JP2024517235A - BLOOD PUMP DEVICE HAVING LEAK-FREE AORTIC ADAPTER ASSEMBLY AND METHOD FOR IMPLANTING THE DEVICE - Patent application - Google Patents

Abstract

A blood pump device with a leak-free aortic adapter assembly, the T-type flow connector catheter includes a catheter insert, a convex neck, and a truss. The catheter insert and the convex neck are connected, and the truss is provided on the catheter insert. The two catheter ends of the catheter insert have gradually thinning walls, the catheter ends are compliantly matched to the artery at the implant site, and the proximal end of the convex neck is connected to the inlet adapter of the blood pump. The aortic adapter assembly is provided with a quick connector type coupler and installation method to provide insertable flow communication between the ventricular assist device and the systemic circulation.
[Selection diagram] None

Description

The present invention relates to a ventricular assist device (VAD), and in particular to a left ventricular assist device (LVAD) based on the principle of counterpulsation support, and to a blood pump device having a leak-free aortic adapter assembly and a method for implanting the device.

Mechanical circulatory support systems, especially left ventricular assist devices (LVADs), are evolving into the standard of care for end-stage heart failure emergency care. Based on the mechanical design of the blood pump, LVAD systems can generally be divided into continuous flow and pulsatile flow pumps. Continuous flow devices are built on a rotary machine driven by an axial or centrifugal flow impeller. Pulsatile flow devices, on the other hand, are designed with a displacement blood pump, usually using a diaphragm blood sac to receive and/or transfer blood into and out of the pump.

The inflow/outflow tubes (cannulations) are artificial flow paths for connecting the LVAD system to the human circulatory system in series or parallel. Historically, less attention has been paid to the design of flow tubes compared to blood pump actuators. Surprisingly, a large number of postoperative complications have been found and are related to the failure of inflow/outflow tubes. Among them, poor events such as mispositioning of cannulation, blockage, thrombus formation, and torsional flow cause pump thrombus formation and thromboembolism. In rotary pump implants, the flow path is mainly established by connecting the inflow tube to the cardiac chamber and the outflow tube to the ascending or descending aorta. Up to now, most rotary pump drain tubes are made of Dacron grafts or similar materials, for example, flexible woven materials, and the connection to the artery is made by the end-to-side anastomotic suturing method. This end-to-side technique is technique-dependent, and complications from poor suturing or incorrect aortic flow can lead to thrombus formation in downstream organs and stroke, thromboembolism, or infarction.

Rotary pumps provide fully supported (4-10 L/min) cardiac output and are currently applied in end-stage heart failure. For patients with milder disease, it is appropriate to use a partially supported (2-3 L/min) minimally invasive implantable LVAD. The partially supported LVAD aims to intervene in heart failure early and improve treatment outcomes. Among these new trends in partially supported, minimally invasive LVAD implants, counter-pulsating blood pumps have been the main focus. This is because they have been proven to increase blood perfusion of the myocardium and major organs during systole and diastole. Counter-pulsating pumps must follow strict time control criteria and be guided by the heart rate. Normally, ventricular contraction and withdrawal begin at end diastole, while organ perfusion enhancement occurs at the time of aortic valve closure (pacing cut-out on the aortic pressure waveform). In fact, the therapeutic effect provided is two-fold: first, to reduce ventricular contraction and reduce myocardial oxygen consumption; second, to increase diastolic pressure and increase perfusion of the myocardium, brain and major organs.

The supportive effect of counterpulsation has been clinically proven through the circumferential subcutaneous delivery of intra-aortic balloon pumps. However, peripheral delivery in the descending aorta is often plagued by vascular complications at the insertion site, making long-term use difficult. In the treatment of heart failure, it is highly meaningful to extend counterpulsation support from acute (less than one week) to long-term (several months to several years). To meet this demand, new surgical methods and equipment innovations are needed to achieve this long-term counterpulsation support goal.

The counterpulsatile support of the present invention was performed through a left thoracotomy in which the device was connected to the thoracic artery. This counterpulsatile invention, called a paraaortic blood pump, was implanted through an access hole formed in the descending aortic wall from the lateral pump connection. In contrast to an intraaortic balloon pump, a paraaortic blood pump does not occlude the blood flow path, so it is more flexible in timing control of the counterpulsatile movement. Animal studies have shown that the hemodynamic support effect of aortic balloon counterpulsatile is superior to that of intraaortic balloon pumps. However, it is unclear whether the support of a paraaortic pump leads to long-term complications. Indeed, the challenge is to construct a flow cannulation, implant it in a minimally invasive manner, and provide safe fluid communication between the pump and the connected artery for a long period of time.

In paraaortic blood pump implantation, a false catheter needs to be created to allow blood inflow and outflow. Usually, the pump needs to be filled rapidly to effectively reduce the left ventricular afterload and generate strong vortex washout to prevent pump thrombus formation. Such rapid pump filling often experiences transient low pressures that cause the flow tube to bend and collapse. Therefore, traditional Dacron graft type cannulation is not possible because woven catheters cannot withstand the compressive forces and the collapsed graft may block the inflow during the pump filling phase. In addition, the high and low pressure pulses from rapid pump injection and filling can lead to bleeding problems at the suture site, especially in the acute phase when the surgical anastomosis has not yet healed. Therefore, it is crucial to design the flow tube to overcome the specific flow characteristics associated with aortic counterpulsation.

The hemodynamic characteristics of paraaortic blood pump flow are not physiological. The blood flow of a suction or jet connected blood pump is not laminar in the arterial direction, as is usually seen in the native aorta. This roll-type artificially generated pump flow is highly turbulent and complex. During the pump filling stage, the blood flow makes a sharp turn into the blood pump, resulting in a diversion zone and a recirculation zone at the turning angle. During the pump jet stage, the blood flow accelerates and enters the descending aorta. The blood flow is characterized by an impinging flow with extreme high pressure and shear stress on the opposing aortic wall region. The flow characteristics induced by this device include a low-velocity recirculation zone and a high-pressure, high-shear impinging jet, which can cause endothelial cell erosion, lipid infiltration, and smooth muscle cell proliferation, which may cause aortic wall stenosis and thrombus formation in the long term, or aortic dissection due to hypertension. Therefore, the design of chronically implanted anti-pulsatile devices requires innovative prosthetic flow intubation to avoid or mitigate the pathophysiological flow phenomena caused by the above-mentioned devices, resulting in vascular maladaptation and adverse events such as thrombus formation or aortic dissection.

Safety of surgical anastomosis is another requirement for the design of flow tubes in para-aortic counterpulsatile pumps. Traditional graft sutures may encounter challenging bleeding complications when subjected to excessive high pressure pulses associated with counterpulsatile flow. High and low pressure cycling is the main driving force for material fatigue failure, especially when the implanted aorta develops lesions (atherosclerosis and calcification), degeneration (thinning of the wall) or aging (hardening of the wall). It is important to note that the aortic wall structure can adapt to the applied stress conditions. Cells and tissues around the implant site are remodeled during the surgical wound healing process and further developed by the non-physiological mechanical environment induced by the equipment. Short-term implantation success does not guarantee long-term graft failure due to poor cellular and morphological adaptation of the vessel wall during the postoperative process.

Up until now, all the viable solutions for connecting the artificial graft to the arterial side and the contralateral side have been based on flexible woven tubes by suture method. In the instrument industry, there is no long-term cannulation solution that can solve the problems related to the flow rate and pressure fluctuation level caused by counterpulsation aortic bypass blood pump. For more than fifty years, the application of intra-aortic balloon pumps has already achieved a medium- to short-term (several days to a month) counterpulsation effect in clinical practice. This typical air energy transportation to drive the inflation/outgassing of the balloon is realized using a thin catheter delivered percutaneously from a remote peripheral artery, with a lumen diameter in the range of 6-10. The arteries at such transport sites are mostly occluded, which usually leads to serious bleeding complications and ischemia of the downstream limbs or arms, preventing the long-term use of percutaneous counterpulsation. Aortic transposition is a new method, which aims to extend the counterpulsation treatment to a longer time range. However, long-term aortic balloon counterpulsation needs to build a flow communicator to solve the above implant problems. Such flow intubation must be non-breakable, easy to implant, non-bleeding, and biocompatible without the risk of causing pathological vascular maladaptation. The present invention strives to meet all these collective needs by proposing an insertable aortic adaptor, which will be described later.

An embodiment of the present invention is a blood pump device having a leak-free aortic adapter assembly including a T-type flow connector for use in an implantable ventricular assist device, the T-type flow connector catheter includes an insert, a convex neck, and a truss. The catheter insert and the convex neck are connected, and both of the catheter insert and the convex neck have smooth surfaces that contact blood. The T-type flow connector includes a polymer elastomer and is reinforced by the truss, which includes a nickel-titanium alloy material. The catheter insert has two catheter ends with gradually thinning walls and an appropriate distance between the tips of the catheter ends and the outermost boundary of the truss, the catheter ends have a compliant matching effect with the artery at the implant site, and the proximal end of the convex neck is used to connect to the inlet adapter of the blood pump.

In one embodiment, the truss is co-injected with the polymer elastomer of the T-shaped flow connector and embedded into the wall of the catheter insertion portion of the T-shaped flow connector.

In one embodiment, the polymer elastomer of the T-shaped flow connector comprises a silicone resin material.

In one embodiment, the polymer elastomer of the T-type flow connector is cast polyurethane.

In one embodiment, the structural compliance of the embedded truss and the structural compliance of the polymer elastomer are approximately equal.

In one embodiment, the tapered catheter end is sharp.

In one embodiment, the convex neck portion has a shallow slope to match the inlet adapter of the blood pump, and the inner diameter of the convex neck portion is slightly smaller than the inner diameter of the inlet adapter of the blood pump.

In one embodiment, the shallow slope is inclined relative to the centerline of the catheter insertion section.

In one embodiment, the truss has a multi-wave structure.

In one embodiment, the convex neck portion includes a neck portion body and an extension portion provided on the neck portion body, the extension portion protruding from the neck portion body, and the maximum inner diameter of the extension portion is greater than the maximum inner diameter of the neck portion body.

In one embodiment, when the convex neck portion is joined to the inlet adapter of the blood pump, the extension portion is closely attached to the inlet adapter, and the neck portion body is provided around the inlet adapter. An aortic adapter assembly.

In one embodiment, the coupler further includes a flange base, a pair of locking rings rotatably mounted to the flange base, and a latch on one of the locking rings for locking the locking ring, the locking ring having an internal groove that clamps the T-shaped flow connector with controlled compression to seal the T-shaped flow connector.

In one embodiment, each of the locking rings has a flange profile that can simultaneously bond the locking rings to the edge of the flange base.

In one embodiment, the latch is made of a spring, ensuring that the coupler is locked in place and cannot be accidentally disengaged.

In one embodiment, the coupler further includes a hinge head provided on the flange base, and the lock ring is pivotally attached to the hinge head so as to be rotatable relative to the hinge head and the flange base.

In one embodiment, the hinge head is located on a first side of the flange base, and the latch is located on a second side of the flange base, with the first and second sides of the flange base facing each other.

In one embodiment, the latch has a slot that can engage with a slanted portion provided on the opposing lock ring.

In one embodiment, the flange base has a generally circular structure, and each of the lock rings has an arc-shaped structure.

In one embodiment, a blood pump device as an implantable ventricular assist device includes a blood pump having an inlet adapter with a tapered beak, and an aortic adapter assembly, the aortic adapter assembly including a catheter insert and a convex neck, the catheter insert and the convex neck are connected, the catheter insert and the convex neck include a T-type flow connector having a smooth surface that contacts blood, and a truss provided in the catheter insert, the T-type flow connector has a polymer elastomer and is reinforced by the truss having a nickel-titanium alloy material, the catheter insert has two catheter ends of the catheter insert having gradually thinning walls, an appropriate distance between the tip of the catheter end and the outermost boundary of the truss, the catheter end has a compliance matching effect with the artery at the implant site, the proximal end of the convex neck is connected to the inlet adapter of the blood pump, and the tapered beak is aligned with the proximal end of the convex neck.

In one embodiment, the inner diameter of the inlet adapter is slightly larger than the inner diameter of the convex neck portion of the T-shaped flow connector.

Another embodiment of the present invention provides a method of implanting a flow connector assembly, comprising: an aortic adapter assembly as described in claim 1, the T-type flow connector being compressible from an initial configuration to a compressed, deployed configuration; the T-type flow connector being crimped into the compressed, deployed configuration; a size of the T-type flow connector in the compressed, deployed configuration being at least half the diameter of the T-type flow connector in the initial configuration; the T-type flow connector being inserted into an artery having an access hole drilled in the arterial wall; releasing the T-type flow connector in the compressed, deployed configuration to deploy; the released T-type flow connector being allowed to self-expand to its deployed, initial configuration; and coupling the deployed T-type flow connector to a ventricular assist device.

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. 11 is a schematic diagram of a para-aortic blood pump device according to a third embodiment of the present invention. FIG. 11 is a schematic diagram of a para-aortic blood pump device according to a fourth embodiment of the present invention. 1 is a schematic diagram of a para-aortic blood pump device attached to a human body according to a first and second embodiment of the present invention; FIG. 13 is a schematic diagram of a para-aortic blood pump device attached to a human body according to the third and fourth embodiments of the present invention. FIG. 2 is a first schematic diagram of a drive unit according to an embodiment of the present invention. FIG. 2 is a second schematic diagram of the drive unit according to the embodiment of the present invention. FIG. 2 is a schematic diagram of the driver functions and major interconnect signals required for the operation of the present invention. FIG. 13 is a schematic diagram of driver functions and major interconnect signals required for operation of the third embodiment of the present invention. FIG. 11 is a schematic diagram of the main interconnection signals of the driver functions required for operation of the second embodiment of the present invention; 1 illustrates electro-mechanical actuator (EMA) piston position trajectories and timing of trigger detection commands associated with counter pulsation cycle assistance. FIG. 1 is a perspective view of a para-aortic blood pump implant according to the first or third embodiment of the present invention. FIG. 2 is a cross-sectional view of a para-aortic blood pump implant according to the first or third embodiment of the present invention. FIG. 1 is a perspective view of a para-aortic blood pump implant according to the second or fourth embodiment of the present invention. FIG. 1 is a cross-sectional view of a para-aortic blood pump implant according to the second or fourth embodiment of the present invention. Another embodiment of the present invention is a connection between a carrier and a pump housing of a blood pump, the connection of the carrier being made by an inlet on the remote housing. FIG. 16 is a cross-sectional view of the blood pump shown in FIG. 15 . FIG. 13 shows a shallow groove design for running electrical wires from an inlet in the remote housing to the pressure sensing chamber in the proximal housing. 1 shows the outer end of the blood pump using an interface adapter connector and connecting the opposing mating artery to the artery. 1 shows the coupling of the blood pump of the present invention to the artery via an insert type connection method using a T-shaped intravascular connector coupled via an interface adapter. 1 shows a cross-sectional view of a body of revolution of a stem assembly including an integral axisymmetric elliptical blood sac and sac, a proximal stem and a distal stem. Exploded view showing the components that make up the axisymmetric elliptical blood sac and stem assembly (note, the sac is in its original shape prior to being joined and integrated to the proximal and distal stems shown in FIG. 3). FIG. 19 shows the structure of the bent trilobal pouch at the end of the blood pouch injection. FIG. 1 shows a perspective view of a carrier connected to a proximal end housing of a blood pump via an introduction member. A cross-sectional view of the distal end of the blood pump and carrier is shown along section AA of FIG. 23 shows a cross-sectional view of the proximal end of the carrier along section AA of FIG. 22. FIG. 13 shows a cross-sectional view of an exhaust port attached to a proximal end housing according to the first embodiment. 1 shows a cross-sectional view of the pressure sensing chamber and introducer member in the proximal housing of the first embodiment (note: the carrier is not attached and the introducer includes a first portion, an extension of the proximal housing, and a second portion associated with the first portion). FIG. 25B shows a perspective view of the MEMS pressure sensor incorporated in FIG. 1 shows a cross-sectional view of a multi-layer carrier of the present invention, including an inner tube for air transport, a middle tube for electrical signal transmission, a coil, a tether, and an outer tube. FIG. 11 is a cross-sectional view showing a design proposal for a multi-chamber conveying material. 1 shows a typical diagram of the flow characteristics during the pump filling stage. FIG. 2 is a diagram showing a schematic flow characteristic of a pump discharge stage. FIG. 2 is a perspective view of a T-shaped flow connector according to the present embodiment. FIG. 2 is a cross-sectional view of a T-type flow connector according to the present embodiment. This is an expanded plan view of an inserted metal stent made of nickel-titanium alloy. Define the lateral stiffness (LS) of the insertion catheter to measure the metal stent and T-type flow connector made of nickel-titanium alloy. FIG. 2 is an exploded schematic diagram of the inside of a coupler and its components. FIG. 2 is a schematic diagram of a coupler in an open state (disposition). FIG. 2 is a schematic diagram of a coupler in a locked configuration. FIG. 2 is a cross-sectional schematic diagram of an aortic adapter and a para-aortic blood pump connected by a coupling. FIG. 1 is a schematic diagram showing step discontinuities caused by the docking method. FIG. 1 is a schematic diagram showing gap discontinuities caused by the docking method. FIG. 1 is a schematic diagram of an inlet adapter for attachment to a remote location on an aortic bypass pump. FIG. 4 is a cross-sectional view of the inlet adapter. 1 shows the downward tapered beak of the inlet adapter that connects to the angled (slanted) surface of the neck of the T-type flow connector. FIG. 13 is a schematic diagram of a crimp flow connector arranged to be wrapped with a tie (wrapped configuration). 1 shows an insert flow connector inserted through an access hole formed in the aortic wall. 1 shows the coiled flow connector fully inserted into the aortic lumen. 1 shows the compression-style flow connector repositioned with its T-neck facing the aortic access port. 1 shows the flow connector in an expanded, deployed form, with the T-neck popping out of the aortic access hole after the ties are released. FIG. 1 shows an illustration of the steps of implanting an aortic adapter into the aortic segment of interest and connecting it to a blood pump.

Below, we provide four examples for realizing the para-aortic blood pump device according to the present invention, as follows:

Referring to FIG. 1, it is a schematic diagram of a paraaortic blood pump device according to a first embodiment of the present invention. The paraaortic blood pump device 10 includes a blood pump 12, an aortic adapter 14, a driving catheter 16, and a driving unit 18. The blood pump 12 further includes a pump housing and a pressure sensor. The inside of the pump housing is composed of two chambers, one for storing blood and the other for receiving driving air. The two chambers are separated by an egg-shaped flexible film, which is suspended and fixed by a pair of stress-relief stems connected to the pump housing. The pressure sensor is mounted in the pump housing of the blood pump 12 and monitors the blood pressure in the blood pump 12 to generate an electronic blood pressure signal. The aortic adapter 14 is a valveless, T-manifold shaped conduit-shaped flow manifold that connects the blood pump 12 to the human aorta. In the first embodiment, the aortic adapter 14 and the blood pump 12 are integrally molded to have a seamless blood contact surface, and the blood pump 12 is connected to the human aorta via the aortic adapter 14. The aortic adapter 14 is made of a flexible material and can be deformed when inserted into a circular hole formed in the aortic wall and transported. The aortic adapter 14 can expand and deploy by itself after being inserted into the aorta, and has sufficient strength to withstand the radial compressive contact force of the aortic cavity applied to the adapter wall due to oversize fitting. The driving catheter 16 is connected to the housing of the blood pump 12 and supplies air pressure pulses to the blood pump 12 and transmits the blood pressure signal received from the pressure sensor. The driving unit 18 is connected to the driving catheter 16 to receive the transmitted electronic blood pressure signal, and the driving unit 18 includes an electromechanical actuator that generates air pressure pulses based on the electronic blood pressure signal, which is sent to the blood pump 12 via the driving catheter 16. This wearable actuator 18 provides a heart rhythm-linked air pressure pulse control law and operates by driving the blood ejection and filling of the implanted blood pump 12.

The aforementioned drive unit 18 includes a battery power supply system 11 and a backup battery power supply system (same or similar to the battery power supply systems 21, 31, 41 in the following Figures 2 to 4 and Figure 8), among which the backup battery power supply system ensures a continuous power supply for the drive unit 18. When the patient does not need to move, the drive unit 18 can also be conveniently powered via an AC adapter. The device also includes a clinical monitor, shown in Figures 2 to 4. It can be connected to the drive unit 18 and provide a user interface for the clinician to monitor the device, display diagnostic messages, and access drive unit parameters for initial startup setting of patient data and optimization of specific treatment operation mode settings.

1 and 2 show two different designs of blood pumps 12, 22 that are systematically coupled to the same drive catheters 16, 26 and drive units 18, 28. FIG. 2 is a schematic diagram of a para-aortic blood pump device 20 according to a second embodiment of the present invention, which differs from the first embodiment of the present invention in that the para-aortic blood pump device 20 according to the second embodiment further includes a coupler 25, or coupling adapter. The blood pump 22 and the aortic adapter 24 of the second embodiment are not integrally molded, but are detachable, and a coupler 25 is provided for coupling the blood pump 22 to the aortic adapter 24. The coupler must be carefully designed to minimize discontinuities in the connection interface. During implantation of the device, the aortic adapter 24 is first fed into the aorta through a hole cut in the aortic wall. The device can use a specially developed implantation tool to mount the coupler 25 around the T-neck of the aortic adapter 24, so that the blood pump 22 can be connected to the coupler 25. After the blood pump 22 is inserted into the thoracic cavity, the blood pump 22 and the aortic adapter 24 are tightly locked and integrated through the coupler 25. Such a detachable blood pump 22 and aortic adapter 24 design has advantages during and after the implantation procedure. During the implantation of the device, the detachable blood pump design makes it easier to implant the aortic adapter, because the surgical area is clearer and is not interfered with by the pump body. In addition, after the implantation procedure, if the pressure sensor fails or the blood sac ruptures, requiring emergency surgical replacement, the blood pump can be removed and replaced. In this respect, the detachable blood pump design of the second embodiment is advantageous. The aortic adapter can be left in the aorta without being removed, which avoids the troublesome and dangerous redo surgery involved in removing the aortic adapter.

Refer to Figures 1 and 3. These are schematic diagrams of paraaortic blood pump devices 10 and 30 according to the first and third embodiments of the present invention. The third embodiment differs from the first embodiment of the present invention in that the driving catheter 16 of the first embodiment is replaced by the distal driving catheter 37, driving catheter interconnector 33, and proximal driving catheter 39 of the driving catheter 36 of the third embodiment. The distal driving catheter 37 is connected to the driving catheter interconnector 33 so as to transmit the electronic blood pressure signal acquired from the pressure sensor and the air pressure pulse transmitted from the driving unit 38. In addition, the driving catheter controller and the vibrator (for alarm warning purposes) built into the driving catheter interconnector 33 are originally built into the driving unit 18 of the first embodiment, so the driving unit 18 of the first embodiment has an additional driving catheter controller and vibrator (compare with the driving unit 38 of the third embodiment). The driving catheter controller is used to process the electronic blood pressure signal, and the vibrator is used to provide audio warnings and tactile feedback. In other words, the mechanical power transmission of the blood pump, as well as the analog/digital signal conversion and alarm notification, realized by the drive catheter 16 and drive unit 18 in the first embodiment are essentially the same as those realized by the distal drive catheter 37, drive catheter interconnector 33, and proximal drive catheter 39 in the third embodiment.

The first embodiment has a simpler drive catheter arrangement design, and also has an electronic signal processing device installed in the drive section, which can minimize the risk of environmental contamination (water or moisture condensation) on the pressure signal measurement associated with the drive catheter interconnector 33 and air leakage at the joint. However, such long-type drive catheters are prone to contact damage such as abrasion, kinking, and cuts due to contact with foreign objects during daily activities. If the drive catheter 16 of the first or second embodiment is seriously damaged, both electronically and mechanically, it may require replacement of the blood pump through surgery. Given the risk of redoing the surgery and the associated medical costs, this is highly undesirable. The third or fourth embodiment employs an intermediate connector (drive catheter interconnector) to mitigate the disadvantages of blood pump replacement associated with such drive catheter breakage. In general, the distal drive catheter 37 has a shorter length exposed outside the body, and the skin dressing or cover worn by the patient's vest is better protected through the drive catheter interconnector 33. In the extreme case where the drive catheter is severely damaged and cannot be repaired, the proximal drive catheter 39, which is most likely to be damaged, can be easily replaced without resorting to surgery. In addition, the third and fourth embodiments are not affected by electromagnetic interference because analog-to-digital signal conversion has already been completed within the circuitry of the drive catheter interconnector 33. In the third and fourth embodiments, the fidelity of the pressure signal can be better ensured because the digital signal transmission within the proximal drive catheter 39 is not sensitive to electromagnetic interference.

Refer to Figures 3 and 4. These are schematic diagrams of para-aortic blood pump devices 30 and 40 according to the third and fourth embodiments of the present invention, respectively. The third embodiment differs from the fourth embodiment in that the aortic adapter 34 and the blood pump 32 of the third embodiment are integrally molded, but the blood pump 42 and the aortic adapter 44 of the fourth embodiment are removable. The fourth embodiment further includes the blood pump 22, the aortic adapter 24, and the coupler 25 of the second embodiment. The same blood pump 42, the aortic adapter 44, and the coupling 45 of the embodiments will not be described here. The distal driving catheter 47, the driving catheter interconnector 43, and the proximal driving catheter 49 included in the driving catheter 46 are similar to the distal driving catheter 37, the driving catheter interconnector 33, and the proximal driving catheter 39 of the driving catheter 36.

Referring to FIG. 5, a schematic diagram of a para-aortic blood pump device attached to a human body according to an exemplary embodiment of the present invention. The para-aortic blood pump device 90 includes a blood pump 92, an aortic adapter 94, a drive catheter internal segment 991, a drive catheter external segment 993, and a drive unit 98. In another embodiment, the para-aortic blood pump device further includes a coupler. The part of the para-aortic blood pump device 90 implanted in the human body includes the blood pump 92, the aortic adapter 94, and the drive catheter internal segment 991. In another embodiment, the para-aortic blood pump further includes a coupler. In a surgical procedure, the aortic adapter 94 is attached to the aorta 95 to create an exit site EX of the drive catheter at a suitable location on the human body skin. The external segment of the para-aortic blood pump device 90 includes the drive catheter external segment 993 and the drive unit 98. Across the exit site EX, a segment of the actuation catheter's internal segment 991 is covered with fabric velvet, which is used to promote subcutaneous tissue ingrowth and achieve infection control. The implanted velvet portion is optimally placed 2-5 cm subcutaneously away from the exit site EX. The actuation unit 98 is a wearable or portable device.

6, a schematic diagram of a paraaortic blood pump device installed in a human body according to an exemplary embodiment of the present invention. The paraaortic blood pump device 90 includes a blood pump 92, an aortic adapter 94, a distal drive catheter 97 (including a drive catheter internal segment 971 and a drive catheter external segment 973 outside the human body), a drive catheter interconnector 93, a proximal drive catheter 99, and a drive section 98. In another embodiment, the paraaortic blood pump device 90 further includes a coupler. The portion of the paraaortic blood pump device 90 implanted in the human body includes the blood pump 92, the aortic adapter 94, and the drive catheter internal segment 971. In another embodiment, the paraaortic blood pump device further includes a coupler. In a surgical procedure, the aortic adapter 94 is installed in the aorta 95 to create an exit site EX of the drive catheter at a suitable location on the human body epidermis. The portion of the paraaortic blood pump device 90 located outside the human body includes the drive catheter external segment 973, the drive catheter interconnector 93, the proximal drive catheter 99, and the drive section 98. The exit site EX is the boundary of the drive catheter, and the distal drive catheter is divided into a drive catheter internal segment 971, which is covered with velvet for infection control, and a drive catheter external segment 973. The drive section 98 is a wearable or portable device.

The embedded subsystem is explained in more detail below.

Implantation is accomplished through a relatively small chest opening via a left thoracotomy using less invasive surgical (LIS) techniques. For example, a chest incision is made at the 7th rib gap as the main incision to allow implantation of the aortic adapter and blood pump. Two other small incisions are made between the 6th and 8th ribs to introduce the proximal and distal aortic clamps. The area between the aortic clamps allows the aortic adapter to be implanted through a hole in the aortic wall. The aortic adapter is flexible and can be curled and reduced to a small delivery shape before implantation. Once implantation in the aorta is complete, the aortic adapter automatically snaps back to its original shape and has a super tight fit with the lumen at the planned implant site. Therefore, the material of the aortic adapter is important and should be flexible, but have sufficient radial strength to keep the wall of the implanted aortic adapter circular and not cause buckling of the wall. Possible materials for the aortic adapter include elastomers such as silicone and polyurethane, and reinforced polymer structures or metal materials can be embedded to enhance structural strength.

The following provides further explanation of each aortic adapter and its functional requirements in the above examples.

Hemodynamically, the aortic adapter provides a blood flow connection between the blood pump and the systemic circulation of the human body. In addition to this function, the aortic adapter can also be used as a mechanical base to secure the blood pump to the aortic adapter. The structure of the aortic adapter must be resilient, yet anti-buckling, and strong enough to withstand the internal blood pressure and external contact forces that arise when the blood pump comes into contact with the surrounding lung tissue, and due to respiration and thoracic diaphragm movement.

The aortic adapter 54 is embedded in the aorta, and the boundaries between the two catheter ends (catheter end 545, catheter end 645) and the aortic cavity form a host/implant interface in the bloodstream (see Figs. 13B and 14B). In order to morphologically and elastically minimize the discontinuity of the host/implant interface, the two catheter ends (catheter end 545, catheter end 645) are arranged to have a flared inner surface contour and a continuously tapered wall thickness distribution. This catheter end design minimizes the step at the interface and incorporates the compliance matching effect required at the time of connection into the aortic adapter design. Therefore, the possibility of thrombus formation at the interface can be significantly reduced. This is because the rate of clot aggregation at the interface is slower than the rate of natural thrombolysis provided by the human aortic endothelium. The gradually thinner catheter wall structure also makes the catheter ends (catheter end 545, catheter end 645) softer (conformable), allowing the catheter ends to expand and contract in response to pulse pressure, creating a dynamic sealing effect to prevent blood cells from getting trapped in the boundary gaps that are the usual source of clot formation.

The drive units of each of the above embodiments are explained further below.

Figures 7 and 8 are perspective views of the right and left sides of the drive unit 78. The internal modules of this compact drive unit 78 include an electromechanical actuator (EMA), an electronic controller, a pair of main batteries, and a backup battery. The drive unit 78 also includes a user interface panel 73, a battery hatch 71, a drive catheter socket 75, an external power socket 77, and a pair of ventilation holes 79, as shown in Figures 7 and 8.

The user interface panel 73 of the drive 78 displays important messages during operation and warnings for device failures and aortic pressure status. If the main battery runs low, it can be replaced through the battery hatch 71. If the patient is in bed and has long-term power available from a wall socket, a cable is used to power the drive 78 through a connection to an external power socket 77. A proximal drive catheter 99, one end of a drive catheter external segment 993, and a voltage sensor signal and air pressure pulse are connected to the drive 78 through the drive catheter socket 75. A pair of vents 79 are attached to the opposite side of the drive 78 to allow outside air to pass through the inside of the drive 78 for cooling purposes.

The drive unit 78 can be externally coupled to a clinical monitor, allowing the clinical controller to collect and display real-time clinical waveform data and store patient data for long-term condition monitoring and diagnosis. The clinical monitor unit can also provide a user interface for the clinician-patient, displaying device monitoring/diagnostic information and using it to set drive unit parameters, and then start the drive unit 78 to optimize and set the patient-specific personal circulatory support operating mode.

In the present embodiment, the EMA is an air actuator, which includes a brushless servo motor, a ball screw unit, and a piston-cylinder assembly. Air is used as the driving medium to reciprocate the blood pump to achieve the functions of ejection and blood pump filling.

The air actuator device is located in the drive unit and is carried by the person receiving treatment from the device. The electromechanical actuator includes a brushless servo motor, a piston and cylinder assembly, and a ball screw unit. The ball screw unit includes a ball screw post and a ball nut. The piston is fixed to the end of the ball screw post, and the screw post is linearly reciprocated by the rotation of the nut. The servo motor includes a rotor and a stator, and the rotor and the ball nut are integrated. Through the electromagnetic induction action of the motor, the rotor rotates, and the screw post and the piston achieve linear reciprocating stroke motion in the cylinder through clockwise and counterclockwise direction changes. The piston stroke sends the air in the drive cylinder to the blood pump through the drive lead wire, completing the bleeding and filling action of the blood sac.

Driving a blood pump using air as a medium faces two problems: one is the problem of gas leakage, and the other is the problem of water condensation caused by blood permeating the blood sac wall. The former impairs the blood pump's blood injection and congestion performance and the power consumption of the motor, while the latter poses the risk of bacteria being generated inside the driving cord due to moisture. To solve these two problems, the air actuator of this ventricular assist device is specially equipped with a pressure balance valve device on the cylinder wall. This valve allows the air mass formed by the air in the cylinder and the atmosphere to communicate with each other and to circulate between each other during the air pressure balance process. The air actuator is equipped with a position and optical sensor that allows the controller to obtain the position information of the piston, issue a piston drive command from the controller to drive the reciprocating motion of the piston, and operate the pressure balance valve. Therefore, the timing and frequency of opening the pressure balance valve can be programmed and stored in the controller. By operating this pressure balance valve, the air in the cylinder is exchanged with the outside air, achieving the functions of air supply and dry gas, and further ensuring the operational safety and performance of this ventricular assist device.

Referring to FIG. 9, the para-aortic blood pump device according to the embodiment of the present invention is divided into three parts. The first part is mainly installed inside the human body, i.e., implanted, and its outer end communicates with the second part. The first part includes a blood pump (including a blood pump pressure sensor), an aortic adapter, and a distal driving catheter segment, each of which is implanted in the human body. The second part is installed outside the human body and includes a proximal driving catheter and a driving catheter electronic module (also called a driving catheter interconnector). The third part is a driving unit installed outside the human body and includes an electromechanical actuator (EMA) controller circuit, a main battery, and a backup battery.

The blood pump pressure sensor is built into the proximal blood pump housing and immersed in a small pressure sensor chamber filled with a sensing medium, which can continuously monitor the blood pump pressure. The remote driving catheter is connected to the pump housing and supplies counter-pulsating air pressure pulses to perform dynamic ejection and blood sac filling. The distal and proximal driving catheters supply the blood pump with air pressure driving air pulses generated by the EMA inside the actuator. The electronic blood pressure signal generated by the pressure sensor of the sphygmomanometer is transmitted to the driving unit. The driving air circuit (shown by the dashed arrow) and the communication signal path (shown by the solid line) are shown in Figure 9 to explain the functional relationship between the interactive operating modules. The details of the aortic adapter have already been described. The controller circuit includes a motor controller unit that drives the brushless motor and a microcontroller unit as a central processing unit, which can process the received pressure signal and generate a control command for the motor controller so that the piston works.

Referring to Figures 10 and 11, a schematic block diagram of the internal functions of the drive unit and the important interconnection signals required for the activation of the blood pump is shown. A further description of the embodiment proposed by the present invention is shown in Figures 10 and 11. Refer to Figures 10 and 11. To explain the drive relationship between the external drive unit and the implant, it is necessary to refer to the contents of the blood pump, drive catheter, distal drive catheter, proximal drive catheter, aortic adapter and drive catheter interconnect described above.

The drive receives the blood pump pressure signal (electrical signal), processes this signal using a trigger detection algorithm, and generates a trigger signal that cooperates with the heart rate to activate the motor actuator. After receiving the specified trigger time, the microcontroller unit sends a command to the motor controller (unit) to drive the piston and provide reverse pulsation cycle support during the injection to fill or fill to injection process.

The electronic controller architecture consists of three functional blocks: a microcontroller unit (MCU), a motor controller block (or motor controller unit), and a power management unit. The following table provides an overview of each functional block of the drive section 78.

Figures 10 and 11 illustrate the signal collection, transmission, processing, and generation of control logic and instructions for the exemplary embodiment disclosed above, and how the electromechanical actuator operates to generate air pressure pulses to drive the blood pump.

Figure 12 shows the trigger detection command of the piston position of the electromechanical actuator related to counterpulsation assist. In Figure 12, the aortic pressure (AoP) waveform without assistance is shown in dotted line, and the aortic pressure waveform with assistance is shown in solid line. When the drive unit is operating in automatic operation mode, the drive unit operation is activated and the system performs a "fill-inject-fill-inject..." cycle assist. This represents a normal synchronous counterpulsation assist operation. The MCU monitors the blood pump pressure (BPP) signal (electrical signal) and detects the timing of left ventricular end diastole (LVED). Upon detecting the LVED timing, the MCU generates an F_Trig signal. The time interval between two consecutive F_Trig signals represents the instantaneous cardiac cycle interval (or period). Based on the estimated heart rate calculated from the interval of several cycles going forward, the MCU determines the ejection time of the blood pump, i.e., the E_Trig signal. The E_Trig signal provides the timing to command the motor controller unit to drive the electromechanical actuator based on a predefined position, velocity and acceleration curve. Once the ejection stroke is completed and the optimal dwell time has elapsed, the electromechanical actuator is commanded to perform a pre-fill operation at a slow fill speed until the F_Trig signal is issued. After receiving the F_Trig signal, the electromechanical actuator begins to perform the remainder of the fill stroke at the specified piston speed.

If the MCU loses the BPP signal (electrical signal) sent from the blood pump, the MCU automatically activates the flushing mode to drive the electromechanical actuator and operate at a predetermined auxiliary frequency and volume pacing amount of the drive. The flushing mode is an equipment protection mode to prevent clot formation in the blood sac without providing synchronous counterpulsation circulatory support.

In principle, the para-aortic blood pump device of the present invention has a better counterpulsation support effect than an intra-aortic balloon pump (IABP) due to the feature of non-occlusive aortic placement. Unlike bedridden or ambulatory IABP patients who have to stay in hospital, the para-aortic blood pump device of the present invention allows patients to leave the hospital and live a better life at home. Therefore, the para-aortic blood pump device of the present invention can obtain economic benefits from a shorter hospital stay, and can further improve the disease status and quality of life of patients.

In recent years, the trend of LVAD use has become saturated, mainly because its application has only been applied to a small group of patients with end-stage heart failure. The application of early intervention LVAD therapy to patients with mild heart failure has long been a clinical goal of cardiac medicine. If early intervention therapy is possible, it is expected that this will have a significant impact on the expanded use of LVAD therapy and influence the progress of cardiac medicine in the future. Clinical evidence shows that providing LVAD support at the moderate to severe heart failure stage can improve some non-ischemic cardiomyopathy patients by improving cardiac function through reverse remodeling of cardiomyocytes and causing sustained myocardial recovery. However, such early intervention intention must rely on two favorable driving factors: a simple and safe surgical procedure and an effective adaptive circulatory support treatment plan as the disease progresses. Continuous VAD support is non-physiologic and takes the supported heart off its normal path of recovery. However, counterpulsation support is physiological and provides systolic contraction relief and diastolic perfusion gain to promote reverse remodeling of myocardial cells and achieve therapeutic goals. In conclusion, the therapeutic strategy provided by the para-aortic blood pump invention meets the evolving conditions of early intervention trends in cardiology. The beneficial attributes of therapeutic efficacy provided by the para-aortic blood pump device, such as adaptive partial-support, minimally invasive surgery, and counterpulsation therapy, make the present invention a potential candidate to help advance the future treatment of heart failure.

The blood pumps in each of the above embodiments are described further below.

13A and 13B are schematic and cross-sectional views of a paraaortic blood pump device installed in a human body, which are the first and third embodiments of the present invention. The implanted subsystem of the paraaortic blood pump device includes a blood pump 52, an aortic adapter 54, and a drive catheter (or a distal drive catheter) 57 connected to the blood pump 52. The blood pump 52 includes a rigid or semi-rigid housing 52h and a blood sac 529, the proximal end of which is closed and the remote end is open, and is seamlessly coupled to the aortic adapter 54. Among them, the blood sac 529 is composed of an utricle membrane 526, and the blood sac 529 is fixed to the proximal end housing 523 of the housing 52h via a proximal end port 530, and is also fixed to the remote housing 525 of the rigid housing 52h via a remote port 540. Within the housing of the blood pump 52, a blood chamber B and an air chamber A are partitioned by an egg-shaped flexible membrane 526, which is suspended from the rigid housing 52h via a pair of stress-relief stems (proximal end port 530, remote port 540). The blood chamber B is for storing blood, and the air chamber A is for receiving driving air. A pressure-sensing mechanism 527 (or blood pressure sensor) is hermetically fitted in the proximal end housing 523 of the rigid housing 52h, and the detected pump pressure is transmitted through the membrane 526, propagates through the incompressible liquid or gel contained in the sealed pressure-sensing chamber 528, and is finally received by the pressure-sensing mechanism 527. After receiving the detected blood pump pressure, the pressure-sensing mechanism 527 generates an electronic blood pressure signal. The size and shape of the implantable subsystem assembly are designed to be implantable for patients with a body surface area (BSA) of 1.2 square meters or greater.

14A and 14B are schematic and cross-sectional views of a portion of a para-aortic blood pump device attached to a human body according to a second and fourth embodiment of the present invention, respectively. The implantation subsystem of the para-aortic blood pump device includes a blood pump 62, an aortic adapter 64, a coupler 65, and a drive catheter (or a distal drive catheter) 67 connected to the blood pump 62. The coupling 65 connects the blood pump 62 to the aortic adapter 64 for entering the vascular system of the person who will implant the device. The blood pump 62 includes a rigid housing 62h, which further includes a proximal housing 623 and a distal housing 625. The structure of the blood pump 62 is similar to that shown in FIG. 14B, except that the opening OP is separate and independent from the aortic adapter 64. The coupler 65 is placed around the neck 643 of the aortic adapter 64. The design and basic design principles of the blood pump will be further described below using the design disclosed in FIG. 14B.

Referring to FIG. 13B, the blood pump 52 includes a molded rigid housing 52h, which further includes a proximal housing 523 and a distal housing 525. The rigid housing 52h has a single opening OP connected to the aortic adapter 54 and enters the patient's vascular system. The opening OP of the blood pump 52 is seamlessly manufactured with the aortic adapter 54, providing a smooth and continuous boundary transition to the neck of the aortic adapter 54. The joining of such an integrated blood sac 529 and aortic adapter 54 assembly to the blood pump 52 connects the proximal housing 523 and the remote housing 525 by bonding the proximal port 530 and the remote port 540, respectively. The blood sac 529 is fixed to the top of the proximal housing 523 such that the non-flexible disc portion is bonded to the central pressure sensing chamber 528 of the proximal housing 523.

The miniaturized pressure sensing mechanism 527 is housed in the proximal housing 523 and is in fluid communication with the closed pressure sensing chamber 528. This device allows for continuous monitoring of blood pressure in the blood sac 529. Because the pressure sensing mechanism 527 does not come into contact with blood, the protection of the rigid housing 52h ensures long-term sensor reliability and fidelity, as the rigid housing 52h helps to insulate the pressure sensing mechanism 527 and its circuitry from the effects of chemical corrosion and protein adhesion due to direct blood contact.

The end of the actuation catheter 57 connects to the remote housing 525, which provides a counterpulsation dynamic pressure pulse for ejecting or injecting blood from the blood pump 52. The actuation catheter design can be multi-cavity or multi-layered to encase the wires used to transmit the pressure signal. The kink resistance of the distal actuation catheter 57 can be improved by adopting a metal coil or a woven net or net as the catheter wall reinforcement member. The overall geometry of the flow path in this blood pump is relatively wide, and the valveless aortic adapter design and pulsatile blood pump operation provide excellent blood processing characteristics, avoiding hemolysis due to high shear forces and thrombus formation or thromboembolism due to low flow rates.

Due to the innovative design of the blood pump 52 and blood sac 529 mentioned above, the sac membrane 526 is very durable. The blood sac 529 is an egg-shaped membrane rotating body with the center of rotation on the centerline of the blood pump 52, and two polymer ports (proximal end port 530, remote port 540) are connected to both ends of the rigid housing 52h, which are configured in a disk shape or annular shape, respectively, and when connected to the rigid housing 52h, they reduce stress concentration as a bending/tensile stress relief mechanism. During the process of injecting blood from the blood pump, the sac membrane 526 is compressed or folded into a tri-lobe shape, in which the maximum strain usually occurs at the fold near the edge of the port (proximal end port 530, remote port 540). The local high membrane stress/strain caused by such large deformation of the film is absorbed and offset by the flexible suspension deformation of the port edge. It is particularly noteworthy that the position of the tri-lobe folding mode is not fixed, the deformation of the film is affected by the gravity direction, and the fold position occurs randomly. In fact, the patient's body posture and orientation may change from time to time during daily activities, including standing, lying, sitting, and exercising. Therefore, the gravitational effect acting on the volume of blood stored in the blood pump 52 constantly changes direction, resulting in the formation of non-fixed rulings. This randomly formed thin film overlapping characteristic constitutes the unique fatigue resistance characteristic of the present invention. The present blood pump 52 is expected to have a longer durability than conventional fixed folded yarn film designs.

The folding and expansion of the sac membrane is closely related to the vortex structure mode contained in the blood sac 529. The design of the blood sac described above is characterized by the random formation of folding lines, and the vortex structure features occur alternately with the change in the pattern of the folded membrane. Therefore, the cleaning effect in the blood pump 52 is strong, unsteady, and characterized by random wandering vortex motion. Such randomness of the blood pump vortex structure helps to wash the blood contact surface of the entire blood sac without generating fixed low-speed reflux areas near the membrane wall and in the fold area. In animal experiments, it was observed that the blood pump of the present invention has a very strong anti-thrombotic ability.

The distal drive catheters in each embodiment are described in further detail below.

5 and 6. A schematic diagram of the driving catheter 96 connecting the blood pump 92 and the driving unit 98. The telemetry driving catheter 97 (driving the catheter body 991) that drives the catheter 96 connects the blood pump 92 to the electromechanical actuator housed in the driving unit 98 by air and transmits the electrical signal obtained from the blood pump pressure sensor mechanism 527 (see FIG. 13). One end of the telemetry driving catheter 97 (driving catheter internal body segment 991) is connected to the blood pump housing, and the other end has a small external connector for air transmission and electrical communication. The telemetry driving catheter 97 (driving catheter internal body segment 991) exits from the skin through the subcutaneous passage under protection with a protective cap. The outer diameter of the telemetry driving catheter 97 (driving catheter internal body segment 991) is designed to be small, and the tube material is radial, which minimizes the stress at the exit site EX and provides comfort to the patient. A portion of the telemetry actuation catheter 97 (actuation catheter internal segment 991) is covered with a porous fabric to promote tissue growth into the fabric and make the exit site EX infectious. The actuation catheter external segment 973, which exits the epidermis, is fixed a short distance away from the skin exit site EX.

The drive catheter internal segment 991 and the drive catheter external segment 993 and their connectors are designed to withstand the tensile loads imposed during percutaneous surgery. After surgery, the drive catheter internal segment 991 and the drive catheter external segment 993 continue to be subjected to loads due to muscle movement, and the drive catheter internal segment 991 and the drive catheter external segment 993 are designed to withstand these fatigue loads. The exterior of the drive catheter internal segment 991 and the drive catheter external segment 993 are designed to be biocompatible and chemically resistant to cleaning agents and disinfectants during clinical use.

The proximal drive catheters of each of the above embodiments are described further below.

The drive catheter external segment 993 and the proximal drive catheter 99 are for connecting the drive catheter internal segment 991 and the distal drive catheter 97 to the drive unit 98, respectively. The near drive unit catheter 99 has a drive catheter interconnector 93 at one end and a drive unit connector at the other end. The drive catheter interconnector 93 includes a circuit board that converts analog blood pump pressure signals to digital signals and a vibrator that provides tactile feedback in addition to the audio alarm. The drive catheter interconnector 93 has a flat shape to prevent twisting of the drive catheter external segment 973 when secured to the patient's skin. In addition, the drive catheter interconnector 93 and the drive catheter embedded wire are sealed and designed to prevent the ingress of water and moisture. The proximal drive catheter 99 is externally mounted to allow replacement and maintenance when deemed necessary, so that if the proximal drive catheter 99 is damaged beyond repair, surgical replacement of the blood pump is not required.

Valveless blood pumps have two advantages in terms of blood processing characteristics. 1) There is no unpleasant valve sound and valve-induced blood cell disorder, thrombus formation and thromboembolism. 2) They have stronger antithrombogenicity. Bidirectional pulsating flow has a better surface cleaning effect, so that it can minimize protein adhesion and avoid the formation of blood clots associated with boundary discontinuities on artificial surfaces in contact with blood. The flow paths of pumps without valve pulsation are much more uniformly wide than those of valve pulsating or continuous flow rotary pumps. Hemolysis (red cell membrane rupture) usually occurs in narrow flow paths with high flow rate gradients, such as the gap between the annulus and leaflet of valved pulsating pumps. In addition, there are often low-speed recirculation and stagnation areas behind the open valves, which may promote the development of thrombus. In contrast, in a nonvalvular pulsatile blood pump, the shear stress on blood cells is actually several orders of magnitude lower, the slow stagnation zones associated with valve geometry and motion are nearly eliminated, blood cell damage and platelet activation are reduced, clot formation and aggregation are reduced, and lower doses of anticoagulants can be used, resulting in easier and safer postoperative care.

Figures 15-17 show schematic diagrams of a blood pump 62, drive catheter 67, and introducer member 63 according to another embodiment of the present invention. This embodiment emphasizes anatomical adaptability to facilitate easier externalization of the blood pump and delivery.

15 and 16, the drive catheter 67 is connected to the remote housing 625 of the blood pump 62. The blood pump 62 has an elliptical blood sac and stem assembly 659 (including the sac 629 and stems 630, 640), a pump housing 62h (proximal end housing 623 and remote housing 625 with inlet connector 6251), and a pressure sensing system 628 fitted to the proximal end housing 623. The aforementioned components, inlet connector 6251, pressure sensing system 628, and drive catheter 67 are essentially the same or correspond to the assemblies/assemblies of the previously described embodiments, and a detailed description of these assemblies and their functions will not be repeated here.

In this embodiment, the remote housing 625 of the pump housing 62h is provided with an introduction member 63, and the driving catheter 67 is connected to the pump housing 62h. The introduction member 63 is also configured in a body shape adjacent to the remote housing 625, and connects the power transmission member tangentially to the outer surface of the pump. With such a through-body design, the design of the pump housing 62h is adapted to the anatomical space available for the installation of the blood pump. The blood pump 62 is rotatably connected to the interface adapter 501, and the driving catheter 67 can be wired in an optimal direction to facilitate smooth subcutaneous tunneling and skin exit. This is advantageous for anatomical adaptability to various shapes of the implant site.

In this embodiment, the introduction member 63 is remotely fitted into the remote housing 625, and the sensor 6271 (FIG. 25A) and the pressure-sensing chamber 628 are disposed within the proximal end housing 623, so that the signal transmission path and the air pressure communication path are separated, and the blood pump 62 must be sealed and protected, and the ingress of biochemical fluids may impair the fidelity of signal transmission after the device is implanted.

Refer to FIG. 17. The pump housing 62h has a groove 621 formed on the outer surface of the remote housing 625, located above the overlapping joint area da (FIG. 16) of the proximal end housing 623 and the remote housing 625. The surface groove 621 is arranged so that an electric wire extends from the outlet of the inlet 63, and the electric wire reaches the electrode 6274 in the second space 6273 (FIG. 25B) along the surface groove 621 above the overlapping joint area DA. In some embodiments, the groove 621 is sealed by potting a waterproof material and/or an annular cap so as not to irritate or injure the tissue with which it comes into contact.

18A and 18AB, the aortic connector 50 is a connection mechanism typically used to connect a blood pump 62 to a target artery 60 via its interface adapter 501. The distal end 504 of the connector 50 opposite the interface adapter 501 is placed on the vessel wall of the artery 60 and connected to the body circulating fluid. The proximal end of the aortic connector 50 (or interface adapter 501) has a smooth interface transition to geometrically match the inlet configuration of the blood pump 62. A coupler is typically required to integrate the proximal end of the connector 50 (or interface adapter 501) with the inlet of the blood pump 62. In some embodiments, the aortic connector 50 can be used. In FIG. 18A, an end-to-side mating of a sutured Dacron or PTFE graft 502 with a target artery (or blood vessel) 60 is shown, which can be used in vascular surgery. In some other embodiments, such as the embodiment shown in FIG. 18B, an insertable aortic connector 503 is used, such as a T-manifold adapter as disclosed in U.S. Patent Application No. US 2008/0300447A1, entitled "Dual-pulsation bi-Ventricular Assist Device."

19 and 20 respectively show a longitudinally symmetrical elliptical blood sac and stem assembly 650 and its components. Polymeric materials for these components include, but are not limited to, segmented polyurethanes with various suitable hardnesses. 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, in some embodiments, are integrally formed in an axisymmetric shape about a common centerline 62C of the blood pump 62. The proximal stem 630 is located at the proximal end 6291 of the blood sac 629, and the distal stem 640 is located at the distal end 6292 of the blood sac 629.

Figure 19 shows an integrated blood sac and stem assembly 650, with the end of the remote stem 640 wrapped around the remote of the blood sac 629 and connected to the inverted membrane 62A. Figure 20 shows the components before connection. Typically, the blood sac 629 is manufactured by dip molding, and the stems 630, 640 are manufactured by injection molding. There is no preferred azimuth angle for the deformation of the deflection sac. In theory, as shown in Figure 21, when the pressure difference on the membrane sac exceeds a certain threshold, the thin-walled sac with an axisymmetric elliptical structure will bend into a tri-valve shape 6293, and the bending of this membrane is related only to the final tri-valve shape 6293 (eigenmode), and the position of the fold 6294 or turn-back line that occurs there is determined by the initial disturbance that causes the bending instability. The uniformity of the thickness in the cross section cut by the center line (or axis of rotation) 62C of the assembly 650 perpendicular to the blood pump 62 is important. Care must be taken to manufacture the sac with high precision to ensure an axisymmetric shape. In real life, the direction of gravity is a major factor in the initiation of folds. The posture of the recipient of the equipment constantly changes due to the patient's daily activities (e.g., standing, sitting, exercising, sleeping, etc.), and the direction of gravity relative to the blood pump direction also constantly changes. Therefore, the deformed folds 6294 of the sac appear in a random state, and the high strain folds are non-stationarily distributed throughout the sac. Therefore, avoiding leaving high strain areas in fixed positions is an important design criterion for the sac to have a long life.

The embodiment of the present invention innovates the traveling fold line characteristic that causes high strain locations to appear in the membrane in a non-stationary manner to increase the fatigue life of the membrane sac. Therefore, the harmful stress concentration phenomenon often associated with flexed blood sacs is improved. Based on the inherent change in the behavior of this flexing mode, the significant increase in the fatigue life of the membrane is due to this non-stationary fold line formation characteristic, which distributes high strain areas throughout the sac. The beneficial results associated with this non-stationary sac deformation mode also depend on the enhanced vortex washing effect in the blood sac. The surface of the sac is washed more thoroughly and irregular walking-like vortices are formed and traverse the entire vortex flow. This significantly reduces the possibility of constant low-speed recirculation areas and fold line creases in the near-wall region, resulting in a long-life and anti-thrombogenic blood pump design.

Figures 22, 23A and 23B show an exemplary embodiment of how the integration method can be employed to attach the blood sac and stem assembly 650 to the pump housing 62h and connect the drive catheter 67 to the proximal end housing 623.

The blood sac 629 is fixed to the pump housing 62h, which includes the proximal end housing 623 and the remote housing 625, to facilitate the filling and discharging action of the pump. In general, the flexure characteristics of the blood sac 629 and the pump housing 62h are significantly different. To complete the design of the blood sac for long-term use, it is necessary to attach an intermediate suspension so that the pump assembly is continuous in the transition of structural characteristics (especially the bending deformation of the membrane). A pair of flexible stems (proximal end stem 630 and remote stem 640) are used as a suspension mechanism that integrates the blood sac 629 and the pump housing 62h. As shown in FIG. 23A, the proximal end stem 630, which is disk-shaped, is connected to the proximal end housing 623, and the remote stem 640, which is ring-shaped, is connected to the remote housing 625. Mechanically, the proximal and distal stems 630, 640 not only hold the blood sac 629 within the pump housing 62h, but also act as a stress relieving suspension mechanism to avoid stress concentrations at the interface attachment points and extend the life of the blood sac 629.

23A, the remote housing 625 includes an extension of the inlet connector 6251, which is coupled to the aortic connector 14. The inlet connector 6251 has a first end 6252 attached to the blood sac 629 and a second end 6253 shaped like a beak, which is coupled to the interface adapter 501 (shown in FIGS. 18A and 18B). The first end 6252 of the inlet connector 6251 mates smoothly with the remote of the blood sac 629. However, the opposing second end 6253 is configured to mate with the interface adapter 501, and the coupling design goal is to minimize discontinuities in the interface to avoid the formation of clots. The beak has a flange structure, which is located in the mid-region of the inlet connector 6251 and acts as a locking component that is received by the interface adapter 501.

During surgery, the closed-end pouch design of the valveless blood pump 62 draws air into the pouch apex by buoyancy and collects air bubbles. See Figures 22 and 24. The proximal end casing 623 is fitted with or provided with an exhaust port 66, and a narrow passageway 661 is provided above the integral pouch pedicle septum 6301 between the proximal end stalk 630 and the blood pouch 629. In some embodiments, the passageway 661 extends along the centerline 62C. After matching the blood pump 62 to the target artery 60, the trapped air emerges under the pressure of arterial blood pressure and collects in the apical space of the blood pouch 629. A thin needle passes through the exhaust port 66, the passageway 661, the pouch pedicle septum 6301, and reaches the interior of the blood pouch 629 to expel the accumulated air. Because the entire capsule stalk 6301 below the exhaust port 66 is relatively rigid and not bent, the capsule 629 holding the perforation is not further structurally destroyed, and structural destruction due to the progression of cracks in the gaps of the perforation can be avoided when subjected to periodic pulse pressure or the expansion and contraction and folding of adjacent capsules.

As shown in Figures 23A and 25A, the pressure sensing mechanism 627 is embedded in the proximal housing 623. Figure 23A shows a cross-sectional detail of the integrated proximal housing 623, introduction material 63, and drive catheter 67. Figure 25A shows the profile of the proximal housing 623 connecting the introduction material 63 extending from the dome of the proximal housing 623 to communicate air and signals to the drive catheter 67.

25A and 25B, the pressure sensing mechanism 627 includes a sensor 6271 hermetically housed in a metal can, and has a first space 6272 for fluid communication and a second space 6273 for housing a microelectromechanical system (MEMS) pressure sensor and associated electronic circuitry. A plurality of electrodes 6274 extend from the bottom of the second space 6273 and are connected to the wires 6702 of the carrier 67 (FIG. 26). The second space 6273 is closer to the driving catheter 67 than the first space 6272. The first space 6272 is open to communicate with the fluid of the sensing medium. A biocompatible fluid or jelly is used as the pressure transmission medium. A chamber or pressure sensing chamber 628 located in the proximal end housing 623 and adjacent to the first space 6272 is formed to contain a sensing fluid. The remote end of the pressure sensing chamber 628 is isolated from the blood cavity by a membrane sac 629. The pressure sensing chamber 628 has two side arms, a first arm 6281 for mounting the sensor 6271 and a second arm 6282 for filling and sealing the medium to be sensed. Thus, blood pressure pulses can be transmitted through the transmembrane blood sac 629 and hydraulically communicate with the remote MEMS sensor 6271 located in the second space 6273.

One embodiment of the present invention innovates the pressure-based blood pump control method and sensor design. It employs a micro-MEMS pressure sensor whose electronics are packaged and embedded in the housing wall. In principle, due to its inherent microscale structure, the grains of the MEMS sensor are very durable. In practice, the durability of the sensor depends on the packaging design. This pressure-sensing mechanism 627 is non-blood contacting and isolated from the corrosive biochemistry associated with blood, providing the long-term signal collection and transmission required for long-term implantable assistive devices.

The carrier material 67 functions as a transmitter that converts electrical signals and transmits air pulse pressure between the blood pump 62 and the drive unit 98. Figure 26 shows a representative multi-layer carrier 67 of the present invention. In this embodiment, the carrier material 67 has an air lumen (or inner air tube) 6701, multiple electric wires 6702, an intermediate air tube 673, a coil 674 such as a metal coil, an outer layer tube 675, a tether 676, a silicone sheath 677, a rigid drive unit connector 678, and a protective hollow connector 679.

The central part of the carrier 67 houses an air lumen 6701 (or air passage, called inner tube) with a lumen diameter of 2-5 mm, which allows the choice between low power consumption and low surgical convenience. The electric wire 6702 for signal transmission is embedded in the wall of the carrier 67. Other carrier variant designs can be adopted. In addition to the design of the multi-layer carrier 67 shown in FIG. 26, the carrier 67 may be multi-cavity, for example, as shown in FIG. 27, so that the electric wire 6702 can be embedded in several small tube chambers and pulse air can flow to the larger tube chamber 6701. One of the small lumens can be attached with a tether 676 to limit the elongation of the delivery 67 and protect the electric wire 6702 from damage due to external pulling force.

The inner tube or air lumen 6701 is received by the air tube 673, with a reinforcement sandwiched between them. Between the inner tube 6701 and the air tube 673, a coil 674 (or woven thread or net) can be reflowed to reinforce the wall of the carrier material, making the carrier material 67 flexible but resistant to kinking. The outer layer tube 675 covers the inner and middle air lumen 6701 and the air tube 673, and the helically wound wire 6702 can be used as a protective cover. Note that in some embodiments, a non-expandable tether 676 can be provided between the outer tube 675 of the carrier material 67 and the silicone cover 677 to increase the tensile elasticity required when the carrier material 67 is externalized. Clinically, the silica gel jacket 677 has been proven to be the least irritating to subcutaneous tissue and to have the lowest rate of transport infection.

In this embodiment, the air lumen 6701, the metal coil 674, the air tube 673, the helical wire 6702, the outer layer tube 675, the tether 676, and the silicone sheath 677 are packed in the main body of the carrier 67. The base end 671 of the carrier 67 is inserted into a socket provided in the drive unit 98. The rigid drive unit connector 678 of the carrier 67 is used to be received in a socket in the drive unit 98. The rigid drive unit connector 678 is mounted flush with the four electrodes 6781 (e.g., FIG. 23) of the plurality of electrodes 6781 to which the wires 6702 are soldered. A protective hollow connector 679 (see FIG. 23b, FIG. 26) is disposed at the junction between the drive catheter 67 and the drive connector 678 to prevent the carrier 67 from kinking at the junction. The base end 671 of the carrier 67 includes a drive connector 678 and a hollow connector 679, and is attached so that it can easily exit the skin without causing unintended puncture injury.

22 and 23A, the connection between the carrier 67 and the blood pump 62 is achieved by an introduction member 63. Depending on the anatomical structure in which the blood pump 62 is implanted, the introduction member 63 can be located in the proximal housing 623 or in the distal housing 625. By integrating the introduction member 63 with the pump housing 62h, the overall external blood pump can be repositioned to guide the transport member 67 in a specific direction, such as the externalization route of the transport member, post-operative skin care, equipment availability, etc.

23A and 25A, the inlet 63 has a first part 631 which is an extension of the proximal housing 623, and the air chamber 6701, the tether 676 and the electric wire 6702 of the driving catheter 67 are connected to the first part 631 via the anchor adapter 672. The introduction member 63 further has a second part 632 which is linked to the first part 631 as a hollow connector of the carrier 67. The first part 631 is a position for connecting the electric wire, anchoring the tether, and bonding with the air chamber and sealing with the housing. The electric wire must not be exposed to the tissue at the implant site and must be protected from tension during the externalization process of the power transmission member. In addition, the connection between the air lumen 6701 and the blood pump 62 must be air-free and current-free. The above blood pump integration task is performed by the first part 631. The second portion 632 houses these boundary components and acts as an external protector, shielding the boundary from mechanical stresses and environmental fluids and moisture.

19 to 27 show a modular design of the first embodiment of the blood pump of the present invention. In this embodiment, the blood pump 62 includes 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 housing 623 and a distal housing 625, and a delivery member 67 connected to the blood pump 62, the delivery member 67 including an air lumen 6701 and an electric wire 6702 in its wall. The introduction member 63 is used to realize electrical and pneumatic communication between the carrier 67 and the blood pump 62 to integrate the drive catheter 67 and the pump housing 62h.

As shown in Figures 19 and 20, the design of the blood sac 629 and the stem 630, 640 connection is disclosed in the previous section. The design and manufacturing focus is to maintain precise axial symmetry in the part manufacturing and joining of the sac and stem assembly. The pressure sensitive mechanism 627 and the introduction member 63 are attached to the rigid portion of the proximal housing 623. Figures 22, 23A and 23B show the design of the introduction member 63 in compaction. It can be seen that the compressed inlet 63 allows for a stronger and more fault tolerant wire and connection.

Figures 28A and 28B show some flow patterns associated with aortic balloon counterpulsation. During late diastole and early systole of left ventricular bleeding, the blood pump draws aortic blood flow into the pump through pump filling (Figure 28A). The blood upstream and downstream around the connector is drawn into the blood pump with a sudden 90-degree flow. This causes flow separation and a low-velocity recirculation region T-201. Also, very high shear appears in the corner region of the T-joint. Meanwhile, during diastole after the aortic valve closes, the blood stored in the pump is blown back into circulation, causing a shock flow on the opposite aortic wall (Figure 28B). This roll, collision flow has a very high local pressure at the collision point T-202, a so-called stagnation point, in which the flow velocity is practically zero and all the kinetic energy associated with the flow velocity is converted into potential energy called total pressure. Such high-pressure shock flow can cause vascular maladaptation, such as smooth muscle cell proliferation and resulting vascular wall stenosis, and the risk of aortic dissection due to persistent local hypertension. These non-physiological flow patterns and the induced high pressure, high shear, and low velocity recirculation phenomena are all ubiquitous near the T-tube. These turbulent, complex flow abnormalities are attenuated or diminished within a distance of 3-5 times the diameter of the implanted arterial lumen. The plug-in flow connector is designed with an insertion catheter length of 5-7 cm, covering most of the non-physiological flow area caused by the pump. The aorta at the implant site is blocked by the inserted flow connector, so the biological vessel wall is protected from the pathological stress conditions caused by the pump, and the artery at the implant site is protected from acute or long-term remodeling complications.

In anti-pulsatile support, the pump fills and ejects alternately in synchronization with the heart rate, creating a special T-joint flow as shown in Figures 28A and 28B. The aforementioned insertable aortic adapter 14 is further detailed in the perspective and cross-sectional schematic diagrams of Figures 29 and 30, respectively.

The aortic adapter 14 is mold injected and has an internal blood-contacting surface 141 that is ultra-smooth, continuous, and free of parting lines. The aortic adapter 14 can have a silicone or other polymeric elastomer material. In some embodiments, the aortic adapter 14 has a polymeric elastomer that includes a silicone material, or the polymeric elastomer is mold injectable polyurethane. The aortic adapter 14 includes an aortic adapter catheter (or catheter insert) 142 for insertion into the aorta 95 (see FIG. 5) and a neck (or ridge) 143 for connection to a blood pump. In this embodiment, the convex neck portion 143 has a neck body 1431 and an extension portion 1432 provided on the neck body 1431, the extension portion 1432 protruding from the neck body 1431, and the maximum inner diameter of the extension portion 1432 is larger than the maximum inner diameter of the neck body 1431. When the convex neck portion 143 is connected to the blood pump 62, the extension portion 1432 is in close contact with the inlet adapter 6251 of the blood pump 62, and the neck body 1431 is surrounded by the coupler 25, and the inlet adapter 6251 is integrally connected to the blood pump 62.

The aortic adapter 14 is thin-walled to maximize flow efficiency. To reinforce its thin-walled structure, a pair of nickel-titanium alloy metal stents (called truss rings) 144 are fitted into and surround both ends of the aortic adapter catheter 142 of the aortic adapter 14.

Figure 29 is a schematic diagram showing the fitting position of the nickel-titanium alloy metal stent 144. Also, the wall thickness of the aortic adapter catheter 142 is gradually tapered toward both catheter ends 145. The gradually tapered wall thickness has a double effect. First, it maximally reduces the discontinuity of the graft 502/subject connection, making the rate of interface clot formation much lower than the rate of thrombolysis provided by contact with the endothelium. Second, the compliance of the catheter becomes softer toward the catheter end 145, creating a compliance matching effect when connecting with the aortic lumen.

One of the complications that is troubling in the delivery of large stent grafts 502 is the endoleak problem. Type I endoleak is an incomplete seal between the end of the implant and the implanted arterial endothelial lumen, resulting in a gap between the leading edge of the implant and the arterial lumen. The oozing blood is trapped in the gap and coagulates into a blood clot, and eventually into a fibrous pseudointima that grows uncontrolled over time. The pseudointima not only occludes the implanted artery, but also signals and stimulates the clotting mechanism, attracting platelet adhesion, which can lead to thrombotic failure. The solution to such endoleak problems is to tightly seal the aortic adaptor 14 against the lumen surface to which it is attached. The aortic adaptor 14 of the present disclosure proposes a compliant design concept that allows the semi-rigid catheter (terminal) end 145 to seamlessly attach to the arterial lumen when subjected to pulsatile pressure. As shown in FIG. 30, the outer diameter 146 of the aortic adapter catheter 142 is slightly larger than the diameter of the inner lumen, and the oversize ratio (defined as the ratio of the catheter diameter to the inner lumen diameter) is within the range of 3-10% under a given nominal blood pressure condition (e.g., 120 mmHg). The compliant catheter distal end 145 dynamically expands and contracts in response to pressure pulsations without creating a boundary gap when blood pressure fluctuates between systole and diastole, or under pulsatile pressure with counterpulsatile support.

Thin-walled elastomeric tubes are often flexible and compliant, but the bulk of the device is such that they are not strong enough to withstand the compressive forces applied, causing the wall of the inserted adapter to bend. Therefore, the combination of a nickel-titanium alloy metal stent 144 structure and an elastomeric substrate with moderate hardness is important. As shown in FIG. 30, this design provides radial stiffness through the nickel-titanium alloy metal stent 144 to support the aortic adapter 14 without bending, and there is a distance "x" between the outermost boundary 1441 of the metal stent 144 and the catheter end 145. In some embodiments, the distance x must be evaluated and precisely defined. With the nickel-titanium alloy metal stent 144 as an expandable frame to support the gradually thinning catheter end 145, it abuts the connected lumen wall, keeping it round without collapsing or wrinkling, and dynamically seals with the lumen. The aortic adaptor 14 can expand and contract in response to pressure pulsations, and the aortic adaptor 14 and the wall of the aorta 95 expand and contract together without causing bleeding complications, dynamically achieving a sealing effect that seals the catheter end 145.

Figure 31 shows a representative embodiment of a nickel titanium alloy stent 144, typically made of a laser sculpted nickel titanium alloy, which is further expanded by a series of expansion and heat treatments. Figure 31 shows a flat development of the metal stent 144. (Per row) The metal stent has a multi-corrugated structure. The metal stent 144 is self-expanding and can be folded or curled into small curls, placed in a pre-crushed roll, and self-released back to its original shape after being placed in the required position.

A simple measure of catheter stiffness (the inverse of compliance) can be expressed as the so-called lateral stiffness (LS), and its measurement method is described in Figure 32. Lateral stiffness is defined as the applied force per unit length F divided by the corresponding radial deflection Y. For the current aortic adapter, a suitable lateral stiffness range is 0.01-0.05 Nt/mm2. The embedded nickel-titanium alloy metal stent 144 and the silicone resin or elastomer substrate contribute to the structural compliance of the coinjected aortic adapter 14. The structural compliance of the metal stent 144 and the structural compliance of the polymer elastomer of the aortic adapter 14 are approximately equal. It is preferable to have a uniformly distributed softness so that the tendency of layering between layers due to expansion and contraction of the composite catheter wall is minimized, and the adapter 14 has a long life.

The aortic adapter 14 is arranged to be connected to the blood pump 62 to promote circulatory support. Here, a quick connector type coupler 25 is provided. As shown in FIG. 33, an exploded schematic view of each part of the coupler 25 in which the aortic adapter 14 and the blood pump 62 are integrated is shown. The coupler 25 includes a flange base 252, a pair of locking rings 253, and a hinge (or hinge assembly) 254 connecting the locking ring 253 and the flange base 252. A spring coil (or coil assembly) 255 is loaded into the hinge joint 256 to hold the locking ring 253 in an open position when the coupler 25 is unlocked (FIG. 34A). The locking mechanism relies on a leaf spring latch 257, which is made of a grooved spring piece and is fixed via a plate 2571 welded to one end of the locking ring 253. The flange base 252 has a substantially circular structure, and each locking ring 253 has an arc structure. The hinge joint 256 is located on a first side 252S1 of the flange base 252, and the leaf spring latch 257 is located on a second side 252S2 of the flange base 252 opposite the first side 252S1. The lock ring 253 is pivotally attached to the hinge joint 256 and is rotatable relative to the hinge joint 256 and the flange base 252. In some embodiments, the coupler 25 and the aortic adapter 14 are part of an aortic adapter assembly.

Figure 34b shows a schematic diagram of the coupler 25 in a locked state, with the slits of the leaf spring latch 257 firmly engaged with the ramps 258 to ensure a secure connection without risk of separation. As shown in Figure 35, the integral connection between the aortic adapter 14 and the blood pump 62 acts as a "gasket" between the rigid flange base 252 and the rigid beak flange 81 of the blood pump inlet adapter 6251 connected together via the deformable adapter proximal end 147 (Figure 30, i.e., the end of the neck).

Specifically, as shown in Fig. 34b and Fig. 35, the closed locking ring 253 can easily perform quick-connect locking without worrying about accidental unlocking. The plate spring latch 257 is attached to the end of one of the locking rings 253. While the locking ring is locked, the plate spring latch 257 bends as it slides over the (convex) slope 258 of the opposite locking ring 253 during locking. When the plate spring latch 257 overcomes the top of the slope 258, it lowers to the bottom of the slope 258 by elastic restoring force, which acts as insurance against accidental unlatching and opening of the locking ring due to pump vibration or long-term rocking. In the case of pump explant or replacement requiring module separation, the plate spring latch 257 is bent and lifted upward by the tool used, and an unlocking force is applied to rotate the locking ring 253 open, allowing the blood pump 62 to be detached from the aortic adapter 14.

The butt joint design makes it impossible for two smooth-surfaced pipe joints to be connected in the bloodstream. In many clinical applications, the surface of the connected graft 502 is rough to promote endothelialization, whereby minute boundary discontinuities in the bloodstream are "smoothed out" by the ingrown cells and proteins. The current aortic adapter 14 avoids the occurrence of poor thrombosis by making the surface smooth, but the reason for this has already been explained. As shown in Figures 28A and 28B, the blood flow in the aortic adapter responds to the counter-pulsating pump with a bidirectional jetting and filling action. This strong bidirectional flow and surface washing effect can easily remove any newly formed blood clots on the rough surface. Therefore, it is believed that the smooth surface design is more compatible with the present invention and is safer to use. The boundary of two connected smooth surfaces in the bloodstream requires careful mechanical and hemodynamic design to prevent in situ occurrence of thrombosis events. The principles and design methods related to the invention of such a new joint are disclosed below.

Figures 36A and 36B respectively show two basic boundary discontinuities that exist in butt joint connections, such as a step 101, 102, or gap 103 between adapter AB1 (e.g., adapter of blood pump 62) and adapter AB2 (e.g., aortic adapter 14). Figures 36A and 36B are exaggerated illustrations; typically, in precision machining, such joint discontinuities are within 10-50 microns, which is large enough to cause clots and thrombus formation.

In reality, two separate objects need to be toleranced even if the machining of each object is exactly the same. Figure 36A depicts two misaligned fittings where all parts are correctly completed except for the centerlines not being aligned. A step 101, 102 facing forward and a step 102 facing backwards are created, and stagnant flow in the step areas 101, 102 is the origin of clots or thromboemboli. As shown in Figure 36B, the non-parallel matching of the joints creates a boundary gap 103. The gap 103 attracts and collects blood cells, which then grow into a pseudointima. This intimal growth is often uncontrollable and can result in a thrombus bolus that is shed from the intimal surface as well as blocking the entire blood flow path. The incorrect boundary of the butt joint can be exacerbated if the connected objects are non-rigid. The aortic adapter 14 of the present disclosure is semi-rigid and can be forcefully press-fitted to connect to the fitting (e.g., inlet adapter) and has a deformed structure and expanded boundary discontinuity. Therefore, in order to achieve a connection between the current semi-rigid aortic adapter and the blood pump, it is necessary to invent a new connection method, as follows:

Refer to Figures 37 and 38. In some embodiments, the blood pump 62 has an inlet connector 80 that includes a beak-shaped flange 81, a beak 82, and a connector body 83 as an extension of the housing of the blood pump 62. The connector body 83 includes a number of holes 86 for connecting the inlet connector 80 to the blood pump 62.

The inner diameter 84 of the beak 82 is slightly larger than the inner diameter 148 of the convex neck portion 143 of the aortic adaptor 14 (see FIG. 30). The contact area between the beak 82 and the proximal end (end) 147 of the adaptor is an annular tapered surface (also called a shallow slope or bevel) 149 as shown in FIGS. 30 and 35, so the beak 82 can be said to be a tapered beak. The shallow slope 149 is inclined with respect to the centerline of the catheter insertion section, and the taper angle of the shallow slope is in the range of 30-60 degrees measured from the rotation centerline of the inlet connector 80. In the initial locking engagement, the lock ring 253 with an internal groove 2531 on the inside is loosely hooked onto the flange of the flange base 252 and the flange of the beak flange 81. When the lock ring 253 is locked, the beak-shaped flange 81 (inlet connector 80) and flange base 252 (coupling 25) described above are housed in the internal groove 2531 of the lock ring 253 and pressed, compressing and clamping the middle silicone rubber end 147 (aortic adapter 14), generating a clamping force that firmly connects the inlet connector 80 of the blood pump 62 to the aortic adapter 14.

The clamping force generating mechanism is shown in FIG. 35. The flange base 252 has two steps 2521 and 2522 that generate the clamping force. Before the locking ring 253 closes, the step 2521 should first be engaged with the locking groove 1433 (see FIG. 30) of the convex neck portion 143 of the aortic adaptor 14. This engagement is achieved by first bending and twisting the convex neck portion 143, and then inserting the deformed convex neck portion 143 through the flange base 252, so that the elastic restoring force of the aortic adaptor 14 allows the bent convex neck portion 143 to return to its original circular shape (or shape), and the step 2521 to be engaged with the engagement groove 1433. The height Z of the internal groove 2531 of the locking ring 253 controls the pressing deformation of the adaptor proximal end 147 of the convex neck portion 143. Referring to FIG. 35, in this locked configuration when the locking ring 253 is closed and locked, it can be seen that the gap Z0 (i.e., the thickness Z0 of the pressed adapter proximal end 147 is smaller than the thickness Z3 of the end 147 of the aortic adapter 14 (see FIG. 30), i.e., the thickness Z0 of the pressed adapter proximal end 147 is smaller than the thickness Z3 of the beginning end 147 (Z0<Z3). For the thickness Z0 of the pressed adapter proximal end 147, the following equation can be obtained from FIG. 35:

Z0 = Z - Z1 - Z2...Formula (1)

In formula (1), Z1 and Z2 are the thicknesses for clamping the step 2522 of the mating parts, the mating beak-shaped flange 81 and the flange base 252, as shown in FIG. 35. Typically, the thickness Z3 is greater than the gap Z0, so that the distorted aortic adapter proximal end 147 generates the clamping force required to seal the beak-shaped flange 81 and the annular flange base 252. The distortion of the adapter proximal end 147 is defined as [Z3-Z0]/Z3, and is sufficient to ensure a reliable sealed connection within the range of 10-30%.

Figure 39 shows the interface characteristics of the current connection between the beak 82 and the tapered surface 149 of the adapter. When connected, the beak 82 and its beak leading edge 85 sink into the semi-rigid tapered surface 149, and the aortic adapter 14 has a depth equivalent to the leading edge radius of the beak leading edge 85, typically 30-50 microns. In Figure 39, the dotted line and bracketed numbers indicate the initial contact between the beak 82 and the beak leading edge 85, while the solid line and open bracketed numbers indicate the locked position. Note that the interface discontinuity is reduced by the recession of the tapered surfaces 149, 82 from their original shape (dashed lines). The thickness of the internal groove 2531 controls the tight fit of the connection. As previously mentioned, the elastic aortic adapter proximal end 147 is compressed at a strain of approximately 10-30% to provide the necessary connection strength to counter pulsatile pumping.

The current design of the interface connection between the blood pump 62 and the aortic adapter 14 has two hemodynamic advantages that reduce in situ thrombus formation. First, it has been observed that the conventional docking connection does not actually result in a step or gap type joint discontinuity. Second, the stagnation flow at the boundary of the beak leading edge 85 can be minimized. This allows for faster blood flow through the connection interface and significantly improves the docking connection deficiencies, i.e., the forward or backward steps 101, 102 or gaps 103 that previously occurred at the interface.

The tapered shallow slope 149 of this embodiment is inclined at an inclination angle relative to the flow direction. Such a slope interface design avoids the occurrence of steps or gaps in the joint due to limited manufacturing accuracy or matching eccentricity associated with conventional docking connections. However, this tapered shallow slope 149 has an inherent deficiency in achieving concentric centerline alignment of the mating counterparts. The connection between the aortic adapter 14 and the beak 82 does not have a strict lateral constraint to ensure the mating alignment. In order to concentrically connect the rigid beak 82 and the semi-rigid shallow slope 149, it is important to simultaneously hook the locking ring around the entire periphery of the flange base 252. If the simultaneous locking/locking engagement is not completed, the first locked shallow slope 149 will tend to distort and tilt more than the other free parts or position the remaining contact surface, resulting in an eccentric connection to the pump. Such an eccentric joint is usually the cause of steps and gaps in the interface, which leads to thrombus formation. This deficiency is compensated for by arranging the distal (lower) locking ring contour 259 (FIG. 34A) of the locking ring 253 to include all circumferential contact areas simultaneously with the locking engagement. When locked, the contact edge of the metal beak 82 (in some embodiments) sinks slightly into the shallow bevel 149 of the compressed silicone resin material (in some embodiments) with controlled depth, further reducing the boundary discontinuity when exposed to blood flow. Moderate administration of anticoagulant can significantly reduce or eliminate normal boundary thrombosis.

The deformability of the structure and delivery method of the aortic adapter 14 provide a special design feature of the present invention. Indeed, material elasticity considerations must be carefully incorporated into the current design. In terms of intraoperative safety and long-term reliability, it is difficult to deliver the insertable graft 502 into the aorta through open aortic wall surgery. Therefore, in some embodiments, the material selected for the aortic adapter 14 should have a predetermined memory shape. During device installation, the adapter 14 is first crimped into a small retention mold (crimped adapter 14 as shown in FIG. 40), which ensures fast and safe device implantation. After the curled aortic adapter 14 is positioned at the desired implant site, the retention mold aortic adapter 14 must be released and self-expand to its original memory shape.

Before inserting the aortic adaptor 14, a hole 12-14 mm in diameter should be made in the aortic wall. Care should be taken when making such an access hole to avoid creating edges that may serve as the initiation point for cracks, as the wall must be expanded as the device is inserted. A side-bite aortic perforator, such as that disclosed in U.S. Patent Application No. 17/034036, is an ideal tool for making a large hole in the aorta. A light tap will create an unbroken hole.

The aortic adaptor 14 is morphologically comprised of two connected circular tubes, forming a t-shaped flow connector for performing para-aortic circulation support. The thickness of the catheter wall is usually 1-2 mm, and the material used is a polymer with suitable hardness, such as silicone resin or urethane, with a hardness of Shore A 80-90. The curl-in structure is significantly different from the large stent graft 502 covered with commercial duck nylon or PTFE (polytetrafluoroethylene) fabric. Figure 40 shows the curl/embed configuration of the aortic adaptor 14. (A nickel-titanium alloy metal stent 144 is inserted and deformed together with the aortic adaptor 14 with the polymer matrix). The aortic adaptor 14 is folded by inserting a collapsed catheter into a part of the aortic adaptor catheter 142, as shown in Figure 40, and the t-shaped convex neck portion 143 is accordingly crushed and flattened and sinks into the folded adaptor 14 body. The diameter of the bent aortic adapter 14 is approximately half the diameter of the original deployed circle.

The folded adapter 14 can be held in place by tensioning a rope, as shown in FIG. 40. Three fixing strings can be installed at both edges and at the center of the catheter, but other fixing methods are also possible. FIGS. 41A, 41B, 41C, and 41D show four representative stages of the aortic adapter 14 in the indwelling form. The first stage (see FIG. 41A) shows an initial transmission image of the pre-roll sac (of the aortic adapter 14) passing through the access hole in the form of a curled pre-roll sac when the pre-roll sac (of the aortic adapter 14) is inclined at a certain angle to the axis of the aorta 95. The second stage (FIG. 41B) shows the aortic adapter 14 in the fully inserted form entering the aorta, with the pre-roll sac passing through the access hole and one end rotating and dropping into the access hole. The fully inserted pre-roll sac is then returned to its original position, and the curled neck portion 143 is aligned with the access hole (FIG. 41C). The restrained ties are released to allow the curled, compressed aortic adapter 14 to return to its original shape by elastically self-expanding (FIG. 41D). The released aortic adapter 14 is tightly surrounded by the aortic cavity, as guaranteed by the appropriate oversize ratio selected before placement. The T-shaped convex neck portion 143 can be slightly deformed and the flange base 252 (FIGS. 33-34B) of the coupler 25 can be attached to the T-shaped convex neck portion 143 in preparation for connecting the blood pump 62. To this end, the blood pump connection can be easily achieved by setting the inlet beak 82 on the tapered shallow slope 149, and then closing the two locking rings 253 of the coupler 25 to achieve a firm device connection.

Additional safety measures can be applied to improve hemostasis and stability of the implanted para-aortic blood pump system. Due to the weight of the blood pump 62 and the pumping force from the counter-pulse support, the placement of the blood pump near the aorta necessarily involves lateral forces (perpendicular to the longitudinal direction of the aorta) and torque on the aortic adaptor 14. The external forces associated with this device may affect the long-term reconstruction of the vascular structure at the implant site. A purse-string suture can be placed in the adventitial layer around the hole. The band suture can provide extra tightening of the aortic wall around the inserted adaptor 14 and is used as a protective measure to prevent the enlargement of the access hole. In addition, surgical tape can be wrapped and tensioned on both ends of the aortic adaptor catheter 142 to strengthen the integral bond between the inserted aortic adaptor 14 and the aorta. With proper design and tightening of the annular tape, a double endoleak prevention can be achieved. In some cases, blood pressure may exceed the upper limit without endoleak due to compliance matching. In these extreme circumstances, the surgical tape comes into play and acts as a hard limiter, sealing the ends of the adapter as they are separated, ensuring hemostasis.

As shown in FIG. 42, in some embodiments, the stepwise placement of the aortic adaptor 14 implantation is detailed herein. Before starting the implantation, the aortic adaptor 14 is prepared in a curled pre-rolled pouch/compressed placement format. After exposing the target thoracic artery through a left thoracotomy, a cross clamp distance of approximately 10 cm spanning the implant site is determined. The access hole to be drilled in the aorta is first marked with a peri-hole marker. Then, the sutures of the band are sutured outside the periphery of the hole in the adventitial layer. The aorta can be partially separated from the surrounding connective tissue, and a pair of surgical tapes can be wrapped around the aorta. After the above preparations are completed, the cross clamp and aortic adaptor insertion can be performed. These insertion steps are described in the steps shown in FIG. 17. First, the aorta is cross-clamped to provide an isolated segment without bleeding issues. Then, a customized aortic perforator is used to create a large access hole for the insertion of the aortic adaptor 14. Next, the folded adapter pre-roll sac is inserted and placed on the cross-clamped aortic segment, as shown in Figures 41A, 41B, and 41C. The folded adapter 14 is then released and returned to its original deployed configuration, as shown in Figure 41D. It is then tightened with a ties and tape as additional protection against hypertensive endoleakage. The coupler 25 is then attached and its step 2521 engages with the locking groove 1433 of the convex neck portion 143 of the aortic adapter 14, in preparation for receiving the connected blood pump 62. The self-alignment ability of the coupler 25 described above allows the inlet connector 80 of the blood pump 62 to be accurately positioned and locked to the aortic adapter 14. The remaining implantation procedure is normal, including cross-clip release, blood pump evacuation, and pump start-up assistance. Typically, for a skilled surgeon, the cross-clamp time required to insert the aortic adapter is about 10 minutes. During this cross-clamp period, the abdominal organs are deprived of blood perfusion, which may cause ischemic damage. To reduce potential surgical damage to these organs, femoral-femoral extra-corporeal membrane oxygenation (ECMO) support can be used to perfuse the abdominal organs and lower extremities. However, the decision to use ECMO support is left to the surgeon. Typically, a typical patient can tolerate an ischemic period of 20 minutes.

As described above, according to an embodiment of the present invention, a ventricular assist device is provided that includes a blood pump, a transfer member, and an introduction member. The blood pump includes an axisymmetric elliptical blood sac and stalk assembly, the system including a flexible blood sac, a proximal stalk, and a distal stalk, in which the flexible blood sac is connected to the proximal stalk and the distal stalk to provide a stress relief suspension mechanism. The blood pump further includes a pump housing including a proximal housing and a distal housing, the stress relief suspension mechanism being connected to the pump housing. The blood pump further includes a built-in pressure sensing system in the proximal housing, the pressure sensing system including a blood pump pressure sensor and a pressure sensing chamber filled with an incompressible fluid for pressure transmission. The carrier includes an air chamber and at least one electric wire and tether included in the wall of the carrier, the electric wire and tether being provided in the wall of the carrier. The inlet connects the delivery to the pump housing.

One embodiment of the present invention discloses a pulsating blood pump design that combines the concept of a non-stationary fold line with a long-term blood sac structure, which can substantially extend the durability of a displacement blood pump. A micro pressure-sensing system is also provided that can be used as a reference waveform for real-time pump control and long-term trend analysis based on real-time big data, and disease monitoring and diagnosis based on real-time big data. In addition, the implantable pressure-sensing system is not in contact with blood, so the reliability requirements for building an implantable sensor system are greatly increased.

Embodiments of the present invention have at least any of the following advantages or effects: The inlet connection between the delivery and pump housing provides a compact inlet design for more robust cord and signal transmission and improved fault tolerance. Additionally, the post-pressure introduction design integrates the induction wire, air tube, and blood pump. This compactness attribute is particularly important for implantable devices. It not only simplifies the surgical procedure and reduces the risk of perioperative implantation, but also helps reduce postoperative morbidity associated with delivery infection.

In some embodiments, the introduction member is integrated into the remote housing of the pump housing, and the introduction member has a first portion that is an extension of the remote housing, to which the air chamber, tether, and wires of the power transmission member are connected. The second portion is used in conjunction with the first portion as a hollow connector for the carrier material to achieve the advantages of anatomical adaptability and adaptability to any shape of the implant site.

The present invention provides an aortic adapter assembly capable of implanting a ventricular assist device, the aortic adapter assembly including a T-type flow connector including a catheter insert and a convex neck portion, the catheter insert and the convex neck portion being connected to each other, both of which have smooth surfaces that contact blood, and a metal stent provided in the catheter insert. The T-type flow connector includes a polymer elastomer, and the polymer elastomer is reinforced by a metal stent having a nickel-titanium alloy material. The catheter insert has gradually thin walls at both catheter ends, the ends of the catheter ends being at an appropriate distance to the outermost boundary of the metal stent, the catheter ends having a compliance matching effect with the artery at the implant site, and the proximal end of the convex neck portion is arranged to be connected to an inlet adapter of a blood pump.

The embodiments of the present invention have at least one of the following advantages or effects. The present invention discloses a flow connector assembly that can allow blood flow into and out of a paraaortic ventricular assist device, especially a counterpulsatile blood pump. Unlike conventional flow connectors that adopt many existing surface roughening methods to promote endothelialization and avoid the occurrence of thrombosis maladaptation events, the aortic adapter of the present disclosure adopts the concept of a smooth surface, an insertable pseudograft 502 to construct the flow connector. In addition, a flexible matching design is applied around the end of the inserted catheter, combining the gradually thinning wall characteristics with a thin-walled polymer supported by a superelastic nickel-titanium alloy frame to achieve the requirement of no endoleakage. The abnormal high pressure, high shear and low recirculation flow phenomena associated with paraaortic counterpulsatile pumping are included in the artificial surface to insert the catheter. Therefore, the hemodynamic effects and risk factors of pathological devices are almost eliminated, and the long-term vascular maladaptation events such as endothelial cell erosion, lipid infiltration, smooth muscle cell proliferation, vascular stenosis, and arterial wall peeling are obviously reduced. In order to achieve a good connection between the semi-rigid flow adapter and the blood pump, the present disclosure provides a quick-fitting coupling. The coupler has a self-aligning interface design, minimizing step and gap discontinuities and reducing the possibility of thrombotic failure at the interface junction. In cooperation with the invention of the aortic adapter, a specially designed deployment method can be provided to ensure a fast and safe deployment process. The curled aortic adapter is pre-wrapped/deployed, and the overall size is reduced to half of the initial size. The pre-rolled pouch adapter can be easily inserted into the aorta at the implant site and expands to the initial position by itself, forming an affixed flow connector without the risk of endoleakage. It not only reduces the risk of implantation during surgery, but also helps to reduce postoperative morbidity associated with device flow and vascular maladaptation at the implant site.

Ordinal numbers, e.g., "first", "second", etc., in this specification and in the patent application, have no sequential relationship to one another and are used only to distinguish between two different components that have the same name.

The embodiments described herein are intended to be illustrative, and variations of those embodiments will become apparent to those of ordinary skill in the art after reading the foregoing description. Accordingly, the present invention includes all modifications and equivalents of the subject matter described within the scope of this patent application.

10, 20, 30, 40, 90 Para-aortic blood pump device 11, 21, 31, 41 Battery supply system 12, 22, 32, 42, 52, 62, 92 Blood pump 14, 24, 34, 44, 54, 64, 94 Aortic adapter 16, 26, 36, 46 Drive catheter 18, 28, 38, 48, 78, 98 Drive unit 25, 45, 65 Coupler (coupling adapter)
33, 43, 93 Drive catheter interconnect 37, 47, 57, 67, 97 Distal drive catheter 39, 49, 99 Proximal drive catheter 52h, 62h Rigid housing 523, 623 Proximal end housing 525, 625 Remote housing 526 Sac membrane 529 Blood sac B Blood chamber A Air chamber 527 Pressure sensing mechanism 528 Pressure sensing chamber 530 Proximal end port 540 Remote port 545, 645 Catheter end 101 Step 102 Step 103 Gap 141 Blood contacting surface 142 Aortic adaptor catheter 143 Convex neck 1431 Neck body 1432 Extension 1433 Card slot 144 Metal stent 1441 Outermost boundary 145 Catheter end 146 Outer diameter 147 Adaptor proximal end 148 Inner diameter 149 Shallow bevel 252 Flange base 221 Step 222 Step 252S1 First side 252S2 Second side 253 Lock ring 231 Internal groove 254 Hinge 255 Spring coil 256 Hinge joint 257 Leaf spring latch 571 Plate 258 Ramp 259 Lock ring profile 50 Connector 501 Adapter 502 Implant 503 Connector 504 Remote 63 Introducing material 62A Inversion membrane 621 Groove 623 Proximal end housing 625 Remote housing 6251 Inlet connector 6252 First end 6253 Second end (23A) 6254
627 Pressure sensing mechanism 6271 Sensor 6272) First space 6273 Second space 6274 Electrode 628 Pressure sensing chamber 6281 First arm 6282 Second arm 629 Blood sac 630 Proximal end stalk 6301 Sac stalk septum 631 First section 632 Second section 640 Remote stalk 643 Neck 650 Stem assembly 66 Exhaust port 661 Passage 67 Carrier 6701 Air lumen 6702 Electrical wire 61) Proximal end 672 Anchor adapter 673 Air tube 674 Coil 675 Outer layer tube 676 Tether 677 Silicone sheath 678 Rigid actuator connector 6781 Electrode 679 Hollow connector 60 Artery 6291 Proximal end 6292 Remote 6293 Three-valve type 6294 Fold 62C Centerline 71 Battery hatch 73 User interface panel 75 Drive catheter socket 77 External power socket 79 Vent 80 Inlet connector 81 Beak flange 82 Beak 83 Connector body 84 Inner diameter 85 Beak leading edge 86 Holes 971, 991 Drive catheter internal segment 973, 993 Drive catheter external segment 95 Aorta 96 Drive catheter EX Exit site OP Opening 110 Electromechanical actuator 120 Motor controller unit 130 Microcontroller unit 140 Power management unit 150 Battery module 170 User interface module 190 Bedside monitor T-201 Low speed recirculation area T-202 Impact point AB1 Adapter AB2 Adapter

Claims (21)

1. An aortic adaptor assembly for use in an implantable ventricular assist device, comprising:
a T-shaped flow connector having a catheter insertion portion, a convex neck portion, and a truss provided on the catheter insertion portion;
the catheter insertion section and the convex neck section are connected, and both of the catheter insertion section and the convex neck section have smooth surfaces that contact blood;
the T-shaped flow connector comprises a polymer elastomer and is reinforced by the truss comprising a nickel-titanium alloy material;
The catheter insert has two catheter ends with gradually thinning walls, and has an appropriate distance between the tip of the catheter end and the outermost boundary of the truss, the catheter end has a compliance matching effect with the artery at the implant site, and the proximal end of the convex neck portion is used to be connected to an inlet adapter of a blood pump.
Aortic adapter assembly. 2. The aortic adapter assembly of claim 1,
An aortic adapter assembly, wherein the truss is co-injected with a polymer elastomer of the T-shaped flow connector and fitted into the wall of the catheter insert of the T-shaped flow connector. 3. The aortic adapter assembly of claim 2,
The aortic adaptor assembly, wherein the polymeric elastomer of the T-shaped flow connector comprises a silicone resin material. 2. The aortic adapter assembly of claim 1,
The aortic adapter assembly, wherein the polymeric elastomer of the T-shaped flow connector is cast polyurethane. 3. The aortic adapter assembly of claim 2,
An aortic adaptor assembly, wherein the structural compliance of the nested truss and the structural compliance of the polymeric elastomer are approximately equal. 2. The aortic adapter assembly of claim 1,
The aortic adapter assembly wherein the end of the tapered catheter is sharp. 2. The aortic adapter assembly of claim 1,
the convex neck portion has a shallow slope to mate with an inlet adapter of the blood pump;
An aortic adaptor assembly, wherein the inner diameter of the convex neck portion is slightly smaller than the inner diameter of the inlet adaptor of the blood pump. 8. The aortic adapter assembly of claim 7,
The shallow bevel is inclined relative to the centerline of the catheter insertion section. 2. The aortic adapter assembly of claim 1,
The aortic adaptor assembly, wherein the truss has a plurality of wavy structures. 2. The aortic adapter assembly of claim 1,
an aortic adaptor assembly, wherein the convex neck portion includes a neck body and an extension portion provided on the neck body, the extension portion protruding from the neck body, and a maximum inner diameter of the extension portion is greater than a maximum inner diameter of the neck body. 11. The aortic adapter assembly of claim 10,
An aortic adapter assembly, wherein when the convex neck portion is joined to the inlet adapter of the blood pump, the extension portion is closely attached to the inlet adapter and the neck portion body is fitted around the inlet adapter. 11. The aortic adapter assembly of claim 10,
Further comprising a coupler;
The coupler includes a flange base, a pair of locking rings rotatably mounted to the flange base, and a latch on one of the locking rings for locking the locking ring, the locking ring having an internal groove that clamps the T-shaped flow connector with controlled compression to seal the T-shaped flow connector. An aortic adapter assembly. 13. The aortic adaptor unit assembly of claim 12,
An aortic adapter assembly, wherein each of the locking rings has a flange profile that allows the locking rings to be co-engaged against an edge of the flange base. 13. The aortic adaptor unit assembly of claim 12,
The latch is made of a spring strip to ensure that the coupler is in a locked state so that there is no risk of it coming loose inadvertently. 13. The aortic adapter assembly of claim 12,
The aortic adaptor unit assembly further includes a hinge head attached to the flange base, the lock ring being pivotally attached to the hinge head and rotatably relative to the hinge head and the flange base. 16. The aortic adapter assembly of claim 15,
The aortic adaptor assembly, wherein the hinge head is located on a first side of the flange base and the latch is located on a second side of the flange base, the first side and the second side of the flange base being opposed. 17. The aortic adapter assembly of claim 16,
The latch has a slot that can engage with a beveled portion provided on the opposing lock ring of the aortic adaptor unit assembly. 17. The aortic adapter assembly of claim 16,
The aortic adaptor unit assembly, wherein the flange base has a generally circular configuration and each of the locking rings has an arcuate configuration. 1. A blood pump device as an implantable ventricular assist device, comprising:
a blood pump having an inlet adapter with a tapered beak;
an aortic adapter assembly,
The aortic adapter assembly includes:
a T-shaped flow connector including a catheter insertion section and a convex neck section, the catheter insertion section and the convex neck section being connected to each other, the catheter insertion section and the convex neck section both having smooth surfaces that contact blood;
a truss provided in the catheter insertion portion,
the T-shaped flow connector comprises a polymer elastomer and is reinforced by the truss comprising a nickel-titanium alloy material;
the catheter insert has two catheter ends with gradually tapered walls, a suitable distance between the tips of the catheter ends and the outermost boundary of the truss, the catheter ends having a compliance matching effect with an artery at an implant site, and a proximal end of the convex neck portion is connected to an inlet adapter of a blood pump;
The tapered beak mates with a proximal end of the convex neck portion. 20. The blood pump apparatus of claim 19,
A blood pump apparatus, wherein the inner diameter of the inlet adapter is slightly larger than the inner diameter of the convex neck portion of the T-shaped flow connector. 1. A method of implanting a flow connector assembly, comprising:
2. An aortic adapter assembly according to claim 1,
the T-shaped flow connector is compressible from an initial configuration to a compressed, deployed configuration;
the T-shaped flow connector is crimped into a compressed, indwelling configuration, the size of the T-shaped flow connector in the compressed, indwelling configuration being at least half the diameter of the T-shaped flow connector in an initial configuration;
The T-shaped flow connector of the compressed indwelling type is inserted into an artery having an access hole drilled in the arterial wall;
Releasing the T-shaped flow connector from its compressed, deployed configuration, and allowing the released T-shaped flow connector to self-expand to its initial deployed configuration;
A method of implanting a flow connector assembly that couples the deployed T-shaped flow connector to a ventricular assist device.

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