Ventricular assist device with a durable displacement blood pump with a built-in pressure sensor

Abstract

A ventricular assist device is provided, comprising a blood pump, a drive line, and a feedthrough. The blood pump comprises a pump housing, an axisymmetric elliptical blood sac and stem assembly received within the pump housing, and a pressure sensing system embedded within the pump housing. The drive line comprises a pneumatic lumen, at least one wire and a tether contained within the wall of the drive line, and said wires and said tether are disposed on the pneumatic lumen. The feedthrough connects the drive line to the pump housing.

Description

Technology Field

[0001] The present invention relates to a ventricular assist device (VAD), and more particularly to a left ventricular assist device (LVAD) having a valveless and long-term blood pump to satisfy blood pressure sensing and mechanical circulation support. Background Technology

[0002] Pulsatile circulatory support generally involves moving blood volume into and from a confined space within a blood pump. Energy transfer is associated with the charging and ejecting motions of the blood volume moved within the blood pump. These energy converters, which transfer externally generated kinetic energy to the blood flow to aid human circulation, are known as displacement pumps, and the energized blood flow is often pulsatile. Historically, pulsatile blood pumps were classified into pumps with valves and pumps without valves. The former typically separates the inflow and outflow tubes, each of which is equipped with an artificial heart valve to achieve unidirectional flow delivery within the blood pump.

[0003] To maximize the efficacy of pulsatile assistance, it is better for the employed blood pump to function in synchronization with the heartbeat. Most often, blood pumps deliver blood in a counterpulsive manner when anastomosed to an artery. Counterpulsive circulatory assistance has been clinically proven to be effective and therapeutic, providing systolic decompression to the myocardium during systole and improving coronary and tracheal perfusion during diastole. To date, pulsatile pumps implantable in almost any organ execute counterpulsation based on electrocardiogram (ECG) waveforms as a reference signal to trigger pump ejection and blood filling. For long-term circulatory assistance, reliable ECG waveforms are paramount in commanding pumping action. Often, electrical leads must be implanted and externalized along with pneumatic drivelines. Consequently, the complexity and invasiveness of the surgery increase, leading to increased bore size and rigidity of the drivelines. As a result, the morbidity of mechanical failure or infection associated with the drivelines and ECG leads has worsened.

[0004] A significant portion (20-30%) of patients with advanced heart failure suffer from arrhythmias. Since ECG waveforms are irregular in patients with arrhythmias, it is difficult to determine the correct timing to trigger counterpulsatile pumping. Incorrectly executed counterpulsatile correction can do harm rather than save a weakened heart. Therefore, the use of counterpulsatile assistance has been limited to the population of patients with advanced heart failure, rather than those with arrhythmias.

[0005] The role performed by blood pump pressure is twofold. First, it serves as a standard for real-time pumping control. Second, recorded pressure data can reflect disease development after device intervention and can be used for long-term health monitoring and trending purposes. For real-time pumping control, specific inter-beat filtering procedures and control logic must be incorporated to command pump charging and discharging movements to facilitate anti-pulsational circulation assistance based on instantaneously detected pressure waveforms.

[0006] The pressure detected by the blood pump sensor is a superimposed signal of the intrinsic arterial pressure and the pressure delivered by the driver. When the pumping operation is stopped, the detected pressure is approximately equal to the intrinsic arterial pressure. In long-term health monitoring applications, pumping control can be configured to have specific beats without operating the device; for example, during a one-minute assist period, two consecutive beats can be programmed not to assist so that the actual arterial pressure can be acquired and registered as an ensemble. Thus, 1,440 pressure waveform ensembles will be collected per day, which is sufficient to determine arterial pressure behavioral trends for long-term disease state monitoring and diagnosis.

[0007] Displacement blood pumps have generally been considered less durable than rotary pumps. Among failure modes, rupture of the blood sac membrane has been shown to cause catastrophic failure in most supports. Displacement pumping is facilitated by periodic, alternating filling and draining strokes for the blood volume stored in the sac space. External pumping energy is supplied by a driver through the driveline. Generally, low energy is supplied during the sac filling phase, while high energy is supplied during the draining phase to push the stored blood volume into the human circulatory system against arterial pressure. The fatigue life of the sac is determined by the maximum deformation occurring in the folds of the membrane or in the suspension area where the membrane is attached to a rigid housing. In addition to structural strength criteria for sac design, the hemodynamic vortex flushing effect of the sac is also important. The lower the residual volume during the pump draining phase, the better blood is emptied from the sac, and stronger vortices are formed during the subsequent pump filling phase. However, lower residual volume leads to more severe sac deformation in the fold and / or suspension interface areas. Due to these conflicting design criteria, designers often have to choose a compromise design perspective that favors hemodynamic considerations. This is one of the major causes associated with durability issues in membrane displacement blood pumps. Therefore, finding a way to provide a better ventricular assist device with a blood pump that is highly durable and has anti-thrombotic capabilities is an important issue regarding the safety and efficacy of the assist. Prior art literature

[65535] Patent Document 1: U.S. Patent Application Publication US2020 / 0246528 (August 6, 2020) Patent Document 2: U.S. Patent Application Publication US2008 / 0300447 (December 4, 2008)

[0008] To resolve defects in conventional products, an embodiment according to the present invention provides a ventricular assist device comprising a blood pump, a drive line, and a feedthrough. The blood pump comprises an axisymmetric elliptical blood sac and stem assembly comprising a flexible membrane sac, a proximal stem, and a distal stem, wherein the flexible membrane sac is attached to the proximal stem and the distal stem as a stress-relieving suspension mechanism; a pump housing comprising a proximal shell and a distal shell, wherein the stress-relieving suspension mechanism is coupled to the pump housing; and a pressure sensing system embedded in the proximal shell, wherein the pressure sensing system comprises a pressure sensor and a pressure sensing chamber filled with an incompressible fluid for pressure transmission. The drive line comprises a pneumatic lumen contained within the wall of the drive line, at least one wire, and a tether, wherein the wire and the tether are positioned adjacent to the pneumatic lumen. The feedthrough connects the drive line to the pump housing.

[0009] In some embodiments, an air removal port is installed within the proximal shell.

[0010] In some embodiments, the channel communicates with the air removal port, and the channel extends along the centerline of the axisymmetric elliptical blood sac and stem assembly and is positioned over the flexible membrane sac.

[0011] In some embodiments, the flexible membrane sac has an inverted membrane located at the distal end of the flexible membrane sac, the distal stem is wrapped and coupled with the inverted membrane; and the proximal stem is located at the proximal end of the flexible membrane sac.

[0012] In some embodiments, the ventricular assist device further comprises an aortic connector having an interface adapter, wherein the distal shell has a distal shell adapter to facilitate connection of the blood pump to a human artery; the distal shell adapter has a first end and a second end, wherein the first end interfaces with the inlet of the flexible membrane sac and the second end interfaces with the interface adapter to connect to the human artery.

[0013] In some embodiments, the distal shell adapter has a beak, and the beak has a flange structure that is coupled to the interface adapter.

[0014] In some embodiments, the ventricular assist device further includes a driver, and the drive line connects the blood pump to the driver.

[0015] In some embodiments, the pressure sensor is sealedly housed within a metal canister and also includes a first space for fluid communication.

[0016] In some embodiments, the pressure sensor includes a second space closer to the drive line than the first space, and the second space is configured to accommodate a MEMS (micro electro-mechanical system) pressure transducer and electronic circuit.

[0017] In some embodiments, the pressure sensing chamber is located within the proximal shell and adjacent to the first space, and the pressure sensing chamber is configured to contain a sensing fluid.

[0018] In some embodiments, the pressure sensing chamber has a first arm and a second arm, the first arm is used for installing a pressure sensor, and the second arm is used for filling and sealing a sensing medium.

[0019] In some embodiments, the feedthrough is integrated into a proximal shell; the feedthrough has a first portion as an extension of the proximal shell to which the wires of the pneumatic lumen, tether, and driveline are joined; and the feedthrough further has a second portion coupled to the first portion to act as a bending relief of the driveline.

[0020] In some embodiments, the drive line further comprises an intermediate pneumatic tube and a coil, the pneumatic lumen is received in the intermediate pneumatic tube, and the coil is located between the pneumatic lumen and the intermediate pneumatic tube.

[0021] In some embodiments, the drive line further comprises an outer layer tube, the pneumatic lumen and the intermediate pneumatic tube are accommodated within the outer layer tube, and the electrical wires are covered by the outer layer tube.

[0022] In some embodiments, the tether is placed on the outer layer tube.

[0023] In some embodiments, the drive line further includes a rigid driver connector located at the proximal end of the drive line, and the ventricular assist device further includes a driver, the drive line is connected to the driver, and the rigid driver connector is connected to the driver.

[0024] In some embodiments, the rigid driver connector has a plurality of electrodes soldered to the electrical wire mounted at the same height.

[0025] In some embodiments, the feedthrough is integrated with the distal shell, and the feedthrough has a first portion as an extension of the distal shell to which the wires of the pneumatic lumen, tether, and driveline are joined; and a second portion coupled to the first portion to act as a bending relief of the driveline.

[0026] In some embodiments, the proximal shell and the distal shell have an overlapping coupling region, and the pump housing has a surface trench formed over the overlapping coupling region.

[0027] In some embodiments, the trench is sealed by a potting waterproofing material. Brief explanation of the drawing

[0028] FIG. 1 is a schematic diagram of a para-aortic blood pump device according to a first embodiment of the present invention. FIG. 2 is a schematic diagram of a para-aortic blood pump device according to a second embodiment of the present invention. FIG. 3 is a schematic diagram of a para-aortic blood pump device according to a third embodiment of the present invention. FIG. 4 is a schematic diagram of a para-aortic blood pump device according to a fourth embodiment of the present invention. FIG. 5 is a schematic diagram of a para-aortic blood pump device installed in a human body according to the first and second embodiments of the present invention. FIG. 6 is a schematic diagram of a para-aortic blood pump device installed in a human body according to the third and fourth embodiments of the present invention. FIG. 7 is a first schematic diagram of a driver according to an exemplary embodiment of the present invention. FIG. 8 is a second schematic diagram of a driver according to an exemplary embodiment of the present invention. FIG. 9 is a block diagram of the key interconnection signals and driver functions required for the operation of the present invention. FIG. 10 is a block diagram of key interconnect signals and driver functions required for the operation of the third embodiment of the present invention. FIG. 11 is a block diagram of key interconnect signals and driver functions required for the operation of the second embodiment of the present invention. Figure 12 shows trigger-sensing commands for electromechanical actuator (EMA) piston positions associated with counter-pulsatile pumping. FIG. 13a is a perspective view showing a para-aortic blood pump implant according to a first or third embodiment of the invention. FIG. 13b is a cross-sectional view showing a para-aortic blood pump implant according to a first or third embodiment of the invention. FIG. 14a is a perspective view showing a para-aortic blood pump implant according to a second or fourth embodiment of the invention. FIG. 14b is a cross-sectional view showing a para-aortic blood pump implant according to a second or fourth embodiment of the invention. FIG. 15 shows a drive line connection to the pump housing of a blood pump achieved through a feedthrough on a distal shell according to another embodiment of the present invention. Fig. 16 is a cross-sectional view of the blood pump shown in Fig. 15. FIG. 17 shows a surface trench design for extending electrical wires from a feedthrough placed in the distal shell to a pressure sensing chamber in the proximal shell. FIG. 18a shows the surgical end-to-end anastomosis of the blood pump invention to an artery through the use of an interface adapter coupling. Figure 18b shows the connection of this blood pump to an artery via an insertable connection method using a T-shaped intravascular connector combined by an interface adapter. FIG. 19 is a cross-sectional view of a rotational body of an axisymmetric elliptical sac and stem assembly including a sac, a proximal stem, and a distal stem. FIG. 20 is an enlarged view showing the components used in the construction of an axisymmetric elliptical blood sac and stem assembly (note that the sac is in its original form before being combined and integrated with the proximal and distal stems shown in FIG. 19). FIG. 21 illustrates a buckle-shaped trilobate sac configuration (eigenmode) at the discharge end of the blood sac illustrated in FIG. 19. FIG. 22 is a perspective view of a drive line connected to the proximal shell of a blood pump through a feedthrough. FIG. 23a is a perspective view of the distal part of the drive line and the blood pump along cross-section AA of FIG. 6. FIG. 23b is a cross-sectional view of the proximal end of the drive line along cross-section AA of FIG. 22. FIG. 24 is a cross-sectional view of an air removal port installed within a proximal shell corresponding to the first embodiment. FIG. 25a is a cross-sectional view of a feedthrough and pressure sensing chamber within a proximal shell of a first embodiment. Note that a drive line is not mounted and the feedthrough comprises a first part, an extension of the proximal shell, and a second part that engages with the first part. FIG. 25b is a perspective view of a micro electro-mechanical system (MEMS) pressure sensor integrated in FIG. 25a. FIG. 26 is a perspective view of a multilayer drive line of the present invention, which includes an inner tube and coils for pneumatic air transport, as well as an intermediate tube, tether, and outer tube for electrical signal conversion. FIG. 27 is a cross-sectional view of a multi-lumen driveline design option of a driveline. Figure 28a is a typical diagram of the flow characteristics during the pump filling stage. Figure 28b is a typical diagram of the flow characteristics at the pump discharge stage. FIG. 29 is a perspective view of the present embodiment of a T-shaped aortic adapter. FIG. 30 is a cross-sectional view of the present embodiment of a T-shaped aortic adapter. Fig. 31 is an extended two-dimensional example of an embedded Nitinol truss. Figure 32 defines the lateral stiffness (LS) used to measure the conduit with the embedded nitinol truss of the T-shaped aortic adapter. FIG. 33 is an exploded view of the parts included in the coupler. Figure 34a is an integrated diagram of the coupler in an open configuration. Figure 34b is an integrated diagram of the coupler in the latch configuration. FIG. 35 is a cross-sectional view of an aortic adapter integrated into a paraaortic blood pump using a coupler. Figure 36a shows a step discontinuity created by the butt joint method. Figure 36b shows the gap discontinuity created by the butt joint method. FIG. 37 is a perspective view of an inlet adapter installed at the distal end of a paraaortic blood pump. FIG. 38 is a cross-sectional view of an inlet adapter installed at the distal end of a paraaortic blood pump. Figure 39 shows the sunk beak of the inlet adapter when connected to the inclined surface at the neck of the T-shaped aortic adapter. FIG. 40 is a perspective view of a corrugated aortic adapter when packaged in a delivery configuration with a tying strap. Fig. 41a shows a packaged aortic adapter half-inserted into an access hole made in the aortic wall. Fig. 41b shows a packaged aortic adapter fully inserted into the aortic lumen. Fig. 41c shows a packaged aortic adapter repositioned on its T-neck toward the aortic access hole. Fig. 41d shows an aortic adapter unfolded with a T-neck protruding from the aortic access hole after the strap is untied. Figure 42 shows step-by-step instructions for implanting an aortic adapter into a target aortic portion and connecting it to a blood pump. Specific details for implementing the invention

[0029] As described below, there are four embodiments that can be employed to realize the present invention of the para-aortic blood pump.

[0030] Referring to FIG. 1, which is a schematic diagram of a periaortic blood pump device according to a first embodiment of the present invention, the periaortic blood pump device (10) comprises a blood pump (12), an aortic adapter (14), a drive line (16), and a driver (18). The blood pump (12) further comprises a pump housing and a pressure sensor. The blood pump housing consists of two chambers, one for storing blood and the other for receiving driving air. These two chambers are separated by an elliptical flexible membrane suspended by a pair of stress-relief stems attached to the pump housing. The pressure sensor is installed within the pump housing to monitor blood pressure inside the blood pump (12) and generates an electrical blood pressure signal. The aortic adapter (14) is a valveless, T-manifold type flow communicator coupled to the blood pump (12) and the human aorta. In the first embodiment, the aortic adapter (14) and the blood pump (12) are formed integrally and have a seamless blood contact surface, and the aortic adapter (14) is provided to connect the blood pump (12) to the human aorta. The aortic adapter (14) is made of flexible materials so that the aortic adapter (14) is deformed during insertion from a hole made in the aortic wall. This aortic adapter conduit is self-expandable after insertion and is also strong enough to suppress radial compression applied to the contacted aortic lumen due to the oversized fitting. A drive line (16) is coupled to the pump housing of the blood pump (12) to provide pressure pulses to the blood pump (12) and to transmit an electrical blood pressure signal received from a pressure sensor. A driver (18) is coupled to a drive line (16) to receive a detected electrical blood pressure signal, and the driver (18) also includes an electromechanical actuator to generate a pressure pulse according to the electrical blood pressure signal and provide a controlled pressure pulse to a blood pump (12) through the drive line (16).The wearable driver (18) provides pressure pulses in time, synchronized with the heartbeat, to drive and control the discharge and charging of the implanted blood pump (12).

[0031] The driver (18) includes a battery power system (11) and a redundant battery power system (the battery power systems (21, 31, 41) and 8 of FIGS. 2 through 4 are identical or similar), wherein the redundant battery power system ensures a continuous power supply to the driver (18). Power may also be supplied to the driver (18) by an AC adapter for the convenience of the device recipient when mobility is not required. Furthermore, although not shown in FIGS. 2 through 4, a clinical monitor unit may be connected to the driver (18) to provide a user interface to the clinician to display device monitoring or diagnostic information to initiate and optimize patient-specific operating mode settings and to access driver parameters.

[0032] In FIGS. 1 and 2, which show two different blood pump (12, 22) designs coupled with the same driveline (16, 26) and driver (18, 28) systems, FIG. 2 is a schematic diagram of a periaortic blood pump device (20) according to a second embodiment of the present invention, and the difference between the second embodiment and the first embodiment of the present invention is that the periaortic blood pump device (20) of the second embodiment further includes a coupler (or coupling adapter) (25). The blood pump (22) and the aortic adapter (24) of the second embodiment are not integrally formed as a single unit but are separable from each other; the coupler (25) is provided to couple the blood pump (22) to the aortic adapter (24). Care must be taken in the design of the coupler to minimize interface discontinuities around the connected area. During device implantation, the aortic adapter (24) is first inserted into the aorta through an access hole made in the aortic wall. Using specially developed implantation tools, the coupling adapter (25) is positioned around the protruding neck of the aortic adapter (24), to which the blood pump (22) can be connected. After the blood pump (22) is introduced into the thoracic cavity, the blood pump (22) and the aortic adapter (24) are tightly integrated using the coupler (25). This detachable blood pump (22) and aortic adapter (24) design includes surgical and postoperative advantages. During device implantation, the detachable pump design makes insertion of the aortic adapter easier because the operating room is clearer without interference from the pump body. Furthermore, in the postoperative period, the blood pump can be detached and replaced if emergency surgical replacement is required due to a malfunction of the pressure sensor or rupture of the blood sac. The detachable blood pump design of the second embodiment is advantageous in this regard. Since the aortic adapter can remain in the aorta without extraction, the cumbersome and dangerous reoperation associated with the removal of the aortic adapter can be avoided.

[0033] Referring to FIGS. 1 and FIGS. 3, which are schematic diagrams of a paraaortic blood pump device (10, 30) according to the first and third embodiments of the present invention, the difference between the third embodiment and the first embodiment of the present invention is that the drive line (16) of the first embodiment is replaced by a drive line (36) comprising a distal drive line (37), a drive line interconnector (33), and a proximal drive line (39) of the third embodiment. The distal drive line (37) is coupled to the drive line interconnector (33) to transmit an electrical blood pressure signal obtained from a pressure sensor and a pressure pulse sent from a driver (38); and the drive line controller and vibrator (for alarm warning purposes) included in the drive line interconnector (33) are originally included in the driver (18) of the first embodiment, so the driver (18) of the first embodiment has additional drive line controllers and vibrators (compared to the driver (38) of the third embodiment). The driveline controller is used to process the electrical blood pressure signal, and the vibrator is used to provide an auditory alarm or tactile feedback. In other words, the mechanical power delivery, analog / digital signal conversion, and alarm for charging and discharging the blood pump, respectively achieved by the driveline (16) and driver (18) of the first embodiment, are substantially the same as those achieved by the distal driveline (37), driveline interconnector (33), and proximal driveline (39) of the third embodiment.

[0034] The first embodiment has a clearer driveline configuration and places an electronic signal processor within the driver, thereby minimizing the risk of environmental contamination (moisture penetration or moisture condensation) of the detected pressure signal and air leakage at the joints, both associated with the driveline interconnector (33). Nevertheless, this long driveline is more susceptible to contact damage such as wear, kinking, cutting, and abrasion caused by contact with foreign matter during daily activities. Whether electronic or mechanical, most damage to the driveline (16) of the first or second embodiment can lead to surgical blood pump replacement, which is highly undesirable in terms of the risk of reoperation and associated medical costs. The third or fourth embodiment mitigates these disadvantages of driveline damage-related blood pump replacement by employing an intermediate connector (driveline interconnector). Generally, the length of the externalized distal driveline (37) is short, and the interconnector (33) is better protected by skin dressing coverage and / or a patient vest. In extreme cases where the driveline is damaged beyond repair, the proximal driveline (39) most likely to be damaged can be easily replaced without resorting to surgery. In addition, the third or fourth embodiment is less affected by electromagnetic interference because analog-to-digital signal conversion is already achieved in the circuit of the interconnector (33). Since digital signal transmission in the proximal driveline (39) is less affected by electromagnetic interference, the fidelity of the pressure signals can be better ensured in the third or fourth embodiment.

[0035] Referring to FIGS. 3 and FIGS. 4, which are schematic diagrams of a periaortic blood pump device (30, 40) according to the third and fourth embodiments of the present invention, the difference between the third and fourth embodiments of the present invention is that the aortic adapter (34) and blood pump (32) of the third embodiment are formed integrally, whereas the blood pump (42) and aortic adapter (44) of the fourth embodiment are detachable; and the fourth embodiment further includes a blood pump (42), an aortic adapter (44), and a coupler (45) that are identical to the blood pump (22), aortic adapter (24), and coupler (25) of the second embodiment, and thus the description thereof is not repeated. The drive line (46) including the distal drive line (47), drive line interconnector (43), and proximal drive line (49) of the fourth embodiment is identical to the drive line (36) including the distal drive line (37), drive line interconnector (33), and proximal drive line (39) of the third embodiment.

[0036] Referring to FIG. 5, which is a schematic diagram of a periaortic blood pump device installed in a human body according to an exemplary embodiment of the present invention, the periaortic blood pump device (90) comprises a blood pump (92), an aortic adapter (94), an internal drive line (991), an external drive line (993), and a driver (98). In another embodiment, the periaortic blood pump device further comprises a coupler. The portion of the periaortic blood pump device (90) implanted in the human body comprises a blood pump (92), an aortic adapter (94), and an internal drive line (991). In another embodiment, the periaortic blood pump device further comprises a coupler. After surgery, the aortic adapter (94) is installed in the aorta (95), and an outlet position (EX) is created at an appropriate location on the surface of the human body. The portion of the periaortic blood pump device (90) located outside the human body comprises an external drive line (993) and a driver (98). The exit location (EX) is used as a boundary, and the internal drive line (991) has a portion covered with velour fabric for tissue growth to control infection. The implanted velour portion is located 2-5 cm subcutaneously from the exit location (EX). The driver (98) is a wearable or portable device.

[0037] Referring to FIG. 6, which is a schematic diagram of a periaortic blood pump device installed in a human body according to an exemplary embodiment of the present invention, the periaortic blood pump device (90) comprises a blood pump (92), an aortic adapter (94), a distal drive line (97) (including an internal distal drive line (971) and an external distal drive line (973) outside the human body), a drive line interconnector (93), a proximal drive line (99), and a driver (98). In another embodiment, the periaortic blood pump device (90) further comprises a coupler. The portion of the periaortic blood pump device (90) implanted in the human body comprises the blood pump (92), the aortic adapter (94), and the internal distal drive line (971). In another embodiment, the periaortic blood pump device further comprises a coupler. After surgery, the aortic adapter (94) is installed in the aorta (95), and an outlet position (EX) is created at an appropriate location on the skin of the human body. A portion of the paraaortic blood pump device (90) located outside the human body includes an outer distal drive line (973), a drive line interconnector (93), a proximal drive line (99), and a driver (98). An exit position (EX) is used as a boundary, and the distal drive line is divided into an inner distal drive line (971) and an outer distal drive line (973), which are covered with velour for infection control. The driver (98) is a wearable or portable device.

[0038] The transplant subsystem is explained further below.

[0039] The graft is achieved through a relatively small thoracic incision using less invasive surgical techniques via a left thoracotomy. For example, the thoracic incision is made in the seventh intercostal space to allow for the placement of the aortic adapter and blood pump. Two other small incisions are made in the sixth and eighth intercostal spaces, respectively, to introduce the proximal and distal aortic cross-clamps. The cross-clamped aortic portion allows the aortic adapter to be inserted into the graft site through an extra opening made in the aortic wall. The aortic adapter is flexible and can be folded and constrained into a smaller delivery configuration before insertion. Upon completion of delivery to the aorta, the folded aortic adapter must be unwound and restored to its original shape with a predetermined oversize to fit the lumen diameter of the graft site. Therefore, the material of the aortic adapter is important; it must be flexible but possess sufficient radial strength to ensure that the delivered adapter conduit remains in a circular shape without wall bending. Candidate aortic adapter structures may be made of silicone or polyurethane elastomers, or may include polymer structures reinforced by embedding.

[0040] The functional requirements of each of the aortic adapters of the embodiments mentioned above are further explained below.

[0041] Hemodynamically, the aortic adapter serves as a communication link between the blood pump and the body's systemic circulation. In addition to this role, the aortic adapter also functions as a mechanical base that holds the blood pump in place when connected to the aortic adapter. The structure of the aortic adapter must be elastic but resistant to twisting, and strong enough to withstand internal blood pressure and external contact forces applied through contact with surrounding lung tissue or the diaphragm associated with respiration and chest movement.

[0042] The aortic adapter (54) is implanted inside the aorta with two conduit ends (545, 645) that interface with the aortic lumen, forming a host / graft boundary in the bloodstream (see FIG. 7b and FIG. 8b). To minimize discontinuity at the host / graft interface, the two conduit ends (545, 645) are morphologically and elastically configured to have an internal surface profile that widens progressively and a wall thickness distribution that decreases progressively. This conduit end design not only minimizes the step difference at the interface but also constitutes a compliance-matching effect for coupling the aortic lumen and the aortic adapter. As a result, thrombi at the interface can be eliminated, because the rate of thrombus aggregation at the interface is slower than the natural thrombolytic rate provided by endothelial cells. In addition, the gradually thinning conduit wall structure makes the ends (545, 645) smoother (compliant) so that the ends can expand and contract with pulsatile blood pressure and forms a dynamic sealing effect that prevents blood from getting stuck in the gap of the joint interface, which is an interface that is often the source of blood clot formation.

[0043] The driver for each of the embodiments mentioned above is further explained below.

[0044] FIGS. 9 and 10 show right and left perspective views of the driver (78). This compact driver (78) includes internal modules including an electromechanical actuator (EMA), an electronic controller, a pair of main batteries, and a spare battery. This driver (78) also includes external modules of a user interface panel (73), a battery access door (71), a driveline outlet (75), an AC outlet (77), and a pair of ventilation windows (79), as shown in FIGS. 9 and 10.

[0045] Alarm warnings regarding important information during operation, device malfunctions, and aortic pressure conditions will be displayed on the user interface panel (73) of the driver (78). The primary battery can be replaced via the battery access door (91) when the primary battery power is depleted. An electric cable is used to supply power to the driver (78) via a connection through an AC outlet (77) when the patient is bedridden, and power from the wall outlet can be utilized for an extended period. The end of the proximal drive line (99, 993) is connected to the driver (78) via a drive line outlet (75), and both electric sensor signals and pneumatic pulses are transmitted through this drive line outlet (75). A pair of ventilation windows (79) are installed on opposite sides of the driver (78) to allow air to flow through the inside of the driver (78) for cooling.

[0046] The driver (78) can be externally coupled to a clinical monitor, and the clinical monitor is provided to collect and display real-time clinical waveform data and also to store patient data for long-term situation monitoring and diagnosis. Furthermore, the clinical monitor unit provides a user interface to the clinician to display device monitoring / diagnostic information to initiate and optimize patient-specific operating modes.

[0047] The EMA is a pneumatic actuator consisting of a brushless servo motor, a ball screw unit, a piston, and a cylinder assembly. Atmospheric air is used as the driving medium to repeatedly discharge and charge the blood pump.

[0048] The EMA module is housed within a driver held by the implant recipient. The EMA consists of a brushless servo motor, a piston and cylinder assembly, and a ball screw unit comprising a ball screw rod and nut. The piston is rigidly mounted on the top of the ball screw rod, which is rotatably coupled to the nut of the ball screw unit. The servo motor includes a starter and a rotor, and the rotor is integrated with the nut of the ball screw unit. Through electromagnetic coupling of rotor / starter induction, the rotor can rotate in both clockwise and counterclockwise directions, thereby driving the ball screw rod linearly back and forth, resulting in a reciprocating piston stroke motion within the cylinder. The stroke motion of the piston propels air to and from the implanted blood pump through a driveline connecting the blood pump and the cylinder.

[0049] There are two air-driven problems associated with the design of this EMA pneumatic actuator: namely, air leakage and water vapor condensation penetrating from the blood through the blood sac walls. The former impairs pump discharge and charge functions and driver power consumption, thereby reducing auxiliary efficiency, while the latter poses a risk of bacterial intrusion within the driveline. To address these two issues, this EMA incorporates a pressure equalization valve installed on the cylinder chamber wall for air replenishment and humidity reduction. The pressure equalization valve opens periodically at a predetermined frequency to allow mass air transfer between the cylinder and the surrounding environment until the air pressure inside the cylinder chamber equals atmospheric pressure. The EMA integrates position and optical sensors to acquire reference trajectory signals so that the electronic controller can generate coordinated control commands to drive the piston stroke motion as well as actuate the pressure equalization valve. Thus, the timing and frequency of the pressure equalization valve actuated for air exchange can be programmed into the controller. By using this integrated pressure equalization valve, the driving air medium can be continuously maintained in a dry and fully filled state within the pneumatic actuator to ensure long-term safe and effective pumping support of the blood pump.

[0050] As illustrated in FIG. 11, a paraaortic blood pump device according to one embodiment of the present invention is divided into three parts. A first part has an externalized end for communicating with a second part and is mostly installed inside the human body (implant); the first part includes a blood pump (including a blood pump pressure sensor), an aortic adapter, and distal driveline parts disposed inside and outside the human body, respectively. A second part is installed outside the human body and also includes a proximal driveline and a driveline electronic component module (or known as a driveline interconnector). A third part is installed outside the human body and is a driver including an electromechanical actuator (EMA), a controller circuit, a main battery, and a backup battery.

[0051] The blood pump pressure sensor is embedded in the proximal blood pump shell and contained in a small pressure sensing chamber filled with a sensing medium, enabling continuous monitoring of the blood pump pressure. The distal drive line is attached to the pump housing and provides timed pneumatic pressure pulses to command the ejection and filling of the blood sac. The distal and proximal drive lines provide pneumatically driven pressure pulses, generated by the EMA inside the driver, to the blood pump; and transmit an electrical blood pressure signal, generated by the pressure of the blood pump pressure sensor, to the driver. The driving air path (indicated by the dashed arrow) and the electrical signal path (indicated by the solid line) are illustrated in FIG. 11 to illustrate the functional relationship between the interacting modules. Details of the aortic adapter have been described above. The controller circuit may include a brushless motor and a motor controller unit to drive the microcontroller unit as a central processor for processing the received pressure signal and generating control commands to cause the motor controller to actuate the piston movement.

[0052] Referring to FIGS. 12 and 13, a block diagram of the key interconnect signals and internal driver functions required to operate the blood pump is shown.

[0053] The embodiments of the present invention are further described as illustrated in FIGS. 12 and 13. To explain the driving relationship between the external driver and the implant, it is necessary to refer to the above-mentioned blood pump, drive line, distal drive line, proximal drive line, aortic adapter, and drive line interconnector.

[0054] The driver receives a blood pump pressure signal (electrical signal) and processes the signal using a trigger detection algorithm to generate a trigger signal that commands EMA operation in synchronization with the heartbeat. Upon receiving the assigned trigger timing, the microcontroller unit sends commands to the motor controller to facilitate the procedure of the piston from ejection to charge or from charge to ejection, thereby providing counter-pulsatile circulation assistance.

[0055] The architecture of the electronic controller reflects three functional blocks: a microcontroller unit (MCU), a motor control block (motor controller unit), and a power management unit. The following table provides an overview of the description of each functional block of the driver (78).

[0056] Function Name explanation Microcontroller unit (MCU) (130) The MCU processes the blood pump pressure to generate a trigger that synchronizes the pumping device with the left ventricle. Depending on the trigger timing, the MCU sends motor commands to the motor controller to drive the piston from ejection to charging or from charging to ejection, according to the treatment management settings. It also manages system data storage, system status monitoring and alarms, commands sent from clinical monitors, and the patient user interface (LCD display, LED indicators, and auditory alarms). Motor controller unit (120) A motor servo control system that responds to continuous commands from a motor controller unit, MCU, to drive a motor with specific position, speed, and acceleration parameters. Electromechanical actuator (EMA) (110) The EMA drives the piston inside the cylinder by a brushless servo motor to provide pulsatile driving pressure to the blood pump. blood pressure sensor (527) The blood pressure signal is obtained from a blood pressure sensing element built inside the implanted blood pump. The blood pressure signal supplies the pressure waveform to the MCU to process and determine the timing for commanding EMA discharge and charge operations. User interface panel (170) Provides LED indicators, audio alarm, button, and LCD display Power management circuit (140) The circuit manages system power, including power switching, battery status monitoring, and charge control. Docking battery (150) Provides power to the system when not supported externally. Clinical monitor (190) Collects and displays real-time device waveform data; initiates driver parameter settings and saves patient identification data; performs long-term situation monitoring and diagnostic trend analysis.

[0057] Signal acquisition, transmission, processing, and control logic and command generation and EMA operation for generating pressure pulses to drive a blood pump are illustrated in FIGS. 12 to 13 for previously described exemplary embodiments.

[0058] Figure 14 shows trigger-detection commands for EMA piston positions related to counter-pulsatile pumping. In Figure 14, the unassisted aortic pressure (AoP) waveform is represented by a dotted line, while the solid line represents the assisted aortic pressure waveform. When the driver operates in Auto Run mode, driver operation begins, and the system performs "charge-eject-charge-eject..." circulation pumping, which represents normal synchronous counter-pulsatile operation. The MCU monitors the blood pump pressure (BPP) signal (electrical signal) to detect the left ventricle end-diastole (LVED) timing. Upon detecting LVED timing, the MCU generates an F_Trig signal. The time interval between two consecutive F_Trig signals represents an instantaneous cardiac cycle interval (or period). Based on the estimated heart rate calculated from previous cycle intervals, the MCU determines the timing for blood pump ejection, i.e., the E_Trig signal. The E_Trig signal provides timing that commands the motor controller unit to drive the EMA according to predetermined position, velocity, and acceleration profiles. After the discharge stroke is completed and the optimized residence time has elapsed, the EMA is commanded to perform a pre-charge operation at a smooth charge speed until the F_Trig signal appears. Upon receiving the F_Trig signal, the EMA begins to perform the remaining charge stroke at a specific piston speed.

[0059] When the MCU misses the BPP signal (electrical signal) sent from the blood pump, the flushing mode is automatically initiated by the MCU to drive the EMA so that it operates at a predetermined pumping speed and driver stroke volume. The flushing mode is used to prevent blood clot formation within the blood sac, and this flushing mode is a device protection mode instead of providing synchronous circulation assistance.

[0060] The periaortic blood pump device of the present invention, having non-occlusive periaortic features, in principle has better counterpulsatile support efficacy compared to an intra-aortic balloon pump (IABP). Unlike ambulatory IABP patients who must remain in bed or in a hospital, the portable periaortic blood pump device allows patients to have the ability to walk, enabling them to leave the hospital and lead a better life at home. Therefore, the periaortic blood pump device of the present invention can further improve the patient's disease condition and quality of life, in addition to the economic benefits obtained from a shortened hospital stay.

[0061] Trends in LVAD use have stagnated in recent years, primarily because their application is recommended only for patients with end-stage heart failure. Applying early intervention LVAD therapy to less severe heart failure patients has long been a clinical goal, and this is expected to have a significant impact on future advancements in cardiology provided by the expanded use of LVAD therapy. Clinical evidence demonstrates that certain non-ischemic cardiomyopathy patients managed in the moderate to severe stages of heart failure who are assisted by LVADs can improve through myocardial reverse remodeling for functional enhancement or sustained myocardial recovery. Nevertheless, the intent of this early intervention must be underpinned by two possible factors: a simple and safe surgical procedure, and effective and adaptive support following the onset of the disease. Continuous-flow VAD support is non-physiological, which causes the assisted heart to deviate from the normal recovery process. However, counterpulse support is physiological and meets therapeutic requirements by providing systolic depletion and increased diastolic perfusion to promote myocellular reverse remodeling. In summary, the therapeutic strategies provided by this paraaortic blood pump invention align with the advancing trend of early intervention in cardiology. The beneficial properties provided by the paraaortic blood pump device, such as adaptive partial-assisted, less invasive surgery, and counterpulsatile therapy, collectively make this invention a promising candidate to contribute to the future development of heart failure treatment.

[0062] Each of the blood pumps of the embodiments mentioned above is further described below.

[0063] Referring to FIG. 13a and FIG. 13b, which are schematic and cross-sectional views of a portion of a periaortic blood pump device installed in the human body according to the first and third embodiments of the present invention, the implantable subsystem of the periaortic blood pump device comprises a blood pump (52), an aortic adapter (54), and a drive line (or distal drive line) (57) attached to the blood pump (52). The blood pump (52) comprises a rigid or semi-rigid pump housing (52h) of the blood pump (52) and a blood sac (529) having a closed proximal end and an open distal end, and also seamlessly integrated with the aortic adapter (54). The blood sac (529) is formed by an elliptical sac membrane (526), ​​and furthermore, the blood sac (529) is secured to the proximal shell (523) of the pump housing (52h) via a proximal stem (530) and to the distal shell (525) of the pump housing (52h) via a distal stem (540). The space within the blood pump is divided into a blood chamber (B) and an air chamber (A), separated by an elliptical flexible sac membrane (526) suspended through a pair of stress-relieving stems (proximal stem (530), distal stem (540)) attached to the pump housing (52h). The blood chamber (B) is for storing blood, and the air chamber (A) is for receiving driving air. The pump housing (52h) includes a proximal shell (523) in which a pressure sensor (or blood pressure sensor) (527) is sealedly embedded, and the detected pump pressure is transmitted across a pouch membrane (526), ​​propagated into an incompressible liquid or jelly contained in a closed pressure sensing chamber (528), and finally received by the pressure sensor (527). After receiving the detected pump pressure, the pressure sensor (527) generates an electrical blood pressure signal. The implanted subsystem components have a body surface area (BSA) of 1.2 m² 2 It was designed in a size and shape that can be implanted in patients with abnormalities.

[0064] Referring to FIG. 14a and FIG. 14b, which are schematic and cross-sectional views of a portion of a periaortic blood pump device installed in a human body according to the second and fourth embodiments of the present invention, the implantable subsystem of the periaortic blood pump device comprises a blood pump (62), an aortic adapter (64), a coupler (65), and a drive line (or distal drive line) (67) attached to the blood pump (62). The coupler (65) is used to connect the blood pump (62) to the aortic adapter (64) to access the recipient's vascular system. The blood pump (62) comprises a pump housing (62h) further comprising a proximal shell (623) and a distal shell (625). The configuration of the blood pump (62) is similar to that disclosed in FIG. 13b, except that the distal opening (OP) is separated from and independent of the aortic adapter. The coupler (65) is positioned around the flexible neck (643) of the aortic adapter (64). In the following, the design disclosed in FIG. 13b is used to further explain the blood pump design and the basis of the basic design.

[0065] Referring to FIG. 13b, the blood pump (52) comprises a molded rigid or semi-rigid blood housing (52h) further comprising a proximal shell (523) and a distal shell (525). The housing (52h) has a single opening (OP) connected to an aortic adapter (54) to access the recipient's vascular system. The opening (OP) of the blood pump (52) is seamlessly formed with the aortic adapter (54) (or the blood pump (52) and the aortic adapter (54) are integrally formed) to provide a smooth and continuous interface transition to the neck of the aortic adapter (54). This integral blood pump (529) and aortic adapter (54) assembly is attached to the pump proximal shell (523) and distal shell (525) through coupling to the proximal stem (530) and distal stem (540), respectively. The blood sac (529) is fixed to the upper part of the proximal shell (523), and a small, non-bendable circular portion of the blood sac (529) is positioned next to the pressure sensing chamber (528) inside the shell (523).

[0066] A small pressure sensor (527) is embedded in the pump shell (523) and also fluidly communicates with a sealed pressure sensing chamber (528). This arrangement enables continuous monitoring of blood pressure contained in a blood sac (529). Since the pressure sensor (527) does not come into contact with blood, long-term sensor reliability and fidelity are ensured by the protection of the pump housing (52h), which isolates the sensor (527) and its electrical circuit from the effects of chemical corrosion and protein adhesion resulting from direct blood contact.

[0067] The driveline (57) end is attached to the pump shell (525) to provide timed pneumatic pulses to actuate the discharge or charge stroke of blood out of or into the blood pump (52). The driveline design may be multi-lumen or multi-layered to accommodate wires for transmitting pressure signals. Metal coils or fabric mesh may be adopted as wall reinforcements to enhance the anti-twisting function of the driveline (57). The blood pump has a wide overall geometry of the blood flow passage, and with a valveless aortic adapter design and pulsatile pumping operation, it constitutes excellent blood handling characteristics that prevent high shear-induced hemolysis as well as thrombosis or thromboembolism generated at low flow rates.

[0068] The blood sac (529) of the blood pump (52) incorporates an innovative design to make the sac membrane (526) durable. The blood sac (529) is an elliptical-shaped membrane body that rotates about the centerline of the blood pump (52). There are two polymer stems (proximal stem (530), distal stem (540)) attached to the ends of the two housings (52h), each configured in the form of a circular disc or ring and acting as a bending / stretching relief mechanism to relieve stress concentration when attached to the rigid housing (52h). During pump discharge, the sac membrane (526) will be compressed or folded into a trilobate shape, where the highest deformation often occurs at the corrugated fold line near the edge of the stem attachment (proximal stem (530), distal stem (540)). This localized high membrane stress / strain resulting from large membrane deformation is substantially reduced or absorbed by the deformation of the flexible stem edge as a bendable suspension. Note that the trilobate folding pattern is non-fixed, as the folds vary from place to place due to the influence of the direction of gravity. In fact, the patient's body posture and orientation, including standing, sleeping, sitting, and exercise postures, can change frequently during daily activities. Therefore, the gravitational effect or body force acting on the amount of blood stored in the blood pump (52) continuously changes, resulting in the initiation and formation of non-fixed fold lines. These continuous membrane fold lines constitute the unique fatigue resistance characteristics of the present invention. The blood pump (52) is expected to have a much longer durability compared to conventional fixed-fold membrane designs.

[0069] The folding and expansion of the membrane is closely related to the vortex flow pattern contained within the blood sac (529). The aforementioned sac design features the formation of continuous folding lines that cause the vortex structure pattern to alternate in response to the folded membrane pattern. Thus, the cleaning effect of the blood pump (52) is characterized by strong, non-stationary, and random walking-type vortex flow motion. The randomness in this pump vortex flow structure helps to clean the entire blood-contact surface without creating fixed low-velocity zones near the membrane wall or in the fold area. Current blood pumps have been observed in animal experiments to have very high thrombolytic resistance.

[0070] The (distal) drive line of each of the embodiments mentioned above is further described below.

[0071] Referring to FIGS. 5 and 6 for a schematic diagram of the drive line used to connect the blood pump (92) to the driver (98), the internal (distal) drive line (991, 97) of the drive line pneumatically connects the blood pump (92) to an electromechanical actuator housed inside the driver (98) and also transmits an electrical signal obtained from the blood pump pressure sensor (527) (see FIG. 13b). The (distal) drive line (991, 97) has one end attached to the blood pump housing and the other end having a small external connector for pneumatic and electrical communication. The (distal) drive line (991, 97) tunnels subcutaneously and exits the skin. The outer diameter of the (distal) drive line (991, 97) is designed to be small, and the tube material is flexible to minimize stress at the exit position (EX) for patient comfort. A portion of the (distal) drive line (991, 971) is covered with a porous fabric that promotes tissue growth, making the exit site resistant to infection. The externalized (distal) drive line (991, 973) is fixed at a short distance beyond the skin exit site (EX).

[0072] The (distal) driveline (991, 971) and its connector are designed to withstand tensile loads applied during surgical externalization. After surgery, the (distal) driveline (991, 971) is continuously subjected to loads induced by muscle movement, and the (distal) driveline (991, 971) is designed to withstand these loads for its intended service life. The externalized portion of the (distal) driveline (991, 971) is also designed to be biocompatible and to have chemical resistance to cleaning agents and disinfectants when used in clinical settings.

[0073] The (proximal) drive line (993, 99) is used to connect the (distal) drive line (991, 97) to the driver (98). The (proximal) drive line (99) has a drive line interconnect (93) at one end and a driver connector at the other end. The drive line interconnect (93) encloses a circuit board that converts analog blood pressure pump pressure signals into digital signals and a vibrator that provides tactile feedback in addition to auditory alarms. The drive line interconnect (93) is provided in a flat shape to prevent twisting from occurring in the (distal) drive line (973) when the drive line interconnect (93) is secured to the patient's skin. Furthermore, the drive line interconnect (93) and the drive line outer cover are designed to be sealed to protect against the entry of water or moisture. Since the (proximal) drive line (99) is installed externally, it can be replaced and / or maintained when deemed necessary, thereby eliminating the need for surgical blood pump replacement when the (proximal) drive line (99) is damaged beyond repair.

[0074] Valveless blood pumps have two advantages in terms of blood processing characteristics: 1) there is no annoying valve noise and valve-induced blood cell damage, thrombus formation, or thromboembolism; and 2) higher thrombolytic resistance because the bidirectional pulsatile flow provides a better surface cleaning effect to minimize protein adhesion and prevent interface discontinuity-related thrombus formation on blood-contact artificial surfaces. The flow passages of valveless pulsatile pumps are uniformly much wider than those of valved pulsatile pumps or continuous-flow rotary pumps. Hemolysis (rupture of red blood cell membranes) generally occurs in narrow flow passages with high flow rate gradients, such as the gaps between the valve ring and the leaflet of a valved pulsatile pump. In addition, low-speed recirculation or stagnation zones often exist behind the open valve, which can promote thrombus formation. In stark contrast, in valveless pulsatile blood pumps, the shear stress applied to blood cells is literally smaller in magnitude, and the low-speed stagnation zone associated with valve geometry and movement is substantially eliminated, resulting in less blood cell damage or platelet activation, less thrombus formation and aggregation, reduced anticoagulant dosage, and easier and safer postoperative management.

[0075] FIGS. 15 through 17 show a blood pump (62), a drive line (67), and a feedthrough (63) according to another embodiment of the present invention. This embodiment highlights anatomical adaptability for easier blood pump placement and drive line externalization.

[0076] As illustrated in FIG. 15, the drive line (67) is connected to the distal shell (625) of the blood pump (62). The blood pump (62) has an elliptical blood sac and stem assembly (650) (including a sac (629) and stems (630, 640)), a pump housing (62h) (having a proximal shell (623) and a distal shell (625)), and a pressure sensing system (628) embedded in the proximal shell (623). The design details of the elliptical blood sac and stem assembly (650), the distal shell adapter (6251), the pressure sensing system (628), and the drive line (67) are substantially the same as or corresponding to the embodiments described above, and module and function descriptions are not repeated herein.

[0077] In this embodiment, the feedthrough (63) is positioned in the distal shell (625) of the pump housing (62h) to connect the driveline (67) to the pump housing (62h). ​​Furthermore, the feedthrough (63) is configured to fit snugly into the body adjacent to the distal shell (625), creating a driveline connection tangentially to the outer surface of the pump. This body-fitting feedthrough design provides a pump housing (62h) design that can adapt to the anatomical space available for blood pump placement. The blood pump (62) can be rotatably connected to the interface adapter (501) and also allows the driveline (67) to be routed in an optimal direction that enables smooth subcutaneous tunneling and skin escape. In this way, anatomical adaptability to the implant placement geometry is favored.

[0078] In this embodiment, the feedthrough (63) is located far from the distal shell (625), while the pressure sensor (6271) (see FIG. 25A) and the sensing chamber (628) are located in the proximal shell (623). Further engineering work must be performed to separate the signal conversion path from the pneumatic communication path and to ensure that the blood pump (62) is sealed and protected from biochemical fluid intrusion that could impair the fidelity of the signal conversion after device implantation.

[0079] FIG. 17 shows that the pump housing (62h) has a surface trench (621) formed on the outer surface of the distal shell (625) and on the overlapping coupling area DA (see FIG. 16) of the proximal shell (623) and the distal shell (625). The surface trench (621) is configured so that electrical wires extend from the outlet of the feedthrough (63) and reach the electrodes (6274) of the pressure sensor (6273) along the trench on the overlapping coupling area (see FIG. 25b). In some embodiments, the trench (621) is sealed by a potting waterproof material and / or a ring-shaped cover to maintain a smooth outer surface so as not to irritate or damage the contacted tissue.

[0080] As illustrated in FIGS. 18a and 18b, the aortic connector (50) generally requires an interface adapter (501) that functions as a coupling mechanism for connecting a blood pump (62) to a target artery (60). The distal end (504) of the connector (50) facing the interface adapter (501) is positioned within the wall of the artery (60) and is also fluidly connected to the human circulatory system. However, the proximal end of the aortic connector (50) (interface adapter (501)) has a seamless interface transition to geometrically match the inlet shape of the blood pump (62). A coupler is typically required to integrate the proximal end of the connector (50) (interface adapter (501)) with the inlet of the blood pump (62). There are some embodiments that can be used as the aortic connector (50). What is illustrated in FIG. 18a is a longitudinal anastomosis of a Dacron or polytetrafluoroethylene (PTFE) graft (502) sutured to a target vessel (60) that can be used in vascular surgery. In some other embodiments, such as that illustrated in FIG. 18b, an insertable aortic connector (503) is used, such as the T-manifold type adapter (503) disclosed in US Patent No. 2008 / 0300447A1, titled “Dual-pulsation bi-Ventricular Assist Device”.

[0081] FIGS. 19 and 20 respectively show the assembly and components of the design of a long-term axisymmetric elliptical blood sac and stem assembly (650). The polymer material selected for these components may be, but is not limited to, segmented polyurethane having various suitable durometers. The components of the sac and stem assembly (650) include a flexible membrane sac (blood sac) (629), a proximal stem (630), and a distal stem (640), all of which are made in axisymmetric shapes and are also integrated together to form a body assembly that rotates about the common centerline (62C) of the blood pump (62). The proximal stem (630) is located at the proximal end (6291) of the sac (629), and the distal stem (640) is located at the distal end (6292) of the sac (629).

[0082] FIG. 19 shows an integral pouch and stem assembly (650) in which the end of the distal stem (640) is wrapped and joined by a reverse membrane (62A) at the distal end of the pouch (629). FIG. 20 shows the components before joining. Generally, the pouch (629) is made by dip molding, while the stems (630, 640) are injection molded. There is no preferred azimuth for the pouch deformation to be deflected. Theoretically, a thin-walled pouch constructed in an axisymmetric elliptical shape will bend into a trilobate configuration (6293) when the pressure difference applied to the membrane exceeds a predetermined threshold, as shown in FIG. 21. This membrane bending is only significant in the final trilobate configuration (6293) (inherent mode), while where the folds (6294), or creases, occur is determined by the initial disturbance that triggers the onset of buckling instability. Thickness uniformity in the cross-sectional planes cut perpendicular to the centerline (or axis of rotation) (62C) of the assembly (650) of the blood pump (62) is important. Care must be taken to maintain high-precision pouch fabrication to ensure an axisymmetric shape. In practical applications, the direction of gravity stands out as a major factor in the initiation of the fold line. The postures of the device recipients are constantly changing, and the direction of gravity relative to the blood pump also constantly changes according to the daily activities the patient undergoes, such as standing, sitting, exercising, and sleeping. Therefore, the folds (6294) of the pouch deformation appear as a random gait-like pattern that non-fixedly distributes high deformation folds throughout the entire pouch. Thus, preventing high deformation areas from remaining in a fixed position is a key design guideline for maintaining the pouch for a long period.

[0083] Embodiments of the present invention innovate continuous fold line properties that cause high deformation locations to appear non-fixed on the membrane to extend the fatigue life of the sac. Thus, the harmful stress concentration phenomenon often associated with bending blood sacs is improved. Based on this fundamental change in bending pattern behavior, the fatigue life of the membrane will be significantly increased due to the non-fixed fold line formation properties that distribute high deformation regions throughout the sac. Furthermore, a beneficial result accompanying this non-fixed sac deformation pattern lies in the enhanced vortex washing effect within the blood sac. The sac surface will be washed more thoroughly by random walk-like vortex formation and traversing. Consequently, the likelihood of a constant low-velocity recirculation zone being created in the peri-wall region or the folds of the fold lines is greatly reduced, resulting in a long-lasting, thromboresistant blood pump design.

[0084] FIGS. 22, FIGS. 23a and FIGS. 23b show an exemplary embodiment of how an integration method is adopted to mount the blood sac assembly to the pump housing (62h) as well as to connect the drive line (67) to the proximal shell (623). The distal drive line (67) has a first end connected to the blood pump (62) at the feedthrough (63), and a second end made as a solid connector (678) having electrodes (6781) that are covered by a bending relief (679) and mounted at the same height.

[0085] The blood sac (629) is secured to a pump housing (62h), which includes a proximal shell (623) and a distal shell (625), to facilitate pump charging and discharging operations. Generally, the bending characteristics of the sac (629) and the housing (62h) are very different. To achieve a long-lasting sac design, an intermediate suspension needs to be installed so that the structural characteristic transition of the pump assembly is continuously maintained, particularly in membrane bending deformation. A pair of movable stems (630 and 640) are adopted as suspensions to integrate the blood sac (629) with the housing (62h). ​​As shown in FIG. 23a, the proximal stem (630), configured in a disc shape, is connected to the proximal shell (623); whereas the distal stem (640), in a ring shape, is connected to the distal shell (625). Mechanically, the proximal and distal stems (630, 640) act as stress-relieving suspensions, which not only keep the blood bag inside the pump housing (62h) but also prevent stress concentration occurring at the interface attachment, thereby extending the lifespan of the bag (629).

[0086] As illustrated at the bottom of FIG. 23a, the distal shell (625) includes an extension, referred to as a distal shell adapter (6251), to be coupled with the aortic adapter (14). This distal shell adapter (6251) has a first end (6252) attached to the pouch (629) and a beak-shaped second end (6253) to be coupled with the interface adapter (501) (illustrated in FIG. 18a and FIG. 18b). The first end (6252) of the adapter (6251) fits seamlessly with the distal end of the pouch (629). However, the opposing second end (6253) is configured to be paired with the interface adapter (501), and the coupling design goal is to minimize interface discontinuity to prevent thrombus formation. The adapter (6251) has a flange structure (6254) that is positioned in the middle area of ​​the distal shell adapter (6251) and acts as a locking element received by the interface adapter (501).

[0087] During surgery, the closed-sac design of the valveless blood pump (62) draws in air, and due to buoyancy, air bubbles will aggregate at the top of the sac. As illustrated in FIGS. 22 and 24, an air removal port (66) is thus installed or created within the proximal shell (623), whereby a narrow channel (661) is provided over the stem (630). An integral sac and stem diaphragm (6301) as an extension from the channel (661) is used to allow an air removal needle to penetrate the blood chamber for air removal. In some embodiments, the channel (661) extends along the centerline (62C). After the blood pump (62) is anastomosed to the target artery (60), the trapped air is pushed out by arterial blood pressure and aggregates on the dome space of the sac (629). A thin needle is used to pierce the air removal port (66), across the channel (661), through the sac and stem septum (6301), and finally reach the inside of the blood sac (629) to expel accumulated air. The integral sac and stem septum (6301) below the air removal port (66) is relatively rigid and does not bend, which will maintain the perforated sac (629) without additional structural defects caused by crack propagation starting from the perforated slit when periodic pulse pressure and the stretching and folding of adjacent sacs are applied.

[0088] As illustrated in FIGS. 23a and 25a, a pressure sensing mechanism (627) is embedded in the proximal shell (623). FIG. 23a shows details of the cross-section of the integral proximal shell (623), the feedthrough (63), and the driveline (67). FIG. 25a shows the profile of the proximal shell (623) connecting the feedthrough (63) extending from the dome of the shell (623) for pneumatic and signal communication with the driveline (67).

[0089] As detailed in FIGS. 25a and 25b, the pressure sensing mechanism (627) includes a first space (6272) for fluid communication and a second space (6273) for accommodating a Micro Electro-Mechanical System (MEMS) pressure transducer and associated electronic circuitry, and has a pressure sensor (6271) that is sealedly housed in a metal canister. There are a plurality of electrodes (6274) protruding from the base of the second space (6273) to be connected to the wires (6702) of the drive line (67) (see FIG. 26). The second space (6273) is closer to the drive line (67) than the first space (6272). A small tube in the first space (6272) is open to fluidly communicate with the sensing medium. A biocompatible fluid or gel is used as the pressure transmission medium. A cavity or pressure sensing chamber (628), located within the proximal shell (623) and adjacent to the first space (6272), is created so that a sensing fluid can be contained therein. This pressure sensing chamber (628) has a distal end separated from the blood chamber by a sac membrane (629). The pressure sensing chamber has two lateral arms: a first arm (6281) and a second arm (6282), wherein the arm (6281) is used to install the pressure sensor (6271) and the arm (6282) is used to fill and seal the sensing medium. Thus, a blood pressure pulse can be transmitted across the sac membrane (629) and hydrodynamically communicate with the MEMS sensor (6271) located away from the second space (6273).

[0090] One embodiment of the present invention innovates a pressure-based blood pump control method and sensor design. A small MEMS pressure sensor having an electronic circuit that is packaged and embedded in the housing wall is adopted. In principle, the MEMS sensor die is highly durable due to its inherent micro-scale structure. In practice, sensor durability depends on the packaging design. The pressure sensing system (627) does not come into contact with blood and is isolated from corrosive biochemical effects associated with blood, thereby providing long-term signal acquisition and transmission required for long-term implantable assistive devices.

[0091] The drive line (67) operates as a communicator for converting electrical signals and transmitting pneumatic pulse pressure between the blood pump (62) and the driver (98). A typical multi-layer drive line (67) in the present invention is illustrated in FIG. 26. In this embodiment, the drive line (67) has a pneumatic lumen (or internal pneumatic tube) (6701), a plurality of wires (6702), an intermediate pneumatic tube (673), a coil (674) (such as a metal coil), an outer layer tube (675), a tether (676), a silicone jacket (677), a rigid driver connector (678), and a protective bending relief (679).

[0092] The central portion of the driveline (67) accommodates a pneumatic lumen (6701) (or air passage, inner tube) having a lumen diameter of about 2-5 mm, depending on the preference between a low-profile design for low energy consumption or ease of operation. Wires (6702) for signal transmission are embedded in the walls of the driveline (67). There are variations of the driveline design that can be adopted. Apart from the multi-layered driveline design shown in FIG. 26, the driveline (67) may be multi-luminal to facilitate the embedding of wires (6702) into three or four smaller lumens and also to allow pulsed air to flow into the larger lumen (6701), for example, as shown in FIG. 27. One of the smaller cavities is installed with a tether (676) to limit the extension of the driveline (67) and protect the wires from damage when an external traction force is applied.

[0093] The inner tube or pneumatic lumen (6701) is accommodated by the intermediate tube (673) and a reinforcing material is inserted between them. Between the inner and intermediate tubes (6701, 673), a coil (674) (or fabric thread or mesh) can be reflowed (thermally co-formed using thermal shrinkage) as a reinforcing material for the driveline wall, making the driveline (67) flexible but resistant to twisting.

[0094] The outer layer tube (675) may be employed to cover the inner and intermediate pneumatic tubes (6701 and 673) and also to cover the spirally wound wires (6702) as a protective covering. In some embodiments, a non-inflatable tether (676) may be placed between the outer tube (675) of the driveline (67) and the silicone jacket (677) to reinforce the tensile elasticity required during the externalization of the driveline (67). Clinically, the silicone jacket (677) has been proven to be the least irritating to subcutaneous tissue and to have the lowest driveline infection rate.

[0095] In this embodiment, a pneumatic lumen (6701), a metal coil (674), an intermediate tube (673), spiral wires (6702), an outer tube (675), a tether (676), and a silicone jacket (677) are packaged in the body of the driveline (67). The proximal end (671) of the driveline (67) must be plugged into an outlet built into the driver (98). The rigid driveline connector (678) of the driveline (67) is configured to be received by the proximal driveline interconnector (93) or the outlet of the driver (98). A plurality of electrodes (6781) (e.g., four electrodes (6781) in FIG. 23b) are mounted at the same height on the rigid driver connector (678) and soldered to the wires (6702). To prevent the drive line (67) from twisting at the joint, a protective bending relief (679) (see FIG. 23b and FIG. 26) is placed over the joint portion of the drive line (67) and the driver connector (678). The proximal end of the drive line, including the driver connector (678) and the bending relief (679), is maintained in a low-profile so as to facilitate escaping the skin without creating undesirable tunneling trauma.

[0096] Connecting the drive line (67) to the blood pump (62) is achieved through a feedthrough (63), as illustrated in FIGS. 22 and FIGS. 23a. The feedthrough (63) may be placed in the proximal shell (623) or the distal shell (625), depending on the anatomical structure in which the blood pump (62) is to be implanted. Integrating the feedthrough (63) with the pump housing (62h) can change the overall external blood pump configuration and also guide the drive line (67) in a specific direction to satisfy implantation requirements, including drive line externalization path, postoperative skin care, and device usability.

[0097] As illustrated in FIGS. 23a and 25a, the feedthrough (63) has a first part (631) as an extension of the proximal shell (623) to which the wires (6702) of the pneumatic lumen (6701), tether (676), and driveline (67) are joined via an anchor adapter (672), where the anchor adapter (672) connects to the distal driveline (67) which allows driving air to pass through the pneumatic lumen (6701). The feedthrough (63) further has a second part (632) that is coupled with the first part (631) to act as a bending relief for the driveline (67). The first part (631) is the location where wire connection, tether fixation, and the joining and sealing of the housing and the pneumatic lumen occur. The wires must not be exposed to the tissue at the implantation site and must also be well protected from tensile forces applied during driveline externalization. Additionally, the connection between the pneumatic lumen (6701) and the blood pump (62) must be pneumatically and electrically leak-free. As illustrated in FIG. 25a, these mentioned blood pump integration operations are performed in the first part (631). However, the second part (632) is employed to accommodate and seal these interface joint elements and acts as an external protector to protect the joint from mechanical deformation and the intrusion of environmental fluid or moisture.

[0098] A modular design related to a first embodiment of the present blood pump invention is disclosed in FIGS. 19 through 27. In this illustrated embodiment, the blood pump (62) comprises an axisymmetric elliptical blood sac and stem assembly (650) (including a flexible membrane sac (629), a proximal stem (630), and a distal stem (640); a pump housing (62h) having a proximal shell (623) and a distal shell (625); and a pressure sensing system (627) embedded in the proximal shell (623). The drive line (67) comprises a pneumatic lumen (6701) and wires (6702) contained in the wall thereof, and is connected to the blood pump (62). To integrate the drive line (67) with the pump housing (62h), a feedthrough (63) is used to achieve both electrical and pneumatic communication between the drive line (67) and the blood pump (62).

[0099] As illustrated in FIGS. 19 and 20, a polymer fabrication and bonding method for integrating the blood sac (629) and stems (630, 640) together was disclosed in the previous section. The key to the design and fabrication is to maintain high-precision axial symmetry in the fabrication and bonding of the sac and stem assembly parts. A pressure sensing system (627) and a feedthrough (63) are installed in the rigid portion of the proximal shell (623). As illustrated in FIGS. 22, 23a, and 25a, a compact feedthrough (63) design is shown. It can be observed that the electrical wiring and connections are more robust and resistant to defects due to the compact feedthrough (63).

[0100] FIGS. 28a and 28b depict some flow patterns associated with paraaortic counterpulsatile. During the end-diastolic and early-systolic phases of left ventricular ejection, the blood pump draws pump charge and aortic flow into the pump (Fig. 28a). Both upstream and downstream blood around the connector will be drawn into the blood pump by creating a sharp 90-degree flow turn. This will create flow separation and low-velocity recirculation zones (T-201). Additionally, abnormally high shear forces will appear in the corner regions of the T-junction. Meanwhile, during diastole following aortic valve closure, the blood stored in the pump is discharged back into the circulatory system, creating a flow that collides with the opposite aortic wall (Fig. 28b). This lateral dump, colliding flow entails very high local pressure at the point of impact (T-202), which is a stagnation point where the so-called flow velocity is literally zero and all kinetic energy associated with the flow velocity is converted into potential energy referred to as total pressure. These high-pressure, impinging flows can induce vascular maladaptation, including smooth muscle cell proliferation and consequent wall stenosis, as well as a risk of aortic dissection due to persistent local hypertension. All these non-physiological flow patterns and induced high pressure, high shear, and low-velocity recirculation phenomena are predominant near the T-junction region. This complex turbulent flow anomaly will collapse or diminish at a distance of 3 to 5 times the lumen diameter of the implanted artery. Currently, implantable aortic adapters are designed with an inserted catheter length of 5 to 7 cm, covering most of the pump-induced non-physiological flow region. Since the aorta at the implantation site is shielded by the inserted aortic adapter, the biological vascular wall can be isolated from the effects of pathological stress conditions induced by the pump, thereby protecting the aorta at the implantation site from acute or long-term remodeling complications.

[0101] In counter-pulse support, pump charging and discharging are operated alternately in synchronization with the heartbeat, which creates a special T-junction flow as shown in FIGS. 28a and 28b. This insertable aortic adapter (14) is shown in more detail in FIG. 29, which is a perspective view, and FIG. 30, which is a cross-sectional view.

[0102] The aortic adapter (14) is injection molded so that its internal blood contact surface (141) is manufactured to be very smooth and continuous without dividing lines. Silicone or other polymer elastomers may be used as materials. The aortic adapter (14) comprises a conduit (or a portion of the conduit to be inserted) (142) to be inserted into the aorta (95) (see FIG. 5) and a neck (or a protruding neck portion) (143) connected to a blood pump. In this embodiment, the neck portion (143) has a neck body (1431) and an extension (1432) disposed on the neck body (1431), wherein the extension (1432) protrudes from the neck body (1431) and the maximum inner diameter of the extension (1432) is greater than the maximum inner diameter of the neck body (1431). When the neck portion (142) is connected to the blood pump (62), the extension portion (1432) comfortably wraps around the inlet adapter (6251) of the blood bump (62), and the neck body (1431) is positioned around the coupler (25) that integrates the aortic adapter (6251) into the blood pump (62).

[0103] The entire aortic adapter (14) is thin-walled to maximize flow efficiency. To reinforce the thin-walled structure, a pair of nitinol trusses (or truss rings) (144) are embedded around both ends of the conduit portion (142) of the aortic adapter (14).

[0104] FIG. 29 is a transparency showing the locations where the nitinol truss (144) is embedded. Furthermore, the wall thickness of the conduit (142) gradually thins toward the conduit end (145). The function of the gradually thinning wall thickness is twofold. First, it minimizes graft / host junction discontinuity and ensures that the rate of thrombus formation at the interface is always lower than the thrombolytic rate provided by the contacted endothelium. Second, the compliance of the conduit becomes smoother toward the conduit end (145), resulting in a compliance-matching effect when coupled with the aortic lumen.

[0105] One of the complications plaguing large stent graft implantation is the problem of intimal leakage. Type-I intimal leakage means that the seal of the graft end against the endothelial lumen of the grafted artery is incomplete, causing a gap to form between the graft leading edge and the arterial lumen. Leaked blood becomes trapped in the gaps and hardens into a thrombus, eventually forming a fibrous pseudo-intimal that will grow uncontrollably over time. The pseudo-intimal not only blocks the grafted artery but can also signal and stimulate the coagulation mechanism to induce platelet adhesion, leading to thrombotic side effects. A solution to address this intimal leakage problem is to tightly seal the aortic adapter (14) against the attached lumen surface. The aortic adapter (14) presents a compliance-matching design concept that enables the semi-rigid conduit ends (145) to be seamlessly attached to the arterial lumen when pulsatile blood pressure is applied. The outer diameter (146) of the aortic adapter conduit (142) is slightly larger than the lumen diameter by an oversize ratio in the range of 3-10% (defined as a percentage increase of the conduit diameter (146) relative to the lumen diameter) conditioned at a given nominal blood pressure (e.g., 120 mmHg). Since blood pressure fluctuates between systole and diastolic, or under pulse pressure generated by counterpulse support, the compliance-matching conduit ends (145) will dynamically expand and contract in response to pressure beats without creating interface gaps.

[0106] Generally speaking, thin-walled tubes made of elastomer tend to be flexible and compliance-matching, but often the device size is too large and they are not strong enough to withstand applied compressive forces, causing wall buckling of the inserted adapter and consequent massive bleeding. Therefore, the combined use of a nitinol truss structure (144) and an elastomer substrate having appropriate hardness is important, as illustrated in FIG. 30. In this design, the radial strength provided by the nitinol truss (144) will help support the aortic adapter (14) without buckling, and there is a distance “X” between the outermost boundary (1441) of the truss (144) and the end of the conduit (145). In some embodiments, this distance (X) must be evaluated and correctly defined. When the nitinol truss (144) is supported by an expandable frame, the gradually thinning conduit end (145) will not collapse or wrinkle, thereby maintaining a circular shape against the attached lumen wall and dynamically sealing the lumen. Note that the aortic adapter (14) can expand and contract in response to pressure pulsations, and the sealing effect is achieved in a dynamic manner in which the aortic adapter (14) and the aortic (95) wall expand and contract together as a whole to seal the conduit end (145) without causing hemorrhage complications.

[0107] A representative embodiment of a nitinol truss (144) made of laser-engraved nitinol tubes, typically expanded through a series of expansion and heat treatments, is illustrated in FIG. 31. FIG. 31 shows a two-dimensional enlarged view of the truss ring (144). Each truss has a plurality of wavy structures (1441). The truss (144) is self-expandable, which can be folded or creased into a smaller pre-packed transfer configuration and also expands to recover its original configuration when unwound after being placed in a desired position.

[0108] A convenient measurement of conduit stiffness (the inverse of compliance) can be expressed as the so-called lateral stiffness (LS) shown in FIG. 32. LS is defined as the value obtained by dividing the applied force (F) per unit length by the corresponding radial deflection (Y). For the current aortic adapter, the appropriate LS range is 0.01-0.05 Nt / mm2. Both the embedded nitinol truss (144) and the silicone or elastomer substrate will contribute to the structural compliance of the co-injected aortic adapter (14). It is optimal to have evenly distributed compliance so that the tendency for interlayer delamination due to the elongation and contraction of the composite conduit walls is minimized, thereby increasing the fatigue life of the adapter (14).

[0109] The aortic adapter (14) is configured to be connected to a blood pump (62) to provide circulatory assistance. Here, a quick connector type coupler is invented. An exploded view of the components of a coupler (25) that integrates the aortic adapter (14) and the blood pump (62) together is shown in FIG. 33. The coupler (25) includes a flange base (252), a pair of collars (253), and hinges (or hinge assemblies) (254) that connect the collars (253) together with the flange base (252). Spring coils (or spring coil assemblies) (255) are loaded into the hinge joint (256), and when unlocked, the collars (253) are held in an open position (Fig. 34a). The locking mechanism is a latch (257) made of a slotted leaf spring and secured by a slab (2571) welded to one of the ends of the collars (253). The flange base (252) has a substantially circular structure, and each collar (253) has an arc-shaped structure. A hinge joint (256) is located on the first side (252S1) of the flange base (252), and the collars (253) are pivotally connected 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) belong to parts of the aortic adapter assembly.

[0110] FIG. 34b illustrates a drawing of a coupler (25) in a locking configuration, where a leaf spring latch (257) drops down from a ramp (258) to ensure the coupling is secure without fear of disengagement. The collars (253) are grooved internally as shown in FIG. 35. The integration of the aortic adapter (14) and the blood pump (62) is achieved through a clamping mechanism using a deformable adapter proximal end (147) (Fig. 30) and a blood pump inlet adapter (6251), respectively, which function as a "gasket" between the connected rigid flanges (252 and 81) of the coupler (25) (see FIG. 35).

[0111] In particular, the quick-connect lock can be easily achieved by closing the collars (253) that will be engaged without fear of unintended unlocking, as illustrated in FIGS. 34b and 35. A leaf spring latch (257) is installed at the tip of one collar (253). During collar closure, this latch (257) will bend as it slides over the ramp (258) of the opposing collar (253) during the locking process. When the latch (257) passes over the top of the ramp (258), it will fall to the base of the ramp (258) by means of an elastic restoring force and acts as a safety device to prevent incidental latch unlocking or collar opening that contributes to pump vibration or shaking during long-term use. In pump transplantation or replacement requiring module separation, the latch (257) is bent and lifted upward by a tool to allow unlocking force to be applied to rotate the collars (253) thereby separating the blood pump (62) from the aortic adapter (14).

[0112] Butt joint designs are not feasible for connecting two smooth-surface tube adapters in the bloodstream. In most clinical applications, the connected graft has a rough surface to promote endothelialization, so that small interface discontinuities in the bloodstream are "smoothed out" by endothelial cells and proteins. The aortic adapter (14) adopts a smooth surface approach to prevent thrombotic side effects. The blood flow of the aortic adapter is bidirectional in response to the discharge and charge operations of counterpulsive pumping, as illustrated in FIGS. 28a and 28b. This strong bidirectional flow and surface cleaning effect will easily remove newly formed neon thrombi on the rough surface. Therefore, a smooth surface design is considered more suitable and safe for the present invention. At the interface of the two combined smooth surfaces in the bloodstream, careful mechanical and hemodynamic design is required to prevent thrombotic phenomena occurring in situ. The basis and design methods associated with this novel joint invention are disclosed below.

[0113] FIGS. 36a and FIGS. 36b respectively show two fundamental interface discontinuities present in a butt joint connection, such as steps (101, 102) or gaps (103) created between an adapter (AB1) (such as the adapter (6251) of the blood bump (62)) and an adapter (AB2) (such as the aortic adapter (14)). The discontinuities shown in FIGS. 36a and FIGS. 36b are exaggerated, and typically in precision machining, these joint discontinuities are within 10-50 microns, which is large enough to cause coagulation and thrombus formation.

[0114] In reality, even if the machining of each body is performed perfectly, there is inevitably a tolerance in aligning two individual bodies. FIG. 36a shows a misaligned joint of two bodies, and everything related to the fabrication of the part, except for the centerline misalignment, was performed correctly. Steps (101, 102) facing forward and backward are created, and stagnant flow in the stepped regions (101, 102) becomes a source for the formation of thrombi or thromboembolisms. In FIG. 36b, an interface gap (103) is created due to the non-parallel matching of the connected bodies. The gap (103) attracts and accumulates blood cells and causes them to grow into a pseudo-intimal, and this intimal growth often becomes uncontrollable, blocking the entire blood flow pathway in addition to thromboembolisms flowing from the intimal surface. Interface errors associated with butt joints can be exacerbated when the connected bodies are not rigid. The aortic adapter (14) is semi-rigid so that it can be forced into a butt joint connection having modified configurations and expanded interface discontinuities. Therefore, to achieve a connection between the semi-rigid aortic adapter and a blood pump, a new connection method must be invented as disclosed below.

[0115] As illustrated in FIGS. 37 and 38, in some embodiments, the blood pump (62) has an inlet adapter (80) comprising a flange (81), a beak (82), and a base (83) that form an extension of the housing of the blood pump (62). A plurality of eyelets (86) are mounted to connect the inlet adapter (80) to the blood pump (62).

[0116] The inner diameter (84) of the beak (82) is slightly larger than the inner diameter (148) of the neck (143) of the aortic adapter (14) (see FIG. 30). The contact area between the beak (82) and the proximal end of the adapter (147) is an annular conical surface (or a shallow inclined surface or ramp) (149) (see FIG. 30 and FIG. 35). The conical angle of the surface (149) 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 collars (253) having an inward groove (2531) (Fig. 34a) will loosely grip the flange of the base (252) and the beak flange (81). As the collar (253) closes, the flanges (81, 252) (of the beak (82) and coupler (25)) are received and compressed by the collar grooves (2531), thereby compressing the sandwiched silicone ends (147) and generating the clamping force required for a tight connection.

[0117] The clamping force generation mechanism is graphed in FIG. 35. There are two steps (2521 and 2522) responsible for generating the clamping force. Before collar closure begins, the step (2521) must engage with the slot (1433) of the aortic adapter neck (143) (Fig. 30). This engagement is achieved by first folding the neck (143) and then inserting the deformed neck (143) across the adapter flange (252), thereby allowing the elastic restoring force of the aortic adapter (14) to return the folded neck (143) to its original circular shape and the step (2521) to engage with the slot (1433). The groove height (Z) of the collar (143) controls the compression deformation of the aortic adapter end (147). Referring to FIG. 35, it can be seen that in a fully locked configuration when the collar (253) is closed and engaged, the gap (Z0) is smaller than the aortic adapter end width (Z3) (Fig. 30). The assumption of the assembled geometry shows that Z0 = Z - Z1 - Z2, where Z1 and Z2 are the flange thicknesses of the fixed counterparts as shown in FIG. 35. Generally, Z3 is larger than Z0, and thus the tensioned aortic adapter end (147) will generate the clamping force required for a sealing connection of the beak flange (81) together with the collar flange base (252). The deformation of the silicon flow adapter end (147), defined as (Z3-Z0) / Z3, is sufficient to ensure a reliable sealing connection in the range of 10-30%.

[0118] FIG. 39 illustrates the coupling characteristics of the current connection between the adapter conical surface (149) and the beak (82). When connected, the beak leading edge (85) and the conical surface (82) will sink into the semi-rigid ramp surface (149) of the aortic adapter (14), with a depth comparable to the leading-edge radius of the beak leading edge (85), typically 30-50 microns. In FIG. 39, the dashed line and numbers in parentheses indicate the beak at initial contact, and the solid line and numbers without parentheses indicate the locked relationship. Note that the interface discontinuity is reduced by the sinking of the conical surface from its original form (dashed line). The width of the collar groove (2531) controls the interference fit of the connection. As previously described, the elastic aortic adapter end (147) will be compressed to approximately 10-30% deformation to provide the coupling force necessary for leak-free integration for pulsatile pumping.

[0119] The current design of the interface connection between the blood pump (62) and the aortic adapter (14) has two hemodynamic advantages for reducing thrombus formation in the field. First, joint discontinuities in the form of steps or gaps, which are created as observed in conventional butt joint connections, will not literally occur. Second, stagnant flow located at the interface of the beak leading edge (85) can be minimized. Thus, blood flow through the connected interface is maintained at high speed, substantially improving the disadvantages of butt connections, namely steps (101, 102) or gaps (103) created at the interface facing forward or backward.

[0120] The current conical surface (149) is inclined with an angle of inclination relative to the flow direction. This ramp interface design prevents step or gap from occurring at the joint due to limited manufacturing precision or matching eccentricity associated with conventional butt connections. Nevertheless, this shallow conical ramp (149) has an inherent disadvantage in satisfying the concentric centerline alignment of the coupled counterparts. The current coupling of the inlet beak (82) and the aortic adapter (14) does not have strict lateral restraints to ensure coupling alignment. To concentrically connect the rigid beak (82) with the semi-rigid aortic adapter flange ramp (149), it is important to simultaneously grip the collars around the entire periphery of the flange base (252). When simultaneous gripping / locking engagement is not achieved, the adapter flange ramp (149) engaged first deforms more than the other free parts, creating a tendency to tilt or displace the remaining contact surfaces leading to the eccentric pump connection. Such eccentric connections are often a causative factor in creating a step or gap at the interface that induces thrombus formation. This disadvantage is resolved by having a flange contour (259) (Fig. 34a) of the distal flange of the collars (253) configured such that the locking engagement simultaneously includes all circumferential contact areas. When locked, the contacted edge of the metal beak (82) sinks slightly into the silicone ramp (149) compressed to a controlled depth and further reduces interface discontinuity when exposed to blood flow. Thus, conventional interface thrombi can be substantially minimized or eliminated with the administration of an appropriate anticoagulant.

[0121] The structural deformability and delivery method associated with the aortic adapter impart special design features to the present invention. In practice, material elasticity considerations must be carefully incorporated into the current design. Surgically delivering an implantable graft to the aorta through an incised aortic wall is challenging in terms of pre- and post-operative safety and long-term reliability. The material selected for the aortic adapter (14) must have a pre-memorized shape. During device delivery, the adapter (14) is first folded into a smaller delivery configuration, and this delivery configuration ensures rapid and safe device implantation. After the folded aortic adapter (14) is placed at the intended implantation site, the delivery configuration must be unwound to self-expand into its original memorized shape.

[0122] Before inserting the aortic adapter, a hole with a diameter of 12-14 mm must be made in the aortic wall. When making this access hole, care must be taken not to create cutting edges that could become the starting point of a crack when wall expansion is required for device insertion. 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. A solid hole with crack-free edges can be successfully made with a single punch.

[0123] The aortic adapter (14) structurally comprises two circular tubes joined together by a T-shaped flow conveyor to enhance periaortic circulation support. The tube walls are typically 1-2 mm thick, and the material used is a polymer such as silicone or polyurethane with appropriate hardness, for example, Shore A 80-90. The corrugated conveyor configuration differs substantially from commercially available large stent grafts covered with Dacron or PTFE (Polytetrafluoroethylene) fabric. FIG. 40 shows the corrugated form of the aortic adapter (14) (the nitinol truss (144) is embedded and deformed together with the polymer substrate). The aortic adapter (14) is folded by pushing in the inserted tube portion (142), and accordingly, the T-neck (extruded neck portion) (143) is compressed and flattened so that it sinks into the folded adapter (14) as shown in FIG. 40. Approximately, the folded configuration has a diameter that is about half the diameter of the original unfolded circular form.

[0124] This folded adapter can be secured by tightening straps. For example, as illustrated in FIG. 40. Three securing straps may be placed around the two edges and the center of the conduit, and many other securing methods may also be considered. FIGs. 41a, 41b, 41c, and 41d show four representative steps of the delivery of the aortic adapter (14). The first step (see FIG. 41a) shows the initial penetration of the corrugated prepack across the access opening while the prepack is tilted at an angle with respect to the aortic (95) axis. The second step (Fig. 41b) shows the complete insertion of the delivery configuration into the aorta, where the prepack is pushed through the access opening as its proximal end is rotated and drops directly into the access opening. The fully inserted prepack is retracted so that its corrugated neck (143) aligns with the perforated access opening (Fig. 41c). The binding straps are released, and the pre-packed delivery configuration is restored to its original unfolded configuration by elastic self-expansion (Fig. 41d). The released aortic adapter (14) is comfortably and snugly wrapped around the aortic lumen, ensured by an appropriate oversize ratio selected prior to delivery. By slightly modifying the T-neck (143), the flange (252) of the coupler (25) (Figs. 33 to 34b) can be mounted on the T-neck (143) to which the blood pump (62) is ready to be connected. Thus, the blood pump connection can be easily achieved by seating the inlet beak (82) on the conical surface (149) of the T-neck and then closing the two collars (253) of the coupler (25) for rigid device integration.

[0125] Additional safety measures may be applied to improve hemostasis and stability of the implanted periaortic blood pump system. Periaortic placement of the blood pump inevitably entails lateral forces (perpendicular to the longitudinal direction of the aorta) and torque applied to the aortic adapter (14) due to the weight of the blood pump (62) and the pumping force generated by counterpulsive support. These device-related external forces may affect the long-term remodeling of the vascular structure at the implantation site. A purse-string suture may be placed on the outer membrane layer around the access opening. The purse-string suture may further tighten the aortic wall against the inserted adapter (14) and act as a protective measure to prevent enlargement of the access opening. Furthermore, surgical tapes may be wrapped around both ends of the conduit (142) to tighten and reinforce the integration of the inserted aortic adapter (14) with the aorta as a whole. Compliance-matching design and banding of the wound tapes can double the guarantee of no intimal leakage. Occasionally, blood pressure can rise above the upper limit that no intimal leakage can be guaranteed by compliance-matching. Under these extreme conditions, the surgical tapes act as a hard limiter to seal the released adapter ends and ensure that hemostasis is maintained.

[0126] A step-by-step demonstration of the aortic adapter (14) implantation is detailed in FIG. 42. A pleated prepack form is prepared before the implantation begins. After a left thoracotomy to expose the target thoracic artery, a cross-clamping distance of approximately 10 cm is defined across the implantation site. An access hole to be punched into the aorta is first marked around the identified hole. A purse-string suture is then sewn outside the hole in the epicardial layer. The aorta may be partially incised from the surrounding connective tissue, and a pair of surgical tapes may be wrapped around the aorta in a loop. After the above preparation work is completed, cross-clamping and aortic adapter insertion are performed. These insertion steps are described in the order shown in FIG. 42. First, the aorta is cross-corrected to provide a separated section without the risk of bleeding. Then, a custom aortic punch is used to create a large access hole for the aortic adapter insertion. Subsequently, the folded adapter prepack is inserted and positioned into the cross-clamped aortic portion as illustrated in FIG. 41a, 41b, and 41c, and then the folded adapter is released and restored to its original unfolded form (Fig. 41d). Sutures and tapes are then tightened as additional protective devices to prevent internal leakage in case of hypertension. The coupler is then installed with an adapter flange (2521) that sits in the slot (1433) of the aortic adapter neck, and is prepared to receive the blood pump (22) to be connected. The self-aligning feature of the coupler design allows the beak (80) of the blood pump (62) to be properly positioned and locked together with the aortic adapter (14). The remaining implantation steps, including cross-clamp release, blood pump de-airing, and pump-assist initiation, are conventional. Generally, for a skilled surgeon, the cross-clamping time required for aortic adapter insertion is approximately 10 minutes. During this cross-clamping time, blood perfusion to abdominal organs is blocked, and ischemic damage may potentially occur.To mitigate this potential surgical damage to the organs, partial femoral-femoral extracorporeal membrane oxygenation (ECMO) support may be administered to perfuse the abdominal organs and lower extremities. However, whether or not to use ECMO support is at the surgeon's discretion. Typically, a healthy patient can withstand 20 minutes of ischemia.

[0127] In summary, embodiments of the present invention provide a ventricular assist device comprising a blood pump, a drive line, and a feedthrough. The blood pump comprises an axisymmetric elliptical blood sac and stem assembly comprising a flexible membrane sac, a proximal stem, and a distal stem, wherein the flexible membrane sac is attached to the proximal stem and the distal stem as a stress-relief suspension mechanism; a pump housing comprising a proximal shell and a distal shell, wherein the stress-relief suspension mechanism is coupled to the pump housing; and a pressure sensing system embedded in the proximal shell, wherein the pressure sensing system comprises a pressure sensor and a pressure sensing chamber filled with an incompressible fluid for pressure transmission. The drive line comprises a pneumatic lumen, at least one wire, and a tether, wherein the wires and the tether are positioned on the drive line wall. The feedthrough connects the drive line to the pump housing.

[0128] One embodiment of the present invention for a displacement pump discloses a pulsatile blood pump design that incorporates a non-fixed folding line concept in a long-term blood sac configuration capable of substantially extending the durability of the displacement blood pump. Additionally, a compact pressure sensing system is disclosed that can be used to function as a reference waveform for real-time pump control, as well as for long-term trend analysis, disease monitoring, and diagnosis, based on evidence-based megadata. Furthermore, the embedded pressure sensing system does not come into contact with blood, thereby significantly improving reliability requirements when constructing an implantable sensor system.

[0129] Embodiments of the present blood pump invention have at least one of the following advantages or effects. By feed-through connection of the drive line to the pump housing, a compact feed-through design is provided, making the electrical wiring and signal conversion more robust and resistant to defects. Furthermore, the compact feed-through design integrates sensor wires and pneumatic tubing with the blood pump. This compact property is particularly essential for implantable devices. This not only simplifies surgery and mitigates pre- and post-operative implant risks but also contributes to a reduction in postoperative morbidity associated with drive line infection.

[0130] In some blood pump embodiments, the feedthrough is integrated with the distal shell of the pump housing and, for the advantage of anatomical adaptability and fit to the geometry of the implantation site, the feedthrough has a first part as an extension of the distal shell to which the pneumatic lumen, tether, and driveline wires are coupled, and a second part coupled to the first part that acts as a bending relief for the driveline.

[0131] One embodiment of the present flow communication invention provides an aortic adapter assembly for an implantable ventricular assist device, which comprises a T-shaped aortic adapter including an inserted conduit portion and a protruding neck portion, wherein the inserted conduit portion is coupled to the protruding neck portion and both have a smooth surface in contact with blood; and a truss disposed on the inserted conduit portion; the T-shaped aortic adapter has a polymer elastomer reinforced by a truss having a nitinol material; the inserted conduit portion has walls that gradually thin at two conduit ends of the inserted conduit portion, having an appropriate distance from the tip of the conduit end to the outermost boundary of the truss, and the conduit end has a compliance-matching effect with respect to the aorta at the implantation site; wherein the proximal end of the protruding neck portion is configured to be coupled to an inlet adapter of a blood pump.

[0132] Embodiments of the present flow communication invention have at least one of the following advantages or effects. The present invention discloses a periaortic ventricular assist device (10), in particular a counterpulsatile blood pump and a flow carrier assembly that enables blood flow delivery therefrom. Unlike many existing flow carriers that employ a rough surface approach to promote endothelialization to prevent thrombotic side effects, the present aortic adapter invention adopts a smooth surface, implantable artificial graft concept to construct the flow carrier. Furthermore, a compliance-matching design is implemented around the implanted conduit ends, which combines progressively thinning wall characteristics with a superelastic nitinol-supported thin-wall polymer to achieve intima-leak-free requirements. The artificial surface of the implanted conduit includes abnormal high pressure, high shear, and low-speed recirculation flow phenomena associated with periaortic counterpulsatile pumping. Therefore, pathologically device-induced hemodynamic effects and risk factors are substantially eliminated, and long-term vascular maladaptation-related adverse effects such as endothelial cell erosion, lipid infiltration, smooth muscle cell proliferation, vascular stenosis, and arterial wall dissection are significantly reduced. To securely connect the semi-rigid flow adapter to a blood pump, a quick-connector type coupler has been invented. This coupler features a self-aligning interface design that minimizes step and gap discontinuities, thereby reducing the likelihood of thrombotic adverse effects at the interface joint. This aortic adapter invention involves a specially designed delivery method that ensures a fast and safe delivery procedure. The corrugated aortic adapter is manufactured as a pre-packed delivery configuration in which its overall size is reduced to half of the unfolded configuration. This pre-packed adapter can be easily inserted into the aorta at the implantation site and self-inflates into its original unfolded configuration, resulting in a comfortably snug-fit flow delivery device without concerns regarding intimal leakage.This is not only beneficial for surgeries that mitigate pre- and post-operative graft risks, but also contributes to reducing postoperative morbidity associated with device-induced flow and vascular maladaptation at the graft site.

[0133] The use of ordinal terms such as "first," "second," "third," etc., in claims to modify claim elements does not, in itself, imply priority, precedence, or order of one claim element over another claim element, or the temporal order in which the operations of the method are performed; rather, it distinguishes claim elements by using them merely as labels to distinguish one claim element with a specific name from another element with the same name (provided that ordinal terms are used).

[0134] Embodiments of the present invention have been described, and variations of these embodiments may become apparent to those skilled in the art by reading the foregoing description. Accordingly, the present invention includes all modifications and equivalents of the subject matter cited in the claims.

Claims

Claim 1 In a ventricular assist device, a blood pump, said blood pump comprises: an axisymmetric elliptical blood sac and stem assembly comprising a flexible membrane sac, a proximal stem, and a distal stem, wherein the flexible membrane sac is attached to the proximal stem and the distal stem as a stress-relief suspension mechanism; a pump housing comprising a proximal shell and a distal shell, wherein the stress-relief suspension mechanism is coupled to the pump housing; and a pressure sensing system embedded in the proximal shell, said pressure sensing system comprising a pressure sensor and a pressure sensing chamber filled with an incompressible fluid for pressure transmission; a drive line comprising a pneumatic lumen contained within the wall of the drive line, at least one wire and a tether, said wire and the tether are disposed within the wall of the drive line; and a feedthrough connecting the drive line to the pump housing; The above flexible membrane sac has an inverse membrane located at the distal end of the flexible membrane sac, the distal stem is wrapped and coupled with the inverse membrane; and the proximal stem is located at the proximal end of the flexible membrane sac, ventricular assist device. Claim 2 In claim 1, a ventricular assist device having an air removal port installed within the proximal shell, Claim 3 A ventricular assist device according to claim 2, wherein the channel communicates with the air removal port, and the channel extends along the centerline of an axisymmetric elliptical blood sac and stem assembly and is positioned over the septum of the integrated membrane sac and stem. Claim 4 delete Claim 5 In a ventricular assist device, the blood pump comprises: an axisymmetric elliptical blood sac and stem assembly comprising a flexible membrane sac, a proximal stem, and a distal stem, wherein the flexible membrane sac is attached to the proximal stem and the distal stem as a stress-relief suspension mechanism; a pump housing comprising a proximal shell and a distal shell, wherein the stress-relief suspension mechanism is coupled to the pump housing; and a pressure sensing system embedded in the proximal shell, wherein the pressure sensing system comprises a pressure sensor and a pressure sensing chamber filled with an incompressible fluid for pressure transmission; a drive line comprising a pneumatic lumen contained within the wall of the drive line, at least one wire and a tether, wherein the wires and the tether are disposed within the wall of the drive line; and a feedthrough connecting the drive line to the pump housing; A ventricular assist device further comprising an arterial connector having an interface adapter, wherein the distal shell has a distal shell adapter to facilitate connection of a blood pump to a human artery; wherein the distal shell adapter has a first end and a second end, the first end interfaces with the inlet of the flexible membrane sac and the second end interfaces with the interface adapter to connect to the human artery. Claim 6 In claim 5, the distal shell adapter has a beak, and the beak has a flange structure that is coupled to the interface adapter, a ventricular assist device. Claim 7 A ventricular assist device according to claim 1 or 5, further comprising a driver, wherein the drive line connects the blood pump to the driver. Claim 8 A ventricular assist device comprising: a blood pump, wherein the blood pump comprises an axisymmetric elliptical blood sac and stem assembly comprising a flexible membrane sac, a proximal stem, and a distal stem, wherein the flexible membrane sac is attached to the proximal stem and the distal stem as a stress-relief suspension mechanism; a pump housing comprising a proximal shell and a distal shell, wherein the stress-relief suspension mechanism is coupled to the pump housing; and a pressure sensing system embedded in the proximal shell, wherein the pressure sensing system comprises a pressure sensor and a pressure sensing chamber filled with an incompressible fluid for pressure transmission; a drive line comprising a pneumatic lumen contained within the wall of the drive line, at least one wire and a tether, wherein the wires and the tether are disposed within the wall of the drive line; and a feedthrough connecting the drive line to the pump housing; wherein the pressure sensor is sealedly contained within a metal canister and also comprises a first space for fluid communication. Claim 9 In claim 8, the pressure sensor comprises a second space closer to the drive line than the first space, and the second space is configured to accommodate a MEMS (micro electro-mechanical system) pressure transducer and an electronic circuit, a ventricular assist device. Claim 10 In claim 9, the pressure sensing chamber is located within the proximal shell and adjacent to the first space, and the pressure sensing chamber is configured to contain a sensing fluid therein, a ventricular assist device. Claim 11 In claim 10, the pressure sensing chamber has a first arm and a second arm, the first arm is used for installing a pressure sensor, and the second arm is used for filling and sealing a sensing medium, a ventricular assist device. Claim 12 A ventricular assist device according to any one of claims 1, 5, and 8, wherein the feedthrough is integrated into a proximal shell; the feedthrough has a first portion as an extension of the proximal shell to which the wires of a pneumatic lumen, tether, and driveline are joined; and the feedthrough further has a second portion coupled to the first portion to act as a bending relief of the driveline. Claim 13 In a ventricular assist device, a blood pump, said blood pump comprises: an axisymmetric elliptical blood sac and stem assembly comprising a flexible membrane sac, a proximal stem, and a distal stem, wherein the flexible membrane sac is attached to the proximal stem and the distal stem as a stress-relief suspension mechanism; a pump housing comprising a proximal shell and a distal shell, wherein the stress-relief suspension mechanism is coupled to the pump housing; and a pressure sensing system embedded in the proximal shell, said pressure sensing system comprising a pressure sensor and a pressure sensing chamber filled with an incompressible fluid for pressure transmission; a drive line comprising a pneumatic lumen contained within the wall of the drive line, at least one wire and a tether, said wire and the tether are disposed within the wall of the drive line; and a feedthrough connecting the drive line to the pump housing; A ventricular assist device, wherein the drive line further comprises an intermediate pneumatic tube and a coil, the pneumatic lumen is received in the intermediate pneumatic tube, and the coil is positioned between the pneumatic lumen and the intermediate pneumatic tube. Claim 14 In claim 13, the drive line further comprises an outer layer tube, the pneumatic lumen and the intermediate pneumatic tube are accommodated within the outer layer tube, and the wires are covered by the outer layer tube, a ventricular assist device. Claim 15 In claim 14, the tether is a ventricular assist device disposed on the outer layer tube. Claim 16 In claim 15, the drive line further comprises a rigid driver connector located at the proximal end of the drive line, and the ventricular assist device further comprises a driver, the drive line is connected to the driver, and the rigid driver connector is connected to the driver, ventricular assist device. Claim 17 In claim 16, the rigid driver connector is a ventricular assist device in which a plurality of electrodes soldered to the wires are mounted at the same height. Claim 18 A ventricular assist device according to any one of claims 1, 5, 8 and 13, wherein the feedthrough is integrated with the distal shell, and the feedthrough has a first portion as an extension of the distal shell to which the wires of the pneumatic lumen, tether and driveline are joined; and the feedthrough further has a second portion coupled to the first portion to act as a bending relief of the driveline. Claim 19 A ventricular assist device according to claim 18, wherein the proximal shell and the distal shell have an overlapping coupling region, and the pump housing has a surface trench formed over the overlapping coupling region. Claim 20 In claim 19, the above trench is sealed by a potting waterproof material, a ventricular assist device.

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