Blood pump device having a leak-free aortic adapter assembly

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 coupler and installation method to provide insertable flow communication between the ventricular assist device and the systemic circulation.

Description

Technical Field

[0001] The present invention relates to a ventricular assist device (VAD), and more particularly to a left ventricular assist device (LVAD) based on the principle of counterpulsation support, a blood pump device having a leak-free aortic adapter assembly, and a method for implanting the device.

Background Art

[0002] Mechanical circulatory support systems, particularly left ventricular assist devices (LVADs), are evolving into treatment standards for emergency treatment of end-stage heart failure. Due to the mechanical design of the blood pump, LVAD systems can generally be divided into continuous flow pumps and pulsatile flow pumps. Continuous flow devices are constructed based on rotating machinery propelled by axial or centrifugal flow impellers. On the other hand, pulsatile flow devices are designed using displacement blood pumps, usually receiving with a diaphragm blood sac to allow the inflow of blood into the pump and / or the outflow from the pump.

Summary of the Invention

Problems to be Solved by the Invention

[0003] The inflow / outflow tube (cannula) is an artificial flow path for connecting the LVAD system to the human circulatory system in series or in parallel. Historically, compared with blood pump actuators, the design of the flow tube has received less attention. Surprisingly, a large number of postoperative complications have been discovered and are related to the failure of the inflow / outflow tube. Among them, adverse events such as misplacement of the cannula, occlusion, thrombosis, and resulting torsional flow cause pump thrombosis and thromboembolism. In the implant of a rotary pump, mainly the inflow tube is connected to the cardiac chamber, and the outflow tube is connected to the ascending aorta or the descending aorta to establish the flow path. So far, most drain tubes of rotary pumps have adopted dacron grafts or similar materials, for example, made of flexible woven materials, and the connection to the artery adopts the end-to-side anastomotic suturing method. This end-to-side anastomosis method is technology-dependent, and complications caused by poor suturing or incorrect aortic flow may cause thrombosis in downstream organs, stroke, thromboembolism, and infarction.

[0004] Rotary pumps provide a cardiac output of complete support (4 - 10 L / min) and are currently applied to patients with end-stage heart failure. For patients with mild conditions, it is appropriate to use a less invasive implantable LVAD with partial support (2 - 3 L / min). The partial support LVAD aims to intervene in heart failure early to improve the treatment outcome. Among such new trends of partial support and less invasive LVAD implants, the counterpulsating blood pump has always been the main focus. This is because it has been proven to increase blood perfusion of the myocardium and major organs during the systolic and diastolic phases of the heart. The counterpulsating pump must follow strict time control criteria and refer to the heart rate. Usually, ventricular contraction and detachment start at the end of diastole, but enhanced organ perfusion is performed at the time of aortic valve closure (pacing cut-out on the aortic pressure waveform). In fact, the provided therapeutic effects are in two aspects. The first is to reduce ventricular contraction and decrease myocardial oxygen consumption, and the second is to increase diastolic pressure and increase perfusion of the myocardium, brain, and major organs.

[0005] The supporting effect of counterpulsation has been clinically proven through peripheral subcutaneous transport of the balloon pump within the aorta. However, peripheral transport of the descending aorta is often troubled by vascular complications at the insertion site and is difficult for long-term use. In the treatment of heart failure, it is highly significant to extend counterpulsation support from acute (less than 1 week) to long-term (months to years). To meet this requirement, new surgical methods and equipment innovation are necessary to achieve this long-term counterpulsation support goal.

Means for Solving the Problems

[0006] The counterpulsation support of the present invention is performed by left thoracotomy, in which the device is connected to the thoracic artery. This counterpulsation invention, called the paracorporeal blood pump, is implanted through an access hole formed in the descending aortic wall from the lateral pump connection part. In contrast to the intra-aortic balloon pump, the paracorporeal blood pump does not occlude the blood flow path, so it is more flexible in timing control of counterpulsation. According to animal studies, the hemodynamic support effect of aortic balloon counterpulsation is superior to that of the intra-aortic balloon pump. However, it is unclear whether the support of the paracorporeal pump causes long-term complications. In fact, the problem is to construct a fluid cannula, implant it in a less invasive way, and safely establish fluid communication over a long period between the artery connected to the pump.

[0007] In the implantation of a paracorporeal blood pump, it is necessary to create a prosthetic catheter to permit blood inflow and outflow. Usually, it is necessary to rapidly fill the pump to effectively reduce the afterload of the left ventricle and generate strong vortex washing to prevent pump thrombosis. Such rapid pump filling often experiences transient low pressures that cause bends and depressions in the flow tubes. Therefore, conventional Dacron graft type cannulation is not possible. This is because woven catheters cannot withstand compressive forces and the depressed grafts may block inflow during the pump filling stage. In addition, the high and low pressure pulses due to rapid pump jet and filling bring bleeding problems to the suture site, especially in the acute phase when the surgical anastomosis has not yet healed. Therefore, it is extremely important to design the flow tube to overcome the specific flow characteristics associated with aortic counterpulsation.

[0008] The hemodynamic characteristics of the paracorporeal blood pump flow are not physiological. The blood flow in the blood pump connected by inhalation or ejection is not laminar in the arterial direction as commonly seen in the natural aorta. This roll-type artificially generated pump flow is highly turbulent and complex. During the filling stage of the pump, the blood flow whirls rapidly into the blood pump, and diversion zones and recirculation zones occur at the turning angles. During the pump jet stage, the blood flow accelerates and enters the descending aorta. The characteristics of the blood flow are that the impact flow has ultra-high pressure and shear stress and acts on the opposing aortic wall regions. The flow characteristics induced by this device include low-speed recirculation regions and high-pressure, high-shear impact jets, which can cause endothelial cell erosion, lipid infiltration, and smooth muscle cell proliferation, and may cause long-term aortic wall stenosis, thrombosis, or aortic dissection due to hypertension. Therefore, the design of a long-term implantable counterpulsation device requires innovative prosthetic flow cannulation to avoid or mitigate the pathophysiological flow phenomena caused by the above-mentioned device, the resulting vascular maladaptation and thrombosis or adverse events of aortic dissection.

[0009] The safety of surgical anastomosis is another requirement for the flow tube design of the intra-aortic counterpulsation pump. When receiving excessive high-pressure fluctuations related to counterpulsatile flow, traditional graft suturing may encounter challenging bleeding complications. High and low-pressure circulations are the main driving forces for material fatigue failure, especially when lesions (atherosclerosis and calcification), degeneration (thinning of the wall thickness), or aging (hardening of the wall) occur in the implanted aorta. It should be noted that the aortic wall structure can adapt to the applied stress conditions. The cells and tissues around the implantation site are reconstructed along with the surgical injury healing process and further develop due to the non-physiological mechanical environment induced by the equipment. Even if short-term transplantation is successful, long-term transplantation failure due to poor cell and morphological adaptation of the blood vessel wall may not be avoided during the postoperative process.

[0010] So far, all feasible solutions for connecting artificial grafts on the arterial side and the opposite side have been based on flexible fabric tubes by suture methods. In the medical device industry, there is no long-term cannulation solution that can solve the problems related to the flow rate and pressure fluctuation levels by the counterpulsatile aortic bypass blood pump. For more than fifty years, the application of the intra-aortic balloon pump has already achieved a clinical counterpulsation effect in the medium and short term (from several days to about one month). This typical air energy transportation for driving the inflation / outgassing of the balloon is realized using a thin catheter sent percutaneously from a remote peripheral artery, and the lumen diameter is within the range of 6 - 10. Most of the arteries at such transportation sites are occluded, which usually leads to serious bleeding complications and ischemia of the downstream limb or arm, preventing long-term use by percutaneous counterpulsation. Aortic transposition is a new method aimed at extending the counterpulsation treatment to a longer time range. However, long-term aortic balloon counterpulsation requires constructing a flow communication device to solve the above implant problems. Such a flow cannula must be non-fracturing, easy to implant, bloodless, and biocompatible without causing pathological vascular maladaptation. The present invention endeavors to meet all population needs by proposing an insertable aortic adapter, which will be described later.

[0011] Embodiments of the present invention are used in an implantable ventricular assist device and are blood pump devices having a leak-free aortic adapter assembly including a T-shaped flow connector, wherein the T-shaped flow connector catheter includes an insertion portion, a convex neck portion, and a truss. The catheter insertion portion and the convex neck portion are connected, and both the catheter insertion portion and the convex neck portion have smooth surfaces that come into contact with blood. The T-shaped flow connector has a polymer elastomer and is reinforced by the truss made of a nickel-titanium alloy material. The catheter insertion portion has a wall in which the two catheter ends gradually become thinner, has an appropriate distance between the tip of the catheter end and the outermost boundary of the truss, and the catheter end has a compliant matching effect on the artery at the implant site. The proximal end of the convex neck portion is used to be connected to the inlet adapter of the blood pump.

[0012] In one embodiment, the truss is co-injected with the polymer elastomer of the T-shaped flow connector and is fitted into the wall of the catheter insertion portion of the T-shaped flow connector.

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

[0014] In one embodiment, the polymer elastomer of the T-shaped flow connector is a cast polyurethane aortic adapter assembly.

[0015] In one embodiment, the structural compliance of the fitted truss and the structural compliance of the polymer elastomer are substantially equal.

[0016] In one embodiment, the gradually thinning ends of the catheter are sharp.

[0017] In one embodiment, the convex neck portion has a shallow slope so as to fit 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.

[0018] In one embodiment, the shallow slope is inclined with respect to the center line of the catheter insertion portion.

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

[0020] In one embodiment, the convex neck portion includes a neck portion main body and an extension portion provided on the neck portion main body. The extension portion protrudes from the neck portion main body, and the maximum inner diameter of the extension portion is larger than the maximum inner diameter of the neck portion main body.

[0021] In one embodiment, when the convex neck portion is joined to the inlet adapter of the blood pump, the extension portion is in close contact with the inlet adapter, and the neck portion main body is an aortic adapter assembly provided around the inlet adapter.

[0022] In one embodiment, it further includes a coupler. The coupler includes a flange base, a pair of locking rings rotatably attached to the flange base, and a latch provided on one of the locking rings for locking the locking ring. The locking ring has an internal concave groove, and the concave groove clamps the T-shaped flow connector by controlled compression to seal the T-shaped flow connector.

[0023] In one embodiment, each of the locking rings has a flange contour, and the flange contour can simultaneously join the locking ring to the edge of the flange base.

[0024] In one embodiment, the latch is made of a spring piece, ensuring that the coupler is in a locked state so as to prevent accidental detachment.

[0025] In one embodiment, the coupler further includes a hinge head provided on the flange base, and the locking ring is pivotally attached to the hinge head so as to be rotatable with respect to the hinge head and the flange base.

[0026] In one embodiment, the hinge head is located on a first side of the flange base, the latch is located on a second side of the flange base, and the first side and the second side of the flange base face each other.

[0027] In one embodiment, the latch has a shear that can engage with an inclined portion provided on one of the opposing locking rings.

[0028] In one embodiment, the flange base has a substantially circular structure, and each locking ring has an arcuate structure.

[0029] In one embodiment, a blood pump device as an implantable ventricular assist device, including a blood pump having an inlet adapter with a tapered beak, and a aortic adapter assembly, the aortic adapter assembly includes a catheter insertion portion and a convex neck portion, the catheter insertion portion and the convex neck portion are connected, the catheter insertion portion and the convex neck portion are both T-shaped flow connectors having smooth surfaces that contact blood, and the Ka a truss provided on the catheter insertion portion, the T-shaped flow connector has a polymer elastomer and is reinforced by the truss made of a nickel-titanium alloy material, the catheter insertion portion has a wall where two catheter ends of the catheter insertion portion gradually become thinner, there is an appropriate distance between the tip of the catheter end and the outermost boundary of the truss, the catheter end has a compliant matching effect on the artery at the implant site, the proximal end of the convex neck portion is connected to the inlet adapter of the blood pump, and the tapered beak is aligned with the proximal end of the convex neck portion.

[0030] 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.

[0031] Another embodiment of the present invention is a method of implanting a flow connector assembly, comprising providing the aortic adapter assembly according to claim 1, wherein the T-shaped flow connector can be compressed from an initial form to a compressed and retained form, the T-shaped flow connector is pressed against the compressed and retained form, the size of the T-shaped flow connector in the compressed and retained form is at least half of the diameter of the T-shaped flow connector in the initial form, the compressed and retained T-shaped flow connector is inserted into an artery having an access hole pierced in its arterial wall, the T-shaped flow connector in the compressed and retained form is released and deployed, and the released T-shaped flow connector can self-expand to its initial form of deployment, and providing a method of implanting a flow connector assembly that couples the deployed T-shaped flow connector to a ventricular assist device.

Brief Description of the Drawings

[0032]

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Best Mode for Carrying Out the Invention

[0033] Four embodiments for realizing the aortic blood pump device according to the present invention are provided below, as follows.

[0034] Refer to FIG. 1. It is a schematic diagram of a blood pump device near the aorta according to the first embodiment of the present invention. The blood pump device 10 near the aorta includes a blood pump 12, an aortic adapter 14, a drive catheter 16, and a drive 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. These two chambers are separated by an oval flexible film, and this film body is suspended and fixed by a pair of stress-relief stems connected to the pump housing. The pressure sensor is installed in the pump housing of the blood pump 12, monitors the blood pressure in this blood pump 12, and generates an electronic blood pressure signal. This aortic adapter 14 is a valveless, T-manifold shaped conduit flow manifold that connects this blood pump 12 and the human aorta. In the first embodiment, this aortic adapter 14 and the blood pump 12 are integrally formed, have a seamless blood contact surface, and the blood pump 12 and the human aorta are connected via the aortic adapter 14. This aortic adapter 14 is formed of a flexible material and can be deformed when inserted into and transferred through a circular hole formed in the aortic wall. This aortic adapter 14 can expand and deploy itself after inserting into the aorta, and has sufficient strength against the radial compression contact force of the aortic lumen applied to the adapter wall surface from an oversize fitting. This drive catheter 16 is connected to the housing of the blood pump 12, supplies an air pressure pulse to the blood pump 12, and transmits the blood pressure signal received from the pressure sensor. This drive unit 18 is connected to the drive catheter 16 so as to receive the transmitted electronic blood pressure signal. This drive unit 18 includes an electromechanical actuator that generates an air pressure pulse based on the electronic blood pressure signal, and is sent to the blood pump 12 via the drive catheter 16. This wearable actuator 18 provides an air pressure pulse control rule linked to the heartbeat rhythm, and drives and operates the ejection and filling of the blood in the implanted blood pump 12.

[0035] The aforementioned drive unit 18 includes a battery power supply system 11 and a backup battery power supply system (the same or similar to the battery power supply systems 21, 31, 41 in Figures 2 to 4 and Figure 8 to be described later). Among them, the backup battery power supply system ensures continuous power supply to the drive unit 18. When the patient does not need to move, power can also be supplied to this drive unit 18 through an AC adapter for convenience. In addition, this device has a clinical monitor, as shown in Figures 2 to 4. It is connected to the drive unit 18 and can provide a user interface for clinicians to display device monitoring and diagnostic messages and access drive unit parameters for initial startup settings of patient data and optimization of specific treatment operation mode settings.

[0036] Figures 1 and 2 show the designs of two different blood pumps 12, 22 systematically coupled to the same drive catheters 16, 26 and drive units 18, 28. Figure 2 is a schematic diagram of the paracorporeal blood pump device 20 according to the second embodiment of the present invention. The difference between the second embodiment and the first embodiment of the present invention is that the paracorporeal blood pump device 20 according to the second embodiment further includes a coupler 25 or a coupling adapter. The blood pump 22 and the aortic adapter 24 of the second embodiment are not integrally formed and are detachable, and a coupler 25 is provided for coupling the blood pump 22 to the aortic adapter 24. The coupler needs to be carefully designed to minimize the discontinuity of the connection interface. During the implantation of the device, the aortic adapter 24 is first fed into the aorta through the indwelling hole notched in the aortic wall. The device can use a specially developed implantation tool to attach the coupler 25 around the T-neck portion of the aortic adapter 24 so that the blood pump 22 can be connected to the coupler 25. After the blood pump 22 is placed in the thoracic cavity, the blood pump 22 and the aortic adapter 24 are firmly locked and integrated via the coupler 25. Such a detachable design of the blood pump 22 and the aortic adapter 24 has advantages in terms of implantation surgery and after-surgery. 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 not interfered with by the pump body. Also, after the implantation surgery, if a failure occurs in the pressure sensor or the blood sac ruptures and an emergency surgical replacement is required, the blood pump can be removed and replaced. In this regard, the detachable blood pump design of the second embodiment is advantageous. The aortic adapter can be left in the aorta without being removed, avoiding the cumbersome and dangerous redo surgery involved in the removal of the aortic adapter.

[0037] Refer to FIGS. 1 and 3. These are schematic diagrams of the aortic blood pump devices 10 and 30 of the first and third embodiments of the present invention. The difference between the third embodiment and the first embodiment of the present invention is that the drive catheter 16 of the first embodiment is replaced by the distal drive catheter 37, the drive catheter interconnect 33, and the proximal drive catheter 39 of the drive catheter 36 of the third embodiment. The distal drive catheter 37 is connected to the drive catheter interconnect 33 so as to transmit the electronic blood pressure signal acquired from the pressure sensor and the pneumatic pulse transmitted from the drive unit 38. Also, since the drive catheter controller and the vibrator (for alarm warning purposes) built into the drive catheter interconnect 33 were originally built into the drive unit 18 of the first embodiment, the drive unit 18 of the first embodiment has the drive catheter controller and the vibrator added (compared with the drive unit 38 of the third embodiment). The drive catheter controller is used to process the electronic blood pressure signal, and the vibrator is used to provide an audible alarm and tactile feedback. That is, the mechanical power transmission of the blood pump, and the analog / digital signal conversion and alarm notification realized by the drive catheter 16 and the drive unit 18 in the first embodiment are basically the same as those realized by the distal drive catheter 37, the drive catheter interconnect 33, and the proximal drive catheter 39 in the third embodiment.

[0038] The first embodiment has a more concise arrangement design of the drive catheter. Also, since the electronic signal processing device is installed inside the drive unit, it is possible to minimize the environmental contamination (condensation of water or moisture) in pressure signal measurement related to the drive catheter interconnector 33 and the risk of air leakage at the joint. However, such a long-type drive catheter is liable to suffer contact damage such as abrasion, kinking, and cuts due to contact with foreign objects during daily activities. If there is serious damage to the drive catheter 16 of the first or second embodiment, it may be necessary to replace the blood pump surgically, both electronically and mechanically. Considering the risk of re-surgery and the medical costs associated with it, this is highly undesirable. The third or fourth embodiment reduces the drawbacks of blood pump replacement related to such damage of the drive catheter by adopting an intermediate connector (drive catheter interconnector). Generally, the distal drive catheter 37 has a shorter length exposed outside the body, and the skin dressing and the cover of the patient's gown are better protected via 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 relying on surgery. Also, in the third or fourth embodiment, since the analog-digital signal conversion has already been completed within the circuit of the drive catheter interconnector 33, it is not affected by electromagnetic interference. In the third or fourth embodiment, since the digital signal transmission in the proximal drive catheter 39 is not sensitive to electromagnetic interference, the fidelity of the pressure signal can be ensured better.

[0039] Refer to FIGS. 3 and 4. These are schematic diagrams of the aortic blood pump devices 30 and 40 according to the third and fourth embodiments of the present invention, respectively. The difference between the third embodiment and the fourth embodiment is that the aortic adapter 34 and the blood pump 32 of the third embodiment are integrally formed, while 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. For the same blood pump 42, aortic adapter 44, and coupling 45 in the embodiments, duplicate explanations are omitted here. The distal drive catheter 47, drive catheter interconnect 43, and proximal drive catheter 49 included in the drive catheter 46 are the same as the distal drive catheter 37, drive catheter interconnect 33, and proximal drive catheter 39 of the drive catheter 36.

[0040] Refer to FIG. 5. This is a schematic diagram of an aortic blood pump device worn on a human body according to an exemplary embodiment of the present invention. The aortic blood pump device 90 includes a blood pump 92, an aortic adapter 94, a drive catheter in-body segment 991, a drive catheter out-of-body segment 993, and a drive unit 98. In other embodiments, the aortic blood pump device further includes a coupler. The portion of the aortic blood pump device 90 embedded in the human body includes the blood pump 92, the aortic adapter 94, and the drive catheter in-body segment 991. In other embodiments, the 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 an appropriate position on the human body epidermis. The out-of-body segment of the aortic blood pump device 90 includes the drive catheter out-of-body segment 993 and the drive unit 98. With the exit site EX as the boundary, a segment of the in-body segment 991 of the drive catheter is covered with fabric velour and used to promote tissue ingrowth into the subcutaneous tissue to achieve infection control. The embedded velour portion is optimally placed 2 to 5 centimeters subcutaneous from the exit site EX. The drive unit 98 is a wearable or portable device.

[0041] Refer to FIG. 6. It is a schematic diagram of a periaortic blood pump device worn on a human body according to an exemplary embodiment of the present invention. The periaortic blood pump device 90 includes a blood pump 92, an aortic adapter 94, a distal drive catheter 97 (including a drive catheter in-body segment 971 and a drive catheter out-of-body segment 973 outside the human body), a drive catheter interconnector 93, a proximal drive catheter 99, and a drive unit 98. In other embodiments, the periaortic blood pump device 90 further includes a coupler. The portion of the periaortic blood pump device 90 embedded in the human body includes the blood pump 92, the aortic adapter 94, and the drive catheter in-body segment 971. In other embodiments, the periaortic blood pump device further includes a coupler. In a surgical operation, the aortic adapter 94 is mounted in the aorta 95 to create an exit site EX of the drive catheter at an appropriate position on the human body epidermis. The portion of the periaortic blood pump device 90 located outside the human body includes the drive catheter out-of-body segment 973, the drive catheter interconnector 93, the proximal drive catheter 99, and the drive unit 98. The exit site EX is the boundary of the drive catheter, and the distal drive catheter is divided into a drive catheter in-body segment 971 covered with velvet for infection control and a drive catheter out-of-body segment 973. The drive unit 98 is a wearable or portable device.

[0042] Hereinafter, the embedded subsystem will be described in detail.

[0043] Implantation is achieved via a relatively small thoracotomy through left thoracotomy using less invasive surgical (LIS) techniques. For example, to permit implantation of the aortic adapter and blood pump, a thoracotomy is made through the space between the 7th ribs as the main opening. Two other small incisions are made between the 6th and 8th ribs to introduce proximal and distal aortic clamps. The area between the aortic clamps is made such that the aortic adapter can be implanted through a hole opened in the aortic wall. The aortic adapter is flexible and can be curled before implantation and reduced to a small delivery shape. When the implantation into the aorta is complete, the aortic adapter automatically springs back to its original shape and has super-adhesion 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 prevent buckling of the wall. As candidates for the material of the aortic adapter, silicone or polyurethane elastomer materials can be selected, or a reinforced polymer structure or metal material can be implanted to strengthen the structural strength.

[0044] The following further describes each aortic adapter and its functional requirements in the above embodiments.

[0045] Hemodynamically, the aortic adapter provides the role of connecting the blood flow 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 mechanism base for fixing the blood pump to the aortic adapter. The structure of the aortic adapter is elastic but anti-buckling and requires sufficient strength to withstand internal blood pressure and external contact forces, which are generated when the blood pump contacts the surrounding lung tissue or due to respiration and diaphragmatic movement of the chest.

[0046] The aortic adapter 54 is implanted within the aorta, and the boundaries between its two catheter ends (catheter ends 545, catheter end 645) and the aortic lumen form a host / graft interface in the blood flow (see FIGS. 13B and 14B). To morphologically and elastically minimize the discontinuity of the host / graft interface, the two catheter ends (catheter ends 545, catheter end 645) are arranged to have a flared inner surface contour and a continuously tapering wall thickness distribution. This catheter end design minimizes the step at the boundary and incorporates the required compliance matching effect during connection into the design of the aortic adapter. Therefore, the possibility of thrombosis at the boundary can be significantly reduced. This is because the rate of clot aggregation at the boundary is slower than the rate of natural thrombolysis provided by the human aortic endothelium. Also, the gradually thinning catheter wall structure softens (fits) the catheter ends (catheter ends 545, catheter end 645), and the catheter tip expands and contracts in response to the pulse pressure, constituting a dynamic sealing effect so that blood cells are not caught in the gap at the boundary which is usually the source of thrombosis.

[0047] Hereinafter, the drive units of the above-described respective embodiments will be further described.

[0048] FIGS. 7 and 8 are perspective views of the right and left sides of the drive unit 78. The internal module of this compact drive unit 78 includes 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 vents 79 as shown in FIGS. 7 and 8.

[0049] On the user interface panel 73 of the drive unit 78, important messages during operation, as well as warnings regarding device malfunctions and aortic pressure conditions, are displayed. When the remaining amount of the main battery is insufficient, the main battery can be replaced by the battery hatch 71. If the patient is lying in the bed and can use the power from the wall-mounted socket for a long time, power is supplied to the drive 78 through the connection of the external power socket 77 using a cable. One end of the proximal drive catheter 99 and the extracorporeal segment 993 of the drive catheter are connected to the drive 78 via the drive catheter socket 75, and a voltage sensor signal and a pneumatic pulse are connected via the drive catheter socket 75. On the opposite side of the drive unit 78, a pair of ventilation openings 79 are attached so that outside air can pass through the inside of the drive unit 78 for cooling purposes.

[0050] The drive unit 78 is externally coupled to a clinical monitor, and the clinical controller can collect and display real-time clinical waveform data and store patient data for long-term status monitoring and diagnosis. Also, the clinical monitor unit can provide a user interface for clinicians and patients, display the monitoring / diagnostic information of the device, be used for setting the drive unit parameters, start the drive unit 78 for the first time, and optimize and set the patient-specific individual circulatory assist operation mode.

[0051] In the embodiment of the present case, the EMA is an air actuator and 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 injection and blood pump filling.

[0052] The air actuator device is located inside the drive unit and is carried by the person receiving treatment with this 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 shaft and a ball nut. The piston is fixed to the end of the ball screw shaft, and the rotation of the nut linearly reciprocates the screw shaft. 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 through the change of the clockwise and counterclockwise directions, a linear reciprocating stroke motion in the cylinder of the screw post and the piston is achieved. The piston stroke drives the air in the drive cylinder to the blood pump through the drive lead wire, completing the bleeding and congestion operations of the blood sac.

[0053] Driving the blood pump using air as the medium faces two problems. One is the problem of gas leakage, and the other is the problem of moisture condensation due to blood penetrating the blood sac wall. The former damages the blood injection and congestion performance of the blood pump and the power consumption of the motor, and the latter causes the risk of bacteria generating in the driving code due to moisture. To solve these two problems, the air actuator of this ventricular assist device is particularly equipped with a pressure balance valve device attached to the cylinder wall. With this valve, the mass of air formed by the mutual communication of the air in the cylinder and the atmosphere can flow mutually during the air pressure balance process. The air actuator is equipped with a controller to obtain the position information of the piston, a piston drive command is issued from the controller to drive the reciprocating motion of the piston, and a position and optical sensor for operating the pressure balance valve are installed. Therefore, the timing and frequency of opening the pressure balance valve can be programmed and stored in the controller. By the operation of 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 operation safety and performance of this ventricular assist device.

[0054] Refer to FIG. 9. The 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, that is, embedded, 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 drive catheter segment, each of which is embedded in the human body. The second part is attached outside the human body and includes a proximal drive catheter and a drive catheter electronic module (also called a drive catheter interconnector). The third part is attached outside the human body and is a drive unit including an EMA (electromechanical actuator) controller circuit, a main battery, and a backup battery.

[0055] The blood pump pressure sensor is built into the proximal end blood pump housing and immersed in a small pressure sensor chamber filled with a sensing medium, and can continuously monitor the blood pump pressure. The remote drive catheter is connected to the pump housing and supplies a counterpulsation pneumatic pulse to perform dynamic injection and blood sac filling. The distal and proximal drive catheters supply the pneumatic drive pneumatic pulse generated by the EMA inside the actuator to the blood pump. The electronic blood pressure signal generated by the pressure sensor of the sphygmomanometer is transmitted to the drive unit. The drive air circuit (indicated by a dashed arrow) and the communication signal path (indicated by a solid line) are shown in FIG. 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 for driving a brushless motor and a microcontroller unit as a central processing unit, and can process the received pressure signal and generate a control command for the motor controller so that the piston works.

[0056] Referring to FIGS. 10 and 11, a schematic block diagram of the functions inside the drive unit and the important interconnection signals necessary for the startup of the blood pump is shown. For a further description of the embodiments proposed by the present invention, it will be as shown in FIGS. 10 and 11. Refer to FIGS. 10 and 11. To explain the drive relationship between the external drive unit and the implant, it is necessary to refer to the content of the aforementioned blood pump, drive catheter, distal drive catheter, proximal drive catheter, aortic adapter, and drive catheter interconnect.

[0057] The drive unit receives the pressure signal (electrical signal) of the blood pump, processes this signal using a trigger detection algorithm, and generates a trigger signal that commands the motor actuator to operate in cooperation with the heart rate. 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 process from injection to filling or from filling to injection.

[0058] The architecture of the electronic controller 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 shows an overview of each functional block of the drive unit 78. TIFF0007714254000001.tif156127

[0059] In FIGS. 10 and 11, for the exemplary embodiments previously disclosed, explain how the electromechanical actuator operates to generate pneumatic pulses for signal collection, transmission, processing, and control logic and command generation, and to drive the blood pump.

[0060] Figure 12 shows the trigger detection command for the piston position of the electromechanical actuator related to reverse pulsation assist. In Figure 12, the aortic pressure (AoP) waveform without assistance is shown by the dotted line, and the solid line shows the aortic pressure waveform during assistance. When the drive unit is operating in the automatic operation mode, the operation of the drive unit is activated, and the system executes the "fill - eject - fill - eject..." cycle assist. This represents the normal synchronous reverse pulsation assist operation. The MCU monitors the blood pump pressure (BPP) signal (electrical signal) and detects the timing of the left ventricular end - diastolic (LVED). When the LVED timing is detected, the MCU generates the F_Trig signal. The time interval between two consecutive F_Trig signals represents the instantaneous heartbeat cycle interval (or period). Based on the estimated heart rate calculated from the intervals of the previous several cycles, the MCU determines the ejection time of the blood pump, that is, the E_Trig signal. The E_Trig signal provides the timing for the motor controller unit to drive the electromechanical actuator based on a predetermined position, speed, and acceleration curve. When the ejection stroke is completed and the optimal dwell time has elapsed, until the F_Trig signal appears, the electromechanical actuator is commanded to perform the pre - fill operation at a slow filling speed. After receiving the F_Trig signal, the electromechanical actuator starts to execute the remaining filling stroke at the specified piston speed.

[0061] When the MCU loses the BPP signal (electrical signal) sent from the blood pump, the MCU automatically activates the cleaning mode to drive the electromechanical actuator and operates at a predetermined assist frequency and the volume pacing amount of the drive unit. The cleaning mode is a facility protection mode for preventing thrombus formation in the blood sac without providing synchronous reverse pulsation circulation support.

[0062] In principle, the extracorporeal blood pump device of the present invention has a better counterpulsation support effect than an intra-aortic balloon pump (IABP) due to the characteristic of non-occlusive aorta indwelling. Different from the situation where bedridden or ambulatory IABP patients have to stay in the hospital, the extracorporeal blood pump device allows patients to leave the hospital and live a better life at home. Therefore, the extracorporeal blood pump device of the present invention can obtain economic benefits from a short hospital stay, and can further improve the patient's disease condition and quality of life.

[0063] In recent years, the usage trend of LVAD has been saturated mainly because its application is only applicable to a small number of end-stage heart failure patient groups. Applying early-intervention LVAD treatment to patients with mild heart failure has long been a clinical goal in cardiology. If early-intervention treatment can be achieved, it is expected to have a significant impact on the expanded use of LVAD treatment and the future progress of cardiology. According to clinical evidence, providing LVAD support at moderate to severe heart failure stages can improve some patients with non-ischemic cardiomyopathy through reverse remodeling of cardiomyocytes or continuous myocardial recovery. However, such an intention for early intervention must depend 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 disengages the supported heart from the normal health recovery path. However, counterpulsation support is physiologic and promotes reverse remodeling of cardiomyocytes by providing systolic contraction removal and diastolic perfusion benefits to achieve the treatment goal. From the above, it can be seen that the treatment strategy provided by the invention of this paracorporeal blood pump meets the development conditions of the early-intervention trend in cardiology. The beneficial attributes of treatment efficacy, such as adaptive partial-support, minimally invasive surgery, and counterpulsation therapy provided by the paracorporeal blood pump device, make the present invention a potential candidate for contributing to the future progress of heart failure treatment.

[0064] The following further describes the blood pumps of the above respective embodiments.

[0065] Figures 13A and 13B are the first and third embodiments of the present invention, respectively, showing a schematic diagram and a cross-sectional view of a paracorporeal blood pump device worn inside a human body. The implant subsystem of the paracorporeal blood pump device includes a blood pump 52, an aortic adapter 54, and a drive catheter (or 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. Its proximal end is closed, the distal end is open, and it is seamlessly coupled to the aortic adapter 54. Among them, the blood sac 529 is composed of an oval membrane 526. The blood sac 529 is fixed to the proximal housing 523 of the housing 52h via a proximal port 530 and is also fixed to the distal housing 525 of the rigid housing 52h via a distal port 540. Inside the housing of the blood pump 52, a blood chamber B and an air chamber A are partitioned by an oval flexible membrane 526. The membrane 526 is suspended from the rigid housing 52h via a pair of stress-relief stems (proximal port 530, distal 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 into the proximal housing 523 of the rigid housing 52h. The detected pump pressure is transmitted through the membrane 526 and propagates through a non-compressible liquid or gel contained in a sealed pressure sensing chamber 528, and finally received by the pressure sensing mechanism 527. After receiving the detected pressure of the blood pump, the pressure detection mechanism 527 generates an electronic blood pressure signal. The size and shape of the implant subsystem assembly are designed in an implantable form and are applicable to patients with a body surface area (BSA) of 1.2 square meters or more.

[0066] Figures 14A and 14B are respectively a schematic diagram and a cross-sectional view of a part of a human body-mounted paraaortic blood pump device according to the second and fourth embodiments of the present invention. The implanted subsystem of the paraaortic blood pump device includes a blood pump 62, an aortic adapter 64, a coupler 65, and a drive catheter (or distal drive catheter) 67 connected to the blood pump 62. The coupling 65 connects the blood pump 62 to the aortic adapter 64 for entry into the vasculature of the implanted person of 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 present blood pump 62 is similar to the structure revealed in Figure 14B, except that the opening OP is separated from the aortic adapter 64 and is independent. The coupler 65 is placed around the neck portion 643 of the aortic adapter 64. Hereinafter, using the design disclosed in Figure 14B, the design of the blood pump and the basic design principle will be further described.

[0067] Refer to Figure 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 vasculature. The opening OP of the blood pump 52 is manufactured seamlessly with the aortic adapter 54, providing a smooth and continuous boundary transition to the neck portion of the aortic adapter 54. The assembly of such an integrated blood sac 529 and the aortic adapter 54 and the joining of the blood pump 52 connect the proximal housing 523 and the distal housing 525 by the adhesion of the proximal end 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 adhered to the pressure sensing chamber 528 at the center of the proximal housing 523.

[0068] The miniaturized pressure sensing mechanism 527 is built into the proximal housing 523 and communicates with a pressure sensing chamber 528 closed with fluid as a medium. With this device, the blood pressure in the blood sac 529 can be continuously monitored. Since 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, which helps the rigid housing 52h to shield the pressure sensing mechanism 527 and its circuit from the effects of chemical corrosion and protein adhesion due to direct blood contact.

[0069] The end of the drive catheter 57 is connected to the remote housing 525 to supply a reverse pulsatile pressure pulse for injecting or infusing blood from the blood pump 52. The design of the drive catheter can be made into multiple cavities or multiple layers to enclose the electric wires used for transmitting the pressure signal. By adopting a metal coil, a fabric net or a net as a catheter wall reinforcing member, the kink resistance of the distal drive catheter 57 can be enhanced. The overall geometry of the flow path in this blood pump is relatively wide, and the design of the valveless aortic adapter and the operation of the pulsatile blood pump constitute excellent blood treatment characteristics, avoiding hemolysis due to high shear force and thrombus formation or thromboembolism due to low flow rate.

[0070] Due to the innovative designs of the aforementioned blood pump 52 and blood sac 529, the bladder membrane 526 is highly durable. The blood sac 529 is an oval membrane rotator with the center line of the blood pump 52 as the center of rotation. Two polymer ports (proximal port 530, distal port 540) are connected to both ends of the rigid housing 52h, each configured in a disc or annular shape. When connected to the rigid housing 52h, they reduce stress concentration as a bending / tensile stress release mechanism. During the process of injecting blood from the blood pump, the cyst membrane 526 is compressed or folded into a tri-lobe shape, and the maximum strain usually occurs at the fold near the edges of the normal ports (proximal port 530, distal port 540). Such local high membrane stress / strain caused by the large deformation of the film is absorbed and offset by the flexible suspension deformation of the port edge. Of particular note is that the position of the three-lobe folding mode is not fixed, the deformation of the film is affected by the direction of gravity, and the fold position occurs randomly. In fact, the patient's body posture and orientation can change from time to time during daily activities, including standing, lying, sitting, and exercising postures. Therefore, the gravitational effect acting on the volume of blood stored in the blood pump 52 constantly changes direction, forming a non-fixed grid line. Such randomly formed overlapping line characteristics of the thin film constitute the unique fatigue resistance characteristics of the present invention. This blood pump 52 is expected to have a longer durability than the conventional fixed folding thread film design.

[0071] The folding and inflation of the bladder membrane are closely related to the vortex structure mode contained in the blood sac 529. The characteristic of the aforementioned blood sac design is that the folding lines are randomly formed, and the characteristics of the vortex structure occur alternately with the change of the pattern of the folded membrane. Therefore, the cleaning effect in the blood pump 52 is strong and unsteady, characterized by a random wandering type of vortex motion. Such randomness of the blood pump vortex structure helps to wash away the blood contact surface of the entire blood sac without generating low-speed recirculation regions fixed near the membrane wall or in the fold area. In animal experiments, it was observed that the blood pump of the present invention has a very strong antithrombotic ability.

[0072] Hereinafter, each distal-side drive catheter in each embodiment will be further described.

[0073] Refer to FIGS. 5 and 6. It is a schematic diagram of a drive catheter 96 that connects a blood pump 92 and a drive unit 98. A remote measurement drive catheter 97 (driving a catheter body 991) that drives the catheter 96 air-connects the blood pump 92 to an electromechanical actuator housed in the drive unit 98 and transmits an electrical signal acquired from a blood pump pressure sensor mechanism 527 (see FIG. 13). One end of the remote measurement drive catheter 97 (the in-catheter 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 remote measurement drive catheter 97 (the in-catheter body segment 991) passes through the skin subcutaneously under protection with a protective cap. The outer diameter of the remote measurement drive catheter 97 (the in-catheter body segment 991) is designed to be small, and the tube material is flexible, minimizing the stress at the exit site EX and giving comfort to the patient. A part of the remote measurement drive catheter 97 (the in-catheter body segment 991) is covered with a porous fabric, promoting the growth of tissue into the fabric and making the exit site EX infectious. The out-of-catheter segment 973 of the drive catheter emerging from the epidermis is fixed at a distance slightly away from the skin exit site EX.

[0074] The in-catheter body segment 991, the out-of-catheter segment 993 of the drive catheter, and its connector are designed to be able to withstand the tensile load applied during percutaneous surgery. After the surgery, the in-catheter body segment 991 and the out-of-catheter segment 993 of the drive catheter continue to be affected by the load due to muscle movement, and the in-catheter body segment 991 and the out-of-catheter segment 993 of the drive catheter are designed to be able to withstand these fatigue loads. The outside of the in-catheter body segment 991 and the out-of-catheter segment 993 of the drive catheter is designed to have biocompatibility and has chemical resistance to cleaning agents and disinfectants during clinical use.

[0075] Next, the proximal drive catheter of each of the above embodiments will be further described.

[0076] The drive catheter extracorporeal segment 993 and the proximal drive catheter 99 are for connecting the drive catheter intracorporeal segment 991 and the distal drive catheter 97 to the drive unit 98, respectively. The near drive unit catheter 99 has a drive catheter interconnect 93 at one end and a drive unit connector at the other end. The drive catheter interconnect 93 includes a circuit board that converts the pressure signal of the analog blood pump into a digital signal, and a vibrator that provides tactile feedback in addition to an audible alarm. The drive catheter interconnect 93 has a flat shape to prevent torsion from occurring in the drive catheter extracorporeal segment 973 when it is fixed to the patient's skin. Also, the drive catheter interconnect 93 and the drive catheter embedded wire are sealed and designed to prevent the intrusion of water and moisture. Since the proximal drive catheter 99 is externally attached so that it can be replaced or maintained when necessary, if the proximal drive catheter 99 is damaged beyond repair, it is not necessary to replace the blood pump surgically.

[0077] The valveless blood pump has two advantages in terms of blood treatment characteristics. 1) There is no unpleasant valve sound, blood cell damage, thrombus formation, and thromboembolism caused by valve induction. 2) It has stronger antithrombotic properties. The bidirectional pulsatile flow has a better surface cleaning effect, so it can minimize protein adhesion and avoid forming blood clots related to boundary discontinuities on the artificial surface in contact with blood. The flow path of the pump without valve pulsation is much wider and more uniform than that of the valve pulsation or continuous flow rotary pump. Hemolysis (rupture of red blood cell membranes) usually occurs in narrow flow paths with high flow velocity gradients, such as the gap between the valve ring and the leaflet of the pulsatile pump with valves. Also, there are often low-speed recirculation and stagnant areas on the back side of the open valve, which may promote thrombus formation. In contrast, in the valveless pulsatile blood pump, the shear stress on blood cells is actually several orders of magnitude smaller, the low-speed stagnant areas related to the geometry and movement of the valve are almost eliminated, the damage to blood cells and platelet activation are reduced, the formation and aggregation of blood clots are reduced, it is converted to a lower dose of anticoagulant for use, and simpler and safer postoperative care is carried out.

[0078] Figures 15 to 17 show schematic views of a blood pump 62, a drive catheter 67, and an introduction member 63 according to another embodiment of the present invention. In this embodiment, anatomical adaptability is emphasized to make the externalization of the blood pump and delivery easier.

[0079] As shown in FIGS. 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 a stem assembly 659 (including a sac 629 and stems 630, 640), a pump housing 62h (a remote housing 625 having a proximal end housing 623 and an inlet connector 6251), and a pressure sensing system 628 fitted into the proximal end housing 623. The aforementioned components, the inlet connector 6251, the pressure sensing system 628, and the drive catheter 67 are basically the same or correspond to the assemblies / fittings of the aforementioned embodiments, and the detailed description of these assemblies and their functions will not be repeated here.

[0080] In this embodiment, an introduction member 63 is provided in the remote housing 625 of the pump housing 62h, and the drive catheter 67 is connected to the pump housing 62h. Further, the introduction member 63 is configured in a body shape adjacent to the remote housing 625, and a power transmission member is tangentially connected to the outer surface of the pump. With such a main body penetration design, the design of the pump housing 62h adapts to the anatomical space that can be used for the installation of the blood pump. The blood pump 62 is rotatably connected to the interface adapter 501, and the drive catheter 67 can be wired in an optimal orientation so that the subcutaneous tunnel and the skin exit can be performed smoothly. This is advantageous for the anatomical adaptability to the geometry of the implant site.

[0081] In this embodiment, the introduction member 63 is remotely fitted to the remote housing 625, and the sensor 6271 (FIG. 25A) and the pressure sensing chamber 628 are arranged in the proximal end housing 623, so that the signal transmission path and the pneumatic communication path are separated, and it is necessary to seal and protect the blood pump 62, and the fidelity of the signal transmission after the device is implanted may be impaired by the intrusion of biochemical fluids.

[0082] Referring to FIG. 17, the pump housing 62h has a groove 621 formed on the outer surface of the remote housing 625 and is located above the overlapping joint region da [FIG. 16] of the proximal end housing 623 and the remote housing 625. The surface groove 621 is arranged such 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 region DA. In some embodiments, the groove 621 is potted and sealed with a waterproof material and / or an annular cap so as not to stimulate or damage the contacting tissue.

[0083] As shown in FIGS. 18A and 18AB, the aortic connector 50 is a connection mechanism typically used to connect the blood pump 62 to the target artery 60 via its interface adapter 501. The distal end 504 of the connector 50 facing the interface adapter 501 is placed on the vessel wall of the artery 60 and is connected to the human circulatory fluid. The proximal end of the aortic connector 50 (or the interface adapter 501) has a smooth interface transition to geometrically match the inlet configuration of the blood pump 62. A coupler is usually required to integrate the proximal end of the connector 50 (or the interface adapter 501) with the inlet of the blood pump 62. In some embodiments, it can be used as the aortic connector 50. FIG. 18A shows the end-to-end alignment of a Dacron or PTFE graft 502 sutured to the target artery (or blood vessel) 60, which can be used in vascular surgery. In some other embodiments, for example, the embodiment shown in FIG. 18B, an insertable aortic connector 503 is used, such as the T-shaped manifold adapter disclosed in U.S. Patent Application No. US 2008 / 0300447A1, titled "Dual-pulsation bi-Ventricular Assist Device".

[0084] FIGS. 19 and 20 show a longitudinally symmetric elliptical blood sac and a stem assembly 650 and its components, respectively. Examples of polymer materials used for these components include, but are not limited to, segmented polyurethanes with various appropriate hardnesses. The components of the sac and stem assembly 650 include a flexible membrane sac (blood sac) 629, a proximal end stem 630, and a distal stem 640. In some embodiments, they are all formed and integrated in an axially symmetric shape with respect to the common centerline 62C of the blood pump 62. The proximal end 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.

[0085] FIG. 19 shows the integrated blood sac and stem assembly 650, with the end of its distal stem 640 remotely encapsulated by the blood sac 629 and coupled to the inverted membrane 62A. FIG. 20 shows the components before coupling. Generally, 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. Theoretically, as shown in FIG. 21, when the pressure difference applied to the membrane sac exceeds a certain threshold, the thin-walled sac with an axisymmetric elliptical structure bends into a three-valve type 6293, and this bending of the membrane is only related to the final three-valve type 6293 (eigenmode), and the position of the fold 6294 or the fold line generated there is determined by the initial disturbance that causes buckling 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. In order to ensure an axisymmetric shape, attention must be paid to the manufacture of the highly accurate sac. In real life, the direction of gravity is a major factor at the start of the fold line. The posture of the recipient of the equipment constantly changes due to the patient's daily activities (such as 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 distributed non-steadily throughout the sac. Therefore, avoiding leaving the high-strain regions at fixed positions is an important design criterion for giving the sac a long lifespan.

[0086] Embodiments of the present invention innovated the running fold line characteristics that non-steadily appear high strain positions in the membrane to extend the fatigue life of the membrane sac. Therefore, the harmful stress concentration phenomenon closely related to the bent blood sac was improved. Based on such an essential change in the behavior of the bending mode, the significant increase in the fatigue life of the membrane is due to such non-steady fold line formation characteristics, which disperse high strain regions throughout the sac. Also, the beneficial results associated with such non-steady sac deformation modes depend on the enhanced vortex washing effect within the blood sac. The surface of the sac is more thoroughly cleaned, and vortices such as irregular walking are formed, crossing the entire vortex. This significantly reduces the possibility of the occurrence of a certain low-speed recirculation region and the crease of the fold line in the region near the wall, realizing a blood pump design with a long life and anti-thrombosis ability.

[0087] Figures 22, 23A, and 23B show exemplary embodiments of how the integration method is employed to attach the blood sac and the stem assembly 650 to the pump housing 62h and connect the drive catheter 67 to the proximal end housing 623.

[0088] The blood sac 629 is fixed to a pump housing 62h that includes a proximal end housing 623 and a distal housing 625 to facilitate the filling and discharging operations of the pump. Generally, the flexure characteristics of the blood sac 629 and the pump housing 62h are significantly different. In order 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 with the transition of structural characteristics (especially the bending deformation of the membrane). A pair of flexible stems (proximal end stem 630 and distal stem 640) are used as a suspension mechanism that integrates the blood sac 629 and the pump housing 62h. As shown in Figure 23A, the disk-shaped proximal end stem 630 is connected to the proximal end housing 623, and the ring-shaped distal stem 640 is connected to the distal housing 625. Mechanically, the proximal and distal stems 630, 640 not only hold the blood sac 629 within the pump housing 62h but also function as a stress removal suspension mechanism that avoids stress concentration at the boundary attachment location and extends the life of the blood sac 629.

[0089] As shown in the lower part of FIG. 23A, the remote housing 625 includes an extension of the inlet connector 6251 and 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, and the second end 6253 is coupled to the interface adapter 501 (shown in FIGS. 18A and 18B). The first end 6252 of the inlet connector 6251 is smoothly aligned with the remote of the blood sac 629. However, the opposing second end 6253 is configured to pair with the interface adapter 501, and the coupling design goal is to minimize interface discontinuity to avoid clot formation. The beak has a flange structure, which is disposed in the middle region of the inlet connector 6251 and functions as a locking component received by the interface adapter 501.

[0090] During surgery, the closed-end sac design of the valveless blood pump 62 sucks air into the sac top by buoyancy and collects air bubbles. Refer to FIGS. 22 and 24. An exhaust port 66 is attached to or provided in the proximal casing 623, and a narrow passage 661 is provided above the integrated sac stem diaphragm 6301 between the proximal stem 630 and the blood sac 629. In some embodiments, the passage 661 extends along the center line 62C. After the blood pump 62 is aligned with the target artery 60, the trapped air is pushed out by the arterial blood pressure and accumulates in the top space of the blood sac 629. The accumulated air is discharged by passing through the exhaust port 66 with a thin needle, through the passage 661, the sac stem diaphragm 6301, and into the interior of the blood sac 629. The entire sac stem membrane 6301 below the exhaust port 66 is relatively rigid and not bent, so the sac 629 with the perforation is not further structurally damaged, and when subjected to periodic pulse pressure or the expansion and contraction or folding of adjacent sacs, structural damage due to the progression of cracks generated in the perforation gap can be avoided.

[0091] As shown in FIGS. 23A and 25A, the pressure sensing mechanism 627 is embedded in the proximal housing 623. FIG. 23A shows the details of the cross-section of the integrated proximal housing 623, the introducer 63, and the drive catheter 67. FIG. 25A shows the contour of the proximal housing 623 connecting the introducer 63 extending from the dome of the proximal housing 623 to communicate air and signals to the drive catheter 67.

[0092] As shown in FIGS. 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 circuits. A plurality of electrodes 6274 extend from the bottom of the second space 6273 and are connected to the electric wire 6702 of the carrier 67 (FIG. 26). The second space 6273 is closer to the drive 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 within the proximal housing 623 and adjacent to the first space 6272 is formed to confine the sensing fluid. The distal 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 attaching the sensor 6271 and a second arm 6282 for filling and sealing the medium to be detected. Therefore, the blood pressure pulse can be transmitted through the transmembrane blood sac 629 and hydraulically communicate with the remote MEMS sensor 6271 located in the second space 6273.

[0093] One embodiment of the present invention has revolutionized the pressure-based blood pump control method and sensor design. A micro MEMS pressure sensor is adopted, and its electronic circuit is packaged and embedded in the housing wall. In principle, due to its inherent microscale structure, the crystal grains of the MEMS sensor are very durable. In fact, the durability of the sensor depends on the packaging design. This pressure-sensing mechanism 627 is non-blood-contact and is isolated from corrosive biochemical actions related to blood, thus providing the long-term signal collection and transmission required for long-term implantable assistive devices.

[0094] The carrier 67 functions as a transmitter that converts electrical signals and transmits air pulse pressure between the blood pump 62 and the drive unit 98. Fig. 26 shows a typical multi-layer carrier 67 in the present invention. In this embodiment, the carrier 67 has an air lumen (or internal air tube) 6701, a plurality of electric wires 6702, an intermediate air tube 673, a coil 674 such as a metal coil, an outer tube 675, a tether 676, a silicon sheath 677, a rigid drive unit connector 678, and a protective hollow connector 679.

[0095] An air lumen 6701 (or air passage, also called inner tube) with an inner lumen diameter of 2-5 mm is accommodated in the central part of the carrier 67, and either low power consumption or low surgical simplicity can be selected. The electric wires 6702 for signal transmission are embedded in the wall of the carrier 67. Other modified designs of the carrier can be adopted. In addition to the design of the multi-layer carrier 67 shown in Fig. 26, the carrier 67 may, for example, as shown in Fig. 27, have a plurality of cavities where the electric wires 6702 are embedded in several small tube chambers and pulsed air can flow through a larger tube chamber 6701. One of the small lumens can be attached together with the tether 676 to limit the elongation of the delivery 67 and protect the electric wires 6702 from damage by external tensile forces.

[0096] The inner tube or air lumen 6701 is received within the air tube 673 with a reinforcement sandwiched therebetween. Between the inner tube 6701 and the air tube 673, a coil 674 (or a woven thread or net) can be reflowed to reinforce the wall of the delivery member, and the delivery member 67 is flexible but resistant to kinking. The outer layer tube 675 covers the inner and intermediate air lumen 6701 and air tube 673, and the helically wound wire 6702 can be used as a protective cover. Note that in some embodiments, a non-extendable tether 676 may be provided between the outer tube 675 of the delivery member 67 and the silicone cover 677 to increase the tensile elasticity required when the delivery member 67 is externalized. Clinically, it has been proven that the silica gel jacket 677 has the least irritation to the subcutaneous tissue and the lowest transport infection rate.

[0097] 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 packaged in the main body of the delivery member 67. The proximal end portion 671 of the delivery member 67 is inserted into a socket provided in the drive unit 98. The rigid drive unit connector 678 of the delivery member 67 is used to be received within the socket in the drive unit 98. The rigid drive unit connector 678 is mounted flush with four electrodes 6781 (for example, in FIG. 23) of a plurality of electrodes 6781 to which the wire 6702 is soldered. The protective hollow connector 679 (see FIGS. 23b and 26) is disposed at the junction between the drive catheter 67 and the drive connector 678 so that the delivery member 67 does not kink at the junction. The proximal end portion 671 of the delivery member 67 includes the drive unit connector 678 and the hollow connector 679 and is mounted so as to be easily withdrawn from the skin without causing an unintended puncture injury.

[0098] As shown in FIGS. 22 and 23A, the connection between the transport member 67 and the blood pump 62 is realized by the introduction member 63. Depending on the anatomical structure in which the blood pump 62 is implanted, the introduction member 63 can be arranged within the proximal end housing 623 or the distal housing 625. By integrating the introduction member 63 with the pump housing 62h, the arrangement of the entire external blood pump can be changed, and the transport member 67 can be guided in a specific direction, such as the externalization path of the transport member, postoperative skin care, and equipment availability.

[0099] As shown in FIGS. 23A and 25A, the inlet 63 has a first portion 631 that is an extension of the proximal end housing 623, and the air chamber 6701, tether 676, and electric wire 6702 of the drive catheter 67 are connected to the first portion 631 via the anchor adapter 672. The introduction member 63 further has a second portion 632 that interlocks with the first portion 631 as a hollow connector of the transport member 67. The first portion 631 is a position for making an electric wire connection, 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 so that no tensile force is applied during the externalization process of the power transmission member. Also, for the connection between the air lumen 6701 and the blood pump 62, it is necessary to have no air and no current leakage. The above blood pump integration task is performed in the first part 631. The second portion 632 houses these boundary components, functions as an external protector, and protects the boundary from mechanical stress, environmental fluid, and moisture.

[0100] Figures 19 through 27 disclose a module design related to a first embodiment of the blood pump of the present invention. In this embodiment, the blood pump 62 has a pump housing 62h having an axially symmetric elliptical blood sac and a stem assembly 650 (including a flexible membrane sac 629, a proximal stem 630, and a distal stem 640), a proximal housing 623, and a distal housing 625. A delivery member 67 is connected to the blood pump 62, and the delivery member 67 includes an air lumen 6701, wires 6702 within its wall. The introduction member 63 is used to achieve electrical and pneumatic communication between the delivery member 67 and the blood pump 62 in order to integrate the drive catheter 67 with the pump housing 62h.

[0101] As shown in FIGS. 19 and 20, the connection design between the blood sac 629 and the stems 630, 640 has been disclosed in the previous section. The key points of the design and manufacturing are to maintain a high-precision axial symmetry in the component manufacturing and the joining of the sac and the stem assembly. The pressure sensing mechanism 627 and the introduction member 63 are attached to the rigid portion of the proximal housing 623. FIGS. 22, 23A, and 23B show the design of the compact introduction member 63. It can be seen that the compressed inlet 63 enables a more robust and fault-tolerant connection with the wires.

[0102] Figures 28A and 28B show several flow patterns related to intra-aortic balloon counterpulsation. During the end-diastolic and early systolic phases of left ventricular ejection, the blood pump draws aortic blood flow into the pump through pump filling (Figure 28A). Blood upstream and downstream of the connector is drawn into the blood pump in a sharp 90-degree flow. As a result, flow separation and a low-velocity recirculation region T-201 occur. Also, very high shear appears in the corner region of the T-joint. On the other hand, during the diastolic phase after aortic valve closure, the blood stored in the pump is ejected back into circulation, creating an impact flow against the opposite aortic wall (Figure 28B). This roll, impact flow has a very high local pressure, a so-called stagnation point, at the impact point T-202, where the flow velocity is actually zero and all the kinetic energy related to the flow velocity is converted into potential energy called the total pressure. Such high-pressure impact flow can cause vascular maladaptation, such as smooth muscle cell proliferation and resulting vessel 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 all commonly exist near the T-tube. This turbulent, complex flow anomaly attenuates or decreases within a distance of 3-5 times the diameter of the implanted arterial lumen. This plug-in flow connector is designed to have an insertion catheter length of 5-7 cm, covering most of the non-physiological flow region caused by the pump. Since the aorta at the implant site is blocked by the inserted flow connector, 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.

[0103] In counterpulsation support, pump filling and ejection are alternately driven in synchronization with the heart rate, creating the special T-joint flow as shown in Figures 28A and 28B. The aforementioned insertion-type aortic adapter 14 is further detailed in the perspective schematic diagram of Figure 29 and the cross-sectional schematic diagram of Figure 30, respectively.

[0104] The aortic adapter 14 is mold injected, and its internal blood contact surface 141 is ultrasmooth and continuous without parting lines. The aortic adapter 14 can be made of silicone resin or other polymer elastomer materials. In some embodiments, the aortic adapter 14 has a polymer elastomer containing a silicone resin material, or the polymer elastomer is a mold injectable polyurethane. The aortic adapter 14 includes an aortic adapter catheter (also referred to as a catheter insertion part) 142 for insertion into the aorta 95 (see FIG. 5), and a neck part (also referred to as a raised part) 143 connected to the blood pump. In this embodiment, the convex neck part 143 has a neck part main body 1431 and an extension part 1432 provided on the neck part main body 1431. The extension part 1432 protrudes from the neck part main body 1431, and the maximum inner diameter of the extension part 1432 is larger than the maximum inner diameter of the neck part main body 1431. When the convex neck part 143 is connected to the blood pump 62, the extension part 1432 is in close contact with the inlet adapter 6251 of the blood pump 62, and the neck part main body 1431 is surrounded by the coupler 25 and integrally connects the inlet adapter 6251 to the blood pump 62.

[0105] The aortic adapter 14 is thin-walled to maximize flow efficiency. To strengthen its thin-walled structure, a pair of nickel-titanium alloy metal trusses (also referred to as metal truss rings) 144 are fitted around and surround both ends of the aortic adapter catheter 142 of the aortic adapter 14.

[0106] Figure 29 is a schematic diagram showing the fitting position of the nickel-titanium alloy metal stent 144. Also, the wall thickness of the catheter 142 of the aortic adapter gradually becomes thinner toward both catheter ends 145. The gradually decreasing wall thickness has a dual effect. First, it minimizes the discontinuity of the connection between the implant 502 / main body, making the formation rate of interfacial clots much lower than the thrombolysis rate provided in contact with the endothelium. Second, the compliance of the catheter becomes softer toward the catheter end 145, resulting in a compliance matching effect when connecting to the aortic lumen.

[0107] One of the complications troubling the transportation of the large stent implant 502 is the endoleak problem. Type I endoleak means that the seal between the embedded end and the embedded arterial endothelial lumen is incomplete, and a gap occurs between the leading edge of the embedding and the arterial lumen. The oozing blood is trapped in the gap, coagulates into a blood clot, and finally becomes a fibrous false intima, growing uncontrollably over time. The false intima not only blocks the embedded artery but also emits signals, stimulates the blood coagulation mechanism to attract platelet adhesion, and may cause thrombotic adverse effects. A solution to such an endoleak problem is to tightly seal the lumen surface equipped with the aortic adapter 14. The aortic adapter 14 of the present disclosure proposes a concept of a compliance adaptation design that allows the semi-rigid catheter (terminal) end 145 to be seamlessly attached to the arterial lumen when receiving pulse pressure. As shown in Figure 30, the outer diameter 146 of the catheter 142 of the aortic adapter is slightly larger than the diameter of the inner lumen. Under the condition of a predetermined nominal blood pressure (for example, 120 mmHg), the excess ratio (defined as the ratio of the catheter diameter to the diameter of the inner lumen) is within the range of 3 - 10%. The compliance-compatible catheter end 145 expands and contracts dynamically in response to pressure pulsations without creating a boundary gap when the blood pressure fluctuates between systole and diastole or under the pulse pressure by counterpulsation support.

[0108] Elastomeric thin-walled tubes are flexible and often compliant, but due to their overly large size, they lack sufficient strength to withstand the compressive forces applied, and the walls of the inserted adapters often bend. Therefore, a combination of a structure using a nickel-titanium alloy metal stent 144 and an elastomeric substrate with appropriate hardness is important. As shown in Figure 30, in this design, the radial rigidity is ensured by the nickel-titanium alloy metal stent 144, so it can support the aortic adapter 14 without bending it, and there is a distance "x" between the outermost boundary 1441 of the metal stent 144 and the catheter end 145. In some embodiments, it is necessary to evaluate and accurately define the aforementioned distance x. By supporting the nickel-titanium alloy metal stent 144 as an expandable frame, the gradually thinning catheter end 145 abuts against the connected lumen wall, which holds it circularly without collapsing or wrinkling, and is dynamically sealed with the lumen. The aortic adapter 14 can expand and contract in response to pressure pulsations, and the aortic adapter 14 and the wall of the aorta 95 expand and contract integrally without causing bleeding complications, and the sealing effect of sealing the catheter end 145 is dynamically realized.

[0109] Figure 31 shows a representative embodiment of a nickel-titanium alloy stent 144, generally laser-engraved from nickel-titanium alloy, which is further expanded by a series of expansions and heat treatments. Figure 31 shows a plan view of the metal stent 144 unfolded. (For each column) The metal stent has a plurality of waveform structures. The metal stent 144 is self-expandable and can be bent or slightly curled, placed in a pre-crushed roll, and after being placed in the required position, it can self-release and return to its original shape.

[0110] The simple measurement of catheter stiffness (the reciprocal of compliance) can be represented by so-called lateral stiffness (LS), and the measurement method will be described in Fig. 32. The lateral stiffness is defined as the value obtained by dividing the acting force F per unit length by the corresponding radial deflection Y. In the current aortic adapter, the appropriate lateral stiffness range is 0.01 - 0.05 Nt / mm2. The embedded nickel-titanium alloy metal stent 144 and the silicon resin or elastomer substrate contribute to the structural compliance of the aortic adapter 14 co-injected. The structural compliance of the metal stent 144 and 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 stratification between layers is minimized by the expansion and contraction of the wall of the composite catheter, and the life of the adapter 14 is prolonged.

[0111] The aortic adapter 14 is arranged to be connected to the blood pump 62 so as to promote circulatory support. Here, a quick-connector type coupler 25 is provided. As shown in FIG. 33, a disassembled schematic diagram of each part of the coupler 25 integrating the aortic adapter 14 and the blood pump 62 is shown. This coupler 25 includes a flange base 252, a pair of lock rings 253, and a hinge (or hinge assembly) 254 connecting the lock ring 253 and the flange base 252. A spring coil (or coil assembly) 255 is loaded on the hinge joint 256, and when the coupler 25 is unlocked, the lock ring 253 is held in the open position (FIG. 34A). The locking mechanism depends on a leaf spring type latch 257, which is made of a grooved spring piece and is fixed through a plate 2571 welded to one end of the lock ring 253. The aforementioned flange base 252 has a substantially circular structure, and each lock ring 253 has an arc structure. The hinge joint 256 is located on the first side 252S1 of the flange base 252, and the leaf spring type latch 257 is located on the second side 252S2 of the flange base 252 opposite to the first side 252S1. The lock ring 253 is pivotally attached to the hinge joint 256 and is rotatable with respect to the hinge joint 256 and the flange base 252. In some embodiments, the coupler 25 and the aortic adapter 14 belong to a part of the aortic adapter assembly.

[0112] FIG. 34b shows a schematic diagram of the coupler 25 in the locked state, where the split of the leaf spring type latch 257 engages firmly with the inclined surface 258, ensuring the safety of the connection without worry of separation. As shown in FIG. 35, the integral connection between the aortic adapter 14 and the blood pump 62 functions as a "gasket" between the rigid flange base 252 and the rigid beak-shaped flange 81 (described later) of the blood pump inlet adapter 6251, which are connected to each other through the deformable adapter proximal end 147 (FIG. 30, i.e., the end of the neck portion).

[0113] Specifically, as shown in FIGS. 34b and 35, the closure lock ring 253 can easily perform a quick-connection type lock, and there is no worry of accidentally unlocking. The leaf spring latch 257 is attached to the end of one of the lock rings 253. While the lock ring is locked, the leaf spring latch 257 bends when sliding on the (convex) inclined surface 258 of the opposite lock ring 253 during locking. When the leaf spring latch 257 crosses the top of the inclined surface 258, it descends to the bottom of the inclined surface 258 by the elastic restoring force, and functions as an insurance to prevent accidental latch release or opening of the lock ring due to the vibration of the pump or long-term rocking. In the case of pump explantation or replacement that requires module separation, the plate spring latch 257 is bent by a tool and lifted upward, and a unlocking force is applied to rotatably open the lock ring 253, so that the blood pump 62 can be detached from the aortic adapter 14.

[0114] The design of the butt joint is impossible for two smooth surface pipe joints to be connected in the blood flow. In many clinical applications, the surface of the connected implant 502 is rough, which promotes endothelialization, so that the minute boundary discontinuities in the blood flow are "smoothed out" by the cells and proteins growing inward. The current aortic adapter 14 avoids the occurrence of poor thrombus formation by smoothing the surface, and the reason has been explained above. As shown in FIGS. 28A and 28B, the ejection and filling operations of the blood flow in the aortic adapter in response to the counterpulsation pump are bidirectional. Due to this strong bidirectional flow and surface cleaning effect, any newly formed blood clot on the rough surface can be easily removed. Therefore, a smooth surface design is considered to be more suitable and safer to use according to the present invention. The boundary of two connected smooth surfaces in the blood flow requires careful mechanical and hemodynamic design to prevent the occurrence of in-situ thrombus events. The principles and design methods related to the invention of such a new joint are disclosed below.

[0115] Figures 36A and 36B each show two basic boundary discontinuities existing in the connection of the butt joint. For example, steps 101, 102, or gaps 103 occur between adapter AB1 (such as the adapter of blood pump 62) and adapter AB2 (such as the aortic adapter 14). Figures 36A and 36B are in an exaggerated drawing method. Usually, in precision machining, such discontinuities of the joint are within 10 - 50 microns, which is large enough to cause clotting and thrombus formation.

[0116] In fact, even if the processing of each object is exactly the same, it is necessary to match the tolerances of two individual objects. In Figure 36A, two misaligned joints are depicted where everything related to part manufacturing is correctly completed except that the center lines do not match. Forward-facing steps and backward-facing steps 101, 102 occur, and the stagnant flow in the step regions 101, 102 becomes the starting point for clot or thromboembolism formation. As shown in Figure 36B, a boundary gap 103 is formed due to the non-parallel matching of the joined bodies. The gap 103 sucks and aggregates blood cells and further grows on the pseudo-intima. This growth of the intima is often uncontrollable, and in addition to the thrombus bolus that has fallen off from the intima surface, it can block the entire blood flow path. When the connected objects are non-rigid, the wrong boundaries of the butt joint may deteriorate. The aortic adapter 14 of the present disclosure is semi-rigid and can be forcibly press-fitted and connected to the joint (for example, the inlet adapter), and has a deformed structure and an enlarged boundary discontinuity. Therefore, in order to realize the connection between the current semi-rigid aortic adapter and the blood pump, it is necessary to invent new connection means as follows.

[0117] Refer to Figures 37 and 38. In some embodiments, the blood pump 62 has an inlet connector 80, and the inlet connector 80 includes a beak-shaped flange 81, a beak (part) 82, and a connector body 83 as an extension of the housing of the blood pump 62. The connector body 83 is provided with a plurality of holes 86 for connecting the inlet connector 80 and the blood pump 62.

[0118] The inner diameter 84 of the beak portion 82 is slightly larger than the inner diameter 148 of the convex neck portion 143 of the aortic adapter 14 (see Fig. 30). As shown in Figs. 30 and 35, the contact area between the beak portion 82 and the proximal end (end portion) 147 of the adapter is an annular tapered surface (or a shallow inclined surface, also referred to as an inclined surface), so the beak portion 82 can be called a tapered beak. The shallow inclined surface 149 is inclined with respect to the center line of the catheter insertion portion, and the taper angle of the shallow inclined surface is within the range of 30 - 60 degrees measured from the rotation center line of the inlet connector 80. In the initial lock engagement, the lock ring 253 having the inner internal groove 2531 is loosely hooked to the flange of the flange base 252 and the flange of the beak-shaped flange 81. With the locking of the lock ring 253, the aforementioned beak-shaped flange 81 (inlet connector 80) and the flange base 252 (coupling 25) are housed in the internal groove 2531 of the lock ring 253 and pressed, so that the end portion 147 (aortic adapter 14) of the intermediate silicone rubber is compressed and clamped, and a clamping force for firmly connecting is generated. In this way, the inlet connector 80 of the blood pump 62 is firmly connected to the aortic adapter 14.

[0119] The above-described clamping force generating mechanism is shown in Fig. 35. The flange base 252 has two steps 2521 and 2522 for generating a 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 adapter 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. Due to the elastic restoring force of the aortic adapter 14, the bent convex neck portion 143 returns to its original circular (or shape), allowing the step 2521 to engage with the engagement groove 1433. The height Z of the internal groove 2531 of the locking ring 253 controls the pressing deformation of the adapter proximal end 147 of the convex neck portion 143. Referring to Fig. 35, in this locked configuration when the locking ring 253 closes and locks, the gap Z0 (i.e., the thickness Z0 of the pressed adapter proximal end 147) is smaller than the thickness Z3 of the end portion 147 of the aortic adapter 14 (see Fig. 30), that is, it can be seen that the thickness Z0 of the pressed adapter proximal end 147 is smaller than the thicker Z3 of the starting end 147 (Z0 < Z3). For the thickness Z0 of the pressed adapter proximal end 147, the following formula can be obtained from Fig. 35.

[0120] Z0 = Z - Z1 - Z2…Formula (1)

[0121] In formula (1), Z1 and Z2 are the thicknesses for clamping the mating beak-shaped flange 81 and the step 2522 of the flange base 252, which are mating parts, respectively, as shown in Fig. 35. Usually, since the thickness Z3 is larger than the gap Z0, the distorted aortic adapter proximal end 147 generates the clamping force required to seal-connect 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 within the range of 10 - 30%, which is sufficient to ensure a reliable sealed connection.

[0122] Figure 39 shows the joining characteristics of the current connection between the beak portion 82 and the shallow inclined (taper) surface 149 of the adapter. When connected, the beak portion 82 and its beak front edge 85 sink into the semi-rigid shallow inclined surface 149, and the aortic adapter 14 has a depth corresponding to the front edge radius of the beak front edge 85, usually 30 - 50 microns. In Figure 39, the numbers in the dotted line and parentheses indicate the initial contact between the beak portion 82 and the beak front edge 85, and the solid line and the numbers without parentheses indicate the state in the locked position. Note that the boundary discontinuity is reduced by the indentation from the original shape (dashed line) of the aforementioned shallow inclined surfaces 149 and 82. The thickness of the internal groove 2531 controls the tight fit of the connection between the two. As described above, the proximal end portion 147 of the elastic aortic adapter is compressed with a strain of about 10 - 30% to provide the bonding force necessary to counter pulsatile pumping.

[0123] The current design of the interface connection between the blood pump 62 and the aortic adapter 14 has two hemodynamic merits in reducing in-situ thrombosis. First, it has been observed that in the conventional docking connection, there is actually no step or gap-type joint discontinuity. Second, the stagnant flow at the boundary of the beak front edge 85 can be minimized. For this reason, the blood flow through the connection interface flows at a high speed, and the defects of the docking connection are significantly improved. That is, the forward or backward steps 101, 102 or gaps 103 that occurred at the interface in the past are significantly improved.

[0124] The gently tapered slope 149 of this embodiment is inclined at an inclination angle with respect 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 gently tapered slope 149 has inherent defects in achieving concentric centering of the mating counterparts. For the connection between the aortic adapter 14 and the beak portion 82, there is no strict lateral restraint to ensure connection alignment. To concentrically connect the rigid beak portion 82 and the semi-rigid shallow slope 149, it is important to surround the entire peripheral edge of the flange base 252 and simultaneously engage the locking ring. If the simultaneous locking / locking engagement is not completed, the initially locked shallow slope 149 tends to distort, tilt, or position the remaining contact surfaces more than the other free parts, resulting in a connection to an eccentric pump. Such an eccentric joint is usually a factor causing steps or gaps at the boundary and can cause thrombosis. This drawback is compensated for by arranging the locking ring contour 259 (FIG. 34A) on the distal (lower) side of the locking ring 253 to include all circumferential contact areas simultaneously with the locking engagement. During locking, the contact edge of the metal (in some embodiments) beak portion 82 sinks slightly into the shallow slope 149 of the depth-controlled compression silicone resin material (in some embodiments), further reducing the boundary discontinuity when exposed to blood flow. By appropriately administering an anticoagulant, normal boundary thrombi can be significantly reduced or eliminated.

[0125] The deformability and delivery method of the structure related to this aortic adapter 14 give special design features to the present invention. In fact, the consideration of the elasticity of the material needs to be carefully incorporated into the current design. In terms of intraoperative safety and long-term reliability, it is difficult to send the implantable prosthesis 502 into the aorta by means of an incision in the aortic wall. Therefore, in some embodiments, the material selected by the aortic adapter 14 should have a predetermined memory shape. When the device is incorporated, the adapter 14 is first crimped into a small indwelling type (the crimped adapter 14 as shown in FIG. 40), and this indwelling type ensures the rapid and safe implantation of the device. After the curled aortic adapter 14 is placed at the desired implant site, the indwelling aortic adapter 14 needs to be released and self-expand to its original memory shape.

[0126] Before inserting the aortic adapter 14, a hole with a diameter of 12-14 mm is formed in the aortic wall. When creating such an access hole, it is necessary to expand the wall surface when inserting the device, and care should be taken not to generate edges that may become the starting point of cracks. A side-biting aortic perforator as disclosed in U.S. Patent Application No. 17 / 034036 is an ideal tool for making large holes in the aorta. By simply tapping gently, a non-cracked hole can be made.

[0127] The aortic adapter 14 morphologically includes two circular tubes that are connected and form a T-shaped flow connector for performing extracorporeal circulation support. The wall thickness of the catheter is usually 1-2 millimeters, and the material used is a polymer with appropriate hardness, such as silicone resin or urethane, and the hardness is Shore A80-90 or the like. The curl-in structure is significantly different from the large stent graft 502 covered with commercially available duck nylon or PTFE (polytetrafluoroethylene) fabric. Figure 40 shows the curl / implant configuration of the aortic adapter 14. (A metal stent 144 made of nickel-titanium alloy is inserted and deformed together with the aortic adapter 14 with a polymer matrix). As shown in Figure 40, the aortic adapter 14 is folded by inserting the collapsed catheter into a part of the aortic adapter catheter 142, and accordingly, the T-shaped convex neck portion 143 is crushed and flattened and sinks into the folded adapter 14 body. Among these, the diameter of the bent aortic adapter 14 is about half of the original unfolded circular diameter.

[0128] The adapter 14 folded in this way can hold the fixation by stretching a rope. As shown in FIG. 40. Three fixing strings can be installed at both edges and the central position of the catheter, but other fixing methods are also conceivable. FIGS. 41A, 41B, 41C, and 41D show four representative stages of the indwelling aortic adapter 14. The first stage (see FIG. 41A) shows an initial fluoroscopic image passing through the entry hole in the form of a curled pre-roll bladder when the pre-roll bladder (of the aortic adapter 14) is inclined at a certain angle with respect to the axis of the aorta 95. The second stage (FIG. 41B) shows that the aortic adapter 14 in a fully inserted configuration enters the aorta, and the front wrap of it passes through the entry hole, and one end of it rotates and drops into the entry hole. Then, the position of the fully inserted pre-roll bladder is returned, and the curled neck portion 143 is aligned with the entry hole (FIG. 41C). The constrained binding string is released so that the aortic adapter 14 placed in the curled compression configuration elastically self-expands to return to its original shape (FIG. 41D). The released aortic adapter 14 is tightly surrounded by the aortic lumen and is ensured by an appropriate oversized ratio selected before insertion. The T-shaped convex neck portion 143 can be slightly deformed, and the flange base 252 of the coupler 25 (FIGS. 33 to 34B) can be attached to the T-shaped convex neck portion 143 to prepare for connecting the blood pump 62. For this purpose, the inlet beak portion 82 is set on the shallow tapered slope 149, and then the two locking rings 253 of the coupler 25 are closed to realize a strong device connection, so that the connection of the blood pump can be easily realized.

[0129] Additional safety measures can be applied to enhance the hemostasis and stability of the implanted paracorporeal blood pump system. Due to the weight of the blood pump 62 and the pumping force by counterpulsation support, the installation of the blood pump near the aorta inevitably involves lateral forces (perpendicular to the longitudinal direction of the aorta) and torque applied to the aortic adapter 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 inserted into the adventitial layer around the hole. The purse-string suture can tightly compress the aortic wall against the inserted adapter 14 and is used as a protective measure to prevent the enlargement of the access hole. Also, the surgical tape can be wrapped around and tightened at both ends of the aortic adapter catheter 142, strengthening the integral connection between the inserted aortic adapter 14 and the aorta. With an appropriate design and winding tightening of the annular tape, double prevention of endoleakage can be achieved. In some cases, the blood pressure may exceed the upper limit without endoleakage with compliance matching. In such extreme situations, the surgical tape begins to function, acting as a hard limiter to seal the end of the separating adapter and ensure hemostasis.

[0130] As shown in FIG. 42, in some embodiments, the step-by-step installation of the aortic adapter 14 will be described in detail herein. Before starting the implantation, a curled pre-roll balloon type / compressed retention type aortic adapter 14 is prepared. After exposing the target thoracic artery through a left thoracotomy, a cross-clamp distance of about 10 cm spanning the implant site is determined. First, a hole perimeter marker is marked on the entry hole into the aorta. Then, the suture of the band is sutured outside the outer 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. When the above preparations are completed, cross-clamping and insertion of the aortic adapter are performed. These insertion procedures will be described in the procedure shown in FIG. 17. First, the aorta is crossed and clamped to provide an isolated segment without bleeding problems. Then, a large entry hole is created using a customized aortic perforator for inserting the aortic adapter 14. Next, as shown in FIGS. 41A, 41B, and 41C, the pre-roll balloon of the folded adapter is inserted and placed in the crossed and clamped aortic segment. Then, as shown in FIG. 41D, the folded adapter 14 is released and returned to its original deployed form. Then, as additional protection against hypertensive endoleak, it is tightened with a ligature and tape. Then, the coupler 25 is attached, and its step 2521 is engaged with the locking groove 1433 of the convex neck portion 143 of the aortic adapter 14 to prepare to receive the connected blood pump 62. Due to the self-alignment ability of the aforementioned coupler 25, the inlet connector 80 of the blood pump 62 can be accurately positioned and locked to the aortic adapter 14. The remaining implantation procedures are normal and include release of the cross-clip, priming of the blood pump, and assistance with starting the pump. Generally, 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 and may cause ischemic injury. To reduce such potential surgical damage to the organs, the abdominal organs and lower limbs can be perfused using femoral-femoral ECMO (femoral-femoral extra-corporeal membrane oxygenation) support.However, it is up to the surgeon to decide whether to employ ECMO support. Ordinarily, an ordinary patient can endure an ischemic time of 20 minutes.

[0131] As described above, according to an embodiment of the present invention, a ventricular assist device including a blood pump, a transfer member, and an introduction member is provided. The blood pump includes an axially symmetric elliptical blood sac and a stem assembly, and the system includes a flexible blood sac, a proximal stem, and a distal stem, wherein the flexible blood sac is connected to the proximal stem and the distal stem and serves as a stress-relieving suspension mechanism. The blood pump further includes a pump housing including a proximal housing and a distal housing, and a suspension mechanism for removing stress is connected to the pump housing. The blood pump further includes a built-in pressure sensing system in the proximal housing, and the pressure sensing system includes a blood pump pressure sensor and a pressure sensing chamber filled with an incompressible fluid for pressure transmission. The transfer member includes an air chamber, at least one electric wire and a tether included in the wall of the transfer member, and the electric wire and the tether are provided on the wall of the transfer member. The inlet connects the delivery and the pump housing.

[0132] An embodiment of the present invention discloses a pulsatile blood pump design that combines the concept of non-steady folding lines in the structure of a long-term blood sac, and can substantially extend the durability of a replacement blood pump. Also provided is a micro pressure sensing system that can be used as a reference waveform for real-time pump control based on real-time big data, long-term trend analysis, disease monitoring, and diagnosis. In addition, since the embedded pressure sensing system is non-blood-contact, the reliability requirements for constructing the embedded sensor system are significantly increased.

[0133] Embodiments of the present invention have at least one of the following advantages or effects. By connecting the delivery and the inlet of the pump housing, a compact inlet design is provided to make the transmission of codes and signals stronger and enhance fault tolerance. In addition, the introduction design after the compression point integrates the conductive wire, the air tube, and the blood pump. This compactness attribute is particularly important for implant devices. It not only simplifies the surgical operation, reduces the risk of perioperative implantation, but also helps to reduce the postoperative morbidity related to delivery infection.

[0134] In some embodiments, the introduction member is integrated with the remote housing of the pump housing, and the introduction member has a first portion that is an extension of the remote housing, and the air chamber, the tether, and the electric wire of the power transmission member are connected to this first portion. The second portion is used as a hollow connector for the carrier in conjunction with the first portion to achieve the advantages of anatomical adaptability and adaptability to any shape of the implant site.

[0135] Therefore, an object of the present invention is to provide a aortic adapter assembly capable of implanting a ventricular assist device, including a T-shaped flow connector including a catheter insertion portion and a convex neck portion, the catheter insertion portion and the convex neck portion being connected and both having a smooth surface in contact with blood, and a metal stent provided in the catheter insertion portion. The T-shaped flow connector has a polymer elastomer, and the polymer elastomer is reinforced by a metal stent having a nickel-titanium alloy material. The catheter insertion portion has a gradually thinner wall at the catheter ends at both ends thereof, the end of the catheter end has an appropriate distance to the outermost boundary of the metal stent, the catheter end has a compliant matching effect on the artery at the implant site, and the proximal end of the convex neck portion is arranged to be connected to the inlet adapter of the blood pump.

[0136] 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 in and out of a ventricular assist device near the aorta, particularly a counterpulsation blood pump. Different from many conventional flow connectors that employ various roughening methods to promote endothelialization and avoid the occurrence of thromboembolic adverse events, the aortic adapter of the present disclosure adopts a smooth surface and the concept of an insertable dummy graft 502 to construct the flow connector. Also, a flexible matching design is applied around the end of the inserted catheter, combining a wall property that gradually thins and a thin polymer supported by a superelastic nickel-titanium alloy frame to achieve the requirement of no internal leakage. The abnormally high pressure, high shear, and low-speed recirculation flow phenomena associated with the pump transport of counterpulsation near the aorta are included in the artificial surface where the catheter is inserted. Therefore, the influence of the pathological device on hemodynamics and risk factors are substantially eliminated, and long-term vascular maladaptation events such as endothelial cell erosion, lipid infiltration, smooth muscle cell proliferation, vascular stenosis, and arterial wall dissection are significantly reduced. To achieve a good connection between the semi-rigid flow adapter and the blood pump, the present disclosure provides a quick-connect coupling. This coupler has a self-aligning boundary design that minimizes step and gap discontinuities and reduces the possibility of thrombotic problems occurring at the boundary junction. In cooperation with the invention of the aortic adapter, a specially designed indwelling method is provided to ensure a quick and safe indwelling process. The curled aortic adapter is placed in a pre-wrapped / indwelling configuration, and the overall size is reduced to half of the initial size. This pre-rolled bladder adapter can be easily inserted into the aorta at the implant site and can expand itself to the initial placement, forming a flow connector that can be affixed without worry of internal leakage. It not only reduces the implantation risk during surgery but also helps reduce the postoperative morbidity associated with the flow by the device and vascular maladaptation at the implant site.

[0137] Ordinal numbers in this specification and the scope of the patent application, such as "first", "second", etc., have no sequential order relationship with each other and are only used to distinguish two different components with the same name.

[0138] After reading the foregoing description, those skilled in the art will clearly understand the changes in these embodiments. Therefore, the present invention includes all modifications and equivalents of the subject matter described in the claims for patent applications.

Explanation of Signs

[0139] 10, 20, 30, 40, 90 Blood pump device near the aorta 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 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 Staircase 102 Staircase 103 Gap 141 Blood contact surface 142 Aortic adapter catheter 143 Convex neck portion 1431 Neck portion body 1432 Extension portion 1433 Card slot 144 Metal stent 1441 Outermost boundary 145 Catheter end 146 Outer diameter 147 Proximal end of the adapter 148 Inner diameter 149 Shallow slope 252 Flange base 221 Staircase 222 Staircase 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 Slope 259 Lock ring contour 50 Connector 501 Adapter 502 Implant 503 Connector 504 Remote 63 Introducer 62A Inverted membrane 621 Groove 623 Proximal housing 625 Distal 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 stalk 6301 Sac stalk diaphragm 631 Part 1 632 Part 2 640 Remote stem 643 Neck part 650 Stem assembly 66 Exhaust port 661 Passageway 67 Carrier 6701 Air lumen 6702 Electric wire 61) Proximal end 672 Anchor adapter 673 Air tube 674 Coil 675 Outer layer tube 676 Tether 677 Silicon sheath 678 Rigid drive part connector 6781 Electrode 679 Hollow connector 60 Artery 6291 Proximal end 6292 Remote 6293 Tricuspid 6294 Fold 62C Center line 71 Battery hatch 73 User interface panel 75 Drive catheter socket 77 External power socket 79 Ventilation port 80 Inlet connector 81 Beak-shaped flange 82 Beak part 83 Connector body 84 Inner diameter 85 Beak leading edge 86 Hole 971, 991 Drive catheter in-vivo segment 973, 993 Drive catheter ex-vivo 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 Clinical Monitor T-201 Low-Speed Recirculation Region T-202 Impact Point AB1 Adapter AB2 Adapter

Claims

1. An aortic adapter assembly for use in an implantable ventricular assist device, comprising a catheter insertion portion, a convex neck portion, a T-shaped flow connector having a truss provided in the catheter insertion portion, and a coupler, wherein the catheter insertion portion and the convex neck portion are connected, and both the catheter insertion portion and the convex neck portion have a smooth surface that contacts blood, the T-shaped flow connector has a polymer elastomer and is reinforced by the truss made of a nickel-titanium alloy material, the catheter insertion portion has a wall where two catheter ends of the catheter insertion portion gradually become thinner, has an appropriate distance between the tip of the catheter end and the outermost boundary of the truss, the catheter end has a compliant matching effect on the artery at the implant site, and the proximal end of the convex neck portion is connected to the inlet adapter of the blood pump, the convex neck portion includes a neck portion body and an extension portion provided on the neck portion body, the extension portion protrudes from the neck portion body, and the maximum inner diameter of the extension portion is larger than the maximum inner diameter of the neck portion body, the coupler includes a flange base, a pair of locking rings rotatably attached to the flange base, and a latch provided on one of the locking rings for locking the locking ring, the locking ring has an internal concave groove, and the concave groove clamps the T-shaped flow connector by controlled compression to seal the T-shaped flow connector, an aortic adapter assembly.

2. The aortic adapter assembly according to claim 1, wherein the truss is co-injected with the polymer elastomer of the T-shaped flow connector and is fitted into the wall of the catheter insertion portion of the T-shaped flow connector.

3. The aortic adapter assembly according to claim 2, wherein the polymer elastomer of the T-shaped flow connector has a silicone resin material.

4. The aortic adapter assembly according to claim 1, wherein the polymer elastomer of the T-shaped flow connector is cast polyurethane.

5. The aortic adapter assembly according to claim 2, Aortic adapter assembly in which the structural compliance of the embedded truss is substantially equal to the structural compliance of the polymer elastomer.

6. The aortic adapter assembly according to claim 1, An aortic adapter assembly in which the end of the catheter that gradually tapers is sharp.

7. The aortic adapter assembly according to claim 1, wherein the convex neck portion has a shallow slope so as to fit the inlet adapter of the blood pump, An aortic adapter assembly in which the inner diameter of the convex neck portion is smaller than the inner diameter of the inlet adapter of the blood pump.

8. The aortic adapter assembly according to claim 7, An aortic adapter assembly in which the shallow slope is inclined with respect to the center line of the catheter insertion portion.

9. The aortic adapter assembly according to claim 1, An aortic adapter assembly in which the truss has a plurality of wavy structures.

10. The aortic adapter assembly according to claim 1, When the convex neck portion is joined to the inlet adapter of the blood pump, the extension portion is in close contact with the inlet adapter, and the neck portion body is provided around the inlet adapter. An aortic adapter assembly.

11. The aortic adapter assembly of claim 1, Each of the lock rings has a flange contour, and the flange contour can simultaneously join the lock ring to the edge of the flange base. An aortic adapter assembly.

12. The aortic adapter assembly of claim 1, The latch is made of a spring piece, and an aortic adapter assembly that ensures that the coupler is in a locked state so as not to accidentally disengage.

13. The aortic adapter assembly according to claim 1, 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 with respect to the hinge head and the flange base. An aortic adapter assembly.

14. The aortic adapter assembly according to claim 13, The hinge head is located on the first side of the flange base, the latch is located on the second side of the flange base, and the first side and the second side of the flange base face each other, a large artery adapter assembly.

15. The large artery adapter assembly according to claim 14, wherein the latch has a scoring that can engage with a bevel provided on one of the opposing locking rings, a large artery adapter assembly.

16. The large artery adapter assembly of claim 14, wherein the flange base has a substantially circular structure and each of the locking rings has an arcuate structure, a large artery adapter assembly.

17. A blood pump device as an implantable ventricular assist device, a blood pump having an inlet adapter with a tapered beak, and the large artery adapter assembly according to claim 1, wherein the large artery adapter assembly includes a catheter insertion portion and a convex neck portion, the catheter insertion portion and the convex neck portion are connected, and the catheter insertion portion and the convex neck portion both have a smooth surface that contacts blood, a T-shaped flow connector, and a truss provided on the catheter insertion portion. The T-shaped flow connector has a polymer elastomer and is reinforced by the truss having a nickel-titanium alloy material, the catheter insertion portion has a wall in which two catheter ends of the catheter insertion portion gradually become thinner, has an appropriate distance between the tip of the catheter end and the outermost boundary of the truss, the catheter end has a compliant matching effect on the artery at the implant site, and the proximal end of the convex neck portion is connected to the inlet adapter of the blood pump, the tapered beak is aligned with the proximal end of the convex neck portion, a blood pump device.

18. The blood pump device according to claim 17, wherein the inner diameter of the inlet adapter is larger than the inner diameter of the convex neck portion of the T-shaped flow connector, a blood pump device.

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