Patent Description:
The present invention relates to a para-aortic blood pump device. The disclosure relates to a ventricular assist device (VAD), in particular to a left ventricular assist device (LVAD) based on the principle of counterpulsation support.

The disease progression of heart failure is characterized by a vicious cycle. Medical therapy for early-to-moderate stage heart failure is to suppress the neurohormonal compensatory mechanism, which slows down rather than remedies the death spiral of the heart failure viscous cycle. Heart transplant or ventricular assist device (VAD) implantation is intended to be used for the terminal-stage heart failure. There exists an "unmet need" gap region between the medically ineffective and transplant/VAD treatments along the course of heart failure progression. As heart failure deteriorates beyond effective medical treatment stage, often, patients cannot receive any therapy but waiting for heart failure to deteriorate further into the refractory stage where transplant/VAD can be considered. This treatment gap is categorized as Interagency Registry for Mechanical Assisted Circulatory Support (INTERMACS) profiles <NUM>-<NUM>, and patients in this class of heart failure typically show symptoms of apnea, pulmonary hypertension, renal hypoperfusion and exercise intolerance with elevated inflammatory responses.

The best policy for treating heart failure by mechanical circulatory support (MCS) is to implant a left ventricular assist device (LVAD) in time before a failing heart becomes refractory. However, this policy is not practically realizable because the current rotary pumps, or continuous-flow LVADs, are invasive in implantation and involve with considerable post-operative morbidity rates, making the LVADs only indicated for terminal-stage heart failure patients.

Counter-pulsatile support therapeutics has been clinically proven effective for more than fifty years by the bedside intra-aortic balloon pump (IABP) administration. The recent success in extending IABP use from bedside to ambulatory support, via implanting the balloon pump from axillary or subclavian artery, has achieved bridging patients to heart transplant. Balloon counterpulsation together with patient ambulation showed surprisingly good benefit in improving the heart functional status. Nevertheless, this balloon delivery site improvement is only a temporary solution. Patients treated by ambulatory IABP are still bounded in hospital and the axillary or subclavian insertion cannot be used safely in a long-term period.

Currently, there exist no clinically approved partial-support (blood pump providing <NUM>-<NUM> Liter per minute pump flow), early intervention devices that are specifically indicated for less-sick advanced heart failure patients. A great portion of such less-sick heart failure patients will deteriorate and develop acute myocardial ischemia or cardiogenic shock requiring emergent mechanical circulatory support. The life-sustaining, acute circulatory support systems such as IABP, extracorporeal membrane oxygenation (ECMO), and percutaneous micro-axial catheter pumps could be emergently administered, however, the mortality rate of the supported patients is generally high in the range of <NUM>-<NUM>%. Moreover, it is often difficult for patients, families, and clinicians to decide if invasive and expensive rotary pumps should be used after the exhaustion of the acute circulatory support devices. Therefore, it is a main subject for related manufacturers to overcome the therapeutic gap in heart failure treatment, by innovating new long-term implantable VAD design that encompasses the early intervention and less traumatic surgery concept.

An early intervention partial-support blood pump, also known as a para-aortic blood pump device, was presently invented according to the counterpulsation support principle. The para-aortic blood pump device of the present invention provides a counter-pulsatile circulatory support including augmentation of systemic blood flow during diastole (heart relaxation) to improve myocardial and organ perfusion while reducing left ventricular workload during systole (heart contraction). For mechanical circulatory support device recipients, earlier hemodynamic support and post-operative ambulatory ability is important for survival and disease improvement. Less invasive surgery, a premise of early intervention, hence, constitutes another critical element aside from the hemodynamic therapeutics the partial-support LVADs can provide. With the aid of specially designed surgical tools, the para-aortic blood pump device of the present invention can be safely implanted via a less traumatic surgical procedure (such as a beating-heart mini-left thoracotomy). Evidences gathered from several contemporary clinical trials performed for late-stage heart failure patients have shown that chronic LVAD-supported heart, either by displacement type blood pump or by rotary pump, can encourage the nonischemic myocardium undergo cellular reverse remodeling, resulting in a significant portion of device-treated patients gaining robust myocardial recovery or remission. This finding implies that if early ambulatory support feature can be included in LVAD design and instituted in treating heart failure, the mechanical unloading would be potentially more effective and beneficial for salvaging the less-sick (INTERMACS profiles <NUM>-<NUM>) patients.

It is plausible and beneficial that before heart transplant or implanting highly invasive rotary pumps, patients can be treated by the less traumatic para-aortic blood pump device for an extended period. Some patients may recover with the circulatory assistance offered by the para-aortic blood pump device, executed via therapeutic counterpulsation for months or longer (<NUM>-<NUM> years) in the form of ambulatory support. For those hemodynamically stabilized but non-recoverable recipients, the para-aortic blood pump device can be used in a long-term manner as a destination therapy device or as a bridging to transplant. The para-aortic blood pump device, hence, can serve as an interim salvage modality before heart transplant or an expensive and invasive rotary pump implantation. The role of the para-aortic blood pump device in heart failure treatment is multi-faceted, it can be used as a means for bridge-to-recovery (or remission), bridge-to-decision (rotary pump), bridge-to-transplant, or alternative-to-transplant (destination therapy). A prior art example is known from e.g. <CIT>).

It is one of primary objectives of the present disclosure to overcome the shortcomings of the prior art by disclosing a para-aortic blood pump device comprising: a blood pump, an aortic adapter, a driveline, and a driver. The blood pump comprises a pump housing, a blood sac, and a pressure sensor, and the pressure sensor is installed in the pump housing of the blood pump for monitoring blood pressure inside the blood pump. The sensed blood pressure is transduced into an electrical signal to be transmitted through a driveline into a driver. The aortic adapter is a T-manifold shaped flow conduit coupled to the blood pump and used for integrating the blood pump with the human aorta. The aortic adapter is thin-walled and may have structural reinforcement embedded for strength enhancement. The driveline is coupled to the housing of the blood pump, providing a pneumatic communication to the blood pump and transmitting the electrical blood pressure signal received from the pressure sensor. The driver is coupled to the driveline for receiving the electrical blood pressure signal. The driver comprises an electro-mechanical actuator commanded by a controller in the driver to generate counter-pulsatile pneumatic pressure pulse according to the sensed electrical blood pressure signal. The pressure pulse is sent to and from the blood pump through the driveline, fulfilling the ejection and filling action of the blood pump when supporting the human circulation.

In some embodiments, the driver further comprises a driveline controller and a vibrator, and the driveline controller is used for processing the electrical blood pressure signal, and the vibrator is used for providing an audible alarm or a tactile feedback.

In some embodiments, the driveline and a part of the driver are substituted by a distal driveline, a driveline interconnector and a proximal driveline, and the distal driveline is provided for transmitting the electrical blood pressure signal and the pressure pulse to the blood pump, and the driveline interconnector comprises a driveline controller and a vibrator, and the driveline controller is used for processing the electrical blood pressure signal, and the vibrator is used for providing an audible alarm or a tactile feedback.

In some embodiments, the blood pump and the aortic adapter are integrally formed.

In some embodiments, the para-aortic blood pump device comprises a coupler, and the coupler includes a coupling adapter installed at a neck of the aortic adapter for coupling the blood pump to the aortic adapter.

In some embodiments, two gradually flared ends of the aortic adapter constitute a smooth transition of an elastic property, being progressively softer in proportion to a wall thickness toward the end.

In some embodiments, the pump housing of the blood pump is installed with the blood sac therein, and the blood sac is an oval-shaped membrane body of revolution to a centerline of the blood pump, at two ends there are two polymeric stems that are bonded with the membrane body, working as a flexing/stretching relief mechanism to alleviate stress concentration when attached onto the pump housing of the blood pump.

In some embodiments, the pump housing of the blood pump has an opening, and the opening continuously integrates with the aortic adapter that provides a smooth interface with a neck of the aortic adapter.

In some embodiments, the electro-mechanical actuator comprises a pressure equalization valve connected to an air chamber, and the pressure equalization valve is opened periodically so that an air pressure in the air chamber can be set to be in equilibrium with atmospheric pressure.

In some embodiments, the para-aortic blood pump device provides a counter-pulsatile augmentation of systemic blood flow during diastole (heart relaxation) to improve myocardial and organ perfusion while reducing left ventricular workload during systole (heart contraction).

In some embodiments, the driver further comprises a trigger-detection micro controller unit, and the electrical blood pressure signal provides the sensed pressure waveform within the blood pump for the trigger-detection micro controller unit to compute and determine the eject and fill timings for the electro-mechanical actuator.

In some embodiments, the electro-mechanical actuator comprises a motor and ball screw/nut unit that drives a reciprocating piston within a cylinder of electro-mechanical actuator; the movement of the piston pushes and withdraws air via the driveline coupled to the blood pump.

In some embodiments, wherein the electro-mechanical actuator is a pneumatic actuator includes: a brushless servo motor and a ball screw piston/cylinder assembly, wherein atmospheric air is used as a driving medium to reciprocally eject/fill the blood pump; and a pressure equalization valve is equipped on the electro-mechanical actuator to solve the problems of air leakage at a piston ring and condensation of vapor permeated out from a blood sac membrane.

In some embodiments, the driver receives the electrical blood pressure signal and processes the electrical blood pressure signal using trigger detection algorithm to generate a trigger signal that commands a driver actuation in synchronization with heart rhythm.

In some embodiments, upon receiving the assigned trigger timing, the micro controller unit sends commands to a motor controller to drive a piston of the electro-mechanical actuator, from eject-to-fill or from fill-to-eject positions, to provide counter-pulsatile circulatory support including contraction unloading during the cardiac systolic phase and perfusion augmentation during the cardiac diastolic phase, respectively.

In some embodiments, the driver further comprises a user interface, and the user interface comprises an indicator, an audio alarm, a button and a liquid crystal display (LCD).

In some embodiments, when the micro controller unit loses the electrical blood pressure signal from the blood pump, a washout mode is launched automatically by the micro controller unit to drive the electro-mechanical actuator, operating at a predetermined pumping rate and driver stroke volume.

In some embodiments, the washout mode is used to prevent the formation of thrombus in the blood pump, which is a device protection mode instead of providing circulatory support.

In some embodiments, the blood sac is anchored via a proximal stem to a proximal shell of the pump housing and via a distal stem to a distal shell of the pump housing.

There are four embodiments that can be employed to realize the present para-aortic blood pump invention, as described below.

With reference to <FIG> for the schematic view of a para-aortic blood pump device in accordance with the first embodiment of the present invention, the para-aortic blood pump device <NUM> comprises: a blood pump <NUM>, an aortic adapter <NUM>, a driveline <NUM> and a driver <NUM>. The blood pump <NUM> further comprises a pump housing and a pressure sensor. The blood pump housing consists of two chambers, one for blood storage and another to receive driving air. These two chambers are separated by an oval-shaped flexible membrane suspended via a pair of stress-relief stems attached to the pump housing. The pressure sensor is installed in the pump housing for monitoring blood pressure inside the blood pump <NUM> to generate electrical blood pressure signal. The aortic adapter <NUM> is a valveless, T-manifold shaped flow communicator coupled with the blood pump <NUM> and human aorta. In the first embodiment, the aortic adapter <NUM> and the blood pump <NUM> are integrally formed, having a seamless blood contacting surface, and the aortic adapter <NUM> is provided for connecting the blood pump <NUM> with the human aorta. The aortic adapter <NUM> is made of flexible materials, allowing the aortic adapter <NUM> be deformed during insertion delivery from a hole made on the aortic wall. This aortic adapter conduit, after inserted, is self-expandable and strong enough to withhold the radial compression exerted from oversized fitting to the contacted aortic lumen. The driveline <NUM> is coupled to the pump housing of the blood pump <NUM> for providing a pressure pulse to the blood pump <NUM> and transmitting the electrical blood pressure signal received from the pressure sensor. The driver <NUM> is coupled to the driveline <NUM> for receiving the sensed electrical blood pressure signal, and the driver <NUM> comprises an electro-mechanical actuator to generate a pressure pulse according to the electrical blood pressure signal and provides a regulated pressure pulse to the blood pump <NUM> through the driveline <NUM>. The wearable driver <NUM> provides timed air pressure pulses in synchronization to cardiac rhythm to drive and control the eject and fill of the implanted blood pump <NUM>.

The driver <NUM> comprises a battery power system <NUM> and a redundant battery power system (the battery power systems <NUM>, <NUM>, <NUM> in the <FIG>, and <FIG> are the same or similar), wherein the redundant battery power system ensures a continuous power supply of the driver <NUM>. Power can also be supplied to the driver <NUM> by an AC adapter for the convenience of the device recipient when mobility is not required. Further, a clinical monitor unit, not shown in <FIG>, can be connected to the driver <NUM> to provide a user interface to the clinician for displaying device monitoring or diagnostic information and for accessing to driver parameters in order to initiate and optimize a patient-specific operational mode setting.

In <FIG> and <FIG> illustrated two different blood pump <NUM>, <NUM> designs coupled with the same driveline <NUM>, <NUM> and driver <NUM>, <NUM> system. <FIG> is the schematic view of a para-aortic blood pump device <NUM> in accordance with the second embodiment of the present invention, and the difference between the second embodiment and the first embodiment of the present invention resides on that the para-aortic blood pump device <NUM> of the second embodiment further comprises a coupler (or coupling adapter) <NUM>. The blood pump <NUM> and the aortic adapter <NUM> of the second embodiment are not integrally formed, but are detachable from each other; and the coupler <NUM> is provided for coupling the blood pump <NUM> to the aortic adapter <NUM>. Care must be exercised in the coupler design to minimize the interface discontinuity around the connected region. During device implantation, the aortic adapter <NUM> is first inserted into the aorta via an access hole made in the aortic wall. Using specially developed implantation tools, a coupling adapter <NUM> is disposed around the protruded neck of the aortic adapter <NUM>, to which the blood pump <NUM> can be connected. Following the introduction of the blood pump <NUM> into the thoracic cavity, the blood pump <NUM> and the aortic adapter <NUM> are firmly integrated using the coupler <NUM>. Such detachable blood pump <NUM> and aortic adapter <NUM> design encompasses surgical and post-operative advantages. During device implantation, the detachable pump design renders the aortic adapter insertion easier because the surgical field is clearer without the interference of the pump body. Further, in the post-operative period, blood pump can be detached and exchanged in case of pressure sensor malfunction or blood sac rupture requiring emergent surgical replacement. The detachable blood pump design of the second embodiment is advantageous in this regard. The aortic adapter can stay in the aorta without explantation, avoiding troublesome and risky redo surgery associated with aortic adapter removal.

With reference to <FIG> and <FIG> for the schematic views of the para-aortic blood pump devices <NUM>, <NUM> in accordance with the first and the third embodiment of the present invention respectively, the difference between the third embodiment and the first embodiment of the present invention resides on that the driveline <NUM> of the first embodiment are substituted by the driveline <NUM> including the distal driveline <NUM>, the driveline interconnector <NUM> and the proximal driveline <NUM> of the third embodiment. The distal driveline <NUM> is coupled to the driveline interconnector <NUM> for transmitting the electrical blood pressure signal acquired from the pressure sensor and the pressure pulse sent from the driver <NUM>; and the driveline controller and the vibrator (for alarm warning purpose) included in the driveline interconnector <NUM> are originally included in the driver <NUM> of the first embodiment, so that the driver <NUM> of the first embodiment has the additional driveline controller and vibrator (compared with the driver <NUM> of the third embodiment). The driveline controller is used for processing the electrical blood pressure signal, and the vibrator is used for providing an audible alarm or a tactile feedback. In other words, the mechanical power transmission to fill and eject the blood pump and the analog/digital signal conversion and alarm annunciation, achieved respectively by the driveline <NUM> and the driver <NUM> of the first embodiment, are substantially the same as those by the distal driveline <NUM>, the driveline interconnector <NUM> and the proximal driveline <NUM> of the third embodiment.

The first embodiment has a cleaner driveline configuration and disposes electronic signal processor in the driver, hence minimizing the risk of environmental contamination (water ingress or moisture condensation) of the sensed pressure signal and the air leak incurred at the joint, both associated with the driveline interconnector <NUM>. Nevertheless, this long driveline is more vulnerable to contact damage such as wear, kink, cut, abrasion arising from the contact with foreign objects in daily activities. Any major damage to the driveline <NUM> of the first or second embodiment, either electronically or mechanically, may warrant surgical blood pump replacement that is highly undesirable in view of surgical redo risk and the associated medical costs. The third or fourth embodiment, by employing a mid-way connector (driveline interconnector), mitigates such driveline damage-related blood pump replacement drawback. In general, the length of externalized distal driveline <NUM> is short and the interconnector <NUM> is protected better by the coverage of the skin dressing and/or the patient vest. In the extreme case of severe damage of the driveline beyond repairable, the most possibly damaged proximal driveline <NUM> can be easily exchanged without resorting to surgery. In addition, the third or fourth embodiment is more immune to electromagnetic interference because the analog-to-digital signal conversion is already accomplished in the circuitry in the interconnector <NUM>. The fidelity of pressure signals can be better assured in the third or fourth embodiment because digital signal transmission in the proximal driveline <NUM> is less susceptible to the electromagnetic interferences.

With reference to <FIG> and <FIG> for the schematic views of the para-aortic blood pump device <NUM>, <NUM> in accordance with the third and the fourth embodiment of the present invention respectively, the difference between the third embodiment and the fourth embodiment resides on that the aortic adapter <NUM> and the blood pump <NUM> of the third embodiment are integrally formed; whereas the blood pump <NUM> and the aortic adapter <NUM> of the fourth embodiment are detachable; and the fourth embodiment further comprises a blood pump <NUM>, an aortic adapter <NUM> and a coupler <NUM> which are the same as the blood pump <NUM>, the aortic adapter <NUM>, and the coupler <NUM> of the second embodiment, and thus their descriptions will not be repeated. The driveline <NUM> including the distal driveline <NUM>, the driveline interconnector <NUM> and the proximal driveline <NUM> of the fourth embodiment is the same as the driveline <NUM> including the distal driveline <NUM>, the driveline interconnector <NUM> and the proximal driveline <NUM> of the third embodiment.

With reference to <FIG> for the schematic view of a para-aortic blood pump device installed in human body in accordance with an exemplary embodiment of the present invention, the para-aortic blood pump device <NUM> comprises a blood pump <NUM>, an aortic adapter <NUM>, an internal driveline <NUM>, an external driveline <NUM>, and a driver <NUM>. In another embodiment, the para-aortic blood pump device further comprises a coupler. The part of the para-aortic blood pump device <NUM> implanted into the human body includes the blood pump <NUM>, the aortic adapter <NUM> and the internal driveline <NUM>. In another embodiment, the para-aortic blood pump further comprises a coupler. After a surgical operation, the aortic adapter <NUM> is installed into an aorta <NUM>, and an exit site EX is created at an appropriate position of an epidermis of the human body. The part of the para-aortic blood pump device <NUM> situated outside the human body includes the external driveline <NUM>, and the driver <NUM>. The exit site EX is used as a boundary, and the internal driveline <NUM> has a segment covered by a fabric velour for tissue ingrowth to attain infection control. The implanted velour portion is placed <NUM>-<NUM> subcutaneously from the exit site EX. The driver <NUM> is a wearable or portable device.

With reference to <FIG> for the schematic view of a para-aortic blood pump device installed to human body in accordance with an exemplary embodiment of the present invention, the para-aortic blood pump device <NUM> comprises a blood pump <NUM>, an aortic adapter <NUM>, a distal driveline <NUM> (including an internal distal driveline <NUM>, an external distal driveline <NUM> outside the human body), a driveline interconnector <NUM>, a proximal driveline <NUM> and a driver <NUM>. In another embodiment, the para-aortic blood pump device <NUM> further comprises a coupler. The part of the para-aortic blood pump device <NUM> implanted into the human body includes the blood pump <NUM>, the aortic adapter <NUM> and the internal distal driveline <NUM>. In another embodiment, the para-aortic blood pump device further comprises a coupler. After a surgical operation, the aortic adapter <NUM> is installed into an aorta <NUM>, and an exit site EX is created at an appropriate position of an epidermis of the human body. The part of the para-aortic blood pump device <NUM> situated outside the human body includes an external distal driveline <NUM>, a driveline interconnector <NUM>, a proximal driveline <NUM> and a driver <NUM>. The exit site EX is used as a boundary, and the distal driveline is divided into an internal distal driveline <NUM>, covered with velour for infection control, and an external distal driveline <NUM>. The driver <NUM> is a wearable or portable device.

The implant subsystem is described further below.

Implantation is achieved through a relatively small thoracic opening by using less invasive surgical techniques via a left thoracotomy. A thoracic incision is made, for example, at the 7th intercostal space as the primary opening to allow placement of the aortic adapter and the blood pump. Two other small incisions are made at the 6th and 8th intercostal spaces, respectively, to introduce proximal and distal aortic cross clamps. The cross-clamped segment of aorta allows aortic adapter be inserted into the implant site through an excess hole made in the aortic wall. The aortic adapter is flexible and is able to be crimped and constrained into a smaller delivery configuration prior to insertion. Upon completion of delivery into aorta, the crimped aortic adapter shall be released and restore back to its original form with predetermined oversize to the implant site lumen diameter. The material of the aortic adapter hence is important, which shall be flexible but possessing sufficient radial strength to make the delivered adapter conduit remain circular without wall buckling. Candidate aortic adapter construct may include that made by silicone or polyurethane elastomers, or those polymeric constructs reinforced with embedment.

The functional requirements of the aortic adapter of each of the aforementioned embodiments is further described below.

Hemodynamically, the aortic adapter plays a role of flow communication between the blood pump and human systemic circulation. Besides this role, the aortic adapter also serves as a mechanical base to hold the blood pump in place when connected to the aortic adapter. The structure of the aortic adapter has to be elastic but kink resistant, and strong enough to withstand the internal blood pressure and the external contact forces, exerted via contact with the surrounding lung tissue or diaphragm associated with respiratory and thoracic movement.

The aortic adapter <NUM> is implanted inside the aorta with its two conduit ends <NUM>, <NUM> interfaced with the aortic lumen forming a host/graft boundary in the blood stream (see FIGs. 7B and 8B). In order to minimize the host/graft interface discontinuity, both morphologically and elastically, the two conduit ends <NUM>, <NUM> are configured to have a gradually flared inner surface profile and a continuously reduced wall thickness distribution. Such a conduit end design minimizes the step at the interface as well as constitutes a compliance-matching effect for joining the aortic adapter with the aortic lumen. Thrombus at the interface can thus be annihilated because the rate of interface clot aggregation is slower than the natural thrombolysis rate provided by the endothelium. In addition, the gradually thinning conduit wall structure renders the conduit ends <NUM>, <NUM> softer (compliant) which enables the ends expand and contract in concert with the pulsatile blood pressure, constituting a dynamic seal effect to prevent blood from being jammed into the gap of the joint interface which often is the origin of thrombus formation.

The driver of each of the aforementioned embodiments is further described below.

Illustrated in <FIG> and <FIG> are the right-side and left-side perspective views of the driver <NUM>. This compact driver <NUM> contains internal modules including an electromechanical actuator (EMA), an electronic controller, a pair of main battery and reserve battery. This driver <NUM> also comprises external modules of a user interface panel <NUM>, a battery access door <NUM>, a driveline receptacle <NUM>, an AC receptacle <NUM>, and a pair of ventilation windows <NUM>, as depicted in <FIG> and <FIG>.

Critical information in operation and alarm warning of device malfunction and aortic pressure conditions will be displayed on the user interface panel <NUM> of the driver <NUM>. The primary battery can be exchanged through the battery access door <NUM> when primary battery power is exhausted. An electric cable is used to power the driver <NUM> via a connection through the AC receptacle <NUM> when patient is bedridden and power from wall outlet can be utilized in a long-term manner. The proximal driveline <NUM>, <NUM> end is connected to the driver <NUM> through the driveline receptacle <NUM>, through which both electric sensor signal and pneumatic pressure pulse are communicated. A pair of ventilation windows <NUM> are installed on the opposite sides of the driver <NUM> to allow ambient air to flow through the interior of the driver <NUM> for cooling purpose.

The driver <NUM> can be coupled externally to a clinical monitor, wherein the clinical monitor is provided for collecting and displaying real-time clinical waveform data and stores patient data for long-term condition monitoring and diagnosis. Further, a clinical monitor unit provides a user interface to the clinician for displaying device monitoring/diagnostic information and for accessing to driver parameter settings in order to initiate and optimize a patient-specific operational mode.

The EMA is a pneumatic actuator consisting of a brushless servo motor, a ball screw unit, a piston and a cylinder assembly. Atmospheric air is used as a driving medium to reciprocally eject and fill the blood pump.

The EMA module is housed within the driver carried by the implant recipient. The EMA consists of a brushless servo motor, a piston and cylinder assembly and a ball screw unit which comprises a ball screw rod and a nut. The piston is firmly mounted on top of the ball screw rod which is in rotational coupling with the nut of the ball screw unit. The servo motor includes a stator and a rotor and the rotor is integrated with the nut of the ball screw unit. Through the electromagnetic coupling of the rotor/stator induction, the rotor can be rotated in both clockwise and counter-clockwise directions, thereby driving the ball screw rod back-and-forth in a rectilinear manner to result in a reciprocating piston stroke motion in a cylinder. The stroke motion of the piston drives air to and from the implanted blood pump via a driveline connecting the blood pump and the cylinder.

There are two air driving problems associated with the present EMA pneumatic actuator design; namely, the air leak and the condensation of water vapor permeated from the blood through the blood sac wall. The former will impair the pump eject and fill function and driver power consumption leading to degradation of support effectiveness, and the latter will cause bacteria invasion risk to the driveline interior. To solve these two problems, the present EMA incorporates a pressure equalization valve installed in the cylinder chamber wall for air replenishment and moisture reduction. The pressure equalization valve is opened periodically at a predetermined frequency, allowing air mass transport between the cylinder and the ambient until the air pressure in the cylinder chamber equals the atmospheric pressure. The EMA incorporates position and optical sensors to acquire reference trajectory signals for the electronic controller to generate coordinated control commands to drive the piston stroke motion as well as to operate the pressure equalization valve. Hence, the timing and frequency for pressure equalization valve to be activated for air exchange can be programmed in the controller. With such pressure equalization valve incorporated the driving air medium can be constantly maintained in full and dry in the pneumatic actuator to guarantee a long-term safe and effective pumping support of the blood pump.

As illustrated in <FIG>, a para-aortic blood pump device in accordance with an embodiment of the present invention is divided into three portions. The first portion is largely installed inside the human body (which is an implant) with an externalized end to communicate with the second portion; and the first portion comprises a blood pump (including a blood pump pressure sensor), an aortic adapter and distal driveline segments disposed inside and outside human body, respectively. The second portion is installed outside the human body and comprises a proximal driveline and a driveline electronics module (or known as a driveline interconnector). The third portion is installed outside the human body, which is a driver comprising an electromechanical actuator (EMA), a controller circuit, a main battery, and a reserve battery.

The blood pump pressure sensor is built into the proximal blood pump shell and immersed in a small pressure sensing chamber filled with sensing medium, allowing a continuous monitoring of the blood pump pressure. A distal driveline is attached to the pump housing and provides timed air pressure pulses to command ejection and filling of the blood sac. The distal and proximal drivelines provide a pneumatically driven pressure pulse, generated by the EMA inside the driver, to the blood pump; and transmits an electrical blood pressure signal, generated by the pressure blood pump pressure sensor, to the driver. A driving air path (indicated by a dotted arrow line) and an electrical signal path (indicated by a solid-line) is illustrated in <FIG> to describe the functional relationship among the interacted modules. The detailed content of the aortic adapter has been described above. The controller circuit can include a motor controller unit for driving the brushless motor and a micro controller unit as a central processor to process the received pressure signal and generate control commands for motor controller to actuate piston motion.

With reference to <FIG> and <FIG>, a block diagram of the driver internal functions and the key interconnecting signals that are necessary for actuating the blood pump is shown. The raised embodiments of the present invention are further described, as illustrated in <FIG> and <FIG>. In order to explain the driving relationship between the external driver and the implant, it is necessary to refer to the aforementioned contents of the blood pump, driveline, distal driveline, proximal driveline, aortic adapter and driveline interconnector.

The driver receives blood pump pressure signal (electric signal) and processes the signal using trigger detection algorithm to generate trigger signal that commands the EMA actuation in synchronization with the heart rhythm. Upon receiving the assigned trigger timing, the micro controller unit sends commands to the motor controller unit to drive the piston, from eject-to-fill or from fill-to-eject courses, to provide counter-pulsatile circulatory support.

The architecture of the electronic controller incorporates three functional blocks, namely, a micro controller unit (MCU), a motor control block (motor controller unit), and a power management unit. The following Table provides descriptive outlines for each functional block of the driver <NUM>.

The signal acquisition, transmission, processing, and the control logic and command generation and EMA actuation to produce pressure pulse to drive the blood pump is illustrated in <FIG> and <FIG> for the exemplified embodiments previously elucidated.

<FIG> depicts the trigger-detection commands for EMA piston position in relation to the counter-pulsatile pumping. In <FIG> the unassisted aortic pressure (AoP) waveform is expressed in dotted line whereas solid line represents the assisted aortic pressure waveform. When the driver operates with Auto Run mode, the driver operation is initiated and the system carries out a "fill-eject-fill-eject. " cyclic pumping, which represents a normal synchronous counter-pulsation operation. The MCU monitors blood pump pressure (BPP) signals (electric signal) and detects the left ventricle end-diastole (LVED) timing. Upon detection of LVED timing, the MCU generates a F_Trig signal. The time interval between two consecutive F_Trig signals represents an instantaneous cardiac cycle interval (or period). Based on an estimated heart rate calculated from the preceding cycle intervals, the MCU determines the timing, the E_Trig signal, for blood pump ejection. The E_Trig signal provides the timing to command the motor controller unit to drive the EMA according to the predetermined position, velocity, and acceleration profiles. When the ejection stroke is completed and after an optimized dwell time elapse, the EMA is commanded to perform a pre-fill action with a gentle filling speed until the F_Trig signal shows up. Upon receiving the F_Trig signal, the EMA starts to perform a residual fill stroke at a specified piston speed.

When the MCU loses the BPP signal (electrical signal) sent from the blood pump, a washout mode is launched automatically by MCU to drive the EMA, operating at a predetermined pumping rate and driver stroke volume. The washout mode is used to prevent the formation of thrombus in the blood sac, which is a device protection mode instead of providing synchronous circulatory support.

The para-aortic blood pump device of the present invention, with its non-occlusive para-aortic feature, in principle, has a better counter-pulsatile support efficacy as compared to the intra-aortic balloon pump (IABP). Unlike the bedridden or ambulatory IABP patients who have to stay in hospital, the portable para-aortic blood pump device allows the patients to leave the hospital and have ambulatory capability to live a better life at home. Hence, the para-aortic blood pump device of the present invention may further improve patient's disease conditions and quality of life, in addition to the economic benefits gained from a shorter hospital stay.

The trend of LVAD use has been plateaued in recent years, mainly because its application is only indicated to the terminal-stage heart failure patient cohort. Applying early intervention LVAD therapy to the less-sick heart failure patients has long been a clinical objective, which is expected to imposing substantial impact on the future cardiac medicine advancement provided by the broadened use of LVAD therapy. Clinical evidences have shown that certain non-ischemic cardiomyopathy patients supported by LVAD, administered in moderate-to-severe heart failure stage, can be improved with myocardial reverse remodeling toward functional upgrade or sustained myocardial recovery. Nevertheless, this intention of early intervention must be ushered by two enabling factors: an easy and safe surgical procedure, and an effective and adaptive support scheme accompanying disease development. Continuous-flow VAD support is non-physiologic, which deranges the supported heart away from a normal course of recovery. The counter-pulsatile support, however, is physiologic and meets the therapeutic requirement by providing systolic contraction unloading and diastolic perfusion augmentation to promote myocyte reverse remodeling. In summary, the treatment strategy provided by this para-aortic blood pump invention aligns with the early intervention trend development in cardiac medicine. Salutary attributes provided by the para-aortic blood pump device, such as adaptive partial-support, less invasive surgery and counter-pulsatile therapeutics will collectively make the present invention a prospective candidate to contribute to the future advancement of heart failure treatment.

The blood pump of each of the aforementioned embodiments is further described below.

With reference to <FIG> and <FIG> for the schematic and sectional view of a part of a para-aortic blood pump device installed in human body in accordance with the first and third embodiments of the present invention, the implant subsystem of the para-aortic blood pump device comprises the blood bump <NUM>, the aortic adapter <NUM>, and the driveline (or distal driveline) <NUM> attached to the blood pump <NUM>. The blood pump <NUM> comprises a rigid or semi-rigid pump housing <NUM> of the blood pump <NUM> and a blood sac <NUM> having its proximal end closed and distal end open and seamlessly integrated with the aortic adapter <NUM>. The blood sac <NUM> is formed by an oval-shaped sac membrane <NUM>, moreover, the blood sac <NUM> is anchored via a proximal stem <NUM> to the proximal shell <NUM> of the pump housing <NUM> and via a distal stem <NUM> to the distal shell <NUM> of the pump housing <NUM>. The space in the blood pump is partitioned into a blood chamber B and an air chamber A separated by the oval-shaped flexible sac membrane <NUM> suspended via a pair of stress-relief stems (proximal stem <NUM>, distal stem <NUM>) attached to the pump housing <NUM>. The blood chamber B is for blood storage and the air chamber A is for receiving driving air. The pump housing <NUM> comprises the proximal shell <NUM> in which a pressure sensor (or a blood pressure sensor) <NUM> is hermetically embedded, and the sensed pump pressure is transmitted across the sac membrane <NUM>, propagated in an incompressible liquid or jelly housed in a closed pressure sensing chamber <NUM>, and finally received by the pressure sensor <NUM>. After receiving the sensed pump pressure, the pressure sensor <NUM> generates an electrical blood pressure signal. The implanted subsystem components are designed with a size and a shape that can be implanted in patients with a body surface area (BSA) of <NUM><NUM> or larger.

With reference to <FIG> and <FIG> for the schematic and sectional view of a part of a para-aortic blood pump device installed in human body in accordance with the second and fourth embodiments of the present invention, the implant subsystem of the para-aortic blood pump device comprises the blood pump <NUM>, the aortic adapter <NUM>, the coupler <NUM> and driveline (or distal driveline) <NUM> attached to the blood pump <NUM>. The coupler <NUM> is used to couple the blood pump <NUM> to the aortic adapter <NUM> to access the recipient's vascular system. The blood pump <NUM> comprises a pump housing <NUM>, which further comprises a proximal shell <NUM> and a distal shell <NUM>. The construct of the present blood pump <NUM> is similar to that disclosed in <FIG>, except that the distal opening OP is separated from and independent with the aortic adapter. The coupler <NUM> is placed around the flexible neck <NUM> of the aortic adapter <NUM>. In the following, we use the design disclosed in <FIG> to further explain the blood pump design and the underlying design rationale.

Referring to <FIG>, the blood pump <NUM> comprises a molded rigid or semi-rigid blood housing <NUM> which further comprises a proximal shell <NUM> and a distal shell <NUM>. The housing <NUM> has a single opening OP connected to the aortic adapter <NUM> to access the recipient's vascular system. The opening OP of the blood pump <NUM> is seamlessly manufactured together with the aortic adapter <NUM> (or the blood pump <NUM> and the aortic adapter <NUM> are integrally formed), which provides a smooth and continuous interface transition to the neck of the aortic adapter <NUM>. Such integrated blood sac <NUM> and aortic adapter <NUM> assembly is attached to the pump proximal shell <NUM> and distal shell <NUM> via a bonding to the proximal stem <NUM> and distal stem <NUM>, respectively. The blood sac <NUM> is anchored to the top of the proximal shell <NUM>, such that a small, non-flexing circular portion of the blood sac <NUM> is disposed next to a pressure sensing chamber <NUM> in the shell <NUM>.

A miniaturized pressure sensor <NUM> is built into the pump shell <NUM> and fluid communicated to the enclosed pressure sensing chamber <NUM>. This arrangement allows a continuous monitoring of the blood pressure contained in the blood sac <NUM>. Since the pressure sensor <NUM> is not blood-contacting, the long-term sensor reliability and fidelity is assured by the protection of the pump housing <NUM> that isolates the sensor <NUM> and its electric circuit from the influence of chemical corrosion and protein adherence arising from the direct blood contact.

A driveline <NUM> end is attached to the pump shell <NUM> to provide timed air pressure pulses for actuating the eject or fill stroke of blood out of or into the blood pump <NUM>. The driveline design can be multi-luminal or multi-layered so as to accommodate the electrical wires for pressure signal transmission. Metallic coil or fabric mesh can be adopted as the wall reinforcement to enhance the anti-kink capability of the driveline <NUM>. The overall geometry of the blood flow passage in the present blood pump is wide, along with the valveless aortic adapter design and pulsatile pumping operation, constituting a superior blood handling property that avoids high shear-induced hemolysis as well as low flow speed generated thrombus formation or thromboembolism.

The blood sac <NUM> of the blood pump <NUM> includes an innovative design to make the sac membrane <NUM> durable. The blood sac <NUM> is an oval-shaped membrane body of revolution to the centerline of the blood pump <NUM>. There are two polymeric stems (proximal stem <NUM>, distal stem <NUM>) bonded at both housing <NUM> ends, configured respectively to be in a circular disc or an annulus shape, and working as a flexing/stretching relief mechanism to alleviate stress concentration when attached to the rigid housing <NUM>. During pump ejection, the sac membrane <NUM> will be compressed or folded into a tri-lobe shape where the highest strain often occurs at the creased folding line near the rim of the stem attachment (proximal stem <NUM>, distal stem <NUM>). This local high membrane stress/strain arising from large membrane deformation is substantially reduced or absorbed by the deformation of the flexible stem rim as a bendable suspension. Notice that the tri-lobe folding pattern is non-stationary, with creases changing from place to place as influenced by the gravitational direction. In fact, a patient's body posture and orientation including the positions of standing, sleeping, sitting, exercising, etc. may change from time to time in daily activities. The gravitational effect or the body force acting on the stored blood volume in the blood pump <NUM> is hence constantly changing, resulting in a non-stationary crease line initiation and formation. Such running membrane folding line constitutes a unique fatigue resistance feature of the present invention. It is anticipated that the present blood pump <NUM> will possess a much longer durability than that of the conventional fixed folding line membrane design.

Membrane folding and expansion are intimately related to the vortex flow pattern contained in the blood sac <NUM>. The aforementioned sac design features a running folding line formation that makes the vortex structure pattern alternatingly change in response to the folded membrane pattern. The washout effect in the blood pump <NUM> is thus strong and non-stationary, characterized by a random walk-like vortical flow movement. Such randomness in the pump vortex flow structure helps washout the entire blood-contacting surface without creating any fixed low-speed zone near the membrane wall or in the crease area. It has been observed in animal trails that the present blood pump is very thromboresistant.

The (distal) driveline of each of the aforementioned embodiments is further described below.

Referring to <FIG> and <FIG> for a schematic view of the driveline used to connect the blood pump <NUM> with the driver <NUM>. The intracorporeal (distal) driveline <NUM>, <NUM> of the driveline pneumatically connects the blood pump <NUM> to the electro-mechanical actuator housed within the driver <NUM> and also carries the electrical signals acquired from the blood pump pressure sensor <NUM> (see <FIG>). The (distal) driveline <NUM>, <NUM> has one end attached to the blood pump housing and the opposite end having a small external connector for pneumatic and electrical communication. The (distal) driveline <NUM>, <NUM> is tunneled subcutaneously and exits the skin. The outer diameter of the (distal) driveline <NUM>, <NUM> is designed to be small and the tubing material is flexible to minimize the exit site EX stress to the patient's comfort. A portion of the (distal) driveline <NUM>, <NUM> is covered with a porous fabric to promote tissue ingrowth so as to make the exit site infection resistant. The externalized (distal) driveline <NUM>, <NUM> is secured with a short distance beyond the skin exit site EX.

The (distal) driveline <NUM>, <NUM> and its connector are designed to withstand the tensile loads applied during surgical externalization. Post-operatively, the (distal) driveline <NUM>, <NUM> is constantly influenced by muscular motion-induced loads, and the (distal) driveline <NUM>, <NUM> is designed to withstand these loads for their intended service life. The externalized portion of (distal) driveline <NUM>, <NUM> is also designed to be biocompatible and chemically resistant to cleaning agents and disinfectant in clinical use.

The (proximal) driveline of each of the aforementioned embodiments is further described below.

The (proximal) driveline <NUM>, <NUM> is used to connect the (distal) driveline <NUM>, <NUM> to the driver <NUM>. The (proximal) driveline <NUM> has a driveline interconnector <NUM> at one end and a driver connector at the other end. The said driveline interconnector <NUM> encloses a circuit board, which converts analog blood pump pressure signal into digital signal, and a vibrator that provides a tactile feedback in addition to the audible alarms. The driveline interconnector <NUM> comes with a flat shape to prevent torsion from being generated to the (distal) driveline <NUM> when the driveline interconnector <NUM> is anchored against the patient's skin. Further, the driveline interconnector <NUM> and the driveline outer cover are designed to be sealed and protected against water or moisture ingression. Since the (proximal) driveline <NUM> is installed externally, it can be replaced and/or maintained when deemed necessary, hence eliminating the surgical blood pump replacement required when the (proximal) driveline <NUM> is damaged beyond repairable.

Valveless blood pump has two advantages in blood handling characteristics: <NUM>) having no annoying valve sound and valve-induced blood cell damage, thrombus formation and thromboembolism; <NUM>) being more thromboresistant because the two-way pulsatile flow has better surface cleaning effect to minimize protein adhesion and avert interface discontinuity-related clot formation over the blood-contacting artificial surfaces. The flow passage in the valveless pulsatile pump is uniformly much wider that those in the valved pulsatile or continuous-flow rotary pumps. Hemolysis (rupture of red blood cell membrane) generally takes place at narrow flow passages with high flow velocity gradient, such as the gaps between the valve ring and leaflet of a valved pulsatile pump. In addition, low-speed recirculation or stasis zone often exists in the back side of the opened valve which may encourage thrombus to be generated. In a sharp contrast, in a valveless pulsatile blood pump, the shear stress applied on blood cells is literally order of magnitude smaller, and the low-speed stasis zone associated with valve geometry and motion is substantially eliminated, which leads to less blood cell damage or platelet activation, less clot formation and aggregation and translates to lower dose of anticoagulant use and easier and safer post-operative care.

<FIG> shows the blood pump <NUM>, the driveline <NUM> and the feedthrough <NUM> according to another embodiment of the present invention. This embodiment stresses the anatomic adaptivity for an easier blood pump placement and driveline externalization.

As shown in <FIG> and <FIG>, the driveline <NUM> is connected to distal shell <NUM> of the blood pump <NUM>. The blood pump <NUM> has the oval-shaped blood sac and stem assembly <NUM> (including sac <NUM> and stems <NUM>, <NUM>), the pump housing <NUM> (having the proximal shell <NUM> and the distal shell <NUM>), and the pressure sensing system <NUM> embedded in the proximal shell <NUM>. The design details of the oval-shaped blood sac and stem assembly <NUM>, the distal shell adapter <NUM>, the pressure sensing system <NUM>, and the driveline <NUM> are substantially identical or corresponding to those of the aforementioned embodiment, and the modular and functional descriptions are not repeated herein.

In this embodiment, the feedthrough <NUM> is disposed in the distal shell <NUM> of the pump housing <NUM> for coupling the driveline <NUM> to the pump housing <NUM>. Further, the feedthrough <NUM> is configured in a body-fitted shape adjacent to the distal shell <NUM>, making driveline connection in a tangential direction to the pump outer surface. Such body-fitted feedthrough design renders pump housing <NUM> design adaptive to the anatomic space available for blood pump placement. The blood pump <NUM> can be rotatably connected to the interface adapter <NUM> and allow the driveline <NUM> be routed with best suited orientation to enable a smooth subcutaneous tunneling and skin exit. In this way, it favors to anatomic adaptivity to the implant site geometry.

In this embodiment, the feedthrough <NUM> is remotely suited in the distal shell <NUM> while the pressure sensor <NUM> (see <FIG>) and sensing chamber <NUM> are located in the proximal shell <NUM>. More engineering work has to be performed to separate signal transduction route from pneumatic communication route and assures that the blood pump <NUM> be sealed and protected against biochemical fluid invasion that might damage the fidelity of signal transduction after device implantation.

<FIG> shows the pump housing <NUM> has a superficial trench <NUM> formed on the outer surface of the distal shell <NUM> and above an overlapped bonding area DA (<FIG>) of the proximal shell <NUM> and distal shell <NUM>. The superficial trench <NUM> is configured to let the electric wires be extended from the outlet of feedthrough <NUM>, along the trench above the overlapped bonding area and reach the electrodes <NUM> of the pressure sensor <NUM> (see <FIG>). In some embodiments, the trench <NUM> is sealed by a potting waterproof material and/or a ring-shaped cover to maintain a smooth outer surface in order not to irritate or hurt the contacted tissue.

As illustrated in <FIG> and <FIG>, the aortic connector <NUM> generally requires an interface adapter <NUM> to serve as a coupling mechanism to connect the blood pump <NUM> onto the target artery <NUM>. The distal end <NUM> of the connector <NUM> opposite to the interface adapter <NUM> is placed in the vascular wall of the artery <NUM>, and in fluid connection with the human circulation. The proximal end (the interface adapter <NUM>) of the aortic connector <NUM>, however, has a smooth interface transition to geometrically match with the inlet morphology of the blood pump <NUM>. A coupler is commonly required to integrate together the proximal end of the connector <NUM> (the interface adapter <NUM>) with the inlet of the blood pump <NUM>. There are some embodiments that can be used as the aortic connector <NUM>. The one shown in <FIG> is the end-to-side anastomosis of a Dacron or Polytetrafluoroethylene (PTFE) graft <NUM> sutured with the target vessel <NUM>, which can be used in vascular surgery. In some other embodiments, such as the one shown in <FIG>, an insertion type aortic connector <NUM> is used, such as the T-manifold shaped adapter <NUM> disclosed in U. <CIT>, entitled "Dual-pulsation bi-Ventricular Assist Device.

<FIG> and <FIG> illustrate respectively the assembly and components of the long-duration axi-symmetric oval-shaped blood sac and stem assembly <NUM> design. The polymeric material chosen for these components can be, but not limited to, segmented polyurethane with various appropriate durometers. The constituent parts of sac and stem assembly <NUM> includes a flexible membrane sac (blood sac) <NUM>, a proximal stem <NUM>, and a distal stem <NUM>, which are all made in axi-symmetric shapes and integrated together into an assembly of body of revolution relative to a common centerline 62C of the blood pump <NUM>. The proximal stem <NUM> is located at the proximal end <NUM> of the sac <NUM>, and the distal stem <NUM> is located at the distal end <NUM> of the sac <NUM>.

<FIG> shows the integrated sac and stem assembly <NUM> of which the end of the distal stem <NUM> is wrapped and bonded with the inverted membrane 62A at the distal end of the sac <NUM>. <FIG> illustrates the constituent parts before bonding. In general, the sac <NUM> is made by dip molding while the stems <NUM>, <NUM> are injection molded. There is no preferred azimuthal angle for sac deformation to be biased. In theory, thin-walled sac constructed in axi-symmetric oval shape will buckle into a tri-lobe configuration <NUM>, as shown in <FIG>, when pressure differential across the membrane exceeds certain threshold. Such membrane buckling only matters with the final tri-lobe configuration <NUM> (eigenmode) and where the creases <NUM>, or folding lines, shall take place is decided by the initial perturbation that triggers the initiation of buckling instability. The thickness uniformity in cross-sectional planes cut perpendicularly to the centerline (or the axis of revolution) 62C of the assembly <NUM> of the blood pump <NUM> is critical. Care must be exercised to maintain a high-precision sac manufacturing such that an axi-symmetric shape is assured. In real life application, gravity direction stands out as the dominant factor in folding line initiation. Postures of device recipients constantly change, so is the gravity direction relative to the blood pump orientation, according to the daily activities the patient is undergoing, such as standing, sitting, exercising, sleeping, etc. The creases <NUM> of sac deformation, hence, appear in a random walk-like pattern which disperses the high-strain creases non-stationarily over the entire sac. Avoidance of high-strain region dwelling at fixed location, therefore, is the key design guideline to make the sac long-duration.

The embodiment of the present invention innovates a running folding line attribute which makes the high-strain location appearing non-stationarily in the membrane to prolong sac fatigue life. The detrimental stress concentration phenomenon frequently associated with the flexing blood sac is hence improved. Based on this fundamental change in flexing pattern behavior, the fatigue life of membrane will significantly increase attributable to this non-stationary folding line formation characteristic that disperses the high strain areas all over the sac. Further, a salutary outcome accompanying this non-stationary sac deformation pattern resides on the vortex washout effect enhanced within the blood sac. The sac surface will be washed more thoroughly with a random walk-like vortex formation and traversing. The probability of producing constant low-speed recirculation zone dwelling in the near-wall region or creases of the folding lines, hence, will be greatly reduced, resulting in a long-duration, thrombo-resistant blood pump design.

<FIG>, <FIG> and <FIG> show an exemplary embodiment of how an integration method is adopted for mounting the blood sac assembly onto the pump housing <NUM> as well as joining a driveline <NUM> to the proximal shell <NUM>. The distal driveline <NUM> has its first end connected to the blood pump <NUM> at the feedthrough <NUM> and the second end manufactured as a solid connector <NUM> having electrodes <NUM> flushed mounted and covered by a bend relief <NUM>.

The blood sac <NUM> is anchored onto the pump housing <NUM>, which includes a proximal shell <NUM> and a distal shell <NUM>, to facilitate pump fill and ejection actions. In general, the flexural properties of the sac <NUM> and the housing <NUM> are vastly different. To accomplish a long-duration sac design, it requires an intermediate suspension to be installed to render the pump assembly continuous in structural property transition, in particular the membrane flexural deformation. A pair of flexible stems <NUM> and <NUM> is adopted as the suspension that integrates the blood sac <NUM> with the housing <NUM>. As shown in <FIG>, the proximal stem <NUM>, configured in a disc shape, connects with the proximal shell <NUM>; whereas the distal stem <NUM>, in an annulus form, connects with the distal shell <NUM>. Mechanically, the proximal and distal stems <NUM>, <NUM> work as a stress-relief suspension, which not only holds the blood sac within the pump housing <NUM> but also avoids stress concentration to occur at interface attachment, hence prolonging the service life of the sac <NUM>.

As shown in the bottom of <FIG>, the distal shell <NUM> includes an extension, termed distal shell adapter <NUM>, which is to be coupled with the aortic adapter <NUM>. This distal shell adapter <NUM> has a first end <NUM> attached to the sac <NUM> and the second end <NUM> featured like a beak to be coupled with the interface adapter <NUM> (depicted in <FIG> and <FIG>). The first end <NUM> of the adapter <NUM> matches with the distal end of the sac <NUM> smoothly. The opposite second end <NUM>, however, is configured to be paired with the interface adapter <NUM>, and the coupling design goal is to minimize the interface discontinuity to avert clot formation. The adapter <NUM> has a flange structure <NUM> which is disposed in the middle region of the distal shell adapter <NUM>, working as a lock element to be received by the interface adapter <NUM>.

During surgical operation, the closed-end sac design of the valveless blood pump <NUM> would attract air and agglomerate air bubbles in the sac top due to buoyancy force. As shown in <FIG> and <FIG>, a de-airing port <NUM> is hence installed or made in the proximal shell <NUM>, in which a narrow channel <NUM> is provided above the stem <NUM>. An integrated sac and stem septum <NUM> as an extension from the channel <NUM> is used to allow de-airing needle to pierce through and enter into the blood chamber for air removal. In some embodiments, the channel <NUM> is extends alongside the centerline 62C. After the blood pump <NUM> is anastomosed with the target artery <NUM>, the trapped air will be pushed by the arterial blood pressure and emerges and aggregates over the dome space of the sac <NUM>. A thin needle is used to piece through the de-airing port <NUM>, across the channel <NUM>, penetrating the sac and stem septum <NUM>, and finally reaches the interior of the blood sac <NUM> to discharge the accumulated air. The integrated sac and stem septum <NUM> beneath the de-airing port <NUM> is relatively rigid and non-flexing, which will maintain the pierced sac <NUM> without further structural failure due to crack propagation initiated at the pierced slit when subject to cyclic pulse pressure and the neighboring sac stretching and folding.

As illustrated in <FIG> and <FIG>, the pressure sensing mechanism <NUM> is embedded in the proximal shell <NUM>. <FIG> illustrates the sectional details of an integrated proximal shell <NUM>, feedthrough <NUM>, and driveline <NUM>. <FIG> depicts the profile of the proximal shell <NUM> that connects a feedthrough <NUM> extended from the dome of the shell <NUM> for pneumatic and signal communication with the driveline <NUM>.

The pressure sensing mechanism <NUM>, as detailed in <FIG> and <FIG>, has a pressure sensor <NUM>, which is hermetically housed in a metal canister, including a first space <NUM> for fluid communication and a second space <NUM> for accommodating the micro electro-mechanical system (MEMS) pressure transducer and the associated electronic circuit. There are multiple electrodes <NUM> stuck out from the base of the second space <NUM>, to be connected with the electric wires <NUM> of the driveline <NUM> (<FIG>). The second space <NUM> is closer to the driveline <NUM> than the first space <NUM>. The small tube of the first space <NUM> is open to fluid communicate with the sensing medium. Biocompatible fluid or jelly is used as the pressure transmission medium. A cavity or a pressure sensing chamber <NUM>, situated in the proximal shell <NUM> and adjacent to the first space <NUM>, is created to allow sensing fluid be enclosed in. This pressure sensing chamber <NUM> has a distal end separated with the blood chamber by the sac membrane <NUM>. The pressure sensing chamber has two side arms: first arm <NUM> and second arm <NUM>, wherein the arm <NUM> is used for installation of pressure sensor <NUM>, and the arm <NUM> is used for filling and sealing a sensing medium. The blood pressure pulse, hence, can be transmitted across the sac membrane <NUM> and hydraulically communicated with the MEMS sensor <NUM> remotely located in the second space <NUM>.

An embodiment of the present invention innovates a pressure-based blood pump control method and sensor design. A miniature MEMS pressure sensor is adopted with electronic circuit packaged and embedded in the pump housing wall. In principle, MEMS sensor die is very durable owing to its intrinsic micro-scaled structure. Sensor durability, in fact, depends on the packaging design. The present pressure sensing system <NUM> is non-blood contacting and isolated from the corrosive biochemical effects associated with blood, thus providing long-duration signal acquisition and transmission that is required for long-term implantable assist devices.

The driveline <NUM> works as a communicator for electric signal transduction and pneumatic pulse pressure transfer between the blood pump <NUM> and the driver <NUM>. A representative multi-layered driveline <NUM> in the present invention is shown in <FIG>. In this embodiment, the driveline <NUM> has a pneumatic lumen (or inner pneumatic tubing) <NUM>, a plurality of electric wires <NUM>, a middle pneumatic tubing <NUM>, a coil <NUM> (such as a metallic coil), an outer layer tubing <NUM>, a tether <NUM>, a silicone jacket <NUM>, a rigid driver connector <NUM>, and a protective bend relief <NUM>.

The central portion of the driveline <NUM> accommodates the pneumatic lumen <NUM> (or air passage, inner tubing) with a lumen diameter around <NUM>-<NUM>, depending on the preference of choice between lower energy consumption or low-profile for easiness of surgery. The electric wires <NUM> for signal transmission are embedded in the wall of the driveline <NUM>. There are variants of driveline design that may be adopted. Aside from the multi-layered driveline design shown in <FIG>, the driveline <NUM>, for instance, can also be multi-luminal to facilitate electric wires <NUM> embedment in several smaller lumens and allow pulsed air to flow in a larger lumen <NUM>, as shown in <FIG>. One of the smaller lumens can be installed with tether <NUM> so as to limit the stretching of the driveline <NUM> and protect the electric wires from damage when subject to the pull force of externalization.

The inner tubing, or pneumatic lumen <NUM>, is received by the middle tubing <NUM> with reinforcement being sandwiched in between. Between the inner and middle tubing <NUM> and <NUM>, the coil <NUM> (or fabric thread or mesh) can be reflowed (thermally co-molded using heat shrink) as a reinforcement to the driveline wall, making the driveline <NUM> flexible but kink resistant. The outer layer tubing <NUM> covers the inner and middle pneumatic tubing <NUM> and <NUM>, and can be employed to cover the spirally wrapped electric wires <NUM> as a protective sheath. In some embodiments, a non-distensible tether <NUM> can be disposed between the outer tubing <NUM> and the silicone jacket <NUM> of the driveline <NUM>, to strengthen the tensile resilience required during externalization of the driveline <NUM>. Clinically it has been demonstrated that silicone jacket <NUM> is least irritative to the subcutaneous tissues and has the lowest driveline infection rate.

In this embodiment, the pneumatic lumen <NUM>, the metal coil <NUM>, the middle tubing <NUM>, the spiral electric wires <NUM>, the outer tubing <NUM>, the tether <NUM>, and the silicone jacket <NUM> are packaged into a body of the driveline <NUM>. The proximal end <NUM> of the driveline <NUM> is to be plugged into a receptacle housed in the driver <NUM>. The rigid driveline connector <NUM> of the driveline <NUM> is configured to be received by the receptacle of the proximal driveline interconnector <NUM> or the driver <NUM>. The rigid driver connector <NUM> is flush mounted with a plurality of electrodes <NUM> (for example, four electrodes <NUM> in <FIG>) soldered with the electric wires <NUM>. A protective bend relief <NUM> (see <FIG> and <FIG>) is placed over the joint segment of the driveline <NUM> and the driver connector <NUM>, in order to keep the driveline <NUM> from being kinked at the joint. The proximal end of the driveline, including the driver connector <NUM> and the bend relief <NUM>, is kept in low-profile so that it is easy to exit the skin without creating undesirable tunneling trauma.

The connection of the driveline <NUM> onto the blood pump <NUM> is accomplished via a feedthrough <NUM>, as illustrated in <FIG>, and <FIG>. The feedthrough <NUM> can be placed in the proximal shell <NUM> or in the distal shell <NUM>, depending on the anatomy where the blood pump <NUM> is to be implanted. Integrating feedthrough <NUM> with pump housing <NUM> may change the overall outer blood pump configuration and direct the driveline <NUM> in certain direction to meet the implant requirements including driveline externalization route, post-operative skin care and device usability.

As shown in <FIG>, and <FIG>, the feedthrough <NUM> has a first portion <NUM> as an extension of the proximal shell <NUM> in which the pneumatic lumen <NUM>, the tether <NUM> and the electric wires <NUM> of the driveline <NUM> are coupled via an anchor adapter <NUM>, wherein the anchor adapter <NUM> connects with the distal driveline <NUM> to allow driving air to pass through the pneumatic lumen <NUM>. The feedthrough <NUM> further has a second portion <NUM> which is interlocked with the first portion <NUM> working as a bend relief of the driveline <NUM>. The first portion <NUM> is the location where electric wire connection, tether anchorage, and pneumatic lumen bonding and seal with the housing take place. It is required that electric wires must not be exposed to implant site tissue and have to be well protected against pull force exerted during driveline externalization. Also, the connection of pneumatic lumen <NUM> with blood pump <NUM> need to be pneumatically and electrically leakage-free. As shown in <FIG>, these mentioned blood pump integration tasks are performed in the first portion <NUM>. The second portion <NUM>, however, is employed to accommodate and seal these interface joint elements, working as an outer protector that protects the joint from mechanical straining and environmental fluid or moisture invasion.

The modular design pertaining to the first embodiment of the present blood pump invention has been disclosed in <FIG>. In this illustrated embodiment, the blood pump <NUM> includes an axi-symmetric oval-shaped blood sac and stem assembly <NUM> (including a flexible membrane sac <NUM>, a proximal stem <NUM> and a distal stem <NUM>); a pump housing <NUM> having a proximal shell <NUM> and a distal shell <NUM>; and a pressure sensing system <NUM> embedded in the proximal shell <NUM>. A driveline <NUM> is connected to the blood pump <NUM>, including a pneumatic lumen <NUM> and electric wires <NUM> included in the wall thereof. To integrate the driveline <NUM> with the pump housing <NUM>, a feedthrough <NUM> is used to accomplish both electric and pneumatic communication between the driveline <NUM> and the blood pump <NUM>.

As shown in <FIG> and <FIG>, the polymeric manufacturing and bonding method integrating together the blood sac <NUM> and stems <NUM>, <NUM> has been disclosed in the previous section. The design and manufacturing essentials lie in maintaining a high-precision axi-symmetry in parts manufacturing and bonding of the sac and stem assembly. The pressure sensing system <NUM> and the feedthrough <NUM> are installed in the rigid part of the proximal shell <NUM>. As shown in <FIG>, <FIG>, and <FIG>, a compact feedthrough <NUM> design is illustrated. It can be observed that, by the compact feedthrough <NUM>, the electric wiring and connection can be achieved more robust and fault tolerant.

<FIG> and <FIG> depict some flow patterns associated with para-aortic counterpulsation. Over end-diastolic and early systolic phase of left ventricular ejection, the blood pump undergoes pump fill and draw aortic flow into the pump (<FIG>). Both upstream and downstream blood around the connector will be sucked into blood pump by making a sharp <NUM> degrees flow turn. Flow separation and low-speed recirculation zones T-<NUM> hence will be created. Also, extraordinarily high-shear will appear at the corner region of the T-junction. On the other hand, during diastolic phase after aortic valve closure, the stored blood in the pump will be ejected back into the circulation, creating impinging flow on the opposite aortic wall (<FIG>). This side dump, impinging flow involves very high local pressure at the impingement point T-<NUM>, the so-called stagnation point where flow speed is literally zero and all kinetic energy associated with flow velocity is converted into a potential energy termed total pressure. Such high-pressure impinging flow may induce vascular maladaptation including smooth muscle cell proliferation and the resultant wall stenosis, and risk of aortic dissection due to persistent local hypertension. All these non-physiologic flow pattern and the induced high-pressure, high-shear, low-speed recirculation phenomena prevail in the vicinity of the T-juncture region. This turbulent, complex flow anomaly will decay or diminish in a distance of <NUM>-<NUM> times the implanted arterial lumen diameter. The present insertion type aortic adapter is designed to have an inserted conduit length of <NUM>-<NUM>, which covers most of the pump-induced non-physiologic flow region. As the implant site aorta is shielded by the inserted aortic adapter, the biologic vascular wall will be isolated from the influences of the pump-induced pathological stress conditions, hence protecting the implant site artery from remodeling complications induced in the acute or long-term period.

In counterpulsatile support, pump fill and ejection are alternatingly actuated in synchronization with cardiac rhythm, which generates a special T-juncture flow as shown in <FIG> and <FIG>. This insertion type aortic adapter <NUM> is further detailed in <FIG>, a perspective view, and in <FIG>, a sectional view, respectively.

The aortic adapter <NUM> is mold injected with its internal blood-contacting surface <NUM> being manufactured ultra-smooth and continuous without any parting lines. Silicone or other polymeric elastomers can be used as the material. The aortic adapter <NUM> comprises a conduit (or an inserted conduit portion) <NUM> which is to be inserted in the aorta <NUM> (see <FIG>) and a neck (or an extruded neck portion) <NUM> that connects to the blood pump. In this embodiment, the neck portion <NUM> has a neck body <NUM> and an extension part <NUM> disposed on the neck body <NUM>, wherein the extension part <NUM> protrudes from the neck body <NUM>, and the maximum inner diameter of the extension part <NUM> is larger than the maximum inner diameter of the neck body <NUM>. When the neck portion <NUM> is connected to the blood pump <NUM>, the extension part <NUM> snuggly embraces the inlet adapter <NUM> of the blood bump <NUM>, and the neck body <NUM> is disposed around by the coupler <NUM> which integrates the aortic adapter <NUM> to the blood pump <NUM>.

The entire aortic adapter <NUM> is thin-walled to maximize the flow efficiency. To strengthen the thin-walled structure, a pair of Nitinol truss (or truss rings) <NUM> is embedded around the two ends of the conduit portion <NUM> of the aortic adapter <NUM>.

<FIG> is a transparent view showing the locations where the Nitinol truss <NUM> is embedded. Further, the wall thickness of the conduit <NUM> is gradually thinning toward the conduit end <NUM>. The function of the gradually thinning wall thickness is two-fold. First, it minimizes the graft/host junction discontinuity and renders the interface clot formation rate always be lower than the thrombolysis rate provided by the contacted endothelium. Second, the compliance of the conduit is becoming softer toward the conduit ends <NUM>, resulting in a compliance-matching effect when joined with the aortic lumen.

One of the complications that plagued the large stent graft implantation is the endo-leak problem. Type-I endo-leak means the seal of the graft end to the endothelial lumen of the implanted artery is not complete, causing gap created between the graft leading-edge and the arterial lumen. Leaked blood will be trapped and jammed in the gaps and solidified into clot and finally becomes fibrous pseudo-intima which will grow uncontrollably in time. Not only the pseudo-intima will obstruct the grafted artery, but also it may signal and stimulate coagulation mechanism to attract platelet adhesion and leads to thrombotic adverse events to occur. The solution to resolve such endo-leak problem is to have a tight seal of the aortic adapter <NUM> with respect to the attached lumen surface. The present aortic adapter <NUM> comes up with a compliance-matching design concept that enables the semi-rigid conduit ends <NUM> seamlessly attach to the arterial lumen when subject to pulsatile blood pressurization. The outside diameter <NUM> of the aortic adapter conduit <NUM> is slightly larger than the luminal diameter with an oversize ratio (defined as percentage increment in conduit diameter <NUM> relative to luminal diameter) in the range of <NUM>-<NUM> % conditioned at certain nominal blood pressure (say <NUM> mmHg). As blood pressure fluctuates between systole and diastole, or under the pulse pressure generated by counterpulsatile support, the compliance-matching conduit ends <NUM> will dynamically expand and contract in response to the pressure pulsation without creating interfacial gaps.

Generally speaking, a thin-walled tubing made of elastomer is flexible and tends to be compliance-matching, but it is often not strong enough to withstand the compression force exerted due to device oversizing, causing wall buckling of the inserted adapter and the resultant massive bleeding. Hence, the combined use of Nitinol truss structure <NUM> and the elastomeric substrate with appropriate hardness is important. As shown in <FIG>, in this design, the radial stiffness provided by the Nitinol truss <NUM> will help support the aortic adapter <NUM> without buckling, and there is a distance "X" between the outmost boundary <NUM> of the truss <NUM> and the conduit end <NUM>. In some embodiments, this distance X should be assessed and correctly defined. With the support of the Nitinol truss <NUM> as a distensible frame, the gradually thinning conduit end <NUM> will not collapse or wrinkle, which remains in circular shape against the attached lumen wall and dynamically seals with the lumen. Notice that the aortic adapter <NUM> can expand and contract in response to the pressure pulsation, and the seal effect is accomplished in a dynamic manner that the aortic adapter <NUM> and the wall of the aorta <NUM> expand and contract together as a whole to seal the conduit end <NUM> without causing bleeding complication.

Illustrated in <FIG> is a representative embodiment of the Nitinol truss <NUM> which is typically made from a laser-carved Nitinol tube further enlarged under a sequence of expansion and a heat treatment. <FIG> shows the two-dimensional expanded view of the truss ring <NUM>. Each truss has a plurality of wavy structures <NUM>. The truss <NUM> is self-expandable, which can be folded or crimped into a smaller prepacked delivery configuration and expands to resume its original configuration when released after being placed at the desired location.

A convenient measure of conduit rigidity (inverse of compliance) can be represented by the so-called lateral stiffness (LS) whose measurement method is illustrated in <FIG>. LS is defined by the applied force per unit length F divided by the corresponding radial deflection Y. For the present aortic adapter, suitable LS range is <NUM>-<NUM> Nt/mm2. Both embedded Nitinol truss <NUM> and silicone or elastomer substrate will contribute the structural compliance to the coinjected aortic adapter <NUM>. It is best to have an equally distributed compliance so that the stretch and contraction of the composite conduit wall will result in a minimal interlayer delamination tendency to increase the fatigue life of the adapter <NUM>.

The aortic adapter <NUM> is configured to be connected to the blood pump <NUM> to facilitate circulatory support. A quick-connector type coupler is invented herein. Illustrated in <FIG> are the exploded view of the components of the coupler <NUM> that integrate together the aortic adapter <NUM> and the blood pump <NUM>. The present coupler <NUM> includes a flange base <NUM>, a pair of collars <NUM>, and hinges (or a hinge assembly) <NUM> that join together the collars <NUM> with the flange base <NUM>. Spring coils (or a spring coil assembly) <NUM> are loaded in a hinge joint <NUM>, maintaining the collars <NUM> in an open position when unlocked (<FIG>). The locking mechanism is a latch <NUM>, made of slotted leaf spring and fixed by a slab <NUM> welded to one of the ends of the collars <NUM>. The flange base <NUM> has a substantially circular-shaped structure, and each collar <NUM> has an arc-shaped structure. The hinge joint <NUM> is located at the first side 252S1 of flange base <NUM>, and the collars <NUM> are pivotally connected to the hinge joint <NUM> and rotatable to the hinge joint <NUM> and the flange base <NUM>. In some embodiments, the coupler <NUM> and the aortic adapter <NUM> belong to portions of an aortic adapter assembly.

<FIG> shows the picture of the coupler <NUM> in locked configuration, in which a leaf spring latch <NUM> is dropped down from a ramp <NUM> to assure the coupling is safe without the concern of detachment. The collars <NUM> are grooved internally as shown in <FIG>. Integration of aortic adapter <NUM> to the blood pump <NUM> is accomplished through a clamping mechanism using the deformable adapter proximal end <NUM> (<FIG>) serving as a "gasket" between the connected rigid flanges <NUM>, and <NUM> of the coupler <NUM> (see <FIG>) and the blood pump inlet adapter <NUM>, respectively.

In particular, quick-connection type locking can easily be carried out by closing the collars <NUM> that will be latched without a concern of unintentional unlocking, as depicted in <FIG> and <FIG>. The leaf spring latch <NUM> is installed at the tip of one collar <NUM>. During collar closing, this latch <NUM> will be bent as it slides on a ramp <NUM> on the opposing collar <NUM> in the course of locking. As the latch <NUM> clears the top of the ramp <NUM>, it will drop down to the base of the ramp <NUM> by elastic restoring force, working as a safe for preventing incidental latch unlock or collar opening attributed to pump vibration or rocking in long-term use. For a pump explant or exchange that requires module decoupling, the latch <NUM> can be bent and lifted upward by a tool, permitting an unlocking force to be exerted to rotationally open the collars <NUM> and hence disengage the blood pump <NUM> from the aortic adapter <NUM>.

A butt joint design is not feasible for connecting two smooth-surfaced tubing adapters in a blood stream. In most clinical applications, the connected graft is with rough surface to promote endothelialization so that tiny interface discontinuity in the blood stream will be "smoothed out" by the ingrown cells and proteins. The present aortic adapter <NUM> adopts smooth surface approach to avoid thrombotic adverse events to occur. The blood flow in the aortic adapter is bi-directional in response to the ejection and filling action of the counterpulsatile pumping, as depicted in <FIG> and <FIG>. Such strong bi-directional flow and surface washing effect will easily dislodge any neon blood clot newly formed on the rough surface. Hence, smooth surface design is considered more suitable and safer for the present invention. The interface of two joined smooth surfaces in blood stream requires a careful mechanical and hemodynamic design to avert the thrombotic events to occur in situ. In the following, the rationale and design method associated with such a novel joint invention is disclosed.

<FIG> and <FIG> show respectively the two fundamental interface discontinuities that exist in a butt joint connection, such as the steps <NUM>, <NUM> or the gap <NUM> produced between an adapter AB1 (such as an adapter <NUM> of the blood bump <NUM>) and an adapter AB2 (such as the aortic adapter <NUM>). The discontinuities shown in <FIG> and <FIG> are exaggerated and typically in precision machining such joint discontinuities are within <NUM>-<NUM> microns, which is big enough to cause clot and thrombus formation.

In practice, tolerance inevitably exists in matching two separate bodies even if the machining of each body is perfectly performed. <FIG> depicts a misaligned joint of two bodies, of which everything related to parts manufacturing is correctly done except the centerline misalignment. Forward-facing and backward-facing steps <NUM>, <NUM> will be produced and the stagnation flow in the step regions <NUM>, <NUM> is the origin that clot or thrombo-emboli will be generated. In <FIG>, an interface gap <NUM> is created due to non-parallel matching of the connected bodies. The gap <NUM> attracts blood cells to accumulate and further grow into pseudo-intima and such intima growth often is uncontrollable, resulting in blocking the entire blood flow passage in addition to the thrombo-emboli shed from the intimal surface. Interface errors associated with butt joint may exacerbate when the connected bodies are non-rigid. The present aortic adapter <NUM> is semi-rigid, which can be forced into butt joint connection with deformed configuration and enlarged interface discontinuities. Hence, to accomplish the present semi-rigid aortic adapter connection with blood pump, a novel connection method must be invented, as disclosed in the following.

As shown in <FIG> and <FIG>, in some embodiments, the blood pump <NUM> has an inlet adapter <NUM>, comprising a flange <NUM>, a beak <NUM> and a base <NUM>, forming an extension of the housing of the blood pump <NUM>. Multiple eyelets <NUM> are equipped for joining the inlet adapter <NUM> with the blood pump <NUM>.

The inner diameter <NUM> of the beak <NUM> is slightly greater than the inner diameter <NUM> (see <FIG>) of the neck <NUM> of the aortic adapter <NUM>. The contact area between the beak <NUM> and the adapter proximal end <NUM> is an annular cone surface (or a shallow inclined surface or ramp) <NUM> (<FIG>, and <FIG>). The cone angle of the surface <NUM> is substantially in the range of <NUM>-<NUM> degrees measured from the centerline of revolution of the inlet adapter <NUM>. In the initial locking engagement, the collars <NUM>, which have inward grooves <NUM> (<FIG>), will catch the flange of the base <NUM> and the beak flange <NUM> loosely. Along with collars <NUM> closing, the flanges <NUM>, <NUM> (of the beak <NUM> and the coupler <NUM>) will be received and squeezed by the collar grooves <NUM>, hence compressing the sandwiched silicone end <NUM>, and generating the clamping force required for a firm connection.

The clamping force generation mechanism is graphically shown in <FIG>. There are two steps <NUM> and <NUM> that are responsible for creating clamping forces. Prior to collar closing initiation, the step <NUM> should be engaged in the slot <NUM> of the aortic adapter neck <NUM> (<FIG>). Such engagement is accomplished by folding the neck <NUM> first, and then insert the deformed neck <NUM> across the adapter flange <NUM>, letting the elastic restoring force of the aortic adapter <NUM> to resume the folded neck <NUM> back to its original circular shape and allow the step <NUM> be engaged into the slot <NUM>. The groove height Z of collar <NUM> controls the squeezed deformation of the aortic adapter end <NUM>. Referring to <FIG> it can be found that the clearance Z0, in full locking configuration when collars <NUM> are closed and latched, is smaller than the aortic adapter end width Z3 (<FIG>). Speculation of the assembled geometry shows that Z0=Z-Z1-Z2, wherein the Z1 and Z2 are the flange thicknesses of the clamped counterparts as shown in <FIG>. In general, Z3 is greater than Z0, hence, the strained aortic adapter end <NUM> will generate clamping force required for a sealed connection of the beak flange <NUM> together with the collar flange base <NUM>. The strain of the silicone flow adapter end <NUM>, defined as (Z3-Z0)/Z3, in the range of <NUM>-<NUM>%, suffices to guarantee a reliable sealed connection.

<FIG> illustrates the joint characteristic of the present connection of the beak <NUM> with the adapter cone surface <NUM>. When connected, the beak leading edge <NUM> and the cone surface <NUM> will sink into the semi-rigid ramp surface <NUM> of the aortic adapter <NUM> with a depth comparable to the leading-edge radius, typically <NUM>-<NUM> microns, of the beak leading edge <NUM>. In <FIG>, dotted line and number in parentheses represent the beak in initial contact while solid line and numbers without parentheses represent the locked relationship. Notice that the interface discontinuity is reduced by the cone surface depression from its original shape (the dotted line). The width of the collar groove <NUM> controls the interference fit of the coupling. As previously explained, the elastic aortic adapter end <NUM> will be compressed approximately with an <NUM>-<NUM>% strain to provide the required coupling force for a leak-free integration against pulsatile pumping.

The present design of interface connection between the blood pump <NUM> and the aortic adapter <NUM> has two hemodynamic merits for reducing thrombus formation in-situ. First, there will be literally no step or gap type joint discontinuities generated as observed in the conventional butt joint connection. Second, stasis flow located in the interface of the beak leading-edge <NUM> can be minimized. Hence, blood stream flowing over the connected interface will be maintained with high-speed, substantially improving the butt connection drawback, namely the forward-facing or backward-facing steps <NUM>,<NUM> or the gap <NUM> created at the interface.

The present cone surface <NUM> is ramped with an inclination angle to the stream direction. Such ramp interface design averts step or gap be generated at joint due to limited manufacturing precision or matching eccentricity associated with conventional butt connection. Nevertheless, this shallow, cone-shaped ramp <NUM> has an intrinsic shortcoming in fulfilling a concentric centerline alignment of the joined counterparts. The present coupling of the aortic adapter <NUM> with the inlet beak <NUM> has no strict lateral constraint to assure coupling alignment. To connect a rigid beak <NUM> with a semi-rigid aortic adapter flange ramp <NUM> concentrically, a simultaneous catching of the collars around the entire peripheral rim of flange base <NUM> is critical. When a simultaneous catching/locking engagement fails to be accomplished, the initially caught adapter flange ramp <NUM> will be strained more than other free portion, creating a tendency to tilt or disposition the rest contact surface leading to an eccentric pump connection. Such an eccentric connection often is the causal factor that generates step or gap at the interface that induces thrombus formation. This drawback is remedied by having the flange contour <NUM> (<FIG>) of the distal flange of the collars <NUM> configured in such a way that the locking engagement simultaneously includes all circumferential contact areas. When locked, the contacted edge of the metallic beak <NUM> will sink slightly into the compressed silicone ramp <NUM> with controlled depth and further reduces the interface discontinuity when exposed to blood stream. Hence, the conventional interface thrombus can be substantially minimized or annihilated with a moderate administration of anticoagulant regimen.

Structural deformability and method of delivery involved in the present aortic adapter confers a special design feature of the present invention. Material elasticity consideration, in fact, need to be carefully incorporated in the present design. Surgically deliver an insertion type graft into aorta via incised aortic wall is challenging in the sense of peri-operative safety and long-term reliability. The material chosen for the present aortic adapter <NUM> should have a preset memorized shape. During device delivery, the adapter <NUM> is first crimped into a smaller delivery configuration, and such delivery configuration guarantees a quick and safe device implantation. After the crimped aortic adapter <NUM> is placed at the intended implant site, the delivery configuration shall be released to self-expand to its original memorized shape.

Prior to aortic adapter insertion a hole of diameter <NUM>-<NUM> ought to be made in the aortic wall. In making such an access hole, care must be taken to avoid making any cut edges that may become a crack initiation point when wall distension is required for device insertion. A side-biting aortic punch, as disclosed in <CIT>, is an ideal tool for making a large hole in the aorta. With one bite of punch a sound hole without any fractured edges can be made successfully.

The aortic adapter <NUM> morphologically includes two circular tubes joined together into a T-shaped flow communicator for para-aortic circulatory support enforcement. The conduit wall is typically <NUM>-<NUM> in thickness and the material used is polymer such as silicone or polyurethane with appropriate hardness, for example, of Shore A <NUM>-<NUM>. The crimped delivery configuration differs substantially from the commercially available large stent graft covered by Dacron or PTFE (Polytetrafluoroethylene) fabric. In <FIG> illustrated the crimped form of the aortic adapter <NUM> (Nitinol truss <NUM> is embedded and deformed with the polymer substrate). The aortic adapter <NUM> is folded by crushing the inserted conduit portion <NUM> and the T-neck (extruded neck portion) <NUM> is squeezed and flattened accordingly and sunk into the folded adapter <NUM> as shown in <FIG>. Approximately, the folded configuration has a diameter roughly half of the original unfolded circular diameter.

This folded adapter can be held fixed by string tightening. For example, as shown in <FIG>, three fixation strings can be placed around the two edge and center of the conduit and many other fixation methods can be thought of. <FIG>, <FIG>, <FIG> and <FIG> show the four representative stages of a delivery of the aortic adapter <NUM>. The first stage (see <FIG>) depicts the initial penetration of the crimped prepack across the access hole while the prepack is inclined with an angle to the axis of the aorta <NUM>. The second stage (<FIG>) illustrates the complete insertion of the delivery configuration into the aorta, where the prepack is pushed through the access hole with the proximal end being rotated and just dropped into the access hole. The completely inserted prepack is then retracted to have its crimped neck <NUM> aligned with the punched access hole (<FIG>). The string constraints will be released and the prepacked delivery configuration resumes back to its original deployed configuration by elastic self-expansion (<FIG>). The released aortic adapter <NUM> will snuggly be embraced by the aortic lumen, guaranteed by a proper oversize ratio selected prior to delivery. By slightly deforming the T-neck <NUM> the flange <NUM> of coupler <NUM> (<FIG>) can be mounted onto the T-neck <NUM> ready for the blood pump <NUM> to be connected. The blood pump connection can therefore be easily accomplished by seating the inlet beak <NUM> onto the T-neck cone surface <NUM> followed by closing the two collars <NUM> of the coupler <NUM> for a firm device integration.

Additional safety measures can be applied to enhance the hemostasis and stability of the implanted para-aortic blood pump system. Para-aortic placement of blood pump inevitably involves lateral force (perpendicular to the longitudinal direction of the aorta) and torque exerted on the aortic adapter <NUM> due to the weight of blood pump <NUM> and the pumping forces generated by counterpulsatile support. Such device related external forcing may affect the long-term remodeling of the implant site vascular structure. A purse-string suture can be placed in the adventitial layer around the access hole. The purse-string suture can additionally tighten the aortic wall against the inserted adapter <NUM> and works as a protective measure to prevent the enlargement of the access hole. Moreover, surgical tapes can be looped and tightened around the two ends of the conduit <NUM>, strengthening the integration of the inserted aortic adapter <NUM> and aorta as a whole. Freedom from endo-leak can be doubly assured by the compliance-matching design and the banding of the looped tapes. Sometimes, blood pressure may elevate beyond the upper bound that endo-leak free can be assured by compliance-matching. Under such extreme condition, surgical tapes come into play working as a hard limiter that seals the detached adapter ends and assures hemostasis be maintained.

The step-by-step demonstration of implanting the aortic adapter <NUM> is detailed in <FIG>. A crimped prepack form is prepared prior to the start of implantation. After left thoracotomy to expose the target thoracic artery, a cross-clamping distance around <NUM> spanning the implant site is defined. The access hole to be punched in the aorta is first marked with hole periphery identified. A purse string suture is then sewn outside the hole periphery in the adventitial layer. The aorta can be partially dissected from the surrounding connective tissue and a pair of surgical tapes can be looped around the aorta. Cross-clamping and aortic adapter insertion is performed after the above preparation work is done. These insertion steps are described in a sequential order illustrated in <FIG>. First, the aorta is cross clamped to provide an isolated segment without bleeding concern. Then, a large access hole for aortic adapter insertion is made using a custom aortic punch. The folded adapter prepack is then inserted and placed into the cross-clamped aortic segment as illustrated in <FIG>, B, and C, followed by the release and restoration of the folded adapter to its original deployed form (<FIG>). Purse string and tapes are then tightened as an extra protection against endo-leak at hypertension. The coupler is then installed with its adapter flange <NUM> seated into the slot <NUM> of the aortic adapter neck, ready to receive the blood pump <NUM> to be connected. The self-alignment capability of the coupler design enables the beak <NUM> of the blood pump <NUM> be properly positioned and locked together with the aortic adapter <NUM>. The rest implantation steps are conventional, including cross clamp release, blood pump deairing and the initiation of pump support. In general, for a trained surgeon, the cross-clamping period required for aortic adapter insertion is around <NUM> minutes. During this cross-clamping period the abdominal organs will be deprived of blood perfusion and ischemic injury could potentially be incurred. To mitigate this potential surgical insult to the organs, partial femoral-femoral extra-corporeal membrane oxygenation (ECMO) support can be administered to perfuse the abdominal organs and the lower limb. However, whether or not to employ ECMO support is to the discretion of the surgeon. Usually, an ischemic time of <NUM> minutes can be tolerated by a healthy patient.

In summary, an embodiment of the present invention provides a ventricular assist device, including a blood pump, a driveline and a feedthrough. The blood pump includes an axisymmetric oval-shaped blood sac and stem assembly, including a flexible membrane sac, proximal stem, and a distal stem, wherein the flexible membrane sac is attached with the proximal stem and the distal stem as a stress-relief suspension mechanism; a pump housing, including a proximal shell and a distal shell, wherein the stress-relief suspension mechanism is coupled to the pump housing; and a pressure sensing system, embedded in the proximal shell, wherein the pressure sensing system includes a pressure sensor and a pressure sensing chamber which is filled with an incompressible fluid for pressure transmission. The driveline includes a pneumatic lumen, at least one electric wire and a tether, wherein the electric wires and the tether are disposed in the driveline wall. The feedthrough connects the driveline to the pump housing.

An embodiment of the present displacement pump invention discloses a pulsatile blood pump design that incorporates a non-stationary folding line concept in the construct of a long-duration blood sac that may substantially prolong the durability of a displacement type blood pump. Also, a miniature pressure sensing system is disclosed, which can be used to serve as reference waveform for real-time pump control as well as for long-term trending analysis, disease monitoring and diagnosis, based on evidence-based mega data. Further, the embedded pressure sensing system is non-blood contacting, which, hence, greatly improves the reliability requirements in building an implantable sensor system.

The embodiment of the present blood pump invention has at least one of the following advantages or effects. By the feedthrough connection of the driveline to the pump housing, a compact feedthrough design is provided to make the electric wiring and signal transduction more robust and fault tolerant. Further, a compact feedthrough design integrates the sensory electric wires and the pneumatic tubing with the blood pump. This compactness attribute is particularly essential for implant devices. It not only simplifies surgical operation and mitigates peri-operative implantation risks, but also contributes to the reduction of post-operative morbidity associated with driveline infection.

In some blood pump embodiments, the feedthrough is integrated with a distal shell of the pump housing and the feedthrough has a first portion as an extension of the distal shell in which the pneumatic lumen, the tether and the electric wire of the driveline are coupled, and a second portion being interlocked with the first portion working as a bend relief of the driveline, to the advantage of anatomic adaptivity and fitness to the implant site geometry.

An embodiment of the present flow communication invention provides an aortic adapter assembly, for an implantable ventricular assist device, comprising: a T-shaped aortic adapter, including: an inserted conduit portion, an extruded neck portion, wherein the inserted conduit portion is joined with the extruded neck portion, both having a blood-contacting surface which is smooth; and a truss, disposed in the inserted conduit portion; wherein the T-shaped aortic adapter has a polymeric elastomer reinforced by the truss having a Nitinol material; wherein the inserted conduit portion has a wall which is gradually thinning at two conduit ends of the inserted conduit portion, with a proper distance of a tip of the conduit end to the outmost boundary of the truss, and the conduit end possesses a compliance-matching effect to an implant site artery; wherein a proximal end of the extruded neck portion is configured to be joined with an inlet adapter of a blood pump.

The embodiment of the present flow communication invention has at least one of the following advantages or effects. The present invention discloses a flow communicator assembly that enables blood flow transport into and out of a para-aortic ventricular assist device <NUM>, in particular, a counterpulsatile blood pump. Unlike many existing flow communicators that employ rough surface approach to promote endothelialization so as to avert thrombotic adverse events to occur, the present aortic adapter invention adopts a smooth surface, insertion type prosthetic graft concept to construct the flow communicator. Further, a compliance-matching design is embodied around the inserted conduit ends, which combines the gradually thinning wall characteristic with a super-elastic Nitinol supported thin-walled polymer to accomplish the endo-leak free requirement. Abnormal high pressure, high shear, and low-speed recirculation flow phenomena associated with para-aortic counterpulsatile pumping are contained within the artificial surface of the inserted conduit. Hence, the pathologic device-induced hemodynamic influences and risk factors are substantially eliminated and long-term vascular maladaptation related adverse events such as endothelial cell erosion, lipid infiltration, smooth muscle cell proliferation, vascular stenosis, arterial wall dissection, etc. are significantly reduced. To accomplish a sound connection of the semi-rigid flow adapter to a blood pump, a quick connector type coupler is invented. This coupler has a self-alignment interface design that minimizes the step and gap discontinuity and hence reduces the possibility of thrombotic adverse events to occur at interface joint. Accompanying this aortic adapter invention is a specially designed delivery method that assures a quick and safe delivery procedure. The crimped aortic adapter is made into a prepack delivery configuration whose overall size is reduced into half of its deployed configuration. This prepacked adapter can be inserted into the implant site aorta easily and self-expands into its original deployed configuration, resulting in a snuggly fitted flow communicator without the concern of endo-leak. It is not only beneficial for surgical operations that mitigate peri-operative implantation risks, but also contributes to the reduction of post-operative morbidity associated with device-induced flow and implant site vascular maladaptation.

Use of ordinal terms such as "first", "second", "third", etc., in the claims to modify a claim element does not by itself connote any priority, precedence, or order of one claim element over another or the temporal order in which acts of a method are performed, but are used merely as labels to distinguish one claim element having a certain name from another element having the same name (but for use of the ordinal term) to distinguish the claim elements.

Claim 1:
A para-aortic blood pump device, comprising:
a blood pump (<NUM>), comprising a pump housing, a blood sac, and a pressure sensor (<NUM>), the blood sac disposed in the pump housing, and the pressure sensor being installed in the pump housing for monitoring blood pressure inside the blood pump to generate an electrical blood pressure signal;
an aortic adapter (<NUM>), being a T-manifold shaped conduit, coupled to the blood pump, and provided for integrating the blood pump with human aorta;
a driveline (<NUM>), coupled to the pump housing of the blood pump, for transmitting the electrical blood pressure signal received from the pressure sensor; and
a driver, coupled to the driveline for receiving the electrical blood pressure signal, and comprising an electro-mechanical actuator to generate a pressure pulse according to the electrical blood pressure signal, and providing the pressure pulse to the blood pump through the driveline.