery. The proximal end of the collar portion is configured to be connected to an inlet adapter of a blood pump. The aortic adapter assembly is provided with a quick connector type coupler and a method of installation to achieve insertable flow communication between a ventricular assist device and the human circulatory system.

Description

TECHNICAL FIELD

[0001] The present invention relates to a ventricular assist device (VAD), in particular to a left ventricular assist device (LVAD) based on the principle of counterpulsation support and a blood pump device with a leakage-free aortic adapter assembly and a device implantation method. BACKGROUND

[0002] Mechanical circulatory support systems, in particular left ventricular assist devices (LVADs), have evolved into a standard of care for rescue therapy in advanced heart failure. Depending on the mechanical design of the blood pump, LVAD systems can generally be divided into continuous flow pumps and pulsatile flow pumps. Continuous flow devices are based on a rotating mechanical construction driven by an axial or centrifugal flow impeller. On the other hand, pulsatile flow devices use a displacement blood pump design and typically use a diaphragm bladder to allow blood to flow into and / or out of the pump.

[0003] Inflow / outflow cannulae are artificial flow conduits used to connect LVAD systems in series or in parallel to the human circulatory system. Historically, the flow cannula design has received less attention than the blood pump actuator. Not surprisingly, a large number of postoperative complications have been found to be associated with inflow / outflow cannula malfunctions. Among the adverse events are, for example, incorrect positioning of the cannula, obstruction, thrombus formation, and distorted flow resulting therefrom leading to pump thrombosis and thromboembolism. For rotary pump implantation, the flow channel is established primarily by connecting the inflow cannula to the heart chamber and the outflow cannula to the ascending or descending aorta. To date, almost all rotary pump inflow cannulae have been made of dacron graft or similar materials, such as made of a flexible textile material, and the connection to the aorta has been made using an end-to-side anastomotic suturing method. This end-to-side anastomotic method relies on technique, and complications caused by poor suturing or incorrect aortic flow can lead to thrombosis of downstream organs, stroke, or thromboembolism and infarction.

[0004] Rotary pumps provide full support (4-10 L / min) of cardiac output and are currently indicated for end-stage heart failure. For less ill patients, partial support (2-3 L / min) of less invasive implantable LVADs is more appropriate. Partial support LVADs aim to intervene early in heart failure to improve treatment outcomes. In this new trend of partial support, less invasive LVAD implantation, counterpulsation blood pumps have been the main focus as they have proven efficacy in increasing myocardial and major organ blood perfusion during systole and diastole. Counterpulsation pumps should be referenced to the heart rhythm, following strict timing control criteria. Typically, ventricular contraction unloading starts at the end of diastole, while organ perfusion enhancement occurs at the point of aortic valve closure (dicrotic notch on the aortic pressure waveform). In fact, the therapeutic effect provided is twofold, including: first, reducing ventricular contraction to reduce myocardial oxygen consumption, and second, increasing diastolic pressure to help increase perfusion of the myocardium, brain, and major organs.

[0005] Counterpulsation support efficacy has been clinically demonstrated by intra-aortic balloon pumps through peripheral subcutaneous delivery of the balloon pump. However, peripheral delivery to the descending aorta is often plagued by complications at the insertion site vessel and is difficult to use long term. In heart failure therapy, it would be highly significant to extend counterpulsation support from acute (less than a week) to long-term (months to years) support. To meet this requirement, new surgical methods and device innovations are needed to achieve this long-term counterpulsation support goal.

[0006] The counterpulsation support of the present invention is performed through left thoracotomy, where the device is connected to the thoracic aorta. This counterpulsation invention, known as a paracorporeal aortic pump, is implanted through an access hole formed in the descending aortic wall from a lateral pump connection. In sharp contrast to the intra-aortic balloon pump, the paracorporeal aortic pump does not obstruct the blood flow path, thus providing more flexibility in counterpulsation timing control. Animal studies have shown that the hemodynamic support effect of paracorporeal counterpulsation is superior to that of the intra-aortic balloon pump. However, it is not yet clear whether paracorporeal pump support will lead to long-term complications. In fact, the challenge is to construct a flow cannula for implantation in a less invasive manner, enabling long-term safe fluid communication between the pump and the connected artery.

[0007] Aortic bypass blood pump implantation requires the establishment of artificial conduits to allow blood flow into and out of the blood pump. Rapid filling of the pump is often required to effectively reduce the afterload of the left ventricle as well as to create a strong vortex wash to prevent pump thrombosis. This rapid pump filling often experiences transient low pressures that cause the flow conduit to collapse. Therefore, the use of traditional dacron graft type cannulas is not feasible because fabric conduits cannot withstand compression forces and a collapsed graft can occlude inflow during the pump filling phase. Furthermore, the high and low pressure pulses caused by rapid ejection and filling of the pump can cause bleeding problems at the suture site, especially during the acute phase when the surgical anastomosis has not yet healed. Therefore, it is critical to design a flow cannula to overcome the specific flow characteristics associated with aortic bypass counterpulsation.

[0008] The hemodynamic characteristics of aortic bypass blood pump flow are not physiological. Blood flow into or out of the connected blood pump is not laminar in the arterial direction as is typically seen in the natural aorta. This side-tilting artificially created pump flow is highly turbulent and complex. During the pump filling phase, blood flow is sharply turned into the blood pump, creating a flow separation zone and recirculation zone at the turn corner. During the pump ejection phase, as blood flow accelerates into the descending aorta, the flow characteristics are characterized by a shock flow with ultra-high pressure and shear stress exerted on the opposing aortic wall regions. The device-induced flow characteristics include a low-velocity recirculation zone and a high-pressure, high-shear shock jet, which can lead to endothelial cell erosion, lipid infiltration, smooth muscle cell proliferation, long-term aortic wall stenosis or thrombosis, or aortic dissection due to high blood pressure. Therefore, long-term implantable counterpulsation device design requires an innovative artificial flow cannula to avoid or mitigate the aforementioned device-induced pathophysiological flow phenomena and resulting adverse vascular maladaptations and thrombosis or aortic dissection.

[0009] Safety of surgical anastomosis is another requirement for the flow cannula design of an aortic bypass counterpulsation pump. Traditional graft suturing can experience challenging bleeding complications when subjected to excessively high pressure fluctuations associated with counterpulsation flow. High and low pressure cycles are a major driver of material fatigue failure, especially when the implanted aorta is diseased (atherosclerotic or calcified), degraded (thinning wall thickness), or aged (hardened wall). It is noted that the aortic wall structure can adapt to the applied stress conditions. Cells and tissues around the implant site will remodel with the surgical injury healing process and further develop according to the non-physiological mechanical environment induced by the device. Even if short-term implant success is achieved, it cannot be guaranteed that long-term graft failure will not occur due to gradual cellular and morphological maladaptation of the vascular wall during the postoperative course.

[0010] To date, all viable solutions for the connection of artificial grafts to the contralateral side of the aortic end are based on flexible braided fabric tubes by means of a suture method. In the industry of instruments, there is no long-term cannulation solution that can solve the problems related to the levels of flow and pressure fluctuations generated by the counterpulsation aorto bypass pump. For more than fifty years, through the application of intra-aortic balloon pumps, counterpulsation has been achieved clinically in the short and medium term (a few days to a month or so). This typical pneumatic energy transmission device for driving the inflation / deflation of the balloon is achieved using a thin catheter that is delivered percutaneously from a remote peripheral artery, with a lumen diameter in the range of 6-10 millimeters. Such a delivery site artery is mostly occluded, which usually leads to serious bleeding complications and ischemia of the downstream limb or arm, thus hindering the long-term use of percutaneous balloon counterpulsation. Aorto bypass implantation is a new approach aimed at extending counterpulsation therapy to a longer time frame. However, long-term aorto bypass counterpulsation requires the construction of a flow connector to solve the aforementioned implantation problems. Such a flow cannula should be non-collapsible, easy to implant, non-bleeding and biocompatible, without having to worry about causing pathological vascular maladaptation. The present invention strives to meet all the collective needs by proposing an insertable aortic adapter, which will be disclosed hereinafter. SUMMARY

[0011] An embodiment of the present invention provides a blood pump device with a leak-proof aortic adapter assembly for an implantable ventricular assist device, comprising a T-shaped flow connector. The T-shaped flow connector comprises a catheter insertion portion, a collar portion and a truss. The catheter insertion portion is connected to the collar portion, and both the catheter insertion portion and the collar portion have smooth blood contact surfaces. The truss is disposed in the catheter insertion portion. The T-shaped flow connector has a polymer elastomer and is reinforced by the aforementioned truss having a nickel-titanium alloy material. The catheter insertion portion has a wall that tapers in thickness at its two catheter ends, and there is a suitable distance between the ends of the catheter ends and the outermost boundary of the truss, and the catheter ends have a compliance-matching effect on the implantation site artery. The proximal end of the collar portion is connected to an inlet adapter of a blood pump.

[0012] In an embodiment, the aforementioned truss is co-injected with the polymer elastomer of the T-shaped flow connector and is embedded in the wall of the catheter insertion portion of the T-shaped flow connector.

[0013] In an embodiment, the aforementioned polymer elastomer of the T-shaped flow connector has a silicone material.

[0014] In an embodiment, the aforementioned polymer elastomer of the T-shaped flow connector is a molded polyurethane.

[0015] In an embodiment, the structural compliance of the aforementioned embedded truss and the structural compliance of the polymer elastomer of the T-shaped flow connector are approximately equal to each other.

[0016] In one embodiment, the aforementioned tapered catheter tip is sharp.

[0017] In one embodiment, the aforementioned neck portion has a shallow bevel to fit an inlet adapter of a blood pump, and an inner diameter of the neck portion is slightly smaller than an inner diameter of the inlet adapter of the blood pump.

[0018] In one embodiment, the aforementioned shallow bevel is inclined with respect to an extension direction of the catheter insertion portion.

[0019] In one embodiment, the aforementioned truss has a plurality of wave structures.

[0020] In one embodiment, the aforementioned neck portion includes a neck body and an extension portion disposed on the neck body, wherein the extension portion protrudes from the neck body, and a maximum inner diameter of the extension portion is larger than a maximum inner diameter of the neck body.

[0021] In one embodiment, when the aforementioned neck portion is engaged with an inlet adapter of a blood pump, the extension portion abuts against the inlet adapter, and the neck body is disposed around the inlet adapter.

[0022] In one embodiment, the aforementioned aortic adapter assembly further comprises a coupler including a flange seat; a pair of collars rotatably mounted on the flange seat; and a latch disposed on one of the collars and used to lock the collars; wherein the collars have an inner groove that can clamp the T-shaped flow connector by controlled compression to seal the T-shaped flow connector.

[0023] In one embodiment, the aforementioned each collar has a flange profile that can simultaneously engage the edge of the flange seat with respect to the collars.

[0024] In one embodiment, the aforementioned latch is made of a spring leaf to ensure that the coupler is in a locked state without worrying about accidental disengagement.

[0025] In one embodiment, the aforementioned coupler further comprises a hinge disposed on the flange seat, and the aforementioned collars are pivotally connected to the hinge and can rotate with respect to the hinge and the flange seat.

[0026] In one embodiment, the aforementioned hinge is located on a first side of the flange seat, and the aforementioned latch is located on a second side of the flange seat, wherein the first side and the second side of the flange seat are opposite.

[0027] In one embodiment, the aforementioned latch has a slot that can be engaged with a bevel disposed on the opposite collar.

[0028] In one embodiment, the aforementioned flange seat has a substantially circular structure, and the aforementioned collars have an arc structure.

[0029] In one embodiment, the foregoing blood pump device, as an implantable ventricular assist device, includes a blood pump, an inlet adapter having a tapered beak, and an aortic adapter assembly. The aortic adapter assembly includes a T-shaped flow connector. The T-shaped flow connector includes a catheter insertion portion, a collar portion, and a truss. The catheter insertion portion is connected to the collar portion, and both the catheter insertion portion and the collar portion have smooth blood contacting surfaces. The truss is disposed in the catheter insertion portion. The T-shaped flow connector has a polymeric elastomer and is reinforced by the foregoing truss having a nitinol material. The catheter insertion portion has a wall that tapers in thickness at two catheter ends thereof, and there is a suitable distance between the ends of the catheter ends and an outermost boundary of the truss, and the catheter ends have a compliant fit to an implant site artery. A proximal end of the collar portion is configured to connect to the inlet adapter of the blood pump, and the tapered beak is configured to mate with the proximal end of the collar portion.

[0030] In one embodiment, the foregoing inlet adapter has an inner diameter that is slightly larger than an inner diameter of the collar portion of the T-shaped flow connector.

[0031] Another embodiment of the present invention provides a method of implanting a flow connector assembly, comprising: providing the foregoing aortic adapter assembly, wherein the T-shaped flow connector is compressible from an initial deployed form to a compressed implant configuration form; crimping the T-shaped flow connector into the compressed implant configuration form, wherein the compressed implant configuration form of the T-shaped flow connector has a size that is at least half of a diameter of the initial form of the T-shaped flow connector; inserting the compressed implant configuration form of the T-shaped flow connector into an artery having an access hole punched in a wall of the artery; releasing the compressed implant configuration form of the T-shaped flow connector to deploy, wherein the released T-shaped flow connector self-expands to the initial deployed form thereof; and coupling the deployed T-shaped flow connector to a ventricular assist device. BRIEF DESCRIPTION OF DRAWINGS

[0032] Embodiments of the present invention can be more fully understood from the following detailed description, taken in conjunction with the accompanying drawings, in which:

[0033] Figure 1 is a schematic view of an aortic bypass blood pump device according to a first embodiment of the present invention;

[0034] Figure 2 is a schematic view of an aortic bypass blood pump device according to a second embodiment of the present invention;

[0035] Figure 3 is a schematic view of an aortic bypass blood pump device according to a third embodiment of the present invention;

[0036] Figure 4 is a schematic view of an aortic bypass blood pump device according to a fourth embodiment of the present invention;

[0037] Figure 5 is a schematic view of an aorto bypass blood pump device implanted in a human body according to the first and second embodiments of the present application;

[0038] Figure 6 is a schematic view of an aorto bypass blood pump device implanted in a human body according to the third and fourth embodiments of the present application;

[0039] Figure 7 is a first schematic view of a driver according to an exemplary embodiment of the present application;

[0040] Figure 8 is a second schematic view of a driver according to an exemplary embodiment of the present application;

[0041] Figure 9 is a schematic view of the driver functions and main interconnection signals necessary for the operation of the present application;

[0042] Figure 10 is a schematic view of the driver functions and main interconnection signals necessary for the operation of the third embodiment of the present application;

[0043] Figure 11 is a schematic view of the driver functions and main interconnection signals necessary for the operation of the second embodiment of the present application;

[0044] Figure 12 depicts electro-mechanical actuator (EMA) piston position trajectories and trigger detection command timing in relation to anti-pulsatile circulatory support;

[0045] Figure 13A is a perspective view of an aorto bypass blood pump implant according to the first or third embodiment of the present application;

[0046] Figure 13B is a cross-sectional view of an aorto bypass blood pump implant according to the first or third embodiment of the present application;

[0047] Figure 14A is a perspective view of an aorto bypass blood pump implant according to the second or fourth embodiment of the present application;

[0048] Figure 14B is a cross-sectional view of an aorto bypass blood pump implant according to the second or fourth embodiment of the present application;

[0049] Figure 15 is a connection of a drive catheter to a pump housing of a blood pump according to another embodiment of the present application, the connection of the drive catheter being realized via an introduction piece on a distal housing piece;

[0050] Figure 16 is Figure 15 a cross-sectional view of the blood pump shown in

[0051] Figure 17 The shallow trench design is shown for extending the wire from the inlet located in the distal housing to the pressure sensing chamber in the proximal housing.

[0052] Figure 18A This demonstrates the connection of the external contralateral anastomosis of the blood pump to the artery using an interface adapter connector;

[0053] Figure 18B This demonstrates how the blood pump of the present invention is coupled to an artery via a plug-in connection method using a T-shaped intravascular connector coupled through an interface adapter.

[0054] Figure 19 A cross-sectional view of a monolithic, axisymmetric elliptical blood sac and a rotating body including the sac, proximal port, and distal port assembly is shown.

[0055] Figure 20 An exploded view is shown, illustrating the components used to construct the axisymmetric elliptical sac and port assembly; (Note: in conjunction with integration into...) Figure 3 (The sac is in its original shape before the proximal and distal ports are shown);

[0056] Figure 21 It showed that Figure 19 The curved trilobed sac structure at the end of the ejaculation of the blood sac is shown.

[0057] Figure 22 A perspective view of the drive catheter is shown, which is connected to the proximal housing of the blood pump via an inlet.

[0058] Figure 23A It showed along Figure 22 A cross-sectional view of the distal side of the blood pump and drive catheter in section AA;

[0059] Figure 23B It showed along Figure 22 A cross-sectional view of the proximal end of the driving catheter at section AA in the diagram;

[0060] Figure 24 A cross-sectional view of the exhaust port installed in the proximal housing corresponding to the first embodiment is shown;

[0061] Figure 25A A cross-sectional view of the pressure sensing chamber and the inlet in the proximal housing of the first embodiment is shown. (Note: The drive conduit is not installed; the inlet includes a first portion, an extension of the proximal housing, and a second portion interlocked with the first portion.)

[0062] Figure 25B It shows the combination in Figure 25A A perspective view of a microelectromechanical system (MEMS) pressure sensor in a photograph;

[0063] Figure 26 A cross-sectional view of the multi-layered drive catheter of the present application is shown, including inner tube for pneumatic air transmission, middle tube for electrical signal transduction, and coil, tether, and outer tube;

[0064] Figure 27 A cross-sectional view of the multi-lumen drive catheter design of the drive catheter is shown;

[0065] Figure 28A A representative view of the flow characteristics of the pump fill phase is shown;

[0066] Figure 28B A representative view of the flow characteristics of the pump ejection phase is shown;

[0067] Figure 29 is a perspective view of the T-Flow connector of the present embodiment;

[0068] Figure 30 is a cross-sectional view of the T-Flow connector of the present embodiment;

[0069] Figure 31 is a planar view of the deployed Nitinol stent graft;

[0070] Figure 32 defines the lateral stiffness (LS) of the insertion catheter for measuring the Nitinol stent graft and T-Flow connector;

[0071] Figure 33 is an exploded view of the coupler and its components;

[0072] Figure 34A is a view of the coupler in the open state (configuration);

[0073] Figure 34B is a view of the coupler in the locked state (configuration);

[0074] Figure 35 is a cross-sectional view of the use of the coupler to connect the aortic adapter with the paracorporeal blood pump;

[0075] Figure 36A shows a view of the step discontinuity created by the docking method;

[0076] Figure 36B shows a view of the gap discontinuity created by the docking method;

[0077] Figure 37 is a view of the inlet adapter for installation at the distal end of the paracorporeal blood pump;

[0078] Figure 38 is a cross-sectional view of the aforementioned inlet adapter;

[0079] Figure 39 a downwardly tapered beak of the inlet adapter is shown in connection with a sloped surface at the neck of the T-shaped flow connector;

[0080] Figure 40 a schematic view of the crimped flow connector is shown in a bundled, crimped-in-place configuration;

[0081] Figure 41A an insertion flow connector is shown passing through an access hole formed in the aortic wall;

[0082] Figure 41B a flow connector in a bundled, compressed form is shown fully inserted into the aortic lumen;

[0083] Figure 41C a repositioned, compressed form of the flow connector is shown with its T-shaped neck facing the aortic access hole;

[0084] Figure 41D an expanded, deployed form of the flow connector is shown with its T-shaped neck deployed in the aortic access hole after release of the bundle; and

[0085] Figure 42 a schematic illustration of the individual steps of implanting the aortic adapter into the target aortic segment and connection to the blood pump is shown. DETAILED DESCRIPTION

[0086] Four embodiments of the aorto-pass blood pump device that can be used to implement the present application are provided below, as follows.

[0087] Reference is made to Figure 1FIG. 1 is a schematic diagram of a paracorporeal blood pump device according to a first embodiment of the present application. The paracorporeal blood pump device 10 includes a blood pump 12, an aortic adapter 14, a drive conduit 16, and a drive unit 18. The blood pump 12 further includes a pump housing and a pressure sensor. The interior of the pump housing is divided into two chambers, one for storing blood and the other for receiving drive air. The two chambers are separated by an egg-shaped flexible membrane that is suspended by a pair of stress-relief stems attached to the pump housing. The pressure sensor is mounted within the pump housing of the blood pump 12 for monitoring the blood pressure within the blood pump 12 to generate an electronic blood pressure signal. The aortic adapter 14 is a valveless, T-manifold shaped conduit that couples the blood pump 12 to the aorta of a human body. In the first embodiment, the aortic adapter 14 and the blood pump 12 are integrally formed with seamless blood contact surfaces and the blood pump 12 is connected to the aorta of a human body via the aortic adapter 14. The aortic adapter 14 is made of a flexible material that allows the aortic adapter 14 to deform during insertion delivery through a circular hole made in the aortic wall. The aortic adapter 14 is self-expandable and strong enough to resist radial compression contact forces from the aorta applied to the adapter wall surface from an oversize fitting. The drive conduit 16 is connected to the housing of the blood pump 12 for providing air pressure pulses to the blood pump 12 and transmitting the blood pressure signal received from the pressure sensor. The drive unit 18 is coupled to the drive conduit 16 for receiving the transmitted electronic blood pressure signal. The drive unit 18 includes an electro-mechanical actuator for generating air pressure pulses in response to the electronic blood pressure signal and sending the air pressure pulses to the blood pump 12 through the drive conduit 16. The wearable drive unit 18 provides air pressure pulse control rhythm in coordination with the heart rhythm to drive the implanted blood pump 12 to eject and fill blood.

[0088] The aforementioned drive unit 18 includes a primary battery power system 11 and a backup battery power system (described later Figures 2 to 4 and Figure 8 The battery power systems 21, 31, 41 are the same or similar), wherein the backup battery power system ensures continuous power supply to the drive unit 18. When the patient does not need to move, the drive unit 18 can also be powered by an alternating current adapter for convenience. In addition, the device has a clinical monitor, which is not shown in Figures 2 to 4As shown in Fig. 1, the drive conduit 16 can be connected to the drive 18 to provide a user interface to the clinician for displaying device monitoring or diagnostic information and for accessing drive parameters for patient data initial start-up settings and optimization of specific therapeutic operating mode settings.

[0089] Figure 1 and Figure 2 Two different blood pump 12, 22 designs are shown coupled to the same drive conduit 16, 26 and drive 18, 28 system. Figure 2 Fig. 2 is a schematic view of an aorto- bypass blood pump apparatus 20 according to a second embodiment of the present application. The second embodiment differs from the first embodiment of the present application in that the aorto- bypass blood pump apparatus 20 of the second embodiment further comprises a coupler, or coupling adapter 25. The blood pump 22 and aortic adapter 24 of the second embodiment are not integrally formed, but are detachable, and the coupler 25 is provided to couple the blood pump 22 to the aortic adapter 24. The coupler must be carefully designed to minimize discontinuities at the interface of the connection. During implantation of the apparatus, the aortic adapter 24 is first inserted into the aorta through an implantation hole cut in the aortic wall. The apparatus can use a specially developed implantation tool to mount the coupling adapter 25 around the T-shaped neck of the aortic adapter 24, so that the blood pump 22 can be connected to the coupling adapter 25. After the blood pump 22 is implanted in the thoracic cavity, the blood pump 22 and the aortic adapter 24 are securely locked and integrated together through the coupler 25. This detachable design of the blood pump 22 and the aortic adapter 24 has advantages in implantation surgery and post-surgery. During implantation of the apparatus, the detachable blood pump design makes it easier to implant the aortic adapter, because the surgical area is clearer and is not disturbed by the pump body. In addition, after implantation surgery, if the pressure sensor fails or the blood sac needs to be replaced in emergency surgery, the blood pump can be detached and replaced. In this regard, the detachable blood pump design of the second embodiment is advantageous. The aortic adapter can remain in the aorta without being removed, avoiding the trouble and danger of a redo surgery associated with removal of the aortic adapter.

[0090] Reference is made to Figure 1 and Figure 3, respectively, are schematic illustrations of the aorto- bypass pump devices 10, 30 of the first and third embodiments of the present application. The third embodiment differs from the first embodiment of the present application in that the drive catheter 16 of the first embodiment is replaced by a distal drive catheter 37, a drive catheter interconnector 33 and a proximal drive catheter 39 of the third embodiment. The distal drive catheter 37 is connected to the drive catheter interconnector 33 for transmitting the electronic blood pressure signal acquired from the pressure sensor and the air pressure pulses sent from the driver 38; and the drive catheter controller and the vibrator (for alarm warning purposes) included in the drive catheter interconnector 33 are originally included in the driver 18 of the first embodiment, thus the driver 18 of the first embodiment has additional drive catheter controller and vibrator (compared to the driver 38 of the third embodiment). The drive catheter controller is used for processing the electronic blood pressure signal, and the vibrator is used for providing a sound alarm or haptic feedback. In other words, the mechanical power transmission of the blood pump, as well as the analog / digital signal conversion and alarm notification, achieved by the drive catheter 16 and the driver 18 of the first embodiment are essentially the same as achieved by the distal drive catheter 37, the drive catheter interconnector 33 and the proximal drive catheter 39 of the third embodiment.

[0091] The first embodiment has a more compact drive catheter configuration design, and the electronic signal processor is placed in the driver, thus, the risk of environmental contamination (water ingress or moisture condensation) on the pressure signal measurement and air leakage at the joint, both of which are related to the drive catheter interconnector 33, is minimized. However, this long drive catheter is more susceptible to contact damage, such as wear and tear, kinking, and cutting caused by contact with foreign objects during daily activities. Any significant damage to the drive catheter 16 of the first or second embodiment, whether electronic or mechanical, can require a surgical replacement of the blood pump. This is highly undesirable in view of the risk of reoperation and associated medical costs. The third or fourth embodiment alleviates this blood pump replacement drawback associated with drive catheter damage by employing an intermediate connector (drive catheter interconnector). In general, the distal drive catheter 37 is shorter in length exposed outside the body and is better protected by the coverage of the skin dressing and patient vest. In the extreme case of severe damage to the drive catheter that cannot be repaired, the proximal drive catheter 39, which is most likely to be damaged, can be easily replaced without resorting to surgery. Furthermore, the third or fourth embodiment is less susceptible to electromagnetic interference because the analog-to-digital signal conversion has already been completed in the circuit in the drive catheter interconnector 33. In the third or fourth embodiment, the signal fidelity of the pressure signal can be better guaranteed because the digital signal transmission in the proximal drive catheter 39 is less sensitive to electromagnetic interference.

[0092] Referring to Figure 3 with Figure 4 FIGS. 3 and 4 are schematic diagrams of a paracorporeal aortic pump device 30, 40 according to third and fourth exemplary embodiments of the present application, respectively. The third and fourth embodiments differ in that the aortic adapter 34 and blood pump 32 of the third embodiment are integrally formed, while the blood pump 42 and aortic adapter 44 of the fourth embodiment are detachable. The fourth embodiment also includes the same blood pump 22, aortic adapter 24, and coupler 25 as the second embodiment. The same blood pump 42, aortic adapter 44, and coupler 45 of the embodiments are not repeated here. The drive conduit 46 includes a distal drive conduit 47, drive conduit interconnect 43, and proximal drive conduit 49 that are the same as the distal drive conduit 37, drive conduit interconnect 33, and proximal drive conduit 39 of the drive conduit 36.

[0093] Referring to Figure 5 FIG. 5 is a schematic diagram of a paracorporeal aortic pump device implanted in a human body according to an exemplary embodiment of the present application. The paracorporeal aortic pump device 90 includes a blood pump 92, an aortic adapter 94, a drive conduit intracoφoreal segment 991, a drive conduit extracoφoreal segment 993, and a driver 98. In another embodiment, the paracorporeal aortic pump device also includes a coupler. The portion of the paracorporeal aortic pump device 90 implanted in the human body includes the blood pump 92, the aortic adapter 94, and the drive conduit intracoφoreal segment 991. In another embodiment, the paracorporeal aortic pump also includes a coupler. During surgery, the aortic adapter 94 is installed into the aorta 95 and an exit site EX is created at an appropriate location on the surface of the human body. The extracoφoreal portion of the paracorporeal aortic pump device 90 includes the drive conduit extracoφoreal segment 993 and the driver 98. The drive conduit intracoφoreal segment 991 has a portion covered with fabric velour for tissue ingrowth to achieve infection control, bounded by the exit site EX. The implanted velour portion is optimally placed two to five centimeters below the exit site EX. The driver 98 is a wearable or portable device.

[0094] Referring to Figure 6FIG. 1 is a schematic diagram of a human body with an aorto bypass blood pump device implanted according to an exemplary embodiment of the present application. The aorto bypass blood pump device 90 includes a blood pump 92, an aorta adapter 94, a distal drive catheter 97 (comprising a drive catheter in-body segment 971, a drive catheter out-of-body segment 973), a drive catheter interconnector 93, a proximal drive catheter 99, and a driver 98. In another embodiment, the aorto bypass blood pump device 90 also includes a coupler. The portions of the aorto bypass blood pump device 90 implanted in the human body include the blood pump 92, the aorta adapter 94, and the drive catheter in-body segment 971. In another embodiment, the aorto bypass blood pump device also includes a coupler. During a surgical procedure, the aorta adapter 94 is implanted into the aorta 95, and an exit site EX is created at a suitable location on the surface of the human body. The portions of the aorto bypass blood pump device 90 located outside the human body include the drive catheter out-of-body segment 973, the drive catheter interconnector 93, the proximal drive catheter 99, and the driver 98. The exit site EX is the boundary of the drive catheter, and the distal drive catheter is divided into the drive catheter in-body segment 971, which is covered with a velour for infection control, and the drive catheter out-of-body segment 973. The driver 98 is a wearable or portable device.

[0095] The implanted subsystems are further described below.

[0096] Implantation is achieved through a left thoracotomy via a relatively small chest opening using less invasive surgical (LIS) techniques. For example, a chest incision is made at the 7th intercostal space as the primary opening to allow implantation of the aorta adapter and blood pump. Two additional small incisions are made at the 6th and 8th intercostal spaces to introduce proximal and distal aorta clamps. The site between the aorta clamps allows the aorta adapter to be implanted through an opening in the aorta wall. The aorta adapter is flexible and can be crimped down to a smaller delivery shape prior to implantation. Upon completion of delivery into the aorta, the aorta adapter automatically springs open and returns to its original shape with an ultra-tight fit to the lumen of the intended implantation site. Thus, the material of the aorta adapter is important, it should be flexible but have sufficient radial strength to keep the implanted aorta adapter tube wall circular without tube wall collapse. Candidate aorta adapter materials can be selected from silicone or polyurethane elastomers, or reinforced by embedding reinforced polymers or metal materials to strengthen the structural strength.

[0097] The aorta adapter of each of the above embodiments and its functional requirements are further described below.

[0098] The aortic adapter serves the hemodynamic role of providing blood flow communication between the blood pump and the body's systemic circulation. In addition to this role, the aortic adapter can also serve as a mechanical base to which the blood pump is affixed. The aortic adapter structure needs to be flexible but anti-buckling, and strong enough to withstand internal blood pressure and external contact forces that the blood pump generates by contacting surrounding lung tissue or due to breathing and diaphragm movement of the chest.

[0099] The aortic adapter 54 is implanted within the aorta with its two conduit ends (conduit end 545, conduit end 645) interfacing with the aortic lumen, creating a host / graft interface in the blood stream (see Figure 13B 、 Figure 14B ). To minimize the discontinuity of the host / graft interface in terms of morphology and flexibility, the two conduit ends (conduit end 545, conduit end 645) are configured to have a flared inner surface profile and a continuously decreasing wall thickness distribution. This conduit end design minimizes the step at the interface and incorporates the compliance matching effect required at connection into the aortic adapter design. The likelihood of thrombus formation at the interface can thus be greatly reduced, as the rate of clot aggregation at the interface will be slower than the natural thrombolysis rate provided by the human aortic endothelium. In addition, the tapered conduit wall structure makes the conduit ends (conduit end 545, conduit end 645) softer (compliant), causing the conduit tips to expand and contract with the pulsatile blood pressure, constituting a dynamic sealing effect to prevent blood cells from getting stuck in the interface gap, which is often a source of thrombus formation.

[0100] The driver of each of the above embodiments is further described below.

[0101] Figure 7 、 Figure 8 are right and left perspective views of the driver 78. This compact driver 78 internal module includes an electromechanical actuator (EMA), an electronic controller, a pair of primary and backup batteries. The driver 78 also includes a user interface panel 73, a battery hatch 71, a drive conduit receptacle 75, an external power receptacle 77, and a pair of air vents 79, as shown in Figure 7 、 Figure 8 .

[0102] Key messages of operation and alerts for device malfunction and aortic pressure conditions will be displayed on the user interface panel 73 of the driver 78. The main battery can be replaced through the battery hatch 71 when the main battery is low. The driver 78 can be powered using a connection through an external power outlet 77 using a cable when the patient is in bed and power from a wall outlet is available for long term use. The proximal drive catheter 99, one end of the extracorporeal section 993 of the drive catheter, can be connected to the driver 78 through the drive catheter hub 75, and the voltage sensor signal and the pneumatic pressure pulse are communicated through the drive catheter hub 75. A pair of air vents 79 are installed on opposite sides of the driver 78 to allow ambient air to flow through the interior of the driver 78 for cooling purposes.

[0103] The driver 78 can be externally coupled to a clinical monitor for the clinical monitor to collect and display real-time clinical waveform data and store patient data for long term status monitoring and diagnosis. In addition, the clinical monitor unit can provide a clinician patient user interface for displaying device monitoring / diagnostic information and for driver parameter settings for initial start-up of the driver 78 and for optimizing setting of patient specific individual circulatory assist operation modes.

[0104] In the embodiment, the EMA is a pneumatic actuator, which includes a brushless servo motor and a ball screw unit, and a piston and cylinder assembly. Air is used as a driving medium to reciprocally actuate the blood pump to achieve the functions of ejecting and filling the blood pump.

[0105] The pneumatic actuator device is carried by the patient under treatment in the driver. The electromechanical actuator includes a brushless servo motor, a piston and cylinder assembly, and a ball screw unit. The ball screw unit includes a ball screw rod and a ball nut. The piston is fixed on the top end of the ball screw rod and is driven to move linearly and reciprocally by the rotation of the ball nut. The servo motor includes a rotor and a stator, and the rotor is integrated with the ball nut. Through electromagnetic induction of the motor, the rotor rotates and changes its direction clockwise and counterclockwise to achieve the linear reciprocating stroke motion of the screw rod and the piston in the cylinder. The reciprocating stroke of the piston drives the air in the cylinder to be sent into the blood pump through the drive conduit to complete the blood ejection and filling actuation of the blood sac.

[0106] There are two problems in driving blood pump with air as medium. One is air leakage problem, and the other is water vapor condensation problem caused by blood permeating through the blood bag wall. The former will harm the efficiency of blood pump ejection and filling and the power consumption of the motor, and the latter will cause the risk of bacteria breeding in the driving wire due to moisture. To solve the two problems, the pneumatic actuator of the ventricular assist device is specially installed with a pressure balance valve device on the cylinder wall. The valve can allow the air in the cylinder to communicate with the atmosphere to form air mass flowing in the air pressure balance process. The pneumatic actuator is provided with position and optical sensors for the controller to obtain piston position information, and generates piston driving instructions to drive the reciprocating motion of the piston and operate the pressure balance valve through the controller. Therefore, the opening time and frequency of the pressure balance valve can be programmed and stored in the controller. Through the action of the pressure balance valve, the air in the cylinder can exchange with the outside atmosphere to achieve the functions of air supplement and dry gas, thereby ensuring the safety and efficiency of the operation of the ventricular assist device.

[0107] Reference Figure 9 The aortic bypass blood pump device according to the embodiment of the present application is divided into three parts. The first part is mainly installed inside the human body (i.e. implant), and the outer end communicates with the second part. The first part includes a blood pump (including a blood pump pressure sensor), an aortic adapter and a distal drive catheter segment implanted in the human body respectively. The second part is installed outside the human body and includes a proximal drive catheter and a drive catheter electronic module (or called drive catheter interconnector). The third part is installed outside the human body, which is a driver including an electromechanical actuator (EMA), a controller circuit, a main battery and a backup battery.

[0108] The blood pump pressure sensor is built in the proximal blood pump housing and immersed in a small pressure sensing chamber filled with sensing medium, thereby allowing continuous monitoring of the blood pump pressure. The distal drive catheter is connected to the pump housing and provides counter-pulsatile air pressure pulses to actuate the ejection and filling of the blood bag. The distal and proximal drive catheters provide pneumatic drive air pressure pulses generated by the EMA inside the driver to the blood pump; and transmit the electronic blood pressure signals generated by the sphygmomanometer pressure sensor to the driver. The drive air path (indicated by dashed arrows) and the electrical signal path (indicated by solid lines) are shown in Figure 9 to describe the functional relationship between the interactive action modules. The details of the aortic adapter have been described above. The controller circuit can include a motor controller unit for driving the brushless motor, and a microcontroller unit as a central processor to process the received pressure signals and generate control instructions for the motor controller to actuate the piston movement.

[0109] Reference Figure 10 and Figure 11, showing the functions within the driver and the key interconnection signals necessary for the blood pump to start. Further explanation of the embodiments presented for the present invention are illustrated in Figure 10 As shown in Figure 11 Referring to Figure 10 As shown in Figure 11 To explain the driving relationship between the external driver and the implant, reference is made to the contents of the aforementioned blood pump, drive catheter, distal drive catheter, proximal drive catheter, aortic adapter, and drive catheter interconnector.

[0110] The driver receives the blood pump pressure signal (electrical signal) and processes the signal using a trigger detection algorithm to generate a trigger signal that commands the motor actuator to actuate in synchrony with the heart rhythm. Upon receiving the specified trigger time, the micro controller unit sends a command to the motor controller (unit) to drive the piston, either from ejection to fill or from fill to ejection process, to provide anti-pulsatile cyclic support.

[0111] The architecture of the electronic controller comprises three functional blocks, namely the micro controller unit (MCU), the motor control block (or motor controller unit), and the power management unit. The following table provides a descriptive summary of each functional block of the driver 78.

[0112]

[0113]

[0114] In Figure 10 , 11 The signal acquisition, transmission, processing, and control logic and instruction generation, and how the electromechanical actuator actuates to produce a pressure pulse to drive the blood pump will be explained for the previously elucidated exemplary embodiments.

[0115] Figure 12 The trigger detection instructions for the electromechanical actuator piston position in relation to anti-pulsatile assistance are depicted. In Figure 12In the absence of assistance, the aortic pressure (AoP) waveform is represented by the dashed line, while the solid line represents the aortic pressure waveform under assistance. When the driver is operating in the automatic mode, the driver operation is initiated, and the system performs a "fill- eject-fill-eject..." cycle assistance, which represents the normal synchronous counter-pulsation assistance operation. The MCU monitors the blood pump pressure (BPP) signal (electrical signal) and detects the left ventricular end diastole (LVED) timing. After detecting the LVED timing, the MCU generates the F_Trig signal. The time interval between two consecutive F_Trig signals represents the instantaneous cardiac cycle interval (or period). Based on the estimated heart rate calculated from the preceding number of cycle intervals, the MCU determines the time for the blood pump to eject, i.e., the E_Trig signal. The E_Trig signal provides the timing to command the motor controller unit to drive the electro-mechanical actuator according to the predetermined position, velocity, and acceleration profiles. When the ejection stroke is completed and the optimized dwell time has elapsed, the electro-mechanical actuator is commanded to perform a pre-filling action at a slower fill velocity until the F_Trig signal is present. Upon receiving the F_Trig signal, the electro-mechanical actuator starts to perform the remaining fill stroke at the specified piston velocity.

[0116] When the MCU loses the BPP signal (electrical signal) sent by the blood pump, the MCU automatically initiates the flush mode to drive the electro-mechanical actuator to operate at a predetermined pumping frequency and driver output per beat. The flush mode is used to prevent thrombus formation in the blood sac and is a device protection mode rather than providing synchronous counter-pulsation cycle support.

[0117] The aortic parapump device of the present invention, in principle, has better counter-pulsation support efficacy than the intra-aortic balloon pump (IABP) due to its non-occlusive aortic implantation feature. Unlike the bedridden or ambulatory IABP patients who have to stay in the hospital, the present aortic parapump device allows the patients to leave the hospital and have the ability to live a better life at home. Therefore, the aortic parapump device of the present invention can further improve the disease condition and quality of life of the patients in addition to the economic benefits obtained from the shorter hospitalization time.

[0118] In recent years, the use of LVADs has trended towards saturation, primarily because their application is limited to a small group of end-stage heart failure patients. The long-standing goal of heart medicine has been to apply early intervention LVAD therapy to less ill heart failure patients. Early intervention therapy is expected to have a major impact on the expanded use of LVAD therapy and the advancement of future heart medicine. Clinical evidence suggests that LVAD support given at the moderate to severe heart failure stage can improve heart function in some non-ischemic cardiomyopathy patients by reverse remodeling of myocardial cells or producing sustained myocardial recovery. However, this early intervention intent must rely on two favorable driving factors: simple and safe surgical procedures, and effective adaptive circulatory support treatment plans that follow the progression of the disease. Continuous flow VAD support is non-physiologic, which can remove the supported heart from the normal healthy recovery path process. However, pulsatile support is physiologic, which promotes reverse remodeling of myocardial cells by providing systolic unloading and diastolic perfusion gain to achieve the therapeutic goal. In summary, the present aortic bypass blood pump invention provides a therapeutic strategy that meets the development conditions of the early intervention trend in heart medicine. The beneficial properties of the efficacy provided by the aortic bypass blood pump device, such as adaptive partial support, minimally invasive surgery, and pulsatile therapy, will collectively make the present invention a potential candidate to help the advancement of future heart failure treatment.

[0119] The blood pump of each of the above embodiments is further described below.

[0120] Figure 13A With Figure 13BFIG. 1 and FIG. 2 are schematic and cross-sectional views of a paracorporeal blood pump device implanted in a human body, according to the first and third embodiments of the present application, respectively. The implantable subsystem of the paracorporeal blood pump device comprises a blood pump 52, an aortic adapter 54, and a drive conduit (or distal drive conduit) 57 connected to the blood pump 52. The blood pump 52 comprises a rigid or semi-rigid housing 52h and a blood sac 529, which is closed at its proximal end and open at its distal end, and seamlessly integrated with the aortic adapter 54. The blood sac 529 is formed by an ovoid membrane 526, and is anchored to the proximal shell 523 of the pump housing 52h through a proximal port 530 and to the distal shell 525 of the pump housing 52h through a distal port 540. Inside the blood pump 52 housing, the ovoid flexible membrane 526 separates a blood chamber B and a gas chamber A, and is suspended to the rigid housing 52h by a pair of stress-relief stems (proximal port 530 and distal port 540). The blood chamber B is used to store blood, and the gas chamber A is used to receive driving air. A pressure sensing mechanism 527 (or blood pressure sensor) is sealingly embedded in the proximal shell 523 of the rigid housing 52h, wherein the sensed pump pressure is transmitted through the membrane 526, propagated in an incompressible liquid or gel contained in a closed pressure sensing chamber 528, and finally received by the pressure sensing mechanism 527. Upon receiving the sensed blood pump pressure, the pressure sensing mechanism 527 will generate an electronic blood pressure signal. The implantable subsystem components are sized and shaped to be implantable, and are suitable for patients with a body surface area (BSA) of 1.2 square meters or more.

[0121] Figure 14A With Figure 14B FIG. 3 and FIG. 4 are schematic and cross-sectional views of a portion of a paracorporeal blood pump device implanted in a human body, according to the second and fourth embodiments of the present application, respectively. The implantable subsystem of the paracorporeal blood pump device comprises a blood pump 62, an aortic adapter 64, a coupler 65, and a drive conduit (or distal drive conduit) 67 connected to the blood pump 62. The coupler 65 is used to connect the blood pump 62 to the aortic adapter 64 for access to the blood vessel system of the device recipient. The blood pump 62 comprises a rigid housing 62h, which further comprises a proximal shell 623 and a distal shell 625. The blood pump 62 is constructed similarly to the blood pump 52 disclosed in Figure 14B except that the opening OP is separate and independent from the aortic adapter 64. The coupler 65 is placed around the neck 643 of the aortic adapter 64. In the following, the design disclosed in Figure 14B will be used to further explain the blood pump design and the underlying design principles.

[0122] Referring to Figure 13BThe blood pump 52 includes a molded rigid housing 52h that also includes a proximal shell 523 and a distal shell 525. The rigid housing 52h has a single opening OP to access the patient's vasculature connected to the aortic adapter 54. The opening OP of the blood pump 52 is seamlessly manufactured with the aortic adapter 54 (or said another way, the blood pump 52 and aortic adapter 54 are integrally formed), which provides a smooth and continuous interface transition to the neck of the aortic adapter 54. The integrated blood bladder 529 and aortic adapter 54 assembly is joined to the proximal and distal shells 523, 525, respectively, by gluing to the proximal and distal ports 530, 540. The blood bladder 529 is anchored to the top of the proximal shell 523 such that the non-flexible disc portion of the blood bladder 529 is glued proximate to the pressure sensing chamber 528 in the center of the proximal shell 523.

[0123] A miniature pressure sensing mechanism 527 is built into the proximal shell 523 and in fluid communication with the enclosed pressure sensing chamber 528. This arrangement allows continuous monitoring of the blood pressure in the blood bladder 529. Since the pressure sensing mechanism 527 does not contact the blood, the protection of the rigid housing 52h ensures long-term sensor reliability and fidelity by isolating the pressure sensing mechanism 527 and its circuitry from the effects of chemical corrosion and protein adhesion caused by direct blood contact.

[0124] A drive conduit 57 is connected to the distal shell 525 to provide counter-pulsating air pressure pulses to drive blood out of or into the blood pump 52. The drive conduit design can be multi-lumen or multi-layered to accommodate electrical wires for pressure signal transmission. A metal coil or fabric mesh or netting can be employed as a conduit wall reinforcement to enhance the anti-kinking ability of the distal drive conduit 57. The overall geometry of the flow path in this blood pump, combined with the valveless aortic adapter design and counter-pulsating blood pump operation, results in superior blood handling characteristics that avoid hemolysis caused by high shear forces and thrombus formation or thromboembolism at low flow rates.

[0125] The innovative design of the blood pump 52 and the blood bladder 529 described above makes the bladder membrane 526 very durable. The blood bladder 529 is an egg-shaped membrane rotating body with the center line of the blood pump 52 as the center of rotation, and two polymer ports (proximal port 530, distal port 540) are combined at both ends of the rigid housing 52h, which are configured as disc or annulus, respectively, and serve as a bending / tension stress release mechanism to reduce stress concentration when connected to the rigid housing 52h. During the blood pump ejection, the bladder membrane 526 will be compressed or folded into a tri-lobe shape, where the maximum strain usually occurs at the crease near the edge of the port (proximal port 530, distal port 540). This local high membrane stress / strain caused by large deformation of the membrane is absorbed and offset by the flexible suspension deformation of the port edge. It is particularly noteworthy that the tri-lobe folding pattern position is not fixed, and the membrane deformation is affected by the direction of gravity, and the crease position will occur randomly. In fact, the patient's body posture and orientation, including standing, sleeping, sitting, and exercising, can change from time to time in daily activities. Therefore, the effect of gravity on the blood volume stored in the blood pump 52 will change direction constantly, resulting in the formation of a non-fixed crease line. This randomly formed membrane folding line feature constitutes a unique anti-fatigue feature of the present invention. It is expected that the present blood pump 52 will have longer durability than conventional fixed folding line membrane designs.

[0126] The bladder membrane folding and expansion is closely related to the vortex structure pattern contained in the blood bladder 529. The feature of the foregoing blood bladder design is that the folding line is randomly formed, so that the vortex structure feature alternates with the change of the folding membrane pattern. Therefore, the flushing effect in the blood pump 52 is strong and irregular, which is characterized by random wandering vortex motion. This randomness of the blood pump vortex structure helps to flush the blood contact surface of the entire blood bladder without any fixed low-speed backflow area near the membrane wall or crease area. It has been observed in animal experiments that the blood pump of the present invention has very strong anti-thrombosis ability.

[0127] Each distal drive catheter in the various embodiments is further described below.

[0128] Referring to Figure 5 , Figure 6 , is a schematic view of a drive catheter connecting the blood pump 92 to the driver 98. The distal drive catheter 97 (drive catheter body segment 991) of the drive catheter pneumatically connects the blood pump 92 to the electromechanical actuator housed within the driver 98, and also transmits the pressure sensing mechanism 527 (see Figure 13B) acquired electrical signals. The distal drive catheter 97 (drive catheter intracorporeal segment 991) is connected at one end to the blood pump housing and has a small external connector at the other end for pneumatic transmission and electrical communication. The distal drive catheter 97 (drive catheter intracorporeal segment 991) is percutaneously passed subcutaneously with a protective cap for protection. The distal drive catheter 97 (drive catheter intracorporeal segment 991) is designed with a small outer diameter and the tubing material is flexible to minimize stress at the exit site EX for patient comfort. A portion of the distal drive catheter 97 (drive catheter intracorporeal segment 991) is covered with a porous fabric to promote tissue ingrowth into the fabric to make the exit site EX resistant to infection. The drive catheter extracorporeal segment 973 that passes out of the epidermis is secured a small distance beyond the skin exit site EX.

[0129] The drive catheter intracorporeal segment 991 and the drive catheter extracorporeal segment 993 and its connector are designed to withstand the tensile loads applied during the percutaneous procedure. After the procedure, the drive catheter intracorporeal segment 991 and the drive catheter extracorporeal segment 993 are continually subjected to loads caused by muscle movement and the drive catheter intracorporeal segment 991 and the drive catheter extracorporeal segment 993 are designed to withstand these fatigue loads for their intended useful life. The outer portions of the drive catheter intracorporeal segment 991 and the drive catheter extracorporeal segment 993 are also designed to be biocompatible and chemically resistant to cleaning agents and disinfectants in clinical use.

[0130] The proximal drive catheter of each of the foregoing embodiments is further described below.

[0131] The drive catheter extracorporeal segment 993, the proximal drive catheter 99 are used to connect the drive catheter intracorporeal segment 991, the distal drive catheter 97, respectively, to the driver 98. The proximal drive catheter 99 has a drive catheter interconnector 93 at one end and a driver connector at the other end. The drive catheter interconnector 93 contains a circuit board that converts the analog blood pump pressure signal to a digital signal and a vibrator that provides tactile feedback in addition to audible alarms. The drive catheter interconnector 93 has a flat shape to prevent twisting of the drive catheter extracorporeal segment 973 when the drive catheter interconnector is secured to the patient's skin. In addition, the drive catheter interconnector 93 and the drive catheter embedded wire are designed to be sealed and protected from water or moisture intrusion. Since the proximal drive catheter 99 is mounted externally, it can be replaced or serviced when deemed necessary, so that when the proximal drive catheter 99 is damaged beyond repair, replacement of the blood pump does not need to be performed surgically.

[0132] Valveless blood pumps have two advantages in terms of blood handling properties: 1) there are no annoying valve sounds and valve-induced blood cell damage, thrombosis and thromboembolism; 2) there is a stronger antithrombogenicity because the bidirectional pulsatile flow has a better surface cleaning effect that minimizes protein adhesion and avoids blood clots associated with interface discontinuities on artificial surfaces in contact with blood. The flow channel in a valveless pulsatile pump is much more uniformly wide than in a valved pulsatile or continuous flow rotary pump. Hemolysis (rupture of the red blood cell membrane) usually occurs in narrow flow channels with high flow velocity gradients, such as the gap between the valve ring and the leaflets of a valved pulsatile pump. In addition, there are often low-velocity recirculation or stagnation zones on the back of the open valve, which can promote thrombus formation. In sharp contrast, in a valveless pulsatile blood pump, the shear stress exerted on the blood cells is actually orders of magnitude smaller, and the low-velocity stagnation zones associated with valve geometry and motion are essentially eliminated, reducing damage to blood cells or platelet activation, reducing blood clot formation and aggregation, and translating into lower doses of anticoagulant use, and easier, safer postoperative care.

[0133] Figures 15 to 17 A schematic view of a blood pump 62, drive catheter 67 and introducer 63 according to another embodiment of the present application is shown. This embodiment emphasizes anatomical adaptability to facilitate easier externalization of the blood pump and drive catheter.

[0134] As Figure 15 With Figure 16 the drive catheter 67 is connected to the distal housing 625 of the blood pump 62. The blood pump 62 has an oval shaped blood bladder and port assembly 659 (including bladder 629 and ports 630, 640), a pump housing 62h (with proximal housing 623 and distal housing 625 with inlet connector 6251), and a pressure sensing system 628 embedded in the proximal housing 623. The aforementioned components, inlet connector 6251, pressure sensing system 628 and drive catheter 67 are substantially identical or correspond to the components / elements of the previously described embodiments, and a detailed description of these elements and their functions is not repeated here.

[0135] In this embodiment, the introducer 63 is provided in the distal housing 625 of the pump housing 62h for connecting the drive catheter 67 to the pump housing 62h. Furthermore, the introducer 63 is configured in a body shape adjacent to the distal housing 625, thereby connecting the power drive catheter to the outer surface of the pump in a tangential direction. Such a body through design allows the design of the pump housing 62h to adapt to the anatomical space available for blood pump placement. The blood pump 62 is rotatably connected to the interface adapter 501 and allows the drive catheter 67 to be routed in the most suitable orientation to enable smooth subcutaneous tunneling and skin exit. In this way, anatomical adaptability to the implant site geometry can be facilitated.

[0136] In this embodiment, the lead-in 63 is remotely fitted in the distal housing 625, while the pressure sensor 6271 (as Figure 25A ) and the pressure sensing chamber 628 are located in the proximal housing 623, more engineering work needs to be performed to separate the signal transmission path from the pneumatic communication path, and to ensure that the blood pump 62 is sealed and protected from the invasion of biochemical fluids, which may damage the fidelity of signal transmission after the device is implanted.

[0137] Referring to Figure 17 , the pump housing 62h has a surface groove 621 formed on the outer surface of the distal housing 625 and above the overlapping joint area DA( Figure 16 ) of the proximal housing 623 and the distal housing 625. The surface groove 621 is configured to extend the wire from the outlet of the lead-in 63 along the surface groove 621 above the overlapping joint area DA and to the electrode 6274 of the second space 6273( Figure 25B ). In some embodiments, the foregoing surface groove 621 is sealed by potting waterproof material and / or annular cover to maintain a smooth outer surface to avoid irritating or injuring the contacted tissue.

[0138] As shown in Figure 18A and Figure 18B , the aortic connector 50 generally requires an interface adapter 501 to be fitted therewith to serve as a connection mechanism to connect the blood pump 62 to the target artery 60. The distal end 504 of the connector 50 opposite to the interface adapter 501 is placed in the blood vessel wall of the artery 60 and is in fluid connection with the human body circulation. While the proximal end of the aortic connector 50 (or the interface adapter 501) has a smooth interface transition to geometrically match the inlet form of the blood pump 62. A coupler is generally required to integrate the proximal end of the connector 50 (or the interface adapter 501) and the inlet of the blood pump 62 together. In some embodiments, the aortic connector 50 can be used as the interface adapter 501. Figure 18A As shown in Figure 18B , an end-to-side anastomosis of a Dacron or PTFE graft 502 sewn with the target artery (or blood vessel) 60 can be used for vascular surgery. In some other embodiments, such as the embodiment shown in

[0139] Figure 19 and Figure 20A long-term axisymmetric elliptical sac and port assembly 650 and its components are shown. The polymer material chosen for these components can be, but is not limited to, segmented polyurethane with various suitable hardness testers. The aforementioned sac and port assembly 650 comprises a flexible membrane sac (sac) 629, a proximal port 630, and a distal port 640, which, in some embodiments, are all formed in an axisymmetric shape relative to a common centerline 62C of the blood pump 62 and integrated together. The proximal port 630 is located at the proximal end 6291 of the sac 629, while the distal port 640 is located at the distal end 6292 of the sac 629.

[0140] Figure 19 An integrated blood sac and port assembly 650 is shown, the end of which of the distal port 640 is wrapped at the distal end of the blood sac 629 and combined with the inverted membrane 62A. Figure 20 The components before assembly are shown. Generally, the sac 629 is manufactured by dip molding, while ports 630 and 640 are manufactured by injection molding. There is no preferred azimuth angle for deflecting sac deformation. Theoretically, when the applied pressure differential exceeds a certain threshold, the thin-walled sac with an axisymmetric elliptical structure will bend into a three-lobed configuration 6293, as shown. Figure 21 As shown. This membrane bending is only related to the final three-lobed configuration 6293 (eigenmode), in which the location of the crease 6294 or fold line is determined by the initial perturbation that induces bending instability. The thickness uniformity in the cross-section cut perpendicular to the centerline (or axis of rotation) 62C of the component 650 of the blood pump 62 is crucial. Care must be taken to maintain high-precision pocket manufacturing to ensure an axisymmetric shape. In real life, the direction of gravity is the primary factor initiating the fold line. The recipient's posture changes constantly according to the patient's daily activities (e.g., standing, sitting, exercising, sleeping, etc.), and the direction of gravity relative to the blood pump direction also changes constantly. Therefore, the crease 6294 of pocket deformation appears in a random pattern, causing high-strain creases to be non-stationarily distributed throughout the pocket. Therefore, avoiding leaving high-strain areas in a fixed position is a key design principle for ensuring a long pocket life.

[0141] Embodiments of the present invention innovate the operating fold line property that makes high strain locations appear non-stationary in the membrane to extend the fatigue life of the membrane sac. Thus the deleterious stress concentration phenomenon often associated with curved blood sacs is ameliorated. Based on this essential change in the behavior of the curved mode, the fatigue life of the membrane will be significantly increased due to this non-stationary fold line formation property that will spread the high strain regions throughout the sac. Furthermore, the beneficial consequences that accompany this non-stationary sac deformation mode depend on the enhanced vortex scouring effect within the blood sac. The sac surface will be more thoroughly washed, forming irregular, walking-like vortices and traversing throughout the vortex. In this way, the likelihood of creating constant low speed recirculation zones or fold line creases in the vicinity of the wall will be greatly reduced, thus enabling a long-lasting, thrombosis-resistant blood pump design.

[0142] Figure 22 、 Figure 23A and Figure 23B An exemplary embodiment is shown that illustrates how the integration method is used to mount the blood sac and port assembly 650 onto the pump housing 62h and to connect the drive conduit 67 to the proximal housing piece 623.

[0143] The blood sac 629 is anchored to the pump housing 62h, which includes a proximal housing piece 623 and a distal housing piece 625 to facilitate the filling and emptying actions of the pump. In general, the flexure properties of the blood sac 629 and the pump housing 62h are quite different. To accomplish a long- lasting blood sac design, an intermediate suspension is required to make the pump assembly continuous in the structural property transition, especially the membrane curved deformation. A pair of flexible ports, the proximal port 630 and the distal port 640, are used as the suspension mechanism to integrate the blood sac 629 with the pump housing 62h. As shown in Figure 23A the proximal port 630, which is presented as a disc, is connected to the proximal housing piece 623; while the distal port 640, which is presented as a ring, is connected to the distal housing piece 625. Mechanically, the proximal port 630 and the distal port 640 act as stress-relieving suspension mechanisms that not only hold the blood sac 629 within the pump housing 62h, but also avoid stress concentration at the interface attachment, thus extending the service life of the blood sac 629.

[0144] As shown in the lower portion of Figure 23A the distal housing piece 625 contains an extension of the inlet connector 6251 that is coupled to the aortic connector 14. The inlet connector 6251 has a first end 6252 that is attached to the blood sac 629 and a second end 6253 that is shaped like a beak, the second end 6253 being coupled to the interface adapter 501 (in Figure 18A 、 Figure 18BThe first end 6252 of the inlet connector 6251 smoothly matches the distal end of the blood bladder 629. However, the opposite second end 6253 is configured to mate with the interface adapter 501, and the coupling design aims to minimize interface discontinuities to avoid clot formation. The beak has a flange structure arranged in the middle region of the inlet connector 6251, which serves as a locking element received by the interface adapter 501.

[0145] During a surgical procedure, the closed-end bladder design of the valveless blood pump 62 will attract air in the bladder top and accumulate air bubbles due to buoyancy forces. Referring to Figure 22 With Figure 24 , a vent 66 is installed or provided in the proximal housing 623, where a narrow passage 661 is provided above the integrated bladder port septum 6301 between the proximal port 630 and the blood bladder 629. In some embodiments, the passage 661 extends along the centerline 62C. After the blood pump 62 is anastomosed with the target artery 60, the trapped air will be pushed by the arterial blood pressure and appear to accumulate on the top space of the blood bladder 629. A thin needle is used to pass through the vent 66 and through the passage 661, the blood bladder, and the port septum 6301 to reach the interior of the blood bladder 629 to expel the accumulated air. The overall bladder port septum 6301 below the vent 66 is relatively stiff and does not bend, which will keep the punctured bladder 629 from further structural damage, avoiding structural damage due to crack propagation initiated at the punctured slit when subjected to periodic pulse pressure and stretching and folding of the adjacent bladder.

[0146] As Figure 23A With Figure 25A shown, a pressure sensing mechanism 627 is embedded in the proximal housing 623. Figure 23A A cross-sectional detail of the integrated proximal housing 623, feedthrough 63, and drive catheter 67 is shown. Figure 25A The profile of the proximal housing 623 is then drawn, which is connected to a feedthrough 63 extending from the dome of the proximal housing 623 for pneumatic and signal communication with the drive catheter 67.

[0147] As Figure 25A , Figure 25B shown, the aforementioned pressure sensing mechanism 627 includes a pressure sensor 6271, which is hermetically housed in a metal can and has a first space 6272 for fluid communication and a second space 6273 for housing a microelectro-mechanical system (MEMS) pressure sensor and related electronic circuitry. A plurality of electrodes 6274 extends from the bottom of the second space 6273 to connect with the electrical wires 6702 of the drive catheter 67 (as Figure 26). The second space 6273 is closer to the drive conduit 67 than the first space 6272. The first space 6272 is open to fluid communication with the sensing medium. Biocompatible fluid or jelly is used as the pressure transfer medium. A lumen or pressure sensing chamber 628 located in the proximal housing 623 and adjacent to the first space 6272 is used to allow the sensing fluid to be enclosed therein. The distal end of the pressure sensing chamber 628 is separated from the blood lumen by a membrane bladder 629. The pressure sensing chamber 628 has two side arms: a first arm 6281 for mounting the pressure sensor 6271 and a second arm 6282 for filling and sealing the sensing medium. Thus, the blood pressure pulses can be transmitted across the membrane bladder 629 and in hydraulic communication with the remote MEMS pressure sensor 6271 located in the second space 6273.

[0148] One embodiment of the present application innovates a pressure-based blood pump control method and sensor design. A miniature MEMS pressure sensor is used, with its electronic circuitry packaged and embedded in the pump housing wall. In principle, the MEMS sensor die is very robust due to its inherent micro-scale structure. In practice, the robustness of the sensor depends on the packaging design. The present pressure sensing mechanism 627 is non-blood-contacting and isolated from the blood-related corrosive biochemical actions, thus providing the long-term signal acquisition and transmission required for long-term implantable assist devices.

[0149] The drive conduit 67 serves as a transducer for electrical signal conversion and pneumatic pulse pressure transfer between the blood pump 62 and the driver 98. Figure 26 A representative multi-layered drive conduit 67 in the present application is shown in FIG. 6. In the present embodiment, the drive conduit 67 has a pneumatic inner lumen (or internal pneumatic tube) 6701, a plurality of electrical wires 6702, an intermediate pneumatic tube 673, a coil 674 (e.g., a metal coil), an outer tube 675, a tether 676, a silicone jacket 677, a rigid driver connector 678, and a protective hollow connector 679.

[0150] The central portion of the drive conduit 67 houses a pneumatic inner lumen 6701 (or air channel, inner tube) with a 2-5 mm inner lumen diameter, which can be chosen depending on whether lower energy consumption or lower surgical simplicity is preferred. The electrical wires 6702 for signal transmission are embedded in the wall of the drive conduit 67. Other variant designs of the drive conduit can be used. Figure 26 In addition to the multi-layered drive conduit 67 design shown, the drive conduit 67 can also be multi-lumened, for example, to allow the electrical wires 6702 to be embedded in several smaller tube lumens and to allow pulsed air to flow in a larger tube lumen 6701, as shown in FIG. 7. Figure 27 One of the smaller tube lumens can be installed with the tether 676 to limit the extension of the drive conduit 67 and to protect the electrical wires 6702 from damage under external pulling forces.

[0151] The inner tube or pneumatic lumen 6701 is received by the pneumatic tube 673, sandwiched by a reinforcement. Between the inner tube 6701 and the pneumatic tube 673, a coil 674 (or fabric wire or mesh) can be backflowed (using heat shrink heat co-molding) as reinforcement to the drive catheter wall, resulting in a drive catheter 67 that is flexible but kink resistant. The outer tube 675 covers the inner and middle pneumatic lumens 6701 and the pneumatic tube 673, and can be used to cover the helically wound electrical wire 6702 as a protective sleeve. In some embodiments, a non-expandable tether 676 can be disposed between the outer tube 675 of the drive catheter 67 and the silicone jacket 677 to enhance the tensile elasticity required during externalization of the drive catheter 67. Silicone jacket 677 has been clinically proven to have minimal irritation to subcutaneous tissue and lowest infection rate of delivery.

[0152] In this embodiment, the pneumatic lumen 6701, metal coil 674, pneumatic tube 673, helical electrical wire 6702, outer tube 675, tether 676, and silicone jacket 677 are packaged into the body of the drive catheter 67. The proximal end 671 of the drive catheter 67 will be inserted into a receptacle provided in the driver 98. The rigid driver connector 678 of the drive catheter 67 is intended to be received by the receptacle in the driver 98. The rigid driver connector 678 is flush mounted with the plurality of electrodes 6781 (e.g., four electrodes 6781) that are welded to the electrical wire 6702. A protective hollow connector 679 (see Figure 23B Figure 23B , Figure 26 ) is placed over the junction of the drive catheter 67 and the driver connector 678 to prevent kinking of the drive catheter 67 at the junction. The proximal end 671 of the drive catheter 67, including the driver connector 678 and the hollow connector 679, is designed to be low-profile so as not to create unwanted puncture trauma.

[0153] As shown in Figure 22 and Figure 23A , the connection of the drive catheter 67 to the blood pump 62 is achieved through the introduction piece 63. Depending on the anatomy in which the blood pump 62 is to be implanted, the introduction piece 63 can be placed in the proximal housing 623 or the distal housing 625. Integrating the introduction piece 63 with the pump housing 62h can change the overall external blood pump configuration and direct the drive catheter 67 in a particular direction to meet implantation requirements, including drive catheter externalization route, post-operative skin care, and device usability.

[0154] As shown in Figure 23A and 25A ​As shown, the introducer 63 has a first portion 631 that is an extension of the proximal housing 623, and the pneumatic lumen 6701, tether 676 and electrical wires 6702 of the drive catheter 67 are coupled in the first portion 631 through an anchor adapter 672. The introducer 63 also has a second portion 632 that interlocks with the first portion 631, which is a cushion bend for the drive catheter 67. The first portion 631 is where the electrical wire connections, tether anchoring and pneumatic lumen bonding and sealing to the housing occur. The electrical wires are required to be protected from exposure to the implant site tissue and must be well protected from tension forces applied during the externalization of the power drive catheter. In addition, the connection of the pneumatic lumen 6701 to the blood pump 62 requires no air and no current leakage. The above mentioned blood pump integration tasks are performed in the first portion 631. The second portion 632, on the other hand, is responsible for housing these interface interface elements, serving as an external protector to protect the interfaces from mechanical stress and environmental fluid or moisture intrusion.

[0155] Figures 19 to 27 A modular design related to the first embodiment of the blood pump of the present invention is disclosed in the middle. In this embodiment, the blood pump 62 includes an axisymmetric elliptical blood sac and port assembly 650 (including the flexible membrane sac 629, proximal port 630 and distal port 640); a pump housing 62h with a proximal housing 623 and a distal housing 625; a drive catheter 67 connected to the blood pump 62, which contains a pneumatic lumen 6701 and electrical wires 6702 in its wall. To integrate the drive catheter 67 with the pump housing 62h, an introducer 63 is used to achieve electrical and pneumatic communication between the drive catheter 67 and the blood pump 62.

[0156] As Figure 19 With Figure 20 As shown, the connection design of the blood sac 629 and ports 630, 640 has been disclosed in the previous section. The point of design and manufacture is to maintain high precision of axisymmetry in the part manufacturing and connection of the sac and port assembly. The pressure sensing mechanism 627 and the introducer 63 are installed in the rigid portion of the proximal housing 623. Figure 22 Figure 23A And Figure 23B A compact introducer 63 design is shown. As can be seen, with the compact introducer 63, a more robust and fault-tolerant wire and connection can be achieved.

[0157] Figure 28A And Figure 28B Some flow patterns related to aortic side-pumping are shown. At the end of diastole and early in systole of left ventricular ejection, the blood pump experiences pump filling and draws aortic blood flow into the pump Figure 28A ​). The upstream and downstream blood around the connector will be drawn into the blood pump by a sharp ninety degree flow turn. This will create a flow separation and low speed recirculation zone T-201. In addition, very high shear will occur in the corner region of the T-junction. On the other hand, during diastole after the aortic valve closes, the blood stored in the pump will be ejected back into the circulation, creating a jet flow against the contralateral aortic wall Figure 28B ). This side-ways, impinging jet flow has very high local pressure at the impingement point T-202, the so-called stagnation point, where the flow velocity is practically zero and all kinetic energy associated with the flow velocity is converted into potential energy, called total pressure. This high pressure impinging jet flow can cause vascular maladaptation, including smooth muscle cell proliferation and resulting stenosis of the vessel wall, and risk of aortic dissection due to the persistent local high blood pressure. All these non-physiological flow patterns and induced high pressure, high shear, low speed recirculation phenomena are prevalent in the vicinity of the T-tube. This turbulent, complex flow abnormality will decay or diminish over a distance of 3-5 times the implanted aortic lumen diameter. The present insertable flow connector is designed to have an insert catheter length of 5-7 centimeters, which covers most of the pump-induced non-physiological flow region. The biological vessel wall will be protected from the pump-induced pathologic stress conditions due to the shielding of the implanted site aorta by the inserted flow connector, thus protecting the implanted site aorta from acute or long-term remodeling complications.

[0158] In counterpulsation support, the filling and ejection of the pump are driven in alternation with the heart rhythm, which creates a special T-junction flow as shown in Figure 28A and Figure 28B . The aforementioned insertable aortic adapter 14 is further detailed in the perspective view Figure 29 and the cross-sectional view Figure 30 , respectively.

[0159] The aortic adapter 14 is mold injected, with its internal blood contacting surface 141 being ultra-smooth and continuous, without any parting lines. The aortic adapter 14 can have a material of silicone or other polymeric elastomer. In some embodiments, the aortic adapter 14 has a polymeric elastomer that includes a silicone material, or the polymeric elastomer is a mold injectable polyurethane. The aortic adapter 14 includes a lumen 143 for insertion into the aorta 95 (see Figure 5The aortic adapter conduit (or conduit insertion portion) 142 and the neck portion (or convex neck portion) 143 connected to the blood pump in the embodiment. In this embodiment, the convex neck portion 143 has a neck portion body 1431 and an extension 1432 disposed on the neck portion body 1431, wherein the extension 1432 protrudes from the neck portion body 1431, and the maximum inner diameter of the extension 1432 is greater than that of the neck portion body 1431. When the convex neck portion 143 is connected to the blood pump 62, the extension 1432 abuts the inlet adapter 6251 of the blood pump 62, and the neck portion body 1431 is disposed around the coupling 25, which integrates the aortic adapter 14 to the blood pump 62.

[0160] The overall aortic adapter 14 is thin-walled to maximize flow efficiency. To strengthen its thin-walled structure, a pair of Nitinol metal trusses (or truss rings) 144 are embedded and surround both ends of the aortic adapter conduit 142 of the aortic adapter 14.

[0161] Figure 29 is a schematic diagram showing the embedded position of the Nitinol metal trusses 144. In addition, the wall thickness of the aortic adapter conduit 142 is gradually thinned towards the conduit end portions 145. The gradually thinned wall thickness has a dual effect. First, it minimizes the discontinuity of the graft / host connection and keeps the interface clot formation rate consistently lower than the thrombolysis rate provided by the contact endothelium. Second, the compliance of the conduit becomes softer towards the conduit end portions 145, resulting in a compliance matching effect when connected to the aortic lumen.

[0162] One of the complications that plagues large stent graft delivery is the problem of endoleaks. Type I endoleaks refer to the incomplete sealing of the graft end to the implanted arterial endothelial lumen, resulting in a gap between the graft leading edge and the arterial lumen. The exuded blood becomes trapped in the gap and coagulates into a blood clot, which eventually becomes a fibrous neointima that grows uncontrollably over time. The neointima not only blocks the implanted artery, but can also signal and stimulate the coagulation mechanism to attract platelet adhesion and cause thrombotic adverse events. The solution to this endoleak problem is to make the aortic adapter 14 tightly seal against the attached lumen surface. The aortic adapter 14 of the present disclosure proposes a compliance matching design concept that enables the semi-rigid conduit (end) portion 145 to seamlessly attach to the arterial lumen when subjected to pulsatile blood pressure. As shown in Figure 30As shown, the outer diameter 146 of the aortic adapter conduit 142 is slightly larger than the inner lumen diameter, with an over-size ratio (defined as the ratio of the conduit diameter to the inner lumen diameter) in the range of 3-10% at a given nominal blood pressure (e.g. 120 mmHg). With blood pressure fluctuating between systole and diastole, or under the pulse pressure created by counterpulsation support, the compliantly-matched conduit end 145 will dynamically expand and contract in response to the pressure pulsations without creating an interfacial gap.

[0163] The thin-walled tube made of elastomer is flexible and often compliantly-matched, but its strength is insufficient to withstand the compressive forces exerted due to the over-size of the device, which often results in the bending of the wall of the inserted adapter. Therefore, the use of the combination of the nitinol metal stent 144 structure and the elastomer base material with appropriate hardness is important. As shown, the nitinol metal stent 144 provides radial stiffness that will help support the aortic adapter 14 without bending it, and there is a distance "X" between the outermost boundary 1441 of the metal stent 144 and the conduit end 145. In some embodiments, the aforementioned distance X should be evaluated and properly defined. Under the support of the nitinol metal stent 144 as an expandable frame, the gradually-thinning conduit end 145 will not collapse or wrinkle, and it remains circular, bearing against the connected lumen wall and dynamically sealed with the lumen. It is noted that the aortic adapter 14 can expand and contract in response to the pressure pulsations, and the sealing effect is achieved in a dynamic manner, i.e. the aortic adapter 14 and the wall of the aorta 95 expand and contract as a whole to seal the conduit end 145 without causing bleeding complications. Figure 30

[0164] Figure 31 A representative embodiment of the nitinol metal stent 144 is shown, which is typically made of a laser-engraved nitinol tube, further expanded under a series of expansion and heat treatment. Figure 31 A planar unfolded schematic of the metal stent 144 is shown. The metal stent has multiple wave-like structures. The metal stent 144 is self-unfoldable, which can be folded or crimped into a smaller pre-packaged configuration, and can self-release to recover its original shape after being placed at the desired location.

[0165] A simple measure of conduit stiffness (the inverse of compliance) can be represented by the so-called lateral stiffness (LS), which is measured as the force F per unit length divided by the corresponding radial deflection Y. For the current aortic adapter, a suitable lateral stiffness range is 0.01-0.05 Nt / mm Figure 32 2 ​​The embedded nickel-titanium alloy metal stent 144 and the silicone or elastomer base will both contribute to the structural compliance of the co-injected aortic adapter 14. The structural compliance of the metal stent 144 and the structural compliance of the polymeric elastomer of the aortic adapter 14 are approximately equal to each other. It is desirable to have a uniform distribution of flexibility so that stretching and contraction of the composite catheter wall will result in minimal interlaminar delamination tendencies, thereby increasing the service life of the adapter 14.

[0166] The aortic adapter 14 is configured to be connected to the blood pump 62 to facilitate circulatory support. A quick connector type coupler 25 is provided herein. As shown in Figure 33 , it is an exploded view schematic diagram showing the various components of the coupler 25 that integrates the aortic adapter 14 and the blood pump 62 together. The coupler 25 includes a flange base 252, a pair of collars 253 and a hinge (or pivot assembly) 254 that connects the collars 253 to the flange base 252. A spring coil (or spring ring assembly) 255 is loaded in the hinge joint 256 that holds the collars 253 in the open position when the coupler 25 is unlocked Figure 34A . Its locking mechanism is by means of a leaf spring type latch 257 that is made of a slotted spring leaf and is secured by a plate 2571 that is welded to one end of the collar 253. The flange base 252 has a generally circular structure and each collar 253 has an arcuate structure. The hinge joint 256 is located at a first side 252S1 of the flange base 252 and the leaf spring type latch 257 is located at a second side 252S2 of the flange base 252 that is opposite to the first side 252S1. The collars 253 are pivoted to the hinge joint 256 and are rotatable relative to the hinge joint 256 and the flange base 252. In some embodiments, the coupler 25 and the aortic adapter 14 are part of an aortic adapter assembly.

[0167] Figure 34B A schematic view of the coupler 25 in the locked state is shown, where the slotted opening of the leaf spring type latch 257 is tightly engaged with the beveled portion 258 to ensure that the connection is secure without the worry of disconnection. As shown in Figure 35 , the integrated connection of the aortic adapter 14 and the blood pump 62 is through the deformable adapter proximal end 147 Figure 30 , i.e. the front end of the convex neck portion, that acts as a “gasket” between the rigid flange base 252 and the rigid beak flange 81 (to be described later) of the blood pump inlet adapter 6251 that are connected to each other.

[0168] In particular, the quick connector type locking can be easily performed by closing the collars 253 without the worry of accidental unlocking, as Figure 34B and Figure 35The aforementioned leaf spring latch 257 is mounted at the end of a locking ring 253. During locking of the locking ring, the leaf spring latch 257 will flex as it slides over the (convex) ramp 258 on the other opposing locking ring 253 during the locking process. When the leaf spring latch 257 passes the top of the ramp 258, it will drop to the bottom of the ramp 258 by elastic restoring force, acting as a safety to prevent accidental unlatching or opening of the locking ring due to pump vibration or long term rocking. For pump explant or replacement that requires module separation, the leaf spring latch 257 can be bent and lifted upwards by using a tool, allowing an unlatching force to be applied to rotationally open the locking ring 253, thus disengaging the blood pump 62 from the aortic adapter 14.

[0169] Docking joint designs are not feasible for joining two surfaces of smooth tubes in blood flow. In most clinical applications, the joined graft surfaces are roughened to promote endothelialization, so that the tiny interface discontinuities in blood flow are "smoothed out" by the in-grown cells and proteins. The current aortic adapter 14 employs a smooth surface approach to avoid thrombotic adverse events, as explained earlier. As shown in Figure 28A and 28B The blood flow in the aortic adapter is bidirectional in response to the ejection and filling actions of the counterpulsation pump. This powerful bidirectional flow and surface washing effect will easily remove any newly formed blood clots on the roughened surfaces. Therefore, a smooth surface design is considered more suitable and safer for use in the present invention. The interface of two joined smooth surfaces in blood flow requires careful mechanical and hemodynamic design to prevent thrombotic events in situ. In the following, the principles and design methods related to this new joint invention are disclosed.

[0170] Figure 36A and Figure 36B illustrate the two basic interface discontinuities present in a docking joint connection, such as the steps 101, 102, or gap 103 created between the adapter AB1 (such as the adapter of the blood pump 62) and the adapter AB2 (such as the aortic adapter 14), respectively. Figure 36A and 36B are drawn in exaggerated scale, typically in precision machining, such joint discontinuities are within 10-50 microns, which is large enough to cause clotting and thrombosis.

[0171] In practice, even if the machining of each piece is exactly the same, the tolerances of matching two separate pieces are required. As Figure 36ATwo misaligned joints are depicted, everything is done correctly in relation to the manufacture of the parts, except the centerline misalignment. Forward and rear facing steps 101, 102 will be created, and stagnant flow in the step regions 101, 102 will be the starting point for clot or thrombus emboli to develop. As shown in Figure 36B , an interface gap 103 is created due to the non-parallel mating of the connectors. The gap 103 attracts blood cells to aggregate and further grow into a pseudo-intima, which tends to be uncontrolled, causing the entire blood flow passage to be blocked in addition to thrombus emboli shedding from the intima surface. Misalignment of the interface of the connected objects can be exacerbated when the connected objects are non-rigid. The aortic adapter 14 of the present disclosure is semi-rigid, which can be forced to press into the mating joint (e.g., the inlet adapter), with a deformed configuration and enlarged interface discontinuity. Therefore, to achieve the connection of the current semi-rigid aortic adapter with the blood pump, a novel connection means must be invented, as described below.

[0172] Referring to Figure 37 and 38 , in some embodiments, the blood pump 62 has an inlet adapter 80, which includes the following parts: a beak flange 81, a beak 82, and a connector body 83 as an extension of the housing of the blood pump 62. The connector body 83 is equipped with a plurality of eyelets 86 for connecting the inlet adapter 80 with the blood pump 62.

[0173] The inner diameter 84 of the beak 82 is slightly larger than the inner diameter 148 of the neck portion 143 of the aortic adapter 14 (see Figure 30 ). The contact area between the beak 82 and the proximal end 147 of the adapter is an annular taper (or chamfer, ramp) 149, as shown in Figure 30 and 35 , so the beak 82 can be referred to as a tapered beak. The aforementioned chamfer 149 is inclined with respect to the centerline of the catheter insertion portion, and the taper angle of the chamfer is substantially in the range of 30-60 degrees measured from the centerline of rotation of the inlet adapter 80. In the initial locking engagement, the locking ring 253 with the inward internal groove 2531 will loosely clamp the flange of the flange base 252 and the flange of the beak flange 81. With the locking of the locking ring 253, the aforementioned beak flange 81 (of the inlet adapter 80) and the flange base 252 (of the coupler 25) will be received and squeezed by the internal groove 2531 of the locking ring 253, thereby compressing the end 147 (of the aortic adapter 14) of the silicone in the middle of the clamping and generating a clamping force of a firm connection. In this way, the inlet adapter 80 of the blood pump 62 will be stably connected together with the aortic adapter 14.

[0174] The aforementioned clamping force generation mechanism is as shown in Figure 35As shown. The flange base 252 has two steps 2521 and 2522, which are responsible for generating the clamping force. Before the locking ring 253 is closed, the step 2521 should first engage with the slot 1433 of the convex neck portion 143 of the aortic adapter 14 (see Figure 30 ). This engagement is achieved by first folding and pressing the convex neck portion 143, and then inserting the deformed convex neck portion 143 through the flange base 252, allowing the elastic restoring force of the aortic adapter 14 to restore the folded and pressed convex neck portion 143 to its original circular (or shape), and allowing the step 2521 to engage with the slot 1433. The height Z of the internal groove 2531 of the aforementioned locking ring 253 controls the extrusion deformation of the adapter proximal end 147 of the convex neck portion 143. Referring to Figure 35 , it can be found that the gap Z0 (i.e., the thickness of the extruded adapter proximal end 147) is less than the thickness Z3 of the end 147 of the aortic adapter 14 in the fully locked configuration when the locking ring 253 is closed and locked (see Figure 30 ), that is, the thickness Z0 of the extruded adapter proximal end 147 is less than the thickness Z3 of the initial end 147 (Z0 < Z3). Regarding the thickness Z0 of the extruded adapter proximal end 147, from Figure 35 the following formula can be obtained:

[0175] Z0 = Z - Z1 - Z2… Formula (1)

[0176] In Formula (1), Z1 and Z2 are the thicknesses of the beak flange 81 for clamping the mating part and the step 2522 of the flange base 252 as shown in Figure 35 respectively. Usually, the thickness Z3 is greater than the gap Z0. Therefore, the adapter proximal end 147 of the strained aortic adapter generates the clamping force required to seal-connect the beak flange 81 and the annular flange base 252. The strain of the adapter proximal end 147 of the silicone flow adapter 14, defined as (Z3 - Z0) / Z3, is in the range of 10 - 30%, which is sufficient to ensure a reliable sealed connection.

[0177] Figure 39 shows the engagement characteristics of the aforementioned connection of the beak 82 with the shallow inclined (cone) surface 149 of the adapter. When making the connection, the beak 82 and its beak leading edge 85 will sink into the semi-rigid shallow inclined surface 149. The aortic adapter 14 has a depth equivalent to the leading edge radius of the beak leading edge 85, generally 30 - 50 microns. In Figure 39In this embodiment, the dashed line and the number in parentheses indicate the initial contact of the beak 82 with its leading edge 85, and the solid line and the number without parentheses indicate the state in the locked position. Note that the interface discontinuity is reduced due to the aforementioned shallow bevel 149 and the aforementioned recess of the beak 82 from its original shape (dashed line). The thickness of the internal groove 2531 controls the tight fit of the coupling. As previously mentioned, the proximal end 147 of the compliant aortic adapter will be compressed with approximately 10-30% strain to provide the required coupling force to achieve a leak-free integration against pulsatile pumping.

[0178] The current design of the interface connection between the blood pump 62 and the aortic adapter 14 has two hemodynamic advantages that reduce in situ thrombosis. First, the interface discontinuity in the form of a step or gap type joint is practically not created as observed in conventional butt connections. Second, the stagnation flow in the interface at the leading edge 85 of the beak 85 can be minimized. Thus, the blood flow through the connection interface will remain high velocity flow, significantly improving the butt connection defect, i.e., a significant improvement against the forward or backward step 101, 102 or gap 103 that has been created in the past at the interface.

[0179] The tapered bevel 149 of the present embodiment is inclined at an angle of inclination with respect to the flow direction. This ramp interface design avoids the creation of a step or gap at the joint due to limited manufacturing precision or matching eccentricity associated with conventional butt connections. However, this tapered bevel 149 has an intrinsic defect in achieving concentric centerline alignment of the mating counterparts. The connection of the aortic adapter 14 to the beak 82 does not have a strict lateral constraint to ensure coupling alignment. In order to connect the rigid beak 82 concentrically with the semi-rigid bevel 149, it is essential to snap the lock ring simultaneously around the entire peripheral edge of the flange base 252. When the simultaneous snap / lock engagement cannot be completed, the bevel 149 that snaps first will be more strained than the other free parts, creating a tendency to tilt or arrange the rest of the contact surface, resulting in an eccentric pump connection. This eccentric connection is often a factor in creating a step or gap at the interface, resulting in thrombosis. This shortcoming is remedied by configuring the lock ring profile 259( Figure 34A ) on the distal side (below) of the lock ring 253 in such a way that it includes all the circumferential contact areas in a locked engagement. When locked, the contact edges of the metal (in some embodiments) beak 82 will sink slightly into the bevel 149 of the silicone material (in some embodiments) with a controlled depth of compression and further reduce the interface discontinuity exposed to blood flow. While a moderate anticoagulant regimen can significantly reduce or eliminate conventional interface thrombosis.

[0180] The structural deformability and delivery method involved in the present aortic connector 14 impart special design features to the present invention. In fact, material elasticity considerations need to be carefully incorporated into the current design. The surgical insertion of an endograft into the aorta through an incision in the aortic wall is challenging in terms of safety during surgery and long-term reliability. To this end, in some embodiments, the material chosen for the aortic connector 14 should have a defined memory shape. During device placement, the connector 14 is first crimped into a smaller placement configuration (e.g., the crimped connector 14 shown in FIG. 1 1 ) and this placement configuration guarantees a quick and safe device implantation. After the crimped aortic connector 14 is placed at the intended implantation site, the placement configuration of the aortic connector 14 should be released to self-expand to its original memory shape. Figure 40

[0181] Prior to insertion of the aortic connector 14, a 12-14 mm diameter hole should be formed in the aortic wall. In making such an access hole, care must be taken to avoid creating any cut edges that can become crack initiation points when the device is inserted and the wall needs to be expanded. A side-biting aortic punch, as disclosed in U.S. Patent Application No. 17 / 034036, is an ideal tool for making a large hole in the aorta. With just one punch, a hole can be successfully made without a crack edge.

[0182] The aortic connector 14 morphologically comprises two circular tubes connected together forming a T-shaped flow connector for performing aortic bypass support. The catheter wall is typically 1-2 mm thick and the material used is a polymer with appropriate stiffness, such as silicone or polyurethane, with a hardness of, for example, Shore A 80-90. The crimped placement configuration is very different from the commercially available large stent grafts covered with Dacron or PTFE (polytetrafluoroethylene) fabric. In Figure 40 The crimped / placement configuration of the aortic connector 14 is shown (with the nitinol metal stent 144 embedded and deformed along with the aortic connector 14 with a polymer base). The aortic connector 14 is folded by the flattened catheter insertion portion aortic connector catheter 142 and its T-shaped collar portion 143 is correspondingly squeezed and flattened to sink into the folded connector 14 body as shown in Figure 40 The diameter of the folded configuration of the aortic connector 14 is approximately half of the original deployed circular diameter.

[0183] This folded connector 14 can be held in place by a string tension as shown in Figure 40 Three fixation strings can be provided at the two edges and the center of the catheter, other fixation methods can also be considered. Figure 41A 、 41B ​Figures 41C and 41D show four representative stages of the implantation configuration of the aortic connector 14. Stage 1 (see Figure 41C) Figure 41A The diagram shows the initial penetration of the pre-coated package (aortic connector 14) through the inlet hole when the pre-coated package is tilted at an angle to the axis of the aorta 95. The second stage ( Figure 41B The image shows the aortic connector 14 in a fully inserted configuration inserted into the aorta, with a pre-coated package pushed past the inlet port, one end of which rotates and falls into the inlet port. The fully inserted pre-coated package is then retracted so that its coiled neck portion 143 aligns with the inlet port. Figure 41C The restraining cords are released, allowing the coiled, compressed aortic connector 14 to elastically expand and return to its original shape. Figure 41D The released aortic coupling 14 will be tightly surrounded by the aortic lumen, ensured by selecting an appropriate oversized ratio before insertion. This is achieved by slightly deforming the T-shaped neck portion 143, which is part of the flange base 252 of the coupling 25. Figures 33 to 34B It can be installed on the T-shaped protruding neck portion 143 to prepare for connection of the blood pump 62. Thus, a secure device connection is achieved by placing the inlet beak portion 82 on the tapered shallow bevel 149 and then closing the two locking rings 253 of the coupler 25, thereby easily achieving the blood pump connection.

[0184] Additional safety measures can be applied to enhance hemostasis and stability of the implanted para-aortic pump system. Due to the weight of the pump 62 and the pumping force generated by the counterpulsation support, the para-aortic placement of the pump inevitably involves lateral forces (perpendicular to the longitudinal direction of the aorta) and torque applied to the aortic connector 14. These device-related external forces may affect long-term remodeling of the vascular structure at the implantation site. A purse-string suture can be placed in the adventitia around the access port. The purse-string suture additionally tightens the aortic wall against the inserted aortic connector 14 and acts as a protective measure against enlargement of the access port. Furthermore, surgical tape can be wrapped around both ends of the aortic connector catheter 142 and tightened, reinforcing the overall connection between the inserted aortic connector 14 and the aorta. A well-fitting design and the application of loop tape will provide double protection against endoleak. Sometimes, blood pressure may rise above the upper limit that can be ensured by compliance matching to prevent endoleak. In such extreme cases, surgical tape will come into play as a rigid restrictor, sealing the ends of the joints to be separated and ensuring hemostasis.

[0185] like Figure 42In some embodiments, the step-by-step implantation of the aortic adapter 14 is described in detail. Prior to the start of implantation, the aortic adapter 14 is prepared in a coiled pre-roll form / compressed deployment form. After the left thoracotomy exposes the target thoracic aorta, the cross-clamp distance of about 10 cm across the implant site is determined. The entry hole in the aorta is first marked with a hole perimeter marker. The banding suture is then sutured outside the hole perimeter in the adventitial layer. The aorta can be partially dissected from the surrounding connective tissue, and a pair of surgical tapes can be wrapped around the aorta. After the above preparations are completed, the cross-clamping and aortic adapter insertion are performed. These insertion steps are described in the order shown in FIGS. 21A-21C. First, the aorta is cross-clamped to provide an isolated segment without bleeding problems. Then, a customized aortic punch is used to make a large entry hole for the insertion of the aortic adapter 14. The folded adapter pre-roll is then inserted and placed into the cross-clamped aortic segment, as shown in FIG. 21A. Thereafter, the folded adapter 14 is released and restored to its original expanded form, as shown in FIG. 21B. The banding and taping are then tightened as additional protection against high-pressure leaks. The coupler 25 is then installed so that its step 2521 is engaged in the clamping groove 1433 of the aortic adapter 14 neck portion 143 to prepare for the connection of the blood pump 62. The self-alignment capability of the aforementioned coupler 25 enables the inlet connector 80 of the blood pump 62 to be properly positioned and locked with the aortic adapter 14. The remaining implantation steps are conventional, including cross-clamp release, blood pump venting, and pump start-up support. In general, the cross-clamp time required for the insertion of the aortic adapter is about 10 minutes for a well-trained surgeon. During this cross-clamp period, the abdominal organs will be deprived of blood perfusion and can cause ischemic injury. To mitigate this potential surgical injury to the organs, partial femoral-femoral extra-corporeal membrane oxygenation (femoral-femoral ECMO) support can be used to perfuse the abdominal organs and lower extremities. However, whether to use ECMO support is at the discretion of the surgeon. In general, the average patient can tolerate 20 minutes of ischemic time. Figure 17 Figure 41A 41B Figure 41D

[0186] ​​​​In summary, the present embodiment provides a ventricular assist device, comprising a blood pump, a drive catheter and an introduction piece. The blood pump comprises an axisymmetric elliptical blood sac and port assembly, which comprises a flexible blood sac, a proximal port and a distal port, wherein the flexible blood sac is connected to the proximal port and the distal port as a stress-relieved suspension mechanism. The blood pump further comprises a pump housing, which comprises a proximal housing part and a distal housing part, wherein the stress-relieved suspension mechanism is connected to the pump housing. The blood pump further comprises a pressure sensing system embedded in the proximal housing part, wherein the pressure sensing system comprises a blood pump pressure sensor and a pressure sensing chamber filled with incompressible fluid for pressure transmission. The aforementioned drive catheter comprises a pneumatic lumen, at least one electrical wire and a tether included in the wall of the drive catheter, wherein the electrical wire and the tether are arranged in the wall of the drive catheter. The aforementioned introduction piece connects the drive catheter and the pump housing.

[0187] An embodiment of the present application discloses a pulsatile blood pump design incorporating a non-stationary folding line concept in the construction of a long sac, which can substantially extend the durability of a replacement blood pump. In addition, a miniature pressure sensing system is also provided, which can be used as a reference waveform for real-time pump control based on real-time big data and long-term trend analysis, disease monitoring and diagnosis. In addition, the embedded pressure sensing system is non-blood contact, thus greatly improving the reliability requirements when constructing an implantable sensor system.

[0188] The present embodiment has at least one of the following advantages or effects. By connecting the drive catheter and the pump housing through the introduction piece, a compact introduction design can be provided to make the electrical wire and signal transmission more robust and have a higher fault tolerance. In addition, the compact introduction design integrates the inductive electrical wire and the pneumatic tube with the blood pump. This compact property is particularly important for implantable devices. It not only simplifies the surgical operation and reduces the risk of perioperative implantation, but also helps to reduce the postoperative morbidity related to drive catheter infection.

[0189] In some embodiments, the introduction piece is integrated with the distal housing part of the pump housing, and the introduction piece has a first part as an extension of the distal housing part, and the pneumatic lumen, the tether and the electrical wire of the drive catheter are coupled in the first part; a second part interlocks with the first part and serves as a buffer bend of the drive catheter to achieve the advantages of anatomical adaptability and adaptability to the geometry of the implantation site.

[0190] An embodiment of the present invention provides an aortic adapter assembly for an implantable ventricular assist device, comprising: a T-shaped flow connector including a conduit insertion portion and a collar portion, wherein the conduit insertion portion is connected to the collar portion, both having smooth blood-contacting surfaces; and a stent disposed in the conduit insertion portion. The T-shaped flow connector has a polymeric elastomer reinforced by the stent having a nickel-titanium alloy material. The conduit insertion portion has a tapered wall at its two ends, the conduit ends having a proper distance from the outermost boundary of the metal stent, and the conduit ends have a compliance matching effect to the implantation site artery, and the proximal end of the collar portion is connected to an inlet adapter of a blood pump.

[0191] An embodiment of the present invention has at least one of the advantages or effects as follows. The present invention discloses a flow connector assembly that enables blood flow into and out of a paracorporeal ventricular assist device, particularly a counterpulsation blood pump. Unlike many existing flow connectors that use a rough surface method to promote endothelialization to avoid thrombosis adverse events, the aortic adapter of the present invention uses a smooth surface, insertion-type prosthetic graft concept to construct the flow connector. In addition, a compliant matching design is implemented around the inserted conduit end, combining a tapered wall feature with a thin-walled polymer supported by a super-elastic nickel-titanium alloy stent to achieve a no-leak requirement. The abnormally high pressure, high shear and low recirculation flow phenomena associated with paracorporeal counterpulsation pumping are contained within the artificial surface of the inserted conduit. Therefore, the blood flow dynamics impact and risk factors caused by pathological devices are substantially eliminated, and events such as endothelial cell invasion, lipid infiltration, smooth muscle cell proliferation, vascular stenosis, and arterial wall peeling are significantly reduced. To achieve a good connection between the semi-rigid flow adapter and the blood pump, the present invention provides a quick connector coupler. The coupler has a self-aligning interface design that minimizes the discontinuity of steps and gaps, thereby reducing the likelihood of thrombotic adverse events at the interface joint. In conjunction with the aortic adapter invention, a specially designed implantation method is provided to ensure a quick and safe implantation process. The coiled aortic adapter is made into a pre-coiled package / implantation configuration, which reduces the overall size to half of its original size. This pre-coiled adapter can be easily inserted into the aorta of the implantation site and expand to its original configuration by itself, thereby forming a snugly installed flow connector without worrying about leakage. It not only helps to reduce the implantation risk in surgery, but also helps to reduce the postoperative morbidity associated with device-induced flow and vascular adaptation at the implantation site.

[0192] The ordinal numbers in the specification and claims, such as "first", "second", and the like, do not have a sequential or chronological precedence relationship with each other, and are only used to distinguish between different elements with the same name.

[0193] Variations of the described embodiments will occur to those of ordinary skill in the art upon reading the foregoing description. The embodiments described herein are meant to be illustrative only and are not intended to limit the scope of the application. Accordingly, the application includes all modifications and equivalents thereof.

Claims

1. An aortic connector assembly for an implantable ventricular assist device, comprising: A T-type flow connector, comprising: A catheter insertion portion and a protruding neck portion, wherein the catheter insertion portion is connected to the protruding neck portion, and both the catheter insertion portion and the protruding neck portion have smooth blood contact surfaces; and A truss, disposed within the conduit insertion portion; and A coupler is disposed on the protruding neck portion; The T-type flow connector has a polymer elastomer and is reinforced by a truss made of nickel-titanium alloy material; The catheter insertion portion has a wall that gradually thins at the two catheter ends of the insertion portion. The ends of the catheter ends are at an appropriate distance from the outermost boundary of the truss, and the catheter ends have a compliant matching effect with the artery at the implantation site. The proximal end of the protruding neck portion is used to connect to an inlet connector of a blood pump, and the blood pump is securely connected to the T-type flow connector via the coupler. The coupler includes: One flange seat; A pair of collars, rotatably mounted on the flange seat; and A latch is provided on one of the collars and used to lock the collar; The collar has an internal groove that can be clamped by controlled compression to seal the T-type flow connector.

2. The aortic connector assembly as claimed in claim 1, wherein, The truss and the T-type flow connector are co-injected and embedded in the wall of the conduit insertion portion of the T-type flow connector.

3. The aortic connector assembly as claimed in claim 2, wherein, The polymer elastomer of this T-type flow connector is made of silicone resin.

4. The aortic connector assembly as claimed in claim 2, wherein, The polymer elastomer is a molded polyurethane.

5. The aortic connector assembly as claimed in claim 2, wherein, The structural compliance of the embedded truss and the structural compliance of the polymer elastomer are approximately equal to each other.

6. The aortic connector assembly of claim 1, wherein, The gradually thinning end of the catheter is sharp.

7. The aortic connector assembly of claim 1, wherein, The protruding neck portion has a shallow bevel to mate with the inlet connector of the blood pump, and the inner diameter of the protruding neck portion is slightly smaller than the inner diameter of the inlet connector of the blood pump.

8. The aortic connector assembly of claim 7, wherein, The shallow bevel is inclined relative to the direction of extension of the catheter insertion portion.

9. The aortic connector assembly of claim 1, wherein, The truss has multiple wave-shaped structures.

10. The aortic connector assembly of claim 1, wherein, The protruding neck portion includes: A neck body; and An extension is provided on the neck body, wherein the extension protrudes from the neck body and the maximum inner diameter of the extension is greater than the maximum inner diameter of the neck body.

11. The aortic connector assembly of claim 10, wherein, When the protruding neck portion engages with the inlet connector of the blood pump, the extension is in close contact with the inlet connector, and the neck body is disposed around the inlet connector.

12. The aortic connector assembly of claim 1, wherein, Each of the aforementioned collars has a flange profile that enables the aforementioned collar to engage simultaneously with the edge of the flange seat.

13. The aortic connector assembly of claim 1, wherein, The latch is made of a spring plate to ensure that the coupler is locked in the locked state without worrying about accidental disengagement.

14. The aortic connector assembly of claim 1, wherein, The coupler also includes a hinge joint disposed on the flange seat, and the aforementioned collar pivotally connected to the hinge joint, which is rotatable relative to the hinge joint and the flange seat.

15. The aortic connector assembly of claim 14, wherein, The hinge is located on the first side of the flange seat, and the latch is located on the second side of the flange seat, wherein the first and second sides of the flange seat are opposite each other.

16. The aortic connector assembly of claim 15, wherein, The latch has a slot that engages with an inclined portion disposed on an opposing of the aforementioned collar.

17. The aortic connector assembly of claim 1, wherein, The flange seat has a generally circular structure, and the aforementioned collars have an arc-shaped structure.

18. A blood pump device, as an implantable ventricular assist device, comprising: A blood pump having an inlet connector, wherein the inlet connector has a conical beak; as well as An aortic connector assembly as claimed in any one of claims 1 to 17.

19. The blood pump device as claimed in claim 18, wherein, The inner diameter of the inlet connector is slightly larger than the inner diameter of the protruding neck portion of the T-type flow connector.

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