Paracorporeal blood pump device

Abstract

An aortic bypass blood pump apparatus includes a blood pump, an aortic adapter, a drive conduit, and a drive. The blood pump includes a blood bladder, a pump housing, and a pressure sensor mounted within the pump housing for monitoring blood pressure within the blood pump. The aortic adapter is a T-shaped conduit connected to the blood pump for connecting the blood pump to the aorta of a human body to provide circulatory support. The drive conduit, in addition to transmitting electronic blood pressure signals to the drive, also allows pneumatic communication with the blood pump. The drive receives and processes the electronic blood pressure signals to determine the timing, speed, and duration of blood pump fill and discharge actions to provide counter-pulsation circulatory support to assist the circulatory system of a human body.

Description

Technical Field

[0001] This 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. Priority is claimed under U.S. Provisional Application Nos. 63 / 162,086 and 63 / 162,098, filed March 17, 2021, the contents of which are incorporated herein by reference. Background Technology

[0002] Mechanical circulatory support technology used in the treatment of heart failure.

[0003] The progression of heart failure is characterized by a vicious cycle. Drug therapy for early to moderate heart failure is a compensatory mechanism that suppresses the production of neurohormones; such therapy slows, rather than treats, the continued deterioration of the vicious cycle. Heart transplantation or ventricular assist device (VAD) implantation is only applicable to terminal stage heart failure. During the progression of heart failure, there exists an "unmet need" period between ineffective drug therapy and heart transplantation / VAD treatment, where no treatment is available. As heart failure worsens beyond the effective drug therapy stage, patients often cannot receive any further treatment and must wait for the heart failure to progress further to the terminal stage where heart transplantation / VAD implantation can be considered. This treatment window is categorized in the Interagency Registry for Mechanical Assisted Circulatory Support (INTERMACS) classification of heart failure levels 4–7. Patients with this type of heart failure typically exhibit symptoms of apnea, renal insufficiency, and exercise intolerance, accompanied by elevated systemic inflammation.

[0004] The optimal strategy for treating heart failure with mechanical circulatory support (MCS) is to implant a left ventricular assist device (LVAD) in time before heart failure becomes irreversible. However, this strategy is practically unfeasible because current turbopumps or continuous-flow LVADs are highly invasive during implantation and involve a considerable rate of postoperative complications, making LVADs only permissible for patients with end-stage heart failure.

[0005] Based on nearly fifty years of experience using intra-aortic balloon pumps (IABPs) at the bedside, antipulsation support therapy has proven clinically effective. Recently, by implanting balloon pumps through the axillary or subclavian arteries, the use of IABPs has been expanded from bedridden support to ambulation support, successfully bridging patients to heart transplants. Balloon antipulsation circulatory support and patient ambulation have shown remarkable therapeutic effects in improving cardiac function. However, improvements in implantation site surgery have limited effectiveness, patients still require hospitalization, and the duration of circulatory support cannot be prolonged.

[0006] Currently, there are no clinically approved early interventional ventricular assist devices (VAPs providing a flow rate of 2-3 liters per minute) that offer partial support, and these devices are particularly suitable for patients with milder advanced heart failure. Most of these patients with milder advanced heart failure will continue to deteriorate and develop acute myocardial ischemia or cardiogenic shock, requiring emergency mechanical circulatory support. Intra-aortic balloon pumps (IABP), extracorporeal membrane oxygenation (ECMO), and percutaneous micro-axial catheter blood pumps are emergency life-sustaining systems that can be administered, but the mortality rate for these patients is generally high, ranging from approximately 40-70%. When acute circulatory support fails, patients, families, and clinicians often find it difficult to decide whether to use highly invasive and expensive turbofan blood pumps. Therefore, developing novel long-term implantable ventricular assist devices with innovative designs and minimally invasive surgical concepts to overcome the traditional window of opportunity in heart failure treatment is a major challenge for medical device manufacturers.

[0007] Early intervention partially supports left ventricular assist devices (LVADs).

[0008] This invention discloses an early interventional partial support blood pump system, also known as a para-aortic blood pump device, based on the principle of anti-pulsation support. The para-aortic blood pump device of this invention provides anti-pulsation circulatory support, including increasing systemic blood flow during diastole (cardiac relaxation) to improve myocardial and organ perfusion, while reducing left ventricular workload during systole (cardiac contraction). For patients with mechanical circulatory support devices, early hemodynamic support and postoperative walking ability are crucial for survival and disease improvement. Minimally invasive surgery is a prerequisite for early intervention; therefore, in addition to the hemodynamic therapy provided by left ventricular assist devices, minimally invasive surgery constitutes another key factor. Using specially designed surgical tools, the para-aortic blood pump device of this invention can be implanted through a less invasive surgical procedure (e.g., a minimally invasive left thoracotomy). Clinical evidence shows that the heart supported by a left ventricular assist device may undergo reverse remodeling after long-term systolic unloading therapy. This finding suggests that early walking support for mechanical unloading is particularly beneficial for patients with milder conditions (intermacs classification 4-7) who are transitioning to recovery or remission.

[0009] To date, approved long-term implantable left ventricular assist devices (LVAPs) are primarily continuous-flow turbine pumps and are only suitable for patients with end-stage heart failure. A treatment window exists between the stage of drug ineffectiveness and the final stage of LVAP support or heart transplantation. The para-aortic pump device of this invention fills this treatment gap by proposing an early, partial support approach, and the device can be implanted through a minimally invasive procedure.

[0010] Prior to heart transplantation or implantation of a highly invasive turbo pump, providing prolonged circulatory support through the creation of a smaller para-aortic pump device is a reasonable and beneficial treatment option for patients. Some patients may recover with circulatory support provided by the para-aortic pump device through therapeutic antipulsation support administered in a non-bedridden manner for several months (1-2 years) or longer. For patients with hemodynamically stable but irreversible heart failure, the para-aortic pump device can be used long-term as a destination therapy device or as a bridge to transplantation. Therefore, the para-aortic pump device can serve as a temporary bridging therapy option before heart transplantation or expensive and highly invasive turbo pump implantation. Thus, the role of the para-aortic pump device in the treatment of heart failure is multifaceted; it can bridge to recovery (or remission), bridge to decision-making therapy (turbo pump), bridge to heart transplantation, or be an alternative to heart transplantation (destination therapy). Summary of the Invention

[0011] One of the main objectives of this invention is to overcome the shortcomings of existing technologies by disclosing a para-aortic blood pump device. The device includes a blood pump, an aortic adapter, a driveline, and a driver. The blood pump includes a pump housing, a blood sac, and a pressure sensor. The pressure sensor is installed inside the pump housing to monitor blood pressure inside the pump. The sensed blood pressure is converted into an electrical signal and transmitted to the driver via the driveline. The aortic adapter is a T-shaped catheter connected to the blood pump for connecting the blood pump to the aorta. The aortic adapter is thin-walled and may have a reinforced design, embedded around the catheters on both sides of the aortic adapter to enhance anti-buckling strength. The driveline is connected to the pump housing of the blood pump, providing pneumatic pulses to the blood pump and transmitting the electronic blood pressure signal received from the pressure sensor. The driver is coupled to the pressure sensor to receive the electronic blood pressure signal. The actuator includes an electromechanical actuator (EMA), whose controller generates pneumatic pulse commands based on sensed electronic blood pressure signals. These pneumatic pulses are transmitted back and forth to the blood pump via a drive catheter, supporting the body's circulation while simultaneously performing the pump's ejection and filling functions.

[0012] In some embodiments, the driver further includes a drive catheter controller and a vibrator, wherein the drive catheter controller is used to process electronic blood pressure signals, and the vibrator is used to provide an audible alarm or a tactile feedback.

[0013] In some embodiments, a portion of the drive catheter and driver is replaced by a distal drive line, a drive line interconnector, and a proximal drive line, with the distal drive line used to transmit electronic blood pressure signals and pressure pulses to the blood pump, and the drive line interconnector including a drive line controller and a vibrator, the drive line controller being used to process the electronic blood pressure signals, and the vibrator being used to provide an audible alarm or a tactile feedback.

[0014] In some embodiments, the blood pump and the aortic connector are integrally formed.

[0015] In some embodiments, the para-aortic blood pump device further includes a coupler, and the coupler includes a coupling adapter mounted on the neck of the aortic connector for coupling the blood pump to the aortic connector.

[0016] In some embodiments, the two flared ends of the aortic connector form a smooth transition of elastic properties, gradually softening as the wall thickness distribution of the aortic connector catheter decreases.

[0017] In some embodiments, the blood pump housing contains a blood sac, which is an oval-shaped membrane body of revolution with the blood pump centerline as the center of rotation. It has two polymeric stems at both ends that are attached to the membrane body. When attached to the blood pump housing, the polymeric stems can be used as a flexing / stretching relief mechanism to reduce stress concentration when the membrane body deforms.

[0018] In some embodiments, the pump housing of the blood pump has an opening that is integrally formed with the aortic connector, providing a seamless, smooth interface for the connection between the blood pump and the neck of the aortic connector.

[0019] In some embodiments, the electromechanical actuator includes a pressure equalization valve connected to a cylinder, which opens periodically to equalize the air pressure inside the cylinder with atmospheric pressure.

[0020] In some embodiments, the para-aortic blood pump device provides a counterpulsating gain of systemic blood flow during diastole (cardiac relaxation) to improve myocardial and organ perfusion, while reducing left ventricular workload during systole (cardiac contraction).

[0021] In some embodiments, the driver includes a trigger detection microcontroller unit that provides a sensed pressure waveform within the blood pump from an electronic blood pressure signal, allowing the trigger detection microcontroller unit to calculate and determine the exhaust and filling times of the electromechanical actuator cylinder.

[0022] In some embodiments, the electromechanical actuator includes a motor and a ball screw unit that drives a reciprocating piston in a cylinder of the electromechanical actuator; the reciprocating piston draws in and ejects air through a drive conduit connected to a blood pump.

[0023] In some embodiments, the electromechanical actuator is a pneumatic actuator, comprising: a brushless servo motor and a ball screw unit and a piston / cylinder assembly, reciprocating to inject / fill the blood pump using the atmosphere as a driving medium, and a pressure balancing valve is installed on the cylinder wall of the pneumatic actuator to solve the problems of piston ring leakage and water vapor condensation from the blood sac.

[0024] In some embodiments, the driver receives an electronic blood pressure signal and processes the electronic blood pressure signal using a trigger detection algorithm to generate a trigger signal that commands the driver to actuate in sync with the heart rhythm.

[0025] In some embodiments, upon receiving a specified trigger time, the trigger detection microcontroller unit sends a command to the motor controller to drive the piston of the electromechanical actuator to follow the piston position from injection to filling or from filling to injection to provide anti-pulsating cycle support, including performing systolic unloading during cardiac systole and perfusion gain during cardiac diastole, respectively.

[0026] In some embodiments, the driver further includes a user interface module, which includes an indicator, an audible alarm, a button, and a liquid crystal display.

[0027] In some embodiments, when the trigger detection microcontroller unit loses the electronic blood pressure signal from the blood pump, the trigger detection microcontroller unit automatically starts a washout mode to drive the electromechanical actuator to operate at a predetermined pumping frequency and stroke volume to flush the blood sac and prevent thrombus formation.

[0028] In some embodiments, the flushing mode is used to prevent thrombus formation in the blood pump; this is a device protection mode rather than providing circulatory support.

[0029] In some embodiments, the blood sac is anchored to a proximal shell of the pump housing via a proximal port and to a distal shell of the pump housing via a distal port. Attached Figure Description

[0030] Figure 1 This is a schematic diagram of a para-aortic blood pump device according to a first embodiment of the present invention;

[0031] Figure 2 This is a schematic diagram of a para-aortic blood pump device according to a second embodiment of the present invention;

[0032] Figure 3 This is a schematic diagram of a para-aortic blood pump device according to a third embodiment of the present invention;

[0033] Figure 4 This is a schematic diagram of a para-aortic blood pump device according to a fourth embodiment of the present invention;

[0034] Figure 5 This is a schematic diagram of a blood pump device installed in the human body beside the aorta according to the first and second embodiments of the present invention;

[0035] Figure 6 This is a schematic diagram of a blood pump device installed in the human body beside the aorta according to the third and fourth embodiments of the present invention;

[0036] Figure 7 This is a first schematic diagram of a driver according to an exemplary embodiment of the present invention;

[0037] Figure 8 This is a second schematic diagram of a driver according to an exemplary embodiment of the present invention;

[0038] Figure 9 This is a schematic diagram of the driver functions and main interconnect signals necessary for the operation of this invention;

[0039] Figure 10 This is a schematic diagram of the driver functions and main interconnect signals necessary for the operation of the third embodiment of the present invention;

[0040] Figure 11 This is a schematic diagram of the main interconnect signals of the driver functions necessary for the operation of the second embodiment of the present invention;

[0041] Figure 12 The piston position trajectory of the electro-mechanical actuator (EMA) associated with anti-pulse cycle support and the timing of trigger detection command issuance are depicted.

[0042] Figure 13A This is a perspective view of the para-aortic blood pump implant according to the first or third embodiment of the present invention;

[0043] Figure 13B This is a cross-sectional view of the para-aortic blood pump implant according to the first or third embodiment of the present invention;

[0044] Figure 14A This is a perspective view of the para-aortic blood pump implant according to the second or fourth embodiment of the present invention;

[0045] Figure 14B This is a cross-sectional view of the para-aortic blood pump implant according to the second or fourth embodiment of the present invention;

[0046] Figure 15 This is another embodiment of the invention, showing the connection between the drive catheter and the pump housing of the blood pump. The connection of the drive catheter is achieved through an introduction on the distal housing.

[0047] Figure 16 yes Figure 15 The diagram shows a cross-sectional view of the blood pump.

[0048] Figure 17 The shallow trench design is shown for extending the wire from the inlet located in the distal housing to the pressure sensing chamber in the proximal housing;

[0049] Figure 18A This demonstrates the connection of the external contralateral anastomosis of the blood pump to the artery using an interface adapter connector;

[0050] Figure 18B This demonstrates how the blood pump of the present invention is coupled to an artery via a plug-in connection method using a T-shaped intravascular connector coupled through an interface adapter.

[0051] Figure 19 A cross-sectional view of an integral axisymmetric elliptical blood sac and a rotating body including the sac, proximal port, and distal port is shown.

[0052] Figure 20 An exploded view is shown, illustrating the components used to construct the axisymmetric elliptical sac and port assembly; (Note: in conjunction with integration into...) Figure 19 (The sac is in its original shape before the proximal and distal ports are shown);

[0053] Figure 21 It showed that Figure 19 The curved trilobed sac structure at the end of the ejaculation of the blood sac is shown.

[0054] Figure 22 A perspective view of the drive catheter is shown, which is connected to the proximal housing of the blood pump via an inlet.

[0055] Figure 23A It showed along Figure 6 A cross-sectional view of the distal side of the blood pump and drive catheter in section AA;

[0056] Figure 23B It showed along Figure 22A cross-sectional view of the proximal end of the driving catheter at section AA in the diagram;

[0057] Figure 24 A cross-sectional view of the exhaust port installed in the proximal housing corresponding to the first embodiment is shown;

[0058] Figure 25A A cross-sectional view of the pressure sensing chamber and the inlet in the proximal housing of the first embodiment is shown. (Note: The drive conduit is not installed; the inlet includes a first portion, an extension of the proximal housing, and a second portion interlocked with the first portion.)

[0059] Figure 25B It shows the combination in Figure 25A A perspective view of a microelectromechanical system (MEMS) pressure sensor in a photograph;

[0060] Figure 26 A cross-sectional view of the multilayer drive conduit of the present invention is shown, including an inner tube for pneumatic air transmission, an intermediate tube for electrical signal transduction, and a coil, a tether, and an outer tube.

[0061] Figure 27 A cross-sectional view of a multi-lumen drive catheter design is shown.

[0062] Figure 28A A typical view showing the flow characteristics during the pump filling stage is displayed;

[0063] Figure 28B A typical view showing the flow characteristics during the pump injection phase is displayed;

[0064] Figure 29 This is a perspective view of the T-type flow connector of this embodiment;

[0065] Figure 30 This is a cross-sectional view of the T-type flow connector in this embodiment;

[0066] Figure 31 This is a schematic diagram of the unfolded plan of the embedded nickel-titanium alloy metal support;

[0067] Figure 32 The lateral stiffness (LS) of the insertion catheter for measuring nickel-titanium alloy metal stents and T-type flow connectors was defined.

[0068] Figure 33 This is an exploded schematic diagram of the coupler and its components;

[0069] Figure 34A This is a schematic diagram of a coupler in the open (configured) state;

[0070] Figure 34BThis is a schematic diagram of a coupler in a locked (configured) state;

[0071] Figure 35 This is a cross-sectional schematic diagram showing the connection between the aortic connector and the para-aortic blood pump using a coupler.

[0072] Figure 36A A schematic diagram of the step discontinuity caused by the docking method is shown;

[0073] Figure 36B A schematic diagram of the gap discontinuity caused by the docking method is shown;

[0074] Figure 37 This is a schematic diagram of an inlet connector for installation at the distal end of a blood pump located beside the aorta.

[0075] Figure 38 This is a cross-sectional view of the aforementioned inlet connector;

[0076] Figure 39 The downward-facing conical beak of the inlet connector is shown, which connects to the ramp surface at the neck of the T-type flow connector;

[0077] Figure 40 A schematic diagram of a crimp-fit ​​flow connector is shown, which is wound into an insertion configuration (wrap form) by a rope bundle;

[0078] Figure 41A An insertion flow connector is shown passing through an access hole formed in the aortic wall;

[0079] Figure 41B A flow connector in a wrapped (compressed) form, fully inserted into the lumen of the aorta, is shown.

[0080] Figure 41C The repositioned, compressed flow connector is shown with its T-shaped neck facing the aortic inlet.

[0081] Figure 41D An expanded, unfolded flow connector is shown, its T-shaped neck popping out and unfolding in the aortic inlet after the rope is released; and

[0082] Figure 42 The illustration shows the individual steps of implanting the aortic connector into the target aortic segment and connecting it to the blood pump. Detailed Implementation

[0083] The following are four embodiments of the para-aortic blood pump device that can be used to implement the present invention, as described below.

[0084] See Figure 1This is a schematic diagram of a para-aortic blood pump device according to a first embodiment of the present invention. The para-aortic blood pump device 10 includes: a blood pump 12, an aortic connector 14, a drive catheter 16, and a driver 18. The blood pump 12 also includes a pump housing and a pressure sensor. The pump housing is composed of two chambers, one for storing blood and the other for receiving drive air. These two chambers are separated by an egg-shaped flexible membrane, which is suspended and fixed by a pair of stress-relief stems connected to the pump housing. The pressure sensor is installed inside the pump housing of the blood pump 12 to monitor the blood pressure within the blood pump 12 and generate an electronic blood pressure signal. The aortic connector 14 is a valveless, T-manifold-shaped flow connector coupled to the blood pump 12 and the human aorta. In a first embodiment, the aortic connector 14 and the blood pump 12 are integrally formed, having a seamless blood contact surface, and the blood pump 12 is connected to the human aorta via the aortic connector 14. The aortic connector 14 is made of a flexible material, allowing it to deform during insertion and delivery through a circular hole made in the aortic wall. The aortic connector 14 self-expands after insertion into the aorta and is strong enough to resist radial compressive contact forces of the aortic lumen applied to the connector wall from an oversize fitting. The drive catheter 16 is connected to the housing of the blood pump 12 for providing pneumatic pulses to the blood pump 12 and transmitting blood pressure signals received from a pressure sensor. The actuator 18, coupled to the drive catheter 16, receives the transmitted electronic blood pressure signal and includes an electromechanical actuator to generate pneumatic pulses based on the electronic blood pressure signal, which are delivered to the blood pump 12 via the drive catheter 16. The wearable driver 18 provides a pneumatic pulse control rhythm that is coordinated with the heart rhythm to drive the ejection and filling action of the implanted blood pump 12.

[0085] The aforementioned driver 18 includes a battery power supply system 11 and a backup battery power supply system (hereinafter). Figures 2 to 4 as well as Figure 8 The battery power supply systems 21, 31, and 41 in the device are the same or similar, wherein the backup battery power supply system ensures continuous power supply to the driver 18. For convenience, the driver 18 can also be powered via an AC adapter when the patient does not need to be moved. In addition, this device has a clinical monitor, which is not included in the... Figures 2-4As shown, it can be connected to drive 18 to provide clinicians with a user interface for displaying device monitoring or diagnostic information and for accessing drive parameters, so as to enable initial setup of patient data and optimization of specific treatment operation mode settings.

[0086] Figure 1 and Figure 2 Two different blood pump 12, 22 designs are shown that are coupled to the same drive catheters 16, 26 and actuators 18, 28 system. Figure 2 This is a schematic diagram of a para-aortic blood pump device 20 according to a second embodiment of the present invention. The difference between the second embodiment and the first embodiment is that the para-aortic blood pump device 20 of the second embodiment further includes a coupler 25, or coupling adapter. In the second embodiment, the blood pump 22 and the aortic connector 24 are not integrally formed but are detachable, and the coupler 25 is provided to couple the blood pump 22 to the aortic connector 24. The coupler must be carefully designed to minimize discontinuities at the connection interface. During device implantation, the aortic connector 24 is first inserted into the aorta through an insertion hole cut in the aortic wall. The device can be implanted using a specially developed implantation tool, which mounts the coupling adapter 25 around the T-shaped neck of the aortic connector 24, allowing the blood pump 22 to connect to the coupling adapter 25. After the blood pump 22 is placed in the thoracic cavity, the blood pump 22 and the aortic connector 24 are securely locked and integrated together by the coupler 25. This design of the detachable blood pump 22 and aortic connector 24 offers advantages both during and after implantation. During device implantation, the detachable blood pump design facilitates aortic connector placement because the surgical area is more clearly defined and unobstructed by the pump itself. Furthermore, after implantation, the blood pump can be removed and replaced if the pressure sensor malfunctions or the blood bladder ruptures, requiring emergency replacement. In this respect, the detachable blood pump design of the second embodiment is advantageous. The aortic connector can remain in the aorta without removal, avoiding the hassle and risk of redo surgery associated with aortic connector removal.

[0087] refer to Figure 1 and Figure 3These are schematic diagrams of the para-aortic blood pump devices 10 and 30 according to the first and third embodiments of the present invention, respectively. The difference between the third embodiment and the first embodiment is that the drive catheter 16 of the first embodiment is replaced by the distal drive catheter 37, drive catheter interconnector 33, and proximal drive catheter 39 of the drive catheter 36 of the third embodiment. The distal drive catheter 37 is connected to the drive catheter interconnector 33 for transmitting electronic blood pressure signals acquired from a pressure sensor and pneumatic pulses sent from the driver 38; and the drive catheter controller and vibrator (for alarm purposes) included in the drive catheter interconnector 33 were originally included in the driver 18 of the first embodiment, thus the driver 18 of the first embodiment has an additional drive catheter controller and vibrator (compared to the driver 38 of the third embodiment). The drive catheter controller processes the electronic blood pressure signal, and the vibrator provides audible alarms or tactile feedback. In other words, the mechanical power transmission of the blood pump, as well as the analog / digital signal conversion and alarm notification implemented by the drive catheter 16 and driver 18 of the first embodiment, are substantially the same as those implemented by the distal drive catheter 37, drive catheter interconnector 33, and proximal drive catheter 39 of the third embodiment.

[0088] The first embodiment features a simpler drive catheter configuration design and incorporates the electronic signal processor within the actuator, thus minimizing the risk of environmental contamination (water ingress or moisture condensation) and air leakage at the connector, both of which are associated with the drive catheter interconnect 33. However, this elongated drive catheter is more susceptible to contact damage, such as abrasion, kinking, or cuts caused by contact with foreign objects during daily activities. Any significant damage to the drive catheter 16 of the first or second embodiment, whether electronic or mechanical, may require surgical replacement of the blood pump. This is highly undesirable given the risks of repeat surgery and associated medical costs. The third or fourth embodiment mitigates this disadvantage of blood pump replacement associated with drive catheter damage by employing an intermediate connector (drive catheter interconnect). Generally, the distal drive catheter 37 has a shorter exposed length and is better protected by the drive catheter interconnect 33 and skindressing and patient vest coverage. In extreme cases where the drive catheter is severely damaged beyond repair, the most likely damaged proximal drive catheter 39 can be easily replaced without resorting to surgery. Furthermore, the third or fourth embodiment is less affected by electromagnetic interference because the analog-to-digital signal conversion has already been completed in the circuitry of the drive catheter interconnector 33. In the third or fourth embodiment, because the digital signal transmission in the proximal drive catheter 39 is less sensitive to electromagnetic interference, the signal fidelity of the pressure signal can be better guaranteed.

[0089] Please see Figure 3 and Figure 4 These are schematic diagrams of the para-aortic blood pump devices 30 and 40 according to the third and fourth embodiments of the present invention, respectively. The difference between the third and fourth embodiments is that the aortic connector 34 and blood pump 32 in the third embodiment are integrally formed; while the blood pump 42 and aortic connector 44 in the fourth embodiment are detachable; the fourth embodiment also includes the blood pump 22, aortic connector 24, and coupler 25 as in the second embodiment. The blood pump 42, aortic connector 44, and coupler 45, which are identical in the embodiments, will not be repeated here. The distal driving catheter 47, driving catheter interconnector 43, and proximal driving catheter 49 included in the driving catheter 46 are the same as those in the driving catheter 36.

[0090] Reference Figure 5 This is a schematic diagram of a para-aortic blood pump device implanted in the human body according to an exemplary embodiment of the present invention. The para-aortic blood pump device 90 includes a blood pump 92, an aortic connector 94, an in-body drive catheter segment 991, an external drive catheter segment 993, and an actuator 98. In another embodiment, the para-aortic blood pump device further includes a coupler. The portion of the para-aortic blood pump device 90 implanted in the human body includes the blood pump 92, the aortic connector 94, and the in-body drive catheter segment 991. In another embodiment, the para-aortic blood pump further includes a coupler. During surgery, the aortic connector 94 is implanted into the aorta 95, and an exit site EX of the drive catheter is created at an appropriate location on the human epidermis. The external portion of the para-aortic blood pump device 90 includes the external drive catheter segment 993 and the actuator 98. A section of the in-body drive catheter segment 991, bounded by the exit site EX, is covered with fabric velour to promote tissue ingrowth for infection control. The implanted velvet portion is best placed 2 to 5 centimeters subcutaneously from the exit site EX. The driver 98 is a wearable or portable device.

[0091] Reference Figure 6This is a schematic diagram of a para-aortic pump device implanted in the human body according to an exemplary embodiment of the present invention. The para-aortic pump device 90 includes a pump 92, an aortic connector 94, a distal drive catheter 97 (including an in-body drive catheter segment 971 and an external drive catheter segment 973 outside the human body), a drive catheter interconnector 93, a proximal drive catheter 99, and an actuator 98. In another embodiment, the para-aortic pump device 90 also includes a coupler. The portion of the para-aortic pump device 90 implanted in the human body includes the pump 92, the aortic connector 94, and the in-body drive catheter segment 971. In another embodiment, the para-aortic pump device also includes a coupler. During surgery, the aortic connector 94 is implanted into the aorta 95, and a drive catheter exit site EX is created at an appropriate location on the human skin. The portion of the para-aortic pump device 90 located outside the human body includes the external drive catheter segment 973, the drive catheter interconnector 93, the proximal drive catheter 99, and the actuator 98. The exit point EX marks the boundary of the drive catheter. The distal drive catheter is divided into an internal drive catheter segment 971, covered with velvet for infection control, and an external drive catheter segment 973. The actuator 98 is a wearable or portable device.

[0092] The implanted subsystem is described further below.

[0093] Implantation is achieved via a left thoracotomy through a relatively small thoracic opening using a less invasive surgical (LIS) technique. For example, a thoracic incision is made at the 7th intercostal space as the primary opening to allow for the implantation of the aortic connector and blood pump. Two additional small incisions are made at the 6th and 8th intercostal spaces to introduce proximal and distal aortic clips. The area between the aortic clips allows the aortic connector to be implanted through an opening in the aortic wall. The aortic connector is flexible and can be rolled down to a smaller delivery shape before implantation. Once implanted in the aorta, the aortic connector automatically springs back to its original shape with a super-tight fit to the lumen at the intended implantation site. Therefore, the material of the aortic connector is important; it should be flexible but have sufficient radial strength to keep the implanted aortic connector wall round without wall buckling. Candidate aortic connector materials can be selected from silicone or polyurethane elastic polymers, or the structural strength can be enhanced by embedded reinforced polymer structures or metallic materials.

[0094] The following further describes the aortic connector and its functional requirements for each of the above embodiments.

[0095] In hemodynamics, the aortic coupling provides a blood flow connection between the blood pump and the systemic circulation. In addition to this function, the aortic coupling also serves as a support structure, securing the blood pump to the coupling. The aortic coupling must be structurally flexible yet anti-buckling, and robust enough to withstand internal blood pressure and external contact forces generated by the blood pump through contact with surrounding lung tissue or by respiration and diaphragmatic movement.

[0096] Aortic connector 54 is implanted into the aorta, with its two catheter tips (catheter tip 545 and catheter tip 645) intersecting with the aortic lumen, forming a host / graft interface in the blood flow (see...). Figure 13B , Figure 14B To minimize morphological and elastic discontinuities at the host / implant interface, the two catheter tips (catheter tip 545 and catheter tip 645) are configured with a flared inner surface profile and a continuously decreasing wall thickness distribution. This catheter tip design minimizes the steps at the interface and incorporates the compliance-matching effect required for connection into the aortic connector design. The likelihood of thrombosis at the interface is thus significantly reduced because clot aggregation at the interface will occur at a slower rate than the natural thrombolysis provided by the human aortic endothelium. Furthermore, the gradually thinning catheter wall structure makes the catheter tips (catheter tip 545 and catheter tip 645) more flexible (compliant), causing the catheter tips to expand and contract with pulsating blood pressure, creating a dynamic seal to prevent blood cells from becoming trapped in the interfacial gaps that are typically sources of thrombosis.

[0097] The drivers of the above embodiments are further described below.

[0098] Figure 7 , Figure 8 These are perspective views of the right and left sides of the drive unit 78. The compact internal modules of this drive unit 78 include an electromechanical actuator (EMA), an electronic controller, a pair of main batteries, and a backup battery. The drive unit 78 also includes a user interface panel 73, a battery compartment door 71, a drive conduit socket 75, an external power socket 77, and a pair of vents 79, as shown. Figure 7 , Figure 8 As shown.

[0099] Critical operational information, as well as alarms for equipment malfunction and aortic pressure status, will be displayed on the user interface panel 73 of the actuator 78. The main battery can be replaced through the battery compartment door 71 when its power is low. When the patient is bedridden and can use power from a wall socket for extended periods, the actuator 78 is powered via a cable connected to an external power outlet 77. One end of the proximal drive catheter 99 and the external segment 993 of the drive catheter can be connected to the actuator 78 through a drive catheter socket 75, through which voltage sensor signals and pneumatic pressure pulses are communicated. A pair of vents 79 are installed on opposite sides of the actuator 78 to allow ambient air to flow through the interior of the actuator 78 for cooling purposes.

[0100] The driver 78 can be externally coupled to a clinical monitor to facilitate the collection and display of real-time clinical waveform data and the storage of patient data for long-term status monitoring and diagnosis. Furthermore, the clinical monitor unit provides a user interface for clinicians and patients, displaying device monitoring / diagnostic information and for setting driver parameters for initial startup of the driver 78 and optimizing patient-specific cyclical auxiliary operating modes.

[0101] In this embodiment, the EMA is a pneumatic actuator, comprising a brushless servo motor, a ball screw unit, and a piston / cylinder assembly. Air, as a driving medium, is used to reciprocate and actuate the blood pump to achieve the functions of ejecting and filling the blood pump.

[0102] The pneumatic actuator is housed within the actuator and carried by the patient undergoing treatment. The electromechanical actuator comprises a brushless servo motor, a piston and cylinder assembly, and a ball screw unit. The ball screw unit includes a ball screw rod and a ball nut. The piston is fixed to the top of the ball screw rod, and the rotation of the nut drives the rod rod to reciprocate linearly. The servo motor includes a rotor and a stator, with the rotor and ball nut integrated. Through electromagnetic induction, the rotor rotates, and the clockwise and counterclockwise rotations achieve the linear reciprocating stroke motion of the rod rod and piston within the cylinder. The piston's reciprocating stroke drives air from the cylinder through drive wires into the blood pump, completing the ejection and filling of the blood sac.

[0103] Using air as a medium to drive a blood pump presents two problems: air leakage and condensation caused by blood seeping through the sac wall. The former impairs the pump's pumping and filling efficiency, as well as the motor's power consumption; the latter poses a risk of bacterial growth within the drive wires due to moisture. To address these issues, this ventricular assist device's pneumatic actuator incorporates a pressure balancing valve on the cylinder wall. This valve allows air within the cylinder to exchange with the atmosphere, enabling airflow during pressure balancing. The pneumatic actuator is equipped with position and optical sensors for the controller to acquire piston position information. The controller then generates piston drive commands to drive the piston's reciprocating motion and operate the pressure balancing valve. Therefore, the timing and frequency of the pressure balancing valve's opening can be programmed and stored in the controller. Through the operation of this pressure balancing valve, air within the cylinder exchanges with the outside atmosphere, achieving both air replenishment and drying functions, thereby ensuring the safety and effectiveness of the ventricular assist device's operation.

[0104] Reference Figure 9 According to an embodiment of the present invention, the para-aortic blood pump device is divided into three parts. The first part is mainly installed inside the human body (i.e., an implant), and its outer end communicates with the second part. The first part includes a blood pump (including a blood pump pressure sensor), an aortic connector, and a distal drive catheter segment, all implanted inside the human body. The second part is installed outside the human body and includes a proximal drive catheter and a drive catheter electronic module (or drive catheter interconnector). The third part is also installed outside the human body and is a driver, including an electromechanical actuator (EMA), a controller circuit, a main battery, and a backup battery.

[0105] The blood pump pressure sensor is integrated into the proximal blood pump housing and embedded in a small pressure sensing chamber filled with sensing medium, allowing for continuous monitoring of the blood pump pressure. The distal drive catheter connects to the pump housing and provides counter-pulsating pneumatic pulses to actuate and fill the sac. Both the distal and proximal drive catheters supply the blood pump with pneumatic drive pulses generated by the EMA inside the actuator; and transmit the electronic blood pressure signal generated by the sphygmomanometer pressure sensor to the actuator. The drive pneumatic path (indicated by dashed arrows) and the electrical signal path (indicated by solid lines) are shown below. Figure 9 The diagram illustrates the functional relationships between the interactive actuation modules. Details of the aortic connector have already been described previously. The controller circuitry may include a motor controller unit for driving the brushless motor, and a microcontroller unit serving as a central processing unit to process the received pressure signals and generate control commands for the motor controller to actuate the piston movement.

[0106] Reference Figure 10 and Figure 11This is a schematic block diagram showing the internal functions of the driver and the key interconnect signals necessary for blood pump startup. Further description of embodiments of the present invention is provided, such as... Figure 10 and Figure 11 As shown. (Refer to...) Figure 10 and Figure 11 To explain the driving relationship between the external actuator and the implant, it is necessary to refer to the aforementioned contents of the blood pump, driving catheter, distal driving catheter, proximal driving catheter, aortic connector and driving catheter interconnector.

[0107] The driver receives the blood pump pressure signal (electrical signal) and processes it using a trigger detection algorithm to generate a trigger signal. This trigger signal commands the motor actuator to actuate in coordination with the heart rhythm. Upon receiving the specified trigger time, the microcontroller unit sends a command to the motor controller (unit) to drive the piston from ejection to filling or from filling to ejection, providing anti-pulsation cycle support.

[0108] The architecture of the electronic controller comprises three functional blocks: a microcontroller unit (MCU), a motor control block (or motor controller unit), and a power management unit. The table below provides a descriptive overview of each functional block of the driver 78.

[0109]

[0110] exist Figure 10 , 11 The present invention describes, with reference to the previously illustrated exemplary embodiments, signal acquisition, transmission, processing, control logic and instruction generation, and how the electromechanical actuator is actuated to generate pressure pulses to drive the blood pump.

[0111] Figure 12 The trigger detection command for the piston position of the electromechanical actuator related to anti-pulsation assistance is described. Figure 12In the diagram, the unassisted aortic pressure (AoP) waveform is represented by a dashed line, while the solid line represents the assisted aortic pressure waveform. When the actuator is running in automatic mode, actuator operation is initiated, and the system performs a "fill-eject-fill-eject..." cycle of assistance, representing normal synchronous antipulsation assistance operation. The MCU monitors the blood pump pressure (BPP) signal (electrical signal) and detects the left ventricular end-diastolic (LVED) timing. Upon detecting the LVED timing, the MCU generates an F_Trig signal. The time interval between two consecutive F_Trig signals represents the instantaneous cardiac cycle interval (or cycle). Based on the estimated heart rate calculated from several preceding cycle intervals, the MCU determines the blood pump ejection time, i.e., the E_Trig signal. The E_Trig signal provides timing to command the motor controller unit to drive the electromechanical actuator according to a predetermined position, velocity, and acceleration curve. After the ejection stroke is completed and an optimized dwell time has elapsed, the electromechanical actuator is commanded to perform a pre-filling action at a slower filling rate until the F_Trig signal appears. Upon receiving the F_Trig signal, the electromechanical actuator begins to execute the remaining filling stroke at the specified piston speed.

[0112] When the MCU loses the BPP signal (electrical signal) sent by the blood pump, the MCU automatically activates the flushing mode to drive the electromechanical actuator at a predetermined auxiliary frequency and driver pulse volume. The flushing mode is used to prevent thrombus formation within the blood sac and is a device protection mode, not to provide synchronous anti-pulsation circulation support.

[0113] The para-aortic blood pump device of the present invention, in principle, offers better anti-pulsation support than the intra-aortic balloon pump (IABP) due to its non-obstructive aortic insertion feature. Unlike bedridden or mobile IABP patients who must remain in the hospital, this para-aortic blood pump device allows patients to leave the hospital and enjoy a better life at home. Therefore, in addition to the economic benefits of shorter hospital stays, the para-aortic blood pump device of the present invention can further improve the patient's condition and quality of life.

[0114] In recent years, the use of LVADs has approached saturation, primarily because its application is only suitable for a small group of patients with end-stage heart failure. Early intervention with LVADs in patients with milder heart failure has long been a clinical goal in cardiology. Early intervention is expected to have a significant impact on the expanded use of LVADs and influence future advancements in cardiology. Clinical evidence suggests that LVAD support in moderate to severe heart failure, in some patients with non-ischemic cardiomyopathy, can improve cardiac function or produce sustained myocardial recovery through reverse remodeling of cardiomyocytes. However, this intention for early intervention relies on two favorable driving factors: simple and safe surgical procedures and an effective adaptive circulatory support program that accompanies the progression of the disease. Continuous-flow VAD support is non-physiologic, as it deviates the supported heart from the normal healthy recovery pathway. However, antipulsation support is physiological, promoting reverse myocardial remodeling to achieve therapeutic goals by providing systolic unloading and diastolic perfusion gain. In summary, the treatment strategy provided by this para-aortic blood pump invention aligns with the development trend of early intervention in cardiac medicine. The beneficial efficacy attributes of the para-aortic blood pump device, such as adaptive partial support, minimally invasive surgery, and antipulsation therapy, collectively make this invention a potential candidate to contribute to future advancements in the treatment of heart failure.

[0115] The blood pumps of each of the above embodiments are further described below.

[0116] Figure 13A and Figure 13BThese are schematic diagrams and cross-sectional views of the first and third embodiments of the present invention, showing a para-aortic blood pump device implanted in the human body. The implantable subsystem of the para-aortic blood pump device includes: a blood pump 52, an aortic connector 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, which is closed proximally and open distally, and seamlessly integrated with the aortic connector 54. The blood sac 529 is composed of an egg-shaped capsule 526, and is anchored to the proximal shell 523 of the housing 52h via a proximal port 530, and to the distal shell 525 of the rigid housing 52h via a distal port 540. Within the housing of the blood pump 52, the egg-shaped flexible capsule 526 divides the space into a blood chamber B and an air chamber A, and the capsule 526 is suspended from the rigid housing 52h via a pair of stress-relief stems (proximal port 530, distal port 540). Blood chamber B stores blood, and air chamber A receives driving air. A pressure sensing mechanism 527 (or blood pressure sensor) is hermetically embedded in the proximal housing 523 of the rigid housing 52h. The sensed pump pressure is transmitted through a membrane 526 and propagates in an incompressible liquid or gel contained in a closed pressure sensing chamber 528, and is ultimately received by the pressure sensing mechanism 527. Upon receiving the sensed blood pump pressure, the pressure sensing mechanism 527 generates an electronic blood pressure signal. The implantable subsystem components are designed to be implantable and suitable for patients with a body surface area (BSA) of 1.2 square meters or more.

[0117] Figure 14A and Figure 14B These are schematic and cross-sectional views of portions of the para-aortic blood pump device implanted in the human body according to the second and fourth embodiments of the present invention, respectively. The implantable subsystem of the para-aortic blood pump device includes: a blood pump 62, an aortic connector 64, a coupler 65, and a drive catheter (or distal drive catheter) 67 connected to the blood pump 62. The coupler 65 is used to connect the blood pump 62 to the aortic connector 64 for access to the vascular system of the device implantee. The blood pump 62 includes a rigid housing 62h, which further includes a proximal housing 623 and a distal housing 625. The construction of this blood pump 62 is similar to... Figure 14B The disclosed structure, except that the opening OP is separate and independent from the aortic connector 64, is used. The coupler 65 is positioned around the neck 643 of the aortic connector 64. In the following text, it will be referred to as... Figure 14B The design disclosed in the paper is used to further explain the blood pump design and basic design principles.

[0118] See Figure 13BThe blood pump 52 includes a molded rigid housing 52h, which also includes a proximal housing 523 and a distal housing 525. The rigid housing 52h has a single opening OP that connects to the aortic connector 54 for access to the patient's vascular system. The opening OP of the blood pump 52 is seamlessly manufactured together with the aortic connector 54 (or the blood pump 52 and the aortic connector 54 are integrally molded), which provides a smooth and continuous interface transition to the neck of the aortic connector 54. Engagement of this integrated sac 529 and aortic connector 54 assembly with the blood pump 52 is achieved by bonding the proximal housing 523 and the distal housing 525 to the proximal port 530 and the distal port 540, respectively. The sac 529 is anchored to the top of the proximal housing 523 such that the non-flexible disc portion of the sac 529 is bonded close to the pressure sensing chamber 528 at the center of the proximal housing 523.

[0119] A miniaturized pressure sensing mechanism 527 is integrated within a proximal housing 523 and communicates with a closed pressure sensing chamber 528 via a fluid medium. This device allows for continuous monitoring of blood pressure in a sac 529. Because the pressure sensing mechanism 527 does not come into contact with blood, the rigid housing 52h ensures long-term sensor reliability and fidelity by isolating the pressure sensing mechanism 527 and its circuitry from the effects of chemical corrosion and protein adhesion caused by direct blood contact.

[0120] The drive catheter 57 is connected to the distal housing 525 to provide counter-pulsating pneumatic pulses to drive blood from or into the blood pump 52. The drive catheter may be multi-lumen or multi-layered to cover the wiring used for pressure signal transmission. Metal coils or fabric meshes or nets may be used as catheter wall reinforcements to enhance the anti-kinking capability of the distal drive catheter 57. The overall wide geometry of the flow channels in this blood pump, combined with the valveless aortic connector design and pulsating pump operation, constitutes excellent blood handling characteristics, avoiding hemolysis caused by high shear forces and thrombus formation or thromboembolism caused by low flow rates.

[0121] The innovative design of the aforementioned blood pump 52 and blood sac 529 makes the capsule 526 highly durable. The blood sac 529 is an egg-shaped membrane rotator with the centerline of the blood pump 52 as its rotation center. Two polymer ports (proximal port 530 and distal port 540) are attached to both ends of the rigid housing 52h, respectively constructed as a disc or annulus. When connected to the rigid housing 52h, it serves as a bending / tensile stress relief mechanism to reduce stress concentration. During blood pump ejection, the capsule 526 will be compressed or folded into a tri-lobe shape, with the maximum strain typically occurring at the creases near the edges of the ports (proximal port 530 and distal port 540). This localized high membrane stress / strain caused by large membrane deformation is absorbed and offset by the flexible suspension deformation at the port edges. It is particularly noteworthy that the position of the tri-lobe folding pattern is not fixed; the membrane deformation is influenced by the direction of gravity, and the crease position occurs randomly. In fact, a patient's body posture and orientation, including standing, sleeping, sitting, and exercising postures, may change from time to time during daily activities. Therefore, the gravitational effect acting on the blood volume stored in the blood pump 52 will constantly change direction, resulting in the formation of non-fixed fold lines. This randomly formed membrane fold line characteristic constitutes the unique fatigue-resistant feature of this invention. It is anticipated that this blood pump 52 will have greater durability than conventional fixed-fold-line membrane designs.

[0122] The folding and expansion of the capsule are closely related to the vortex structure pattern contained within the blood sac 529. The aforementioned blood sac design is characterized by randomly formed fold lines, causing the vortex structure features to alternate with changes in the folded membrane pattern. Therefore, the flushing effect in the blood pump 52 is strong and unpredictable, characterized by random, wandering vortex motion. This randomness of the blood pump's vortex structure helps flush the entire blood-contact surface of the blood sac without creating any fixed, low-velocity backflow zones near the membrane wall or in the folded areas. Animal studies have observed that the blood pump of the present invention possesses very strong antithrombotic capabilities.

[0123] Each distal driving catheter in each embodiment is further described below.

[0124] See Figure 5 , Figure 6 The diagram shows a drive conduit 96 that connects the blood pump 92 to the actuator 98. The distal drive conduit 97 (drive conduit body section 991) of the drive conduit 96 pneumatically connects the blood pump 92 to an electromechanical actuator housed within the actuator 98, and also transmits pressure from the blood pump pressure sensing mechanism 527 (see...). Figure 13BThe distal drive catheter 97 (intradermal segment 991) acquires electrical signals. One end of the distal drive catheter 97 is connected to the blood pump housing, and the other end has a small external connector for pneumatic transmission and electrical communication. The distal drive catheter 97 (intradermal segment 991) exits the skin subcutaneously under the protection of a protective cap. The outer diameter of the distal drive catheter 97 (intradermal segment 991) is designed to be small, and the tubing material is flexible to minimize stress on the exit site EX for patient comfort. A portion of the distal drive catheter 97 (intradermal segment 991) is covered with porous fabric to promote tissue ingrowth into the fabric, thereby providing infection resistance to the exit site EX. The external segment 973 of the drive catheter, which exits the epidermis, is secured a short distance outside the skin exit site EX.

[0125] The inner segment 991 and outer segment 993 of the driving catheter, along with their connectors, are designed to withstand tensile loads applied during percutaneous surgery. Postoperatively, the inner and outer segments 991 and 993 are continuously subjected to loads induced by muscle movement, and are designed to withstand these fatigue loads throughout their intended service life. The external portions of the inner and outer segments 991 and 993 are also designed to be biocompatible and chemically resistant to cleaning agents and disinfectants for clinical use.

[0126] The proximal driving catheter of each of the foregoing embodiments is further described below.

[0127] The external segment 993 and proximal drive catheter 99 are used to connect the internal segment 991 and distal drive catheter 97 to the actuator 98, respectively. The proximal drive catheter 99 has a drive catheter interconnect 93 at one end and an actuator connector at the other. The drive catheter interconnect 93 includes a circuit board that converts analog blood pump pressure signals into digital signals and a vibrator that provides tactile feedback in addition to an audible alarm. The drive catheter interconnect 93 has a flat shape to prevent twisting of the external segment 973 of the drive catheter when it is secured to the patient's skin. Furthermore, the drive catheter interconnect 93 and the embedded wiring within the drive catheter are designed to be sealed and prevent water or moisture intrusion. Because the proximal drive catheter 99 is externally mounted, it can be replaced or maintained as deemed necessary, thus eliminating the need for surgical replacement of the blood pump when the proximal drive catheter 99 is damaged beyond repair.

[0128] Valveless blood pumps offer two advantages in blood processing characteristics: 1) they eliminate unpleasant valve noise and valve-induced blood cell damage, thrombosis, and thromboembolism; 2) they exhibit stronger antithrombotic properties due to the superior surface-cleaning effect of bidirectional pulsatile flow, minimizing protein adhesion and preventing the formation of blood clots associated with interfacial discontinuities on artificial surfaces in contact with blood. The flow channels in valveless pulsatile pumps are significantly wider and more uniform than those in valved pulsatile or continuous-flow rotary pumps. Hemolysis (red blood cell membrane rupture) typically occurs in narrow flow channels with high velocity gradients, such as the gap between the valve annulus and leaflets in valved pulsatile pumps. Additionally, low-velocity recirculation or stagnation zones often exist on the back side of an open valve, which can promote thrombus formation. In stark contrast, in a valveless pulsatile blood pump, the shear stress applied to blood cells is actually several orders of magnitude smaller, and the low-velocity stagnation zone associated with valve geometry and movement is essentially eliminated, thereby reducing damage to blood cells or platelet activation, reducing the formation and aggregation of blood clots, and translating to the use of lower doses of anticoagulants, as well as simpler and safer postoperative care.

[0129] Figures 15 to 17 A schematic diagram of a blood pump 62, a drive catheter 67, and an inlet 63 according to another embodiment of the present invention is shown. This embodiment emphasizes anatomical adaptation to facilitate easier externalization of the blood pump and drive catheter.

[0130] like Figure 15 and Figure 16 As shown, the drive conduit 67 is connected to the distal housing 625 of the blood pump 62. The blood pump 62 has an elliptical blood sac and port assembly 659 (including sac 629 and ports 630, 640), a pump housing 62h (having a proximal housing 623 and a distal housing 625 with an inlet connector 6251), and a pressure sensing system 628 embedded in the proximal housing 623. The aforementioned components, inlet connector 6251, pressure sensing system 628, and drive conduit 67 are substantially the same as or correspond to the components / elements of the aforementioned embodiments, and a detailed description of these elements and their functions will not be repeated here.

[0131] In this embodiment, an inlet 63 is disposed in the distal housing 625 of the pump housing 62h for connecting the drive conduit 67 to the pump housing 62h. Furthermore, the inlet 63 is configured in a body shape adjacent to the distal housing 625, thereby connecting the power drive conduit to the outer surface of the pump in a tangential direction. This through-body design allows the pump housing 62h to be adapted to the anatomical space available for pump placement. The pump 62 is rotatably connected to the interface adapter 501 and allows the drive conduit 67 to be routed in the most suitable orientation, enabling smooth subcutaneous tunneling and skin exit. This facilitates anatomical adaptation to the geometry of the implantation site.

[0132] In this embodiment, the inlet 63 is remotely adapted into the remote housing 625, while the sensor 6271 (e.g.) Figure 25A The pressure sensing chamber 628 is located in the proximal housing 623. Further engineering work needs to be performed to separate the signal transmission path from the pneumatic communication path and to ensure that the blood pump 62 is sealed and protected against biochemical fluid intrusion, which may impair the fidelity of signal transmission after device implantation.

[0133] See Figure 17 The pump housing 62h has a surface groove 621 formed on the outer surface of the distal housing 625 and located in the overlapping mating region DA of the proximal housing 623 and the distal housing 625. Figure 16 Above. Surface groove 621 is configured to allow wires to extend from the outlet of inlet 63, the wires being along surface groove 621 above the overlapping engagement area DA and reaching the second space 6273. Figure 25B The electrode 6274. In some embodiments, the aforementioned groove 621 is sealed by a potting waterproof material and / or annular cap to maintain a smooth outer surface so as not to irritate or damage the tissue in contact.

[0134] like Figure 18A and Figure 18B As shown, the aortic connector 50 is typically required via its interface adapter 501 to serve as a connection mechanism for connecting the blood pump 62 to the target artery 60. The distal end 504 of the connector 50, opposite the interface adapter 501, is positioned within the vessel wall of the artery 60 and is in fluid connection with the human circulatory system. The proximal end of the aortic connector 50 (or interface adapter 501) has a smooth interface transition to geometrically match the inlet shape of the blood pump 62. A coupler is typically required to integrate the proximal end of the connector 50 (or interface adapter 501) and the inlet of the blood pump 62. In some embodiments, this can be used as the aortic connector 50. Figure 18A The illustration shows an end-to-side anastomosis of a Dacron or PTFE graft 502 sutured with a target artery (or vessel) 60, which can be used in vascular surgery. In some other embodiments, for example... Figure 18B The illustrated embodiment uses an insertable aortic connector 503, such as the T-shaped manifold adapter disclosed in U.S. Patent Application No. US2008 / 0300447A1, entitled "Dual-pulsation bi-Ventricular Assist Device".

[0135] Figure 19 and Figure 20A long-term axisymmetric elliptical sac and port assembly 650 and its components are shown. The polymer material chosen for these components can be, but is not limited to, segmented polyurethane with various suitable hardness testers. The aforementioned sac and port assembly 650 comprises a flexible membrane sac (sac) 629, a proximal port 630, and a distal port 640, which, in some embodiments, are all formed in an axisymmetric shape relative to a common centerline 62C of the blood pump 62 and integrated together. The proximal port 630 is located at the proximal end 6291 of the sac 629, while the distal port 640 is located at the distal end 6292 of the sac 629.

[0136] Figure 19 An integrated blood sac and port assembly 650 is shown, the end of which of the distal port 640 is wrapped at the distal end of the blood sac 629 and combined with the inverted membrane 62A. Figure 20 The components before assembly are shown. Generally, the sac 629 is manufactured by dip molding, while ports 630 and 640 are manufactured by injection molding. There is no preferred azimuth angle for deflecting sac deformation. Theoretically, when the applied pressure differential exceeds a certain threshold, the thin-walled sac with an axisymmetric elliptical structure will bend into a three-lobed configuration 6293, as shown. Figure 21 As shown. This membrane bending is only related to the final three-lobed configuration 6293 (eigenmode), in which the location of the crease 6294 or fold line is determined by the initial perturbation that induces bending instability. The thickness uniformity in the cross-section cut perpendicular to the centerline (or axis of rotation) 62C of the component 650 of the blood pump 62 is crucial. Care must be taken to maintain high-precision pocket manufacturing to ensure an axisymmetric shape. In real life, the direction of gravity is the primary factor initiating the fold line. The recipient's posture changes constantly according to the patient's daily activities (e.g., standing, sitting, exercising, sleeping, etc.), and the direction of gravity relative to the blood pump direction also changes constantly. Therefore, the crease 6294 of pocket deformation appears in a random pattern, causing high-strain creases to be non-stationarily distributed throughout the pocket. Therefore, avoiding leaving high-strain areas in a fixed position is a key design principle for ensuring a long pocket life.

[0137] Embodiments of this invention innovate the operation of fold lines, which causes high-strain locations to appear non-stationarily within the membrane, thereby extending the fatigue life of the sac. This improves the detrimental stress concentration often associated with bent sacs. Based on this fundamental change in bending mode behavior, the membrane's fatigue life is significantly increased, attributed to this non-stationary fold line formation characteristic, which disperses high-strain regions throughout the sac. Furthermore, the beneficial results accompanying this non-stationary sac deformation mode depend on the enhanced eddy current scouring effect within the sac. The sac surface is more thoroughly cleaned, forming irregular, walking-like eddies that traverse the entire sac. This greatly reduces the likelihood of the formation of constant low-velocity recirculation zones or fold line creases near the wall, thus achieving a long-lasting, thrombosis-resistant pump design.

[0138] Figure 22 , Figure 23A as well as Figure 23B An exemplary embodiment is shown in which an integration method is used to mount the blood sac and port assembly 650 onto the pump housing 62h and to connect the drive conduit 67 to the proximal housing 623.

[0139] The sac 629 is anchored to the pump housing 62h, which includes a proximal housing 623 and a distal housing 625 to facilitate the pump's filling and discharging actions. Generally, the flexural characteristics of the sac 629 and the pump housing 62h differ significantly. To achieve the sac design for long-term use, an intermediate suspension is required to ensure continuity of the pump assembly during structural characteristic transitions, particularly membrane bending deformation. A pair of flexible ports (proximal port 630 and distal port 640) are used as a suspension mechanism to integrate the sac 629 with the pump housing 62h. Figure 23A As shown, the disc-shaped proximal port 630 is connected to the proximal housing 623; while the annular distal port 640 is connected to the distal housing 625. Mechanically, the proximal port 630 and the distal port 640 act as stress-relieving suspension mechanisms, which not only hold the blood sac 629 within the pump housing 62h but also prevent stress concentration at the interface attachment point, thus extending the service life of the blood sac 629.

[0140] like Figure 23A As shown in the lower part, the distal housing 625 includes an extension of the inlet connector 6251, which is coupled to the aortic connector 14. The inlet connector 6251 has a first end 6252 and a second end 6253 that are attached to the sac 629, wherein the second end 6253 is beak-shaped, and the second end 6253 is coupled to the interface adapter 501 (in Figure 18A , Figure 18B(As shown in the diagram) Coupling. The first end 6252 of the inlet connector 6251 smoothly mates with the distal end of the blood sac 629. However, the opposite second end 6253 is configured to mate with the interface adapter 501, and the coupling design aims to minimize interface discontinuities to avoid clot formation. The beak has a flange structure arranged in the middle region of the inlet connector 6251, which serves as a locking element received by the interface adapter 501.

[0141] During surgical procedures, the closed-end balloon design of the valveless blood pump 62 will draw air into the top of the balloon and cause air bubbles to accumulate due to buoyancy. (See reference) Figure 22 and Figure 24 An vent 66 is installed or provided in the proximal housing 623, wherein a narrow channel 661 is provided above the integrated sac port diaphragm 6301 between the proximal port 630 and the sac 629. In some embodiments, the channel 661 extends along a centerline 62C. After the blood pump 62 is anastomosed with the target artery 60, the trapped air is pushed out by arterial blood pressure and accumulates in the top space of the sac 629. A fine needle is used to pass through the vent 66, through the channel 661, and the sac port diaphragm 6301 to reach the interior of the sac 629 to expel the accumulated air. The integrated sac port diaphragm 6301 below the vent 66 is relatively rigid and non-flexible, which will keep the perforated sac 629 from further structural damage and prevent structural damage due to crack propagation induced at the perforation gap when subjected to periodic pulse pressure and stretching and folding of adjacent sacs.

[0142] like Figure 23A and Figure 25A As shown, a pressure sensing mechanism 627 is embedded in the proximal housing 623. Figure 23A Cross-sectional details of the integrated proximal housing 623, feedthrough 63, and drive conduit 67 are shown. Figure 25A The outline of the proximal housing 623 is shown, which is connected to an inlet 63 extending from the dome of the proximal housing 623 for pneumatic and signal communication with the drive conduit 67.

[0143] like Figure 25A , Figure 25B As shown, the aforementioned pressure sensing mechanism 627 includes a sensor 6271, which is hermetically housed in a metal can and has a first space 6272 for fluid communication and a second space 6273 for housing a microelectromechanical system (MEMS) pressure sensor and related electronic circuitry. Multiple electrodes 6274 extend from the bottom of the second space 6273 to connect to wires 6702 of the drive conduit 67 (e.g., ...). Figure 26The second space 6273 is closer to the drive conduit 67 than the first space 6272. The first space 6272 is open to fluid communication with the sensing medium. A biocompatible fluid or jelly is used as the pressure transmission medium. A cavity or pressure sensing chamber 628 is formed in the proximal housing 623 and adjacent to the first space 6272 to allow the sensing fluid to be contained therein. The distal end of the pressure sensing chamber 628 is separated from the blood chamber by a membrane sac 629. The pressure sensing chamber 628 has two side arms: a first arm 6281 and a second arm 6282, wherein the first arm 6281 is used to mount the sensor 6271, and the second arm 6282 is used to fill and seal the sensing medium. Thus, blood pressure pulses can be transmitted across the membrane sac 629 and hydraulically communicated with the remote MEMS sensor 6271 located in the second space 6273.

[0144] One embodiment of the present invention innovates a pressure-based blood pump control method and sensor design. A miniature MEMS pressure sensor is employed, with its electronic circuitry packaged and embedded within the pump housing wall. In principle, MEMS sensor chips are highly durable due to their inherent microscale structure. In practice, the sensor's durability depends on the packaging design. This pressure sensing mechanism 627 is non-blood-contact and isolated from blood-related corrosive biochemical effects, thereby providing the long-term signal acquisition and transmission required for long-term implantable assistive devices.

[0145] The drive catheter 67 serves as a transmitter for converting electrical signals and transmitting pneumatic pulse pressure between the blood pump 62 and the driver 98. Figure 26 The diagram illustrates a representative multilayer drive conduit 67 of the present invention. In this embodiment, the drive conduit 67 has a pneumatic inner cavity (or inner pneumatic tube) 6701, multiple wires 6702, an intermediate pneumatic tube 673, a coil 674 (e.g., a metal coil), an outer tube 675, a tether 676, a silicone sheath 677, a rigid actuator connector 678, and a protective hollow connector 679.

[0146] The central portion of the drive catheter 67 houses a pneumatic inner lumen 6701 (or air channel, inner tube) with an inner diameter of 2-5 mm, which can be selected depending on whether lower energy consumption or lower surgical simplicity is desired. A wire 6702 for signal transmission is embedded in the wall of the drive catheter 67. Other variations of the drive catheter design are also possible. Figure 26 In addition to the multi-layered drive conduit 67 design shown, the drive conduit 67 can also be multi-lumen, for example, to facilitate the embedding of the wire 6702 in several smaller lumens and allow pulsed air to flow in the larger lumen 6701, such as... Figure 27 As shown. One of the smaller lumens can be installed together with the tether 676 to limit the extension of the drive conduit 67 and protect the wire 6702 from damage under external tension.

[0147] The inner tube or pneumatic cavity 6701 is received by the pneumatic tube 673, with reinforcement sandwiched in between. Between the inner tube 6701 and the pneumatic tube 673, a coil 674 (or fabric thread or mesh) can be recirculated (using heat-shrinkable heat-co-molding) as reinforcement for the drive catheter wall, thereby making the drive catheter 67 flexible but kink-resistant. The outer tube 675 covers the inner and intermediate pneumatic cavities 6701 and the pneumatic tube 673, and can be used to cover the spirally wound wire 6702 as a protective sheath. In some embodiments, a non-expandable tether 676 may be disposed between the outer tube 675 of the drive catheter 67 and the silicone sheath 677 to enhance the tensile elasticity required during the externalization of the drive catheter 67. Clinically, the silicone sheath 677 has been shown to cause minimal irritation to subcutaneous tissue and deliver the lowest infection rate.

[0148] In this embodiment, the pneumatic inner cavity 6701, metal coil 674, pneumatic tube 673, spiral wire 6702, outer tube 675, tether 676, and silicone sheath 677 are packaged into the body of the drive conduit 67. The proximal end 671 of the drive conduit 67 is inserted into a socket provided in the actuator 98. The rigid actuator connector 678 of the drive conduit 67 is received by the socket in the actuator 98. The rigid actuator connector 678 is connected to a plurality of electrodes 6781 (e.g., wires 6702 soldered to them). Figure 23B The four electrodes 6781 are flush-mounted. Protective hollow connector 679 (see...) Figure 23B , Figure 26 The drive catheter 67 is placed on the junction of the drive catheter 67 and the actuator connector 678 to prevent the drive catheter 67 from kinking at the junction. The proximal end 671 of the drive catheter 67, including the actuator connector 678 and the hollow connector 679, is designed for easy, low-profile removal from the skin without causing unwanted puncture trauma.

[0149] like Figure 22 and Figure 23A As shown, the connection between the drive catheter 67 and the blood pump 62 is achieved via an introduction 63. Depending on the anatomy in which the blood pump 62 is to be implanted, the introduction 63 can be placed in the proximal housing 623 or the distal housing 625. Integrating the introduction 63 with the pump housing 62h can alter the overall configuration of the external blood pump and guide the drive catheter 67 in a specific direction to meet implantation requirements, including the externalization route of the drive catheter, postoperative skin care, and device availability.

[0150] like Figure 23A and 25AAs shown, the introducer 63 has a first portion 631 extending from the proximal housing 623, and the pneumatic lumen 6701 of the drive catheter 67, the tether 676, and the wire 6702 are coupled to the first portion 631 via an anchoring adapter 672. The introducer 63 also has a second portion 632, which interlocks with the first portion 631 and serves as a hollow connector for the drive catheter 67. The first portion 631 is the location for wire connection, tether anchoring, and pneumatic lumen bonding and sealing with the housing. The wires must not be exposed to the implantation site tissue and must be well protected to prevent tension during the externalization of the power-driven catheter. Furthermore, the connection between the pneumatic lumen 6701 and the blood pump 62 requires no air or current leakage. The aforementioned blood pump integration tasks are performed in the first portion 631. The second portion 632, on the other hand, is responsible for housing these interface elements and acts as an external protector to protect the interface from mechanical stress and intrusion by environmental fluids or moisture.

[0151] Figures 19 to 27 A modular design relating to a first embodiment of the blood pump of the present invention is disclosed. In this embodiment, the blood pump 62 includes an axisymmetric elliptical blood sac and port assembly 650 (including a flexible membrane sac 629, a proximal port 630, and a distal port 640); a pump housing 62h having a proximal housing 623 and a distal housing 625; and a drive conduit 67 connected to the blood pump 62, the drive conduit 67 including a pneumatic cavity 6701 and wires 6702 within its walls. To integrate the drive conduit 67 with the pump housing 62h, an inlet 63 is used to provide electrical and pneumatic communication between the drive conduit 67 and the blood pump 62.

[0152] like Figure 19 and Figure 20 As shown, the connection design of the sac 629 and ports 630 and 640 has been disclosed in the previous sections. A key design and manufacturing consideration is maintaining high-precision axisymmetry in parts manufacturing and the connection of the sac and port assemblies. The pressure sensing mechanism 627 and the inlet 63 are mounted in the rigid portion of the proximal housing 623. Figure 22 , Figure 23A as well as Figure 23B The compacted inlet 63 design is shown. It can be seen that the compacted inlet 63 enables more robust and fault-tolerant wiring and connections.

[0153] Figure 28A and Figure 28B Some flow patterns associated with para-aortic counterpulsation are shown. During the end-diastolic and early systolic phases of left ventricular ejection, the pump undergoes pump filling and draws aortic blood into the pump ( Figure 28ABlood upstream and downstream around the connector is drawn into the pump via a sharp 90-degree flow deflection. This creates flow separation and a low-velocity recirculation zone, T-201. Furthermore, very high shear occurs at the corner of the T-connector. On the other hand, during diastole after aortic valve closure, blood stored in the pump is ejected back into circulation, creating a spurting flow on the contralateral aortic wall. Figure 28B This lateral, impingement flow exhibits very high local pressure at the impact point T-202, a so-called stagnation point where the flow velocity is virtually zero and all kinetic energy associated with the flow velocity is converted into potential energy called total pressure. This high-pressure impingement flow can lead to vascular maladaptation, including smooth muscle cell proliferation and resulting vascular wall stenosis, as well as the risk of aortic dissection due to persistent local hypertension. All these non-physiological flow patterns and induced high-pressure, high-shear, low-velocity recirculation phenomena are prevalent near the T-tube. This turbulent, complex flow anomaly will attenuate or decrease over a distance of 3-5 times the diameter of the implanted artery lumen. This insertable flow connector is designed with an insertion catheter length of 5-7 cm, covering most of the non-physiological flow area induced by the pump. Because the aorta at the implantation site is shielded by the inserted flow connector, the bio-vascular wall will be protected from the pathological stress conditions induced by the pump, thus protecting the implantation site artery from acute or long-term remodeling complications.

[0154] In counterpulsation support, the pump's filling and ejection alternately drive the heart rhythm, producing a effect similar to... Figure 28A and Figure 28B The special T-shaped connector shown flows. The aforementioned insertable aortic coupling 14 is respectively in... Figure 29 3D diagram and Figure 30 Further details are provided in the sectional view.

[0155] The aortic connector 14 is molded, and its internal blood contact surface 141 is ultra-smooth and continuous, without any parting lines. The aortic connector 14 may be made of silicone or other polymeric elastomers. In some embodiments, the aortic connector 14 has a polymeric elastomer comprising a silicone material, or the polymeric elastomer is mold-injectable polyurethane. The aortic connector 14 includes a method for insertion into the aorta 95 (see...). Figure 5The aortic connector catheter (or catheter insertion portion) 142 in the blood pump and the neck (or protruding neck portion) 143 connected to the blood pump. In this embodiment, the protruding neck portion 143 has a neck body 1431 and an extension 1432 disposed on the neck body 1431, wherein the extension 1432 protrudes from the neck body 1431, and the maximum inner diameter of the extension 1432 is larger than the maximum inner diameter of the neck body 1431. When the protruding neck portion 143 is connected to the aforementioned blood pump 62, the extension 1432 is in close contact with the inlet connector 6251 of the blood pump 62, and the neck body 1431 is surrounded by the aforementioned coupler 25, which integrates the inlet connector 6251 into the blood pump 62.

[0156] The integral aortic connector 14 is thin-walled to maximize flow efficiency. To reinforce its thin-walled structure, a pair of nickel-titanium alloy metal trusses (or metal stent rings) 144 are embedded in and surround the ends of the aortic connector catheter 142 of the aortic connector 14.

[0157] Figure 29 This is a schematic diagram showing the insertion position of the nickel-titanium alloy metal stent 144. Furthermore, the wall thickness of the aortic connector catheter 142 gradually decreases towards the two catheter tips 145. This gradually thinning wall thickness serves a dual purpose. First, it minimizes discontinuities in the graft / master graft connection and ensures that the interfacial clot formation rate is consistently lower than the thrombolysis rate provided by the contact endothelium. Second, the catheter's compliance becomes softer towards the catheter tips 145, thus creating a compliance-matching effect when connected to the aortic lumen.

[0158] One of the complications plaguing the delivery of large stent grafts is endoleak. Type I endoleak refers to an incomplete seal between the graft tip and the implanted arterial endothelial lumen, resulting in a gap between the graft tip and the arterial lumen. Leaking blood becomes trapped in this gap and coagulates into a clot, eventually forming a fibrous pseudointimacy that grows uncontrollably over time. This pseudointimacy not only obstructs the transplanted artery but can also signal and stimulate the coagulation mechanism to attract platelet adhesion and lead to thrombotic adverse events. The solution to this endoleak problem is to ensure a tight seal between the aortic connector 14 and the attached lumen surface. The aortic connector 14 disclosed herein presents a compliant matching design concept that allows the semi-rigid catheter (terminal) tip 145 to seamlessly attach to the arterial lumen under pulsating blood pressure. Figure 30As shown, the outer diameter 146 of the aortic connector catheter 142 is slightly larger than the diameter of the inner lumen, with an oversize ratio (defined as the ratio of the catheter diameter to the inner lumen diameter) in the range of 3-10% under a given nominal blood pressure (e.g., 120 mmHg). As blood pressure fluctuates between systole and diastole, or under pulse pressure generated by counterpulsation support, the compliant matching catheter tip 145 will dynamically expand and contract in response to pressure pulsations without creating an interfacial gap.

[0159] Thin-walled tubes made of elastomers are flexible and tend to be compliant, but their strength is insufficient to withstand the compressive forces applied due to oversized devices, which often leads to wall bending of the inserted connector. Therefore, it is important to use a combination of a nickel-titanium alloy metal support structure (144) and an elastomer substrate with appropriate stiffness. Figure 30 As shown, in this design, the radial stiffness provided by the nickel-titanium alloy metal stent 144 helps support the aortic connector 14 without bending it, and a distance “X” exists between the outermost boundary 1441 of the metal stent 144 and the catheter tip 145. In some embodiments, the aforementioned distance X should be evaluated and properly defined. Supported by the nickel-titanium alloy metal stent 144 as an expandable frame, the gradually thinning catheter tip 145 does not collapse or wrinkle, but remains rounded, resting against the connected lumen wall and dynamically sealing with the lumen. It should be noted that the aortic connector 14 can expand and contract in response to pressure pulsations, and the sealing effect is achieved dynamically, i.e., the aortic connector 14 and the wall of the aorta 95 expand and contract together as a whole to seal the catheter tip 145 without causing bleeding complications.

[0160] Figure 31 A representative embodiment of a nickel-titanium alloy metal support 144 is shown, which is typically made of a laser-engraved nickel-titanium alloy tube and further expanded under a series of expansion and heat treatments. Figure 31 A planar unfolded schematic diagram of the metal support 144 is shown. (Each row) of the metal support has multiple wavy structures. The metal support 144 is self-deployable, foldable or rollable into a smaller pre-wrapped insertion configuration, and self-released to return to its original shape after being placed in the desired position.

[0161] A simple measurement of duct stiffness (the reciprocal of compliance) can be expressed by the so-called lateral stiffness (LS), the measurement method of which is described in... Figure 32The lateral stiffness is defined as the force F per unit length divided by the corresponding radial deflection Y. For current aortic connectors, a suitable lateral stiffness range is 0.01–0.05 Nt / mm². The embedded nickel-titanium alloy metal stent 144 and the silicone or elastomer substrate both contribute to the structural compliance of the co-injected aortic connector 14. The structural compliance of the metal stent 144 and the structural compliance of the polymer elastomer of the aortic connector 14 are approximately equal to each other. A uniformly distributed flexibility is preferred so that the stretching and contraction of the composite catheter wall will result in minimal delamination tendency, thereby increasing the service life of the connector 14.

[0162] Aortic connector 14 is configured to connect to blood pump 62 to facilitate circulatory support. A quick-connect type coupler 25 is provided herein. Figure 33 The diagram shown is an exploded view of the components of the coupler 25, which integrates the aortic connector 14 and the blood pump 62. The coupler 25 includes a flange base 252, a pair of locking rings 253, and a hinge (or shaft assembly) 254 connecting the locking rings 253 to the flange base 252. A spring coil (or spring coil assembly) 255 is loaded in the hinge joint 256, which holds the locking rings 253 in the open position when the coupler 25 is unlocked. Figure 34A Its locking mechanism relies on a plate spring latch 257, which is made of a slotted spring plate and is secured by a plate 2571 welded to one end of the locking ring 253. The aforementioned flange base 252 has a generally circular structure, and each locking ring 253 has an arcuate structure. A hinge joint 256 is located on a first side 252S1 of the flange base 252, while the plate spring latch 257 is located on a second side 252S2 of the flange base 252 opposite to the first side 252S1. The locking ring 253 is pivotally connected to the hinge joint 256 and is rotatable relative to the hinge joint 256 and the flange base 252. In some embodiments, the coupler 25 and the aortic connector 14 are part of an aortic connector assembly.

[0163] Figure 34B A schematic diagram of the coupler 25 in a locked state is shown, wherein the slot of the plate spring latch 257 engages tightly with the bevel 258 to ensure a secure connection without fear of disengagement. Figure 35 As shown, the integrated connection between the aortic connector 14 and the blood pump 62 is achieved through the deformable connector proximal end 147 ( Figure 30 The front end of the protruding neck portion serves as a "gasket" between the rigid flange base 252 and the rigid beak flange 81 (described later) of the blood pump inlet connector 6251, which are interconnected.

[0164] Specifically, quick-connect locking can be easily achieved by closing the locking ring 253 without worrying about accidental unlocking, such as... Figure 34B and Figure 35 As shown. The aforementioned plate spring latch 257 is mounted at the end of a locking ring 253. During locking, the plate spring latch 257 bends as it slides on a (convex) ramp 258 on the opposite locking ring 253. As the plate spring latch 257 passes the top of the ramp 258, it descends to the bottom of the ramp 258 by elastic restoring force, acting as a safety device to prevent accidental latch unlocking or ring opening due to pump vibration or prolonged oscillation. For pump implants or replacements requiring modular disassembly, the plate spring latch 257 can be bent and lifted upwards using a tool, allowing an unlocking force to be applied to rotately open the locking ring 253, thereby disengaging the blood pump 62 from the aortic connector 14.

[0165] The butt joint design is not feasible for connecting two smooth-surfaced connectors in the bloodstream. In most clinical applications, the surface of the connected graft is roughened to promote endothelialization, thereby "smoothing out" minute interfacial discontinuities in the bloodstream by inwardly growing cells and proteins. The current aortic connector 14 uses a smooth surface to avoid thrombotic adverse events, for reasons explained earlier. Figure 28A and 28B As shown, the blood flow in the aortic connector responds bidirectionally to the jetting and filling action of the counterpulsation pump. This powerful bidirectional flow and surface-cleaning effect will readily remove any newly formed neon blood clots on the rough surface. Therefore, a smooth surface design is considered more suitable and safer for use in this invention. The interface between the two connected smooth surfaces in the blood flow requires careful mechanical and hemodynamic design to prevent in-situ thrombotic events. The principles and design methods related to this novel connector invention are disclosed below.

[0166] Figure 36A and Figure 36B Two basic interface discontinuities present in the mating connector connection are shown, such as steps 101, 102 or gap 103 generated between connector AB1 (e.g., connector of blood pump 62) and connector AB2 (e.g., aortic connector 14). Figure 36A and 36B In this exaggerated drawing method, such discontinuities at joints, typically within 10-50 micrometers in precision machining, are large enough to lead to clots and thrombosis.

[0167] In practice, even if each object is manufactured exactly the same, it is still necessary to match the tolerances of the two individual objects. For example... Figure 36ATwo misaligned joints are depicted; aside from the misalignment of the centerline, everything else related to part manufacturing is correctly completed. This will create forward-facing and backward-facing steps 101 and 102, and the stagnant flow in the step regions 101 and 102 will be the starting point for clot formation or thromboembolism. Figure 36B As shown, an interfacial gap 103 is created due to the non-parallel matching of the connectors. This gap 103 attracts blood cells to aggregate and further grow into a pseudointima. This intimal growth is often uncontrollable, leading to obstruction of the entire blood flow channel in addition to thrombi detaching from the intimal surface. When the connected objects are non-rigid, the erroneous interface of the mating joint can be exacerbated. The aortic connector 14 disclosed herein is semi-rigid, which can be forcibly pressed into the mating joint for connection (such as an inlet connector), exhibiting a deformable structure and enlarged interfacial discontinuities. Therefore, to achieve the connection between the current semi-rigid aortic connector and the blood pump, a novel connection method must be invented, as described below.

[0168] See Figure 37 and 38 In some embodiments, the blood pump 62 has an inlet connector 80, which 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 to the blood pump 62.

[0169] The inner diameter 84 of the beak portion 82 is slightly larger than the inner diameter 148 of the protruding neck portion 143 of the aortic connector 14 (see...). Figure 30 The contact area between the beak 82 and the proximal end (terminus) 147 of the connector is an annular conical surface (or shallow slope, ramp) 149, such as... Figure 30 and 35 As shown, the beak 82 can therefore be referred to as a conical beak. The aforementioned shallow bevel 149 is inclined relative to the centerline of the catheter insertion portion, and the cone angle of the shallow bevel is substantially in the range of 30-60 degrees measured from the rotation centerline of the inlet connector 80. In the initial locking engagement, the locking ring 253 with the inwardly recessed groove 2531 loosely engages the flange of the flange base 252 and the flange of the beak-shaped flange 81. As the locking ring 253 locks, the aforementioned beak-shaped flange 81 (of the inlet connector 80) and the flange base 252 (coupler 25) are received and compressed by the inner groove 2531 of the locking ring 253, thereby compressing and clamping the intermediate silicone end 147 (aortic connector 14) and generating a clamping force for a secure connection. In this way, the inlet connector 80 of the blood pump 62 will be securely connected to the aortic connector 14.

[0170] The aforementioned clamping force generating mechanism, such as Figure 35As shown. The flange base 252 has two steps 2521 and 2522, which are responsible for generating the clamping force. Before the locking ring 253 is closed, the step 2521 should first engage with the slot 1433 of the convex neck portion 143 of the aortic adapter 14 (see Figure 30 ). This engagement is achieved by first folding and squeezing the convex neck portion 143, and then inserting the deformed convex neck portion 143 through the flange base 252, allowing the elastic restoring force of the aortic adapter 14 to restore the folded and squeezed convex neck portion 143 to its original circular (or shape), and allowing the step 2521 to engage with the slot 1433. The height Z of the internal groove 2531 of the aforementioned locking ring 253 controls the extrusion deformation of the proximal end 147 of the adapter of the convex neck portion 143. Referring to Figure 35 , it can be found that the gap Z0 (i.e., the thickness of the extruded proximal end 147 of the adapter) is less than the thickness Z3 of the end 147 of the aortic adapter 14 in the fully locked configuration when the locking ring 253 is closed and locked (see Figure 30 ), that is, the thickness Z0 of the extruded proximal end 147 of the adapter is less than the thickness Z3 of the initial end 147 (Z0 < Z3). Regarding the thickness Z0 of the extruded proximal end 147 of the adapter, from Figure 35 the following formula can be obtained:

[0171] Z0 = Z - Z1 - Z2 … Formula (1)

[0172] In Formula (1), Z1 and Z2 are respectively the thicknesses of the beak flange 81 for clamping the mating part and the step 2522 of the flange base 252 as shown in Figure 35 . Usually, the thickness Z3 is greater than the gap Z0. Therefore, the strained proximal end 147 of the aortic adapter generates the clamping force required to seal - connect the beak flange 81 and the annular flange base 252. The strain of the proximal end 147 of the silicone - flow adapter 14, defined as (Z3 - Z0) / Z3, within the range of 10 - 30%, is sufficient to ensure a reliable sealed connection.

[0173] Figure 39 Shows the engagement characteristics of the aforementioned connection between the beak 82 and the shallow inclined (conical) surface 149 of the adapter. When making the connection, the beak 82 and its beak leading edge 85 will sink into the semi - rigid shallow inclined surface 149. The aortic adapter 14 has a depth equivalent to the leading - edge radius of the beak leading edge 85, generally 30 - 50 microns. In Figure 39In the diagram, the dashed lines and numbers in parentheses indicate the initial contact between the beak 82 and its leading edge 85, while the solid lines and numbers without parentheses indicate the locked position. It should be noted that the interface discontinuity is reduced by the aforementioned shallow bevel 149 and the recess of the beak 82 from its original shape (dashed lines). The thickness of the internal groove 2531 controls the tight fit of the coupling. As previously described, the proximal end 147 of the resilient aortic connector will be compressed with approximately 10-30% strain to provide the required coupling force for leak-free integration against pulsating pumping.

[0174] The current design of the interface connection between the blood pump 62 and the aortic connector 14 offers two hemodynamic advantages, reducing in-situ thrombosis. First, step or gap-type connector discontinuities are virtually eliminated, as observed in conventional docking connections. Second, stagnant flow at the interface located at the anterior border 85 of the proboscis is minimized. Therefore, blood flow through the connection interface will remain at a high velocity, significantly improving docking connection defects, specifically addressing the forward or backward steps 101, 102 or gaps 103 that were previously present at the interface.

[0175] The tapered shallow bevel 149 of this embodiment is inclined at an angle relative to the flow direction. This bevel interface design avoids steps or gaps at the joint due to limited manufacturing precision or misalignment associated with conventional butt joints. However, this tapered shallow bevel 149 has inherent drawbacks in achieving concentric centerline alignment of the mating counterparts. The connection between the aortic connector 14 and the beak 82 lacks strict lateral constraints to ensure alignment. For concentric connection of the rigid beak 82 and the semi-rigid shallow bevel 149, it is crucial that the locking ring is simultaneously engaged around the entire peripheral edge of the flange base 252. When simultaneous engagement / locking is not achieved, the initially engaged shallow bevel 149 will strain more than the other free portions, tending to tilt or align the remaining contact surfaces, resulting in an eccentric connection to the pump. This eccentric connection is often a factor that creates steps or gaps at the interface, leading to thrombosis. Through the locking ring profile 259 on the distal side (below) of the locking ring 253 ( Figure 34A This drawback is mitigated by a locking engagement configuration that simultaneously includes all circumferential contact areas. When locked, the contact edge of the metal (in some embodiments) beak 82 is slightly recessed into the shallow bevel 149 of the compressed silicone material (in some embodiments) with a controlled depth, further reducing interfacial discontinuities when exposed to blood flow. Furthermore, conventional interfacial thrombosis can be significantly reduced or eliminated through the appropriate application of an anticoagulant regimen.

[0176] The structural deformability and delivery method of this aortic connector 14 give the invention its unique design features. In fact, considerations of material elasticity need to be carefully incorporated into the current design. Delivering an insertable graft into the aorta via a surgical incision of the aortic wall is challenging in terms of surgical safety and long-term reliability. Therefore, in some embodiments, the material chosen for the aortic connector 14 should have a predetermined shape memory. During device insertion, the connector 14 is first crimped into a smaller insertion configuration (e.g., Figure 40 The coiled aortic connector 14 shown in the diagram, and this insertion configuration ensures rapid and safe device implantation. After the coiled aortic connector 14 is placed at the intended implantation site, the inserted aortic connector 14 should be released to self-expand to its original memory shape.

[0177] Before inserting the aortic connector 14, a hole with a diameter of 12-14 mm should be formed in the aortic wall. When creating such an entry hole, care must be taken to avoid any cut edges that could become crack initiation points due to the need to expand the wall surface during device insertion. A side-biting aortic perforator, as disclosed in U.S. Patent Application No. 17 / 034036, is an ideal tool for creating large holes in the aorta. A single, gentle tap can successfully create a hole without any crack edges.

[0178] The aortic connector 14 morphologically comprises two circular tubes joined together to form a T-shaped flow connector for performing aortic paracirculation support. The catheter wall thickness is typically 1-2 mm, and the material used is a polymer with appropriate rigidity, such as silicone or polyurethane, with a hardness of, for example, Shore A 80-90. The coiled insertion structure differs significantly from commercially available large stent grafts covered with dacron or PTFE (polytetrafluoroethylene) fabric. Figure 40 The diagram illustrates a coiled / insertion configuration of the aortic connector 14 (with a nickel-titanium alloy metal stent 144 embedded and deforming along with the polymer-based aortic connector 14). The aortic connector 14 is folded through a flattened catheter insertion portion of the aortic connector catheter 142, and its T-shaped neck portion 143 is correspondingly compressed and flattened into the folded connector 14 body, as shown. Figure 40 As shown. The diameter of the folded aortic connector 14 is approximately half the diameter of the original unfolded circle.

[0179] This folding connector 14 can be held in place by pulling a rope tight. For example... Figure 40 As shown. Three fixing ropes can be installed at the two edges and the center of the catheter, or other fixing methods can be considered. Figure 41A , 41BFigures 41C and 41D show four representative stages of the implantation configuration of the aortic connector 14. Stage 1 (see Figure 41C) Figure 41A The diagram shows the initial penetration of the coiled pre-coil form through the inlet hole when the pre-coil (aortic connector 14) is tilted at an angle to the axis of the aorta 95. The second stage ( Figure 41B The image shows the aortic connector 14 in a fully inserted configuration inserted into the aorta, with a pre-coated package pushed past the inlet port, one end of which rotates and falls into the inlet port. The fully inserted pre-coated package is then retracted so that its coiled neck portion 143 aligns with the inlet port. Figure 41C The restraining cords are released, allowing the coiled, compressed aortic connector 14 to elastically expand and return to its original shape. Figure 41D The released aortic coupling 14 will be tightly surrounded by the aortic lumen, ensured by selecting an appropriate oversized ratio before insertion. This is achieved by slightly deforming the T-shaped neck portion 143, which is part of the flange base 252 of the coupling 25. Figures 33 to 34B It can be installed on the T-shaped protruding neck portion 143 to prepare for connection of the blood pump 62. Thus, a secure device connection is achieved by placing the inlet beak portion 82 on the tapered shallow bevel 149 and then closing the two locking rings 253 of the coupler 25, thereby easily achieving the blood pump connection.

[0180] Additional safety measures can be applied to enhance hemostasis and stability of the implanted para-aortic pump system. Due to the weight of the pump 62 and the pumping force generated by the counterpulsation support, the para-aortic placement of the pump inevitably involves lateral forces (perpendicular to the longitudinal direction of the aorta) and torque applied to the aortic connector 14. These device-related external forces may affect long-term remodeling of the vascular structure at the implantation site. A purse-string suture can be placed in the adventitia around the access port. The purse-string suture additionally tightens the aortic wall against the inserted connector 14 and acts as a protective measure against enlargement of the access port. Furthermore, surgical tape can be wrapped around both ends of the aortic connector catheter 142 and tightened, reinforcing the overall connection between the inserted aortic connector 14 and the aorta. A well-fitting design and the application of loop tape will provide double protection against endoleak. Sometimes, blood pressure may rise above the upper limit that can be ensured by compliance matching to prevent endoleak. In such extreme cases, surgical tape will come into play as a rigid restrictor, sealing the ends of the joints to be separated and ensuring hemostasis.

[0181] like Figure 42In some embodiments, the step-by-step setup of the implanted aortic connector 14 is described in detail here. Before implantation begins, a pre-rolled / compressed insertion configuration of the aortic connector 14 is prepared. After exposing the target thoracic artery via a left thoracotomy, a cross-clamp distance of approximately 10 cm across the implantation site is determined. An entry hole made in the aorta is first marked with a periphery marker. Then, a band suture is sutured around the periphery of the hole in the adventitia. The aorta can be partially separated from the surrounding connective tissue, and a pair of surgical tapes can be wrapped around the aorta. After completing the above preparations, the cross-clamp and aortic connector are inserted. These insertion steps are as follows: Figure 17 The procedure is described as follows. First, the aorta is crossed and clamped to provide segmental isolation without bleeding issues. Then, a large access hole is made using a custom aortic punch for inserting the aortic connector 14. The folded connector pre-roll package is then inserted and placed into the crossed-clamped aortic segment, as shown. Figure 41A , 41B As shown in 41C. Afterwards, the folded connector 14 is released and restored to its initial unfolded form, as... Figure 41D As shown. The cord and tape are then used to tighten the connection as additional protection against hypertensive endoleak. The coupler 25 is then installed, with its step 2521 engaging in the slot 1433 of the protruding neck portion 143 of the aortic connector 14, ready to receive the blood pump 62 to be connected. The self-aligning capability of the aforementioned coupler 25 allows the inlet connector 80 of the blood pump 62 to be correctly positioned and locked with the aortic connector 14. The remaining implantation steps are routine, including cross-clamp release, blood pump venting, and pump start-up support. Generally, the cross-clamping time required for insertion of the aortic connector by a trained surgeon is approximately 10 minutes. During this cross-clamping, abdominal organs will be deprived of blood perfusion, potentially leading to ischemic injury. To mitigate this potential surgical damage to organs, partial femoral-femoral extracorporeal membrane oxygenation (femoral-femoral ECMO) support can be used to perfuse abdominal organs and lower limbs. However, the decision to use ECMO support is at the surgeon's discretion. Normally, patients can tolerate 20 minutes of ischemia.

[0182] In summary, embodiments of the present invention provide a ventricular assist device, including a blood pump, a drive catheter, and an inlet. The blood pump includes an axisymmetric elliptical sac and a port assembly comprising a flexible sac, a proximal port, and a distal port, wherein the flexible sac is connected to the proximal and distal ports as a stress-relieving suspension mechanism. The blood pump further includes a pump housing comprising a proximal housing and a distal housing, wherein the stress-relieving suspension mechanism is connected to the pump housing. The blood pump further includes a pressure sensing system embedded in the proximal housing, wherein the pressure sensing system includes a blood pump pressure sensor and a pressure sensing chamber filled with incompressible fluid for pressure transmission. The aforementioned drive catheter includes a pneumatic cavity, at least one wire and a tether included in the wall of the drive catheter, wherein the wire and tether are disposed in the wall of the drive catheter. The aforementioned inlet connects the drive catheter and the pump housing.

[0183] One embodiment of the present invention discloses a pulsating blood pump design that incorporates a non-stationary folding line concept in the construction of a long-term blood sac, which can substantially extend the durability of a replacement blood pump. Furthermore, a miniature pressure sensing system is also provided, which can be used as a reference waveform for real-time pump control based on real-time big data, as well as for long-term trend analysis, disease monitoring, and diagnosis. Additionally, the embedded pressure sensing system is non-blood-contact, thus significantly improving the reliability requirements when constructing implantable sensor systems.

[0184] The embodiments of the present invention have at least one of the following advantages or effects. By connecting the drive catheter to the inlet of the pump housing, a compact inlet design can be provided, making the wiring and signal transmission more robust and more tolerant. Furthermore, the compact inlet design integrates the sensing wires and pneumatic tubing with the blood pump. This compactness is particularly important for implantable devices. It not only simplifies surgical procedures and reduces the risks of perioperative implantation, but also helps reduce postoperative morbidity associated with drive catheter infection.

[0185] In some embodiments, the introductory element is integral with the distal housing of the pump housing, and the introductory element has a first portion as an extension of the distal housing, in which the pneumatic interior of the power-driven catheter, the tether, and the wire are coupled; a second portion interlocks with the first portion and serves as a hollow connector for driving the catheter to achieve the advantages of anatomical adaptability and adaptability to the geometry of the implantation site.

[0186] One embodiment of the present invention provides an aortic connector assembly for an implantable ventricular assist device, comprising: a T-type flow connector including a catheter insertion portion and a protruding neck portion, wherein the catheter insertion portion and the protruding neck portion are connected, both having smooth blood-contact surfaces; and a metal stent disposed in the catheter insertion portion. The T-type flow connector has a polymer elastomer reinforced by a metal stent made of a nickel-titanium alloy material. The aforementioned catheter insertion portion has gradually thinning walls at its two ends, the distal end of the catheter end being at an appropriate distance from the outermost boundary of the metal stent, and the catheter end having a compliant fit to the artery at the implantation site. The proximal end of the protruding neck portion is configured to connect to an inlet connector of a blood pump.

[0187] Embodiments of the present invention offer at least one of the following advantages or effects. The present invention discloses a flow connector assembly that enables blood flow into and out of aortic paraventricular assist devices, particularly counterpulsation pumps. Unlike many existing flow connectors that employ rough surface methods to promote endothelialization and thus avoid adverse events such as thrombosis, the aortic connector disclosed herein employs a smooth surface, insertable prosthetic graft concept to construct the flow connector. Furthermore, a compliant matching design is implemented around the inserted catheter tip, combining gradually thinning wall characteristics with a thin-walled polymer supported by a superelastic nickel-titanium alloy frame to achieve the requirement of no endoleak. The abnormally high pressure, high shear, and low-velocity recirculation flow phenomena associated with paraventricular counterpulsation pumping are contained within the artificial surface of the inserted catheter. Therefore, hemodynamic effects and risk factors caused by pathological devices are essentially 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 connector and the pump, the present disclosure provides a quick-connect coupling. This coupler features a self-aligned interface design that minimizes discontinuities in steps and gaps, thereby reducing the likelihood of thrombotic adverse events at the interface joint. Combined with the invention of the aortic connector, and a specially designed insertion method, it ensures a rapid and safe insertion process. The coiled aortic connector is manufactured in a pre-coiled package / insertion configuration, reducing its overall size to half its initial size. This pre-coiled connector can be easily inserted into the aorta at the implantation site and self-expands to its initial configuration, forming a close-fitting flow connector without concerns about endoleak. This not only helps mitigate the risks of implantation during surgery but also contributes to reducing postoperative morbidity associated with device-induced flow and vascular maladaptation at the implantation site.

[0188] The ordinal numbers in this specification and claims, such as "first," "second," etc., are not sequential in any particular order; they are only used to distinguish two different elements with the same name.

[0189] The embodiments described herein will be readily apparent to those skilled in the art upon reading the foregoing description. Therefore, this invention includes all modifications and equivalents of the subject matter recited in the claims.

Claims

1. A para-aortic blood pump device, comprising: A blood pump includes a pump housing, a blood sac, and a pressure sensor, wherein the blood sac is disposed inside the pump housing, and the pressure sensor is installed inside the pump housing to monitor the blood pressure inside the blood pump and generate an electronic blood pressure signal. An aortic connector, which is a T-shaped catheter, is connected to the blood pump and is used to connect the blood pump to the human aorta. A drive catheter is coupled to the pump housing of the blood pump and transmits the electronic blood pressure signal received from the pressure sensor. as well as A driver, coupled to the drive catheter to receive the electronic blood pressure signal, includes an electromechanical actuator that generates a pressure pulse based on the electronic blood pressure signal to drive the sac to perform a corresponding action, and provides the pressure pulse to the blood pump through the drive catheter; and A coupler for coupling the blood pump to the aortic connector; The coupler includes a flange base, a pair of locking rings, and a hinge assembly that connects the locking rings to the flange base and connects the blood pump to the aortic connector by closing the locking rings.

2. The paracorporeal blood pump apparatus of claim 1, wherein, The driver also includes a drive catheter controller and a vibrator, wherein the drive catheter controller is used to process the electronic blood pressure signal, and the vibrator is used to provide an audible alarm or a tactile feedback.

3. The paracorporeal blood pump apparatus of claim 1, wherein, The drive catheter and a portion of the actuator are replaced by a distal drive catheter, a drive catheter interconnector, and a proximal drive catheter. The distal drive catheter is used to transmit the electronic blood pressure signal and the pressure pulse to the blood pump. The drive catheter interconnector includes a drive catheter controller and a vibrator. The drive catheter controller is used to process the electronic blood pressure signal, and the vibrator is used to provide an audible alarm or a tactile feedback.

4. The para-aortic blood pump device as described in claim 1, wherein, The blood pump is integrally formed with the aortic connector.

5. The para-aortic blood pump device as described in claim 1, wherein, The two gradually opening ends of the aortic connector form a smooth transition in elastic properties, gradually softening as the wall thickness of the catheter implanted in the aorta decreases.

6. The para-aortic blood pump device as claimed in claim 1, wherein, The blood pump housing contains the blood sac, which is an egg-shaped membrane rotating around the center line of the blood pump. It has two polymer ports at both ends that are connected to the blood sac. When the membrane is attached to the pump housing, the polymer ports can be used as a release mechanism for bending and stretching of the membrane to reduce stress concentration.

7. The para-aortic blood pump device as claimed in claim 1, wherein, The pump housing of the blood pump has an opening, and the opening is integrally formed with the aortic connector to form a seamless and smooth interface for engagement with the neck of the aortic connector.

8. The para-aortic blood pump device as claimed in claim 1, wherein, The electromechanical actuator includes a pressure balancing valve connected to the cylinder of the electromechanical actuator. The pressure balancing valve opens periodically to balance the air pressure in the cylinder with the atmospheric pressure.

9. The para-aortic blood pump device as claimed in claim 1, wherein, This para-aortic blood pump device provides a counter-pulsatile boost to systemic blood flow during diastole (cardiac relaxation) to improve myocardial and organ perfusion, while reducing the workload of the left ventricle during systole (cardiac contraction).

10. The para-aortic blood pump device as claimed in claim 1, wherein, The driver includes a trigger detection microcontroller unit that provides a sensed blood pressure waveform within the blood pump using an electronic blood pressure signal, allowing the trigger detection microcontroller unit to calculate and determine the ejection and refill times of the electromechanical actuator.

11. The para-aortic blood pump device as claimed in claim 1, wherein, The electromechanical actuator includes a motor and a ball screw unit that drives a reciprocating piston in a cylinder of the electromechanical actuator; the movement of the reciprocating piston draws in and ejects air through a drive conduit connected to the blood pump.

12. The para-aortic blood pump device as claimed in claim 1, wherein, This electromechanical actuator is a pneumatic actuator, comprising: A brushless servo motor, a ball screw unit, and a piston and cylinder assembly, wherein the blood pump reciprocates by ejecting / filling air as a driving medium; and A pressure balancing valve is installed on the pneumatic actuator to address the problems of piston ring leakage in the piston and cylinder assembly and water vapor condensation from the blood sac.

13. The para-aortic blood pump device as claimed in claim 10, wherein, The driver receives the electronic blood pressure signal and processes it using a trigger detection algorithm to generate a trigger signal that commands the driver to actuate in coordination with the heart rhythm.

14. The para-aortic blood pump device as claimed in claim 13, wherein, Upon receiving a specified trigger time, the trigger detection microcontroller unit sends a command to a motor controller to drive a piston of the electromechanical actuator to follow the injection to the filling or from the filling to the injection position to provide anti-pulsating cycle support, including systolic unloading during cardiac systole and perfusion enhancement during cardiac diastole, respectively.

15. The para-aortic blood pump device as claimed in claim 1, wherein, The driver also includes a user interface module, which includes an indicator, an audible alarm, a button group and an LCD display.

16. The para-aortic blood pump device as claimed in claim 10, wherein, When the trigger detection microcontroller unit loses the electronic blood pressure signal from the blood pump, the trigger detection microcontroller unit automatically activates a flushing mode to drive the electromechanical actuator to operate at a predetermined pumping frequency and stroke volume.

17. The para-aortic blood pump device as claimed in claim 16, wherein, This flushing mode is designed to prevent blood clots from forming in the blood pump; it is a device protection mode, not a mode to provide circulatory support.

18. The para-aortic blood pump device as claimed in claim 6, wherein, The blood sac is anchored to a proximal shell of the pump housing via a proximal port and to a distal shell of the pump housing via a distal port.

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