DEVICES AND METHODS FOR SUPPORTING CARDIAC FUNCTION
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application claims the benefit of U.S. Provisional Application Serial No. 63/562,549, filed March 7, 2024, the complete disclosure of which is incorporated herein by reference in its entirety for all purposes. This application is also generally related to US Provisional Application Serial No. 63/513,145, filed March 2, 2023; US Non-Provisional Application Serial Nos. 18/177,290, filed March 2, 2023 and 18/056,749, filed November 18, 2022, and 17/252,065, filed December 14, 2020; International Application Serial Nos. US 2022/35177, filed June 28, 2022 and US 2022/35172, filed June 28, 2022; and US Patent Nos. 9,789,236, 8,827,889, 10,293,090, 10,729,834 and 10,172,987, the entire disclosures of which are incorporated herein by reference in their entirety for all purposes.
FIELD
[0002] The present disclosure generally relates to systems and methods for supporting cardiac function and to systems and methods for implanting such devices into an extracardiac site within the patient.
BACKGROUND
[0003] Heart failure (HF) is a complex syndrome that occurs secondary to inherited or acquired abnormalities of cardiac structure and/or function that impair the ability of the left ventricle to eject blood. More than five million people in the United States live with heart failure, and the incidence of heart failure continues to increase, due in part to the expanded aging population and advances in therapeutic management of cardiovascular disease. Transplantation has become the standard treatment for eligible patients with irreversible severe biventricular failure unresponsive to medical or surgical treatment. The supply of donor hearts has decreased in recent years while the demand has increased. As patients become more hemodynamically compromised, there is an increased risk of death prior to transplantation, and a less favorable outcome following transplantation.
[0004] A ventricular assist device (VAD) is a medical device that partially or completely replaces the function of a damaged or failing heart. VADs typically function to reduce myocardial work by reducing ventricular preload while maintaining systemic circulation. A particular VAD may be used to assist the patient’s right ventricle (RVAD), left ventricle (LVAD) or both ventricles (BiVAD), depending on the needs of the patient. VADs may be extracorporeal, paracorporeal, implantable with percutaneous power support, or fully implantable and may provide continuous or pulsatile flow. VADs may be employed on a shortterm or long-term basis, or as permanent destination therapy. Timely VAD therapy may restore hemodynamic stability and end-organ function.
[0005] VADs have an outer casing, which may be a collapsible stent design, and typically include an axial or radial flow pump within the casing to support cardiac function. The casing is typically implanted into one of the lower chambers of the heart, such as the left ventricle, where it receives blood. The pump includes a rotor with impeller blades that rotate and add work to the blood, propelling it from the device to the aorta for distribution to the rest of the body. Recently, systems have been designed to wirelessly power and control the axial pump, thereby obviating the need to implant a power source within the patient. In addition, VADs can be implanted using minimally invasive procedures without the need for open heart surgery. [0006] Although VADs may sometimes be intended for short term use, for example, to provide post-operative assistance to a surgically repaired heart or as a bridge while awaiting a transplant, VADs are increasingly being used as a long-term solution. For example, VADs are now being implanted in patients suffering from congestive heart failure and for destination therapy (DT) for patients with heart failure who are no longer responding to optimal medical management and are not candidates for heart transplant surgery. The broadened use criteria of VADs coupled with a growing imbalance of transplant candidates and available hearts have resulted in an increased frequency of LVAD implantation and longer durations of support. As LVAD utilization grows, expectations of an improved and stable quality of life have become increasingly important as patients desire to return to a normal lifestyle and experience minimal disruptions from their LVADs.
[0007] While blood pumps have been effective for many patients, further improvements that prolong the effectiveness and lifetime of such blood pumps are desired. Implantable blood pumps should be compact so as to facilitate mounting the pump within the patient’s body. They should also provide high reliability in prolonged use within the patient, most typically years, or even decades of service. An implantable blood pump should also be efficient so as to minimize the power required to operate the pump. This is particularly significant where, as in most applications, the pump is powered by a portable battery or other portable power source carried on or in the patient’s body. Moreover, the pump should be designed to minimize damage to the patient’s blood, exhibiting low hemolysis and good resistance to thrombosis. [0008] Blood pumps for LVADs are typically surgically attached to the bottom of the left ventricle of the heart. The pump then draws blood directly from the left ventricle and into the aorta. The left ventricle, however, generates enormous pressure when it contracts and ejects blood into the aorta. For example, the aortic blood pressure may rise to a peak value, called the systolic aortic pressure, of 120 mm Hg or more. This pressure creates stresses and loads on the blood pump and the elements that anchor the blood pump to the left ventricle. These stresses can cause bleeding events, mechanical failure and/or wear on the pump components which decrease the effectiveness of the pump over time.
[0009] To mitigate these issues, researchers have considered implanting LVADs into the left atrium. These devices would be designed to draw freshly oxygenated blood from the left atrium and propel this blood into the aorta, thereby bypassing the left ventricle. Since the left atrium operates at significantly lower pressures than the left ventricle (e.g., typically about 25 mm Hg), this also reduces the stresses and loads applied to the devices that are implanted therein.
[0010] While the concept of implanting VADS in the left atrium shows some promise, this idea also suffers from a number of drawbacks and challenges. For example, the left atrium is positioned within the circulatory system between the lungs and the aorta. Blood in the left atrium flows through the left ventricle and into the aorta. If a blood clot forms on, or within, the implantable pump, it could break away and pass into the aorta and the arteries supplying blood to the brain, resulting in a thrombotic stroke.
[0011] The left atrium also provides challenges due its size constraints and relative positioning in the circulatory system of the heart. The volume of the left atrium is relatively small, leaving very little space for implanting a LVAD pump. In addition, the left atrium is positioned between the pulmonary artery and the left ventricle in the circulatory system. Therefore, it would be difficult to access and deliver a pump into the left atrium in a minimally invasive implantation procedure.
[0012] What is needed, therefore, are improved devices for supporting cardiac function that overcome the challenges and deficiencies with existing devices. It would be particularly desirable to provide intracardiac devices that have a prolonged effectiveness such that these devices may, for example, provide a longer duration of cardiac support. It would also be desirable to provide improved devices and methods for implanting such intracardiac devices. SUMMARY
[0013] The following presents a simplified summary of the claimed subject matter in order to provide a basic understanding of some aspects of the claimed subject matter. This summary is not an extensive overview of the claimed subject matter. It is intended to neither identify key or critical elements of the claimed subject matter nor delineate the scope of the claimed subject matter. Its sole purpose is to present some concepts of the claimed subject matter in a simplified form as a prelude to the more detailed description that is presented later.
[0014] The present description provides systems, devices and methods for supporting cardiac function. The systems and methods are particular useful for longer term implantation in patients that are, for example, suffering from congestive heart failure for destination therapy (DT), bridge to transplant therapy (BTT) and for any patients with heart failure who are no longer responding to optimal medical management and are not candidates for heart transplant surgery, such as end-stage heart failure (HF) patients and patients with HF with preserved ejection fraction (HFpEF). However, it will be recognized that the devices and methods described herein may also be used for shorter term “acute” use as, for example, mechanical circulatory support devices (MSC) to provide hemodynamic support to patients who present with cardiogenic shock and other disorders.
[0015] In one aspect, a system for supporting cardiac function in a patient comprises an elongate housing configured for implantation into the patient exterior to the pericardial cavity. The housing has an inlet and an outlet spaced longitudinally from the inlet. The inlet and the outlet define a primary blood flow path from a left atrium through at least a portion of the housing to an aorta. The system further comprises a motor coupled to the housing and an impeller coupled to the motor and configured to propel blood through the primary blood flow path. The device bypasses the left ventricle by drawing freshly oxygenated blood from the left atrium and propelling this blood directly into the aorta, thereby reducing the pre-load on the left ventricle. Unloading from the left atria may improve right heart failure caused by septal shift that occurs in this patient population when cannulated in the left ventricle, and will potentially allow for pulsatility and decreasing the possibility of GI bleeding. Additionally, avoiding cannulation of the LV apex may improve heart contractility that is often lost due to removal of the apical myocardium that occurs during the coring in conventional LVAD implants.
[0016] In various embodiments, the housing is configured for implantation exterior to the thoracic cavity. The system comprises a first tube that extends from the left atrium, through the pericardium and exterior to the thoracic cavity, where it couples to an inlet of the housing. The system further comprises a second tube that extends from the outlet of the housing to a vessel in the arterial system, such as the aorta or another artery that supplies oxygenated blood to the body. In an exemplary embodiment, the first and second tubes are removably coupled to the inlet and outlet of the pump so that the pump can be removed, replaced and/or repositioned without removing the tubes.
[0017] This configuration provides several advantages. First, implanting the pump outside of the thoracic cavity is less invasive as the pump may be implanted through a minimally invasive opening in the patient, such as a left atrial surgical approach that does not require an open heart procedure and/or a catheter approach through the vascular system. Second, the pump may be easily accessed after surgery without entering the thoracic cavity or the pericardium. This access allows for the pump to be quickly and easily repaired, cleaned, repositioned and or replaced without removing or replacing the tubes that are anchored to sensitive tissue structures within the thoracic cavity.
[0018] In an exemplary embodiment, the first and/or second tubes may comprise a woven graft with a synthetic coating. The synthetic material on the graft replaces the collagen impregnated material found on conventional grafts and may improve blood weeping that typically occurs in the early postoperative period after implantation.
[0019] In another aspect, a system for supporting cardiac function in a patient comprises an elongate housing configured for implantation into the patient and having an inlet and an outlet spaced longitudinally from the inlet. The inlet and the outlet define a primary blood flow path from a left atrium through at least a portion of the housing to an aorta. The system further comprises a motor coupled to the housing and an impeller coupled to the motor and configured to propel blood through the primary blood flow path. The system further comprises a power source, such as a rechargeable battery, coupled to the housing and configured to provide power to operate the pump. The power source may be housed within the housing, or it may be coupled to the housing by a suitable connector and implanted within the patient nearby the housing, or in another location. In either event, the power source is preferably implanted exterior to the thoracic cavity.
[0020] In an exemplary embodiment, the power source is implanted subcutaneously underneath the outer skin surface of the patient and is configured to be recharged directly through the skin surface. This allows the pump to remain fully charged without performing a surgical intervention. [0021] The system may also include a wireless power system (discussed below) for providing primary power to the pump, in which case the battery operates as a backup to the wireless power (i.e., if the wireless power is not available to operate the pump). However, rechargeable batteries have a limited lifetime and must be continuously recharged or replaced entirely. Implanting the battery outside of the thoracic cavity allows for easy access to either recharge or replace the battery, thus ensuring that the pump’s secondary power source is continuously operational. This unique feature provides an important safety measure if and when the wireless power connection between the externally power source and the implanted system is temporarily disrupted. Additionally, this small reserve of implanted power can be used to allow the patient to resume some brief activities of daily living without the need for an external power pack, such as when bathing or dressing.
[0022] In addition, providing wireless power delivery to the pump with a backup battery that can be charged through the skin eliminates the requirement for a transcutaneous driveline that transmits power from an external power source to the pump. This decreases the risk of infection associated with conventional drivelines and improves patient quality of life and mobility.
[0023] In an exemplary embodiment, the housing is configured for implantation in an intercostal space. In this embodiment, the system further comprises an anchor for attaching the housing to one or more rib bones and/or an intercostal muscle of the patient. Anchoring the housing to the rib bone provides a more secure attachment than conventional pumps, which are typically anchored to a heart wall within the more volatile environment of the chest cavity.
[0024] In certain embodiments, the system further includes a tube or other conduit fluidly coupling the outlet of the housing with the aorta, or another artery that supplies oxygenated blood to the body, and an anchor coupled to the tube and configured for securing the tube across a vessel wall within the arterial system. The anchor creates an anastomosis across this wall to provide a fluid pathway from the pump to the arterial system. The term “anchor” as used herein means any device or method for securing the tube to the wall, such as sutures adhesives, self-expanding tubes, screw anchors, hook anchors, balloon anchors and the like.
[0025] In certain embodiments, the system includes a second tube or conduit fluidly coupling the inlet with the left atrium and an anchor coupled to the tube and configured for securing the tube across a wall between the right atrium and the left atrium. The wall may be, for example, an atrial septum. The anchor creates an anastomosis across the atrial septum to provide a fluid pathway from the left atrium to the right atrium.
[0026] In certain embodiments, the housing includes one or more sensors configured for detecting a physiological parameter of the right atrium, such as pressure, temperature or the like. The sensors are coupled to an internal or external controller and configured to transmit data related to the physiological parameter to the controller.
[0027] The anchors on the atrial septum and/or the arterial system may also include sensors configured for detecting one or more physiological parameters of the left atrium and the aorta, respectively. These sensors may also be coupled to the controller. This allows physiological parameters of the right atrium, left atrium and the aorta to be measured and monitored during operation of the pump.
[0028] In various embodiments, the pump comprises an elongate housing having first and second ends, an internal surface, a first inlet for blood disposed between the first and second ends and an outlet spaced longitudinally from the first inlet. The first inlet and the outlet define a primary blood flow path through the housing. The pump includes a rotatable element, such as a rotor, disposed within the housing and spaced from the internal surface to define a clearance therebetween. An impeller is coupled to the rotor for propelling blood from the first inlet to the outlet of the housing along the primary blood flow path. The housing includes a second inlet fluidly coupled to the clearance between the rotor and the housing to define a secondary flow path through the clearance. The blood passing through the secondary flow path continuously flushes the clearance between the rotor and the housing to minimize the formation and/or growth of blood clots and/or to remove heat generated by the rotor. This design, therefore, substantially reduces the risk of thrombosis within the pump or in the patient’s heart or vascular system.
[0029] In certain embodiments, the rotor and the impeller are suspended within the housing such that they are spaced from the inner surfaces of the housing. Thus, the secondary flow path also provides a fluid bearing for the rotor (and in some embodiments, the impeller). In certain embodiments, this fluid bearing is a hydrodynamic bearing. Reducing or eliminating contact surfaces and/or mechanical bearings between the rotating components of the pump and its housing reduces wear on these components, thereby increasing the longevity of the pump. [0030] In certain embodiments, the rotor comprises at least one rotational element, such as a rib, vane, blade or other projection extending into the clearance between the rotor and the housing. The rib(s) are configured to draw blood into the secondary flow path during rotation of the rotor. The rib(s) may extend in a substantially helical or spiral direction around the rotor. [0031] In other embodiments, the rotational element comprises at least one channel, groove, indentation, serration, notch, or the like extending around the outer surface of the rotor. The groove(s) draws blood into the secondary flow path during rotation of the rotor. In certain embodiments, the groove extends in a substantially helical or spiral direction around the outer surface of the rotor.
[0032] The pump may further comprise a motor stator within the housing for rotating the rotor to thereby drive the impeller. The motor stator may, for example, be located in the outer housing and may comprise stator windings that drive the rotor. Alternatively, the rotor may be driven with a motor that is coupled to the housing.
[0033] In another aspect, a method for supporting cardiac function in a patient comprises delivering a pump into an extracardiac location of the patient, generating a fluid path from a left atrium to the pump and generating a fluid path from the pump to the aorta or another artery that supplies oxygenated blood to the body. Blood is drawn from the left atrium into an inlet of the pump through a primary flow path such that the blood flows through an outlet of the pump and into the arterial system.
[0034] In certain embodiments, a tube or other fluid conduit is anchored through a wall between the left atrium and the right atrium, such as the atrial septum, and then extended through the pericardium and the thoracic cavity and coupled to the inlet of the pump. A second tube or fluid conduit is anchored to a wall of the patient’s arterial system, such as the aorta or another artery that supplies oxygenated blood to the body, and fluidly coupled to the outlet of the pump.
[0035] In various embodiments, the method further comprises implanting a battery within the patient and coupling the battery to the pump. In an exemplary embodiment, the battery is implanted subcutaneously within the patient and is recharged by transmitting power to the power source through an outer skin surface of the patient.
[0036] In various embodiments, the method further comprises transmitting a level of power to the pump from a power source located external to the patient and transmitting power from the battery to the pump when the level of power falls below a threshold level. BRIEF DESCRIPTION OF THE DRAWINGS
[0037] The accompanying drawings, which are incorporated in and constitute a part of this specification, illustrate several embodiments of the disclosure and together with the description, serve to explain the principles of the disclosure.
[0038] FIG. l is a schematic depicting an exemplary ventricular assist system;
[0039] FIG. 2 is an enlarged view of the system of FIG. 1;
[0040] FIG. 3 is a side view of an axial flow pump;
[0041] FIG. 4A is a cross-sectional view of the axial flow pump of FIG. 3;
[0042] FIG. 4B is an enlarged cross-sectional view of a rotor of the axial flow pump;
[0043] FIG. 4C is an enlarged perspective view of an impeller of the axial flow pump;
[0044] FIG. 4D is a side cross-sectional view of another embodiment of an axial flow pump;
[0045] FIG. 4E is a perspective view of a rotor for the axial flow pump of FIG. 4D;
[0046] FIG. 5 is another cross-sectional view of the pump of FIG. 3;
[0047] FIG. 6 is a partially transparent perspective view of the pump of FIG. 3;
[0048] FIG. 7 is a partially transparent perspective view of another embodiment of an axial flow pump;
[0049] FIG. 8 is a side view of another embodiment of an axial flow pump;
[0050] FIG. 9 is a cross-sectional view of the pump of FIG. 8;
[0051] FIG. 10 is a schematic view illustrating various components of a ventricular assist system;
[0052] FIG. 11 is a block diagram of an external device that comprises a wireless power transmitter of the system of FIG. 10;
[0053] FIG. 12 is a block diagram of an internal controller and wireless receiver of the system of FIG. 10;
[0054] FIG. 13 is a schematic view illustrating wireless power transfer with the system of FIG. 10;
[0055] FIG. 14 is a schematic block diagram illustrating certain functions of the system of FIG. 10;
[0056] FIGS. 15A and 15B illustrate pressure data for some of the pump embodiments described herein; and
[0057] FIG. 16 is a graph illustrating H-Q curves for some of the pump embodiments described herein. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0058] Particular embodiments of the present disclosure are described hereinbelow with reference to the accompanying drawings; however, it is to be understood that the disclosed embodiments are merely exemplary of the disclosure and that the disclosure may be embodied in various forms. Therefore, specific structural and functional details disclosed herein are not to be interpreted as limiting, but merely as a basis for the claims and as a representative basis for teaching one skilled in the art to variously employ the present disclosure in virtually any appropriately detailed structure. Well-known functions or constructions are not described in detail to avoid obscuring the present disclosure in any unnecessary detail. It should be understood also that the drawings are not drawn to scale and are not intended to represent absolute dimensions or relative size. Instead, the drawings help to illustrate the concepts described herein.
[0059] Systems, devices and methods are provided for supporting cardiac function. In the representative embodiments, the devices are implantable intracardiac devices, such a ventricular assist device (VADs) for assisting or replacing cardiac function, such as in the case of ventricular failure. The intracardiac devices are particularly useful for longer term implantation in patients suffering from congestive heart failure, for destination therapy (DT), bridge to transplant therapy (BTT), bridge to candidacy therapy and destination therapy for any patients with heart failure who are no longer responding to optimal medical management and are not candidates for heart transplant surgery.
[0060] For example, suitable candidates may include patients with American Heart Association (AHA) Stage D/NYHA Class IIIB/IV heart failure who have failed optimal medical management. Failure of optimal medical management is defined herein as: (1) Development of symptoms at rest and reduction in exercise tolerance in previously stable IIIB patients who have received maximally tolerable doses of the four major categories of drugs and are not candidates for, or improved by, CRT; (2) development of drug intolerance in a previously stable IIIB patient due to hypersensitivity (e.g. ACE inhibitors or ARBs), hypotension (e.g. ACE inhibitors, ARBs or beta blockers) or renal impairment (e.g. ACE inhibitors, aldosterone antagonists) resulting in development of some symptoms at rest despite alternative drug therapy; (3) the occurrence of pulmonary congestion, ascites or refractory peripheral edema despite maximal tolerable medical therapy in Class IV patients; and (4) the necessity for intermittent or continuous IV inotropes, intra-aortic balloon pump or hospitalization for heart failure in either Class IIIB or IV patients who have failed maximal tolerable medical therapy or who developed sensitivity to any of the classes of drugs. These patients may be at risk for death from refractory end-stage heart failure.
[0061] However, it will be recognized that the devices of the present disclosure may also be used as mechanical circulatory support devices (MSC) to provide hemodynamic support to patients who present with, for example, cardiogenic shock. In addition, the intracardiac devices may be used in other applications, such as artificial hearts, ECMO devices, implantable heart monitors and defibrillators, pacemakers, or other intracardiac devices.
[0062] In various embodiments, the systems described herein comprise an implanted blood pump, an implanted power receiver and blood pump controller, an externally worn power storage and transmission pack and hand-held control devices used by the patient and/or caregiver to monitor and optimize the blood pump operation. The power storage and transmission pack operates to wirelessly charge the implanted pump, and may also provide diagnostic and monitoring capabilities via device telemetry to optimize pump and controller charging and operation. In certain embodiments, the systems may also include an implanted battery and an external charger that operates as a backup for power to the pump in the event that wireless power is unavailable.
[0063] In the representative embodiments, the implanted pump may include an axial flow pump designed to support cardiac function by pumping blood from the left atrium to the patient’s arterial system. The pump is housed within a casing that may be wirelessly powered and controlled. In some embodiments, the pump may be implanted using minimally invasive procedures without the need for open heart surgery.
[0064] In certain embodiments, the implanted pump may comprise a centrifugal or radial flow pump that moves the fluid by means of the transfer of rotational energy from one or more driven rotors or impellers. Fluid enters the rotating impeller along the pump axis and is driven radially outwards by centrifugal force along its circumference through the impeller’s vane tips. In certain cases, the fluid-flow and turbomadiine efficiencies gained from using centrifugal impellers, as opposed to axial impellers, at selected pressure rise, flow rate, rotational speed, and device diameter (as vell as from the less aggressive interaction between the pump and the blood for a given level of pumping) may outweigh any challenges imposed by the geometry of the centrifugal pump. Levels of pumping that are required in the context of pumping blood can be provided with less input power and less damage to the blood. Operation in-series may result in lower power levels than some axial pumps and may make it possible to reduce the dimensions of the pump. Reducing damage to blood reduces the risk of adverse side-effects during use. Suitable centrifugal pumps for use with the systems described herein may be found in U.S. Patent Nos. 11,813,445, 9, 556,873 and 8535,212, the complete disclosures of which are incorporated herein by reference for all purposes.
[0065] In certain cases, the implanted pump may include inflow and outflow valves that are closeable to seal the pump from a subject’s anatomy. Closing the inflow and outflow valves modulate flow and allow for sealing of the pump, prolonging the life of the pump when not in use.
[0066] In some embodiments, the implanted pump may include a cleaning system configured to introduce and circulate cleaning solutions and therapeutics to the pump. For example, the cleaning system includes an access port that enables rapid circulation of a cleaning solution into the pump. Coupling the cleaning system to the inflow and outflow valves allows for maintenance of the pump while implanted without biological or chemical fouling (such as thrombosis, intimal hyperplasia, encrustation, and the like). In other embodiments, the implanted pump may be accessible by the caregiver through minimally invasive procedures to directly clean and/or repair the pump.
[0067] Referring now to FIGS. 1 and 2, an exemplary ventricular assist system 100 includes a cardiac device 102 that is implanted within the patient exterior to the thoracic cavity. Cardiac device 102 comprises an outer casing 104, an internal lumen running between two open ends and an axial pump (not shown) between the two open ends. Device 102 may also include an internal controller, wireless power receiver and suitable electronics for receiving power from an external source (discussed in more detail below). System 100 may further include a wireless power transmitter (not shown) that includes a transmitting coil to communicate power and data to the controller in device 102 (discussed in more detail below). The power transmitter also may wirelessly communicate information to external patient and clinician devices to enable continuous and remote patient and system monitoring. A more complete description of suitable wireless power systems can be found in U.S. Patent Nos. 11,090,481, 9,919,088 and 9,415,149, the complete disclosures of which are incorporated herein by reference in their entirety for all purposes.
[0068] Intracardiac device 102 may be implanted through open surgical procedures, percutaneously, endoscopically, or through a minimally invasive procedure, for example, by advancing a catheter through the patient’s vascular system. In an exemplary embodiment, device 102 is delivered through an endoscopic or minimally invasive procedure into a suitable target location exterior to the thoracic cavity, such as an intercostal space 120 between two ribs 122, 124. In this embodiment, system 100 may include one or more anchors (not shown) for securing device 102 against ribs 122, 124.
[0069] Device 100 includes an axial pump (discussed below), which includes a motor and an impeller to pump blood from the left atrium into the aorta and throughout the patient’ s body. Thus, the pump essentially bypasses the left ventricle by drawing freshly oxygenated blood from the left atrium and propelling this blood into aorta, thereby reducing the pre-load on the left ventricle. Since the pump is implanted outside of the pericardial cavity, any blood clots that form on the pump will not break away and pass into the aorta and the arteries supplying blood to the brain, thereby eliminating the potential for a thrombotic stroke. In addition, implanting the pump outside of the thoracic cavity is less invasive as the pump may be implanted through a minimally invasive opening in the patient, such as a left atrial surgical approach that does not require an open heart procedure and/or a catheter approach through the vascular system. Also, the pump may be easily accessed after surgery without entering the thoracic cavity or the pericardium. This access allows for the pump to be quickly and easily repaired, cleaned and or replaced without removing or replacing the tubes that are anchored to sensitive tissue structures within the thoracic cavity.
[0070] System 100 further comprises a second housing 130 coupled to device 102 with an electrical connector 132 and including wireless electronics, a receiving coil and a power source, such as a rechargeable battery. Housing 139 is configured for implantation within the patient exterior to the thoracic cavity. In an exemplary embodiment, housing 130 is configured for subcutaneous implantation adjacent to, or near, the outer skin surface of the patient, such as the dermis, the subcutaneous tissue (i.e., the hypodermis), or the area immediately underlying the hypodermis. Housing 130 may be implanted at a location close enough to the outer skin surface to allow electrical power to be transferred through the skin surface to housing 130. This allows the battery to be recharged with an external power source without conducting a surgical intervention in the patient.
[0071] In certain embodiments, the battery within housing 130 is designed to provide the primary power to device 102 and may be recharged periodically to ensure that the pump operates continuously. In other embodiments, system 100 further includes a wireless power delivery system (not shown) that delivers the primary power to device 102 and the battery functions as a backup power source if the wireless power is reduced or lost entirely. Suitable wireless power devices and systems for use with system 102 are discussed below. [0072] In an exemplary embodiment, the battery is configured to store sufficient power to operate the blood pump for a minimum of 120 minutes. This unique feature provides an important safety measure were the wireless power connection between the externally worn power storage and transmission pack and the implanted system briefly disrupted. Additionally, this small reserve of implanted power can be used to allow the patient to resume some brief activities of daily living without the need for an external power pack, such as when bathing or dressing.
[0073] In certain embodiments, system 102 comprises one or more sensors (not shown) within device 102 that detect the power provided to the pump with the wireless delivery system. System 102 may further comprise a processor or controller coupled to the sensors that includes an executable software program for determining whether the power supplied is sufficient to operate the pump within device 102. The processor may be configured to activate the battery when the power falls below a threshold level and to deactivate the battery when the power meets or exceeds the threshold level. This ensures that the pump operates continuously and the battery does not needlessly drain power when not required.
[0074] As shown, system 100 further comprises a first flexible tube 140 extending from an outlet 142 of device 102 and a second flexible tube 150 extending from an inlet 152 of device 102. Second tube 152 extends into the left atrium 160 such that oxygenated blood flows from the left atrium 160 into the pump within device 102. First tube 140 extends into the ascending aorta 162 or through a vessel or vessels that connect to the ascending aorta 162 such that blood flows from the pump into the arterial system of the patient.
[0075] In an exemplary embodiment, the first and/or second tubes 140, 150 may comprise a woven graft with a synthetic coating. The synthetic material on the graft replaces the collagen impregnated material found on conventional grafts and may improve blood weeping that typically occurs in the early postoperative period after implantation.
[0076] In certain embodiments, second tube 150 extends through an opening in the right atrium 170, through the right atrium 170 and into an opening the septal wall (not shown) between the right atrium 170 and the left atrium 160. This opening may, for example, include an anchor or tube that extends through the atrial wall to allow blood to flow from the left atrium through the right atrium and into device 102.
[0077] In other embodiments, second tube 150 includes an extension 154 that extends directly into the left atrium 160. In this embodiment, the blood flows directly from the left atrium '60 through extension 154 and tube 150 and into device 102. [0078] The anchors create an anastomosis across the walls of the heart to provide a fluid pathway from the pump to the arterial system. The term “anchor” as used herein means any device or method for securing the tube to the wall, such as sutures adhesives, self-expanding tubes, screw anchors, hook anchors, balloon anchors and the like. Tubes 150, 140, and the anchors may be configured for deployment through the vascular system such that they are secured to the suitable locations within the patient’s heart. Device 102 may be coupled to tubes 150, 140 in vivo after they have been secured to such locations in the heart.
[0079] Referring now to FIGS. 3-6, an exemplary embodiment of an intracardiac device 200 includes an outer casing 202 having a substantially cylindrical main body 204 with first and second ends 206, 208. Main body 204 preferably has a substantially uniform outer diameter to facilitate insertion of device 200 into an artery or specific delivery device. In some embodiments, however, device 200 may be inserted percutaneously, endoscopically or through an open incision in the patient. In these embodiments, device 200 may have other configurations.
[0080] Main body 204 includes a first inlet 210 located between first and second ends 206, 208, an outlet 212 at, or near, second end 208 and a second inlet 214 at, or near, first end 206. First inlet 210 and outlet 212 are fluidly coupled to each other to define a primary blood flow path 220 through an internal lumen in casing 202. Second inlet 214 is fluidly coupled to either or both of first inlet 210 and outlet 212 to define a secondary blood flow path 230 through an internal lumen of casing 202, as discussed in more detail below.
[0081] First inlet 210 preferably comprises a semi-circular opening in outer casing 202 that extends at least partially around the circumference of casing 202, preferably at least about 25% of the circumference, and more preferably at least about 50%. The exact size and shape of first inlet 210 is designed to provide sufficient flow from a heart chamber surrounding device 200 into primary blood flow path 220. Of course, other configurations are possible. For example, first inlet 212 may comprise one or more openings spaced from each other around the circumference of casing 202. Such openings may have any suitable cross-sectional shape, e.g., circular, square, diamond, rectangular, triangular or the like.
[0082] Device 200 further includes a motor stator (not shown) that is preferably integral with outer casing 202 and may include stator windings and a back iron. A tubular rotatable element 240 is positioned within casing 202 between first and second inlets 210, 214. Rotatable element 240 comprises a rotor portion of the motor and is configured to be rotated (i.e., driven) by the motor stator. In one embodiment, the motor stator includes one or more permanent magnets and rotor 240 includes one or more magnets such that rotor 240 may be rotated around its longitudinal axis by a suitable magnetic field, as is known in the art. Casing 204 may be formed from a magnetically permeable material selected to minimize power losses due to magnetic hysteresis. Electrical conductors (not shown) passing through casing 202 provide power and control signals to the electric motor.
[0083] Rotor 240 is coupled to an impeller 250 that comprises a hub 252 and one or more rotating blades 254 that project from hub 252 for drawing blood through inlet 210. The blades 254 may take any appropriate shape and be of any appropriate number. Blades 254 preferably define a clearance with the inner surface of casing of about 0.1 mm to about 0.8 mm, preferably about 0.2 mm to about 0.4 mm, more preferably about 0.3 mm. In one embodiment, blades 254 have a substantially helical shape such that the blades 254 spiral around hub 252 from the upstream end to the downstream end. Blades 254 may have the same, or a different, pitch. Each blade 254 may have a pitch that varies from hub 252 to the tip of the blade 254.
[0084] As shown in FIG. 4C, impeller 250 may include three blades 254 extending from hub 252 and spaced apart from each other. In certain embodiments, the pitch angle of each of the blades 254 changes in the longitudinal direction such that the angle between the blade surface and the blood flow increases in the downstream direction. Thus, the angle between the blade surface at the hub (where the blood first contacts the blade) is smaller and closer to parallel to the blood flow direction to reduce turbulence and minimize damage to the blood cells upon initial contact with blade 254. As the blood flows along the surface of the blade downstream, this angle increases to provide sufficient power to accelerate the blood flow and propel the blood radially relative to the housing.
[0085] Impeller further comprises a stator 260 that is configured to redirect the flow of the blood from the radial direction to the longitudinal direction towards outlet 212. Stator 260 includes one or more blade-shaped surfaces 266 that have pitch angles that decrease in the downstream direction. Similar to the impeller blades, surfaces 266 are designed to reduce the impact of the radial blood flow at the upstream end of the surface 266 and then to gradually redirect this blood flow in the longitudinal direction. This design reduces turbulence and minimizes damage to blood cells.
[0086] Rotor 240 preferably includes one or more ribs 270 extending from an outer surface 272 of rotor 240. Ribs 270 may comprise blades, vanes or other projections that extend around outer surface 272 and are configured to draw fluid into casing from second inlet 214 as rotor 240 rotates around its longitudinal axis. Ribs 270 preferably have a substantially helical shape with the same orientation as impeller blades 254 such that the flow of blood in secondary blood flow path 230 is in substantially the same direction as primary blood flow path 220.
[0087] Pump 600 provides an efficient design that may pump at least 5 Liters of blood per minute, preferably at least about 6 Liters/minute, at the physiological pressures typically existing within the heart chambers. Applicant has conducted tests of pump 600 to measure the pump’s performance parameters. These tests have shown that pump 600 can pump over 5.5 Liters/minute of water at pressures around 59 mmHG at a rotational speed of about 25.4K RPM, and over 6 Liters/minute (about 6.4 Liters/minute) at pressures around 85 mmHG at a rotational speed of about 27.6K RPM.
[0088] In addition, pump 600 consumes less power than conventional axial flow pumps. Applicant has tested the power consumption of pump 600 in water and has determined that the pump consumes about 30 Watts at 25K RPM and about 32 Watts at 27.6 K RPM.
[0089] Of course, the pumps described herein are not limited to the specific impeller configuration described above and shown in the figures. For example, pump 200 can alternatively employ a fluid actuator that has a shaftless design for the actuation of fluids. The actuator comprises a housing having a plurality of blades. The housing has a hollow, substantially cylindrical shape having a long axis with open ends and an outer and an inner surface. Each of the blades is attached to the inner surface of the housing and extends from opposite ends of housing in a helical pattern. The blades are thereby configured to actuate a fluid by the rotation of the housing along its long axis. The rotation can be achieved by mechanical linkage with a motor, such as by a rim driven connection or an end-driven connection. The rotation can also be achieved by magnetic coupling with external electromagnets or a rotating magnet. The blades may have any suitable cross-section shape, including a substantially parallelogram-like cross-sectional, rectangular, with rounded edges, with sharp edges, and the like. A more complete description of a suitable fluid actuator with a shaftless design can be found in International Patent Application No. PCT/US2019/037047, the complete disclosure of which is incorporated herein by reference in its entirety for all purposes.
[0090] As shown in FIG. 4B, rotor 240 defines a clearance 280 between its outer surface 272 and the inner surface 282 of casing 202. This clearance 280 provides the space for secondary blood flow path 230. Blood flowing through secondary flow path 230 supports rotor 240 within casing 202, thereby providing a fluid bearing for rotor 240 (i.e., with no mechanical bearings). In addition, the blood continuously flushes clearance 280 to minimize the formation and/or growth of blood clots and/or to remove heat generated by the motor and rotor 240. The width of clearance preferably remains substantially constant and is in the range of about 0.1 mm to about 0.8 mm, preferably about 0.2 mm to 0.4 mm, and more preferably about 0.3 mm. This width is preferably substantially the same as the clearance between impeller blades 254 and casing 202.
[0091] The operation of rotor 240 and impeller 250 creates a force that draws these elements forward (i.e., in the direction opposite the blood flow or left to right in FIG. 3). Since device 200 does not contain any mechanical bearings to arrest this movement and prevent the rotor 240 from being drawn so far forward that it contacts the inner surface of casing 202, device 200 includes one or more fluid pressure elements that provides resistance to the flow of blood along secondary flow path 230. This resistance at least partially offsets these axial forces and serves to arrest the forward translation of rotor 240 and impeller 250 within casing 202.
[0092] In one embodiment, the fluid pressure elements comprise an enlarged bulb 292 coupled to, or integral with, rotor 240 and having an outer diameter larger than the outer diameter of rotor 240. Bulb 292 includes an outer surface 296, a first inclined surface 294 adjacent the outer surface of rotor 240 that is transverse to the flow of blood in secondary flow path 230 and a second inclined surface 298 adjacent inlet 210. Outer casing 202 includes a substantially cylindrical inner surface 282 that surrounds rotor 240 to provide clearance 280. This inner surface 282 includes an inclined portion 299 that extends alongside inclined surface 294 of bulb 292 to form a clearance 295 therebetween.
[0093] Clearance 295 has a smaller cross-sectional area than clearance 280. Thus, fluid flowing clearance 295 is compressed creating a higher fluid pressure within this area. This higher fluid pressure applies a force against inclined surface 294 of bulb 292. The force applied against bulb 292 is in the opposite direction of forces applied by impeller 250 and rotor 240 and therefore at least partially resists these axial forces to maintain the axial position of impeller 250 and rotor 240 relative to housing.
[0094] The angle of inclined surface 298 is critical. The larger the angle between inclined surface 298 and the longitudinal axis or the direction of clearance 280, the greater the force that is applied against inclined surface 298 as blood flows therethrough (the relative cross- sectional area of clearance 295 will almost impact these forces). On the other hand, a large change in direction of blood flow through clearance 295 could cause damage to the blood cells. Therefore, Applicant has discovered that the optimal angle for inclined surface is about 5 degrees to about 45 degrees, preferably between about 10 degrees and about 30 degrees. [0095] Of course, it will be recognized that other configurations for providing an offsetting axial force may be included in device 200. For example, the thickness of clearance 280 may be reduced in others places along secondary flow path 230 to create high pressure regions. Alternatively, secondary flow path 230 may include other surfaces or elements, such as projections extending into path 230 from either rotor 240 or casing 204, or a roughened surface on the rotor or casing. In some cases, secondary flow path 230 may be designed to provide a non-linear path through casing 204 to provide additional force vectors in the opposite direction of the flow provided by impeller 250.
[0096] Device 600 may further include an additional magnetic bearing to maintain the axial positions of rotor 240 and impeller 250 if the secondary flow path does not sufficiently resist these forces. In one embodiment, for example, the axial magnetic bearing may comprise a permanent axial housing magnet (not shown) positioning within casing 202 that cooperates with a permanent axial rotor magnet (not shown) positioned in the rotor 240 and/or the impeller 250. In another embodiment, the axial magnetic bearing may include an active magnetic bearing that operates alone or in conjunction with a passive magnetic bearing. In this embodiment, the axial magnetic bearing may comprise, for example, a cylindrical passive magnet designed to counteract the axial forces encountered when rotor 240 is up to speed, surrounded by an active magnet, designed to compensate for additional axial loads, such as those present during pre-load or after-load of impeller 250. In yet another embodiment, permanent magnets may be radially distributed around impeller 250 and/or rotor 240. The attractive force of the magnetic coupling provides axial restraint to impeller 250.
[0097] Device may also include a radial magnetic bearing for stabilizing radial forces against rotor 240 and impeller 250 to minimize contact between these components and casing 202. For example, permanent radial bearing magnets (not shown) may be disposed within casing 202 and designed to cooperate with rotor bearing magnets in rotor 240 and/or impeller 250. The radial bearing magnets allow the rotor 240 and impeller 250 to rotate relative to casing 202 without significant radial contact. In addition, they assist the fluid bearing described above to maintain the annular clearance 280 between rotor 240 and casing 202, as well as the clearance between impeller blades 254 and casing 202.
[0098] Rotor 240 may further include an upstream magnetic bearing 281 positioned at the end of rotor 240 opposite impeller 250 that includes one or more magnets therein (not shown) to form the axial and/or radial magnetic bearings for device 200. [0099] Alternatively, magnetic bearing 281 may function similar to enlarged portion 292 of rotor 240 to provide a relatively high fluid pressure region that creates stabilizing axial forces. For example, bearing 281 is designed with a smaller outer diameter than the remainder of rotor 240 (see FIG. 4B). In this configuration, bearing 281 and rotor 240 define an inclined surface 293 therebetween. Inclined surface 293 may be configured to create a clearance between bearing 281 and the inner surface of housing 202 that has a smaller cross-sectional area than the cross-sectional area of clearance 280. Similar to the above description of enlarged portion 292, this increases the fluid pressure within this clearance and applies a force against inclined surface 293.
[00100] Referring now to FIGS. 4D-4F, 5 and 6, another embodiment of an axial flow pump 500 will now be described. As in the previous embodiment, pump 500 includes an outer housing or casing 502 and a rotor 540 coupled to an impeller 550 and a stator or diffuser 560. Rotor 540 and impeller 550 may be coupled together by any suitable means, such as a threaded screw type connection 509 that allows rotor 540 to rotate impeller 550. Impeller 550 and stator 560 are rotatably coupled to each other with a rotational linkage 511 such that stator 560 remains stationary within housing 502 as impeller 550 rotates.
[00101] In this embodiment, rotor 540 includes one or more grooves, channels or the like 570 extending around an outer surface 572 of rotor 540. Groove 570 preferably extend around outer surface 572 in a spiral or helical direction similar to ribs 270 and function in the same manner to draw blood into a secondary flow path 530 that passes through a clearance 580 between rotor 540 and an inner surface 582 of housing 202.
[00102] Pump 500 comprises an enlarged bulb 592 coupled to, or integral with, rotor 540 and having an outer diameter larger than the outer diameter of rotor 540. Bulb 592 includes a surface 594 adjacent the outer surface of rotor 540 that is transverse to the flow of blood in secondary flow path 530. As in previous embodiments, this compresses the fluid creating a higher fluid pressure within this area. This higher fluid pressure applies a force against inclined surface 594 of bulb 592. The force applied against bulb 592 is in the opposite direction of forces applied by impeller 550 and rotor 540 and therefore at least partially resists these axial forces to maintain the axial position of impeller 550 and rotor 540 relative to housing.
[00103] Referring now to FIGS. 4E and 4F, rotor 540 further comprises a plurality of variable pressure surfaces 520 that are spaced from each other both longitudinally and circumferentially with respect to rotor 540. Variable pressure surfaces 520 each have a first end 522 adjacent groove 570 and a second end 524 circumferentially spaced away from groove 570 such that the surfaces 520 extend from groove 570 to a portion of outer surface 572 between adjacent spirals of the groove 570.
[00104] Variable pressure surfaces 520 are at least partially recessed from the outer surface 572 of rotor 550. Specifically, surfaces 520 are angled in the circumferential direction such that second end 524 is substantially parallel with the outer surface of 572 of rotor and first end 522 extends inwardly at an angle relative to outer surface 572. Thus, first end 522 is recessed from outer surface 572 and gradually angles upward relative to surface 572 until it joins with the outer surface and is no longer recessed. This creates a greater cross-sectional area between the inner surface 582 of housing 502 at first end 522 of pressure surface 520 than the cross- sectional area between inner surface 582 of housing and second end 524 of housing. Also, the cross-sectional area between first end 522 and inner surface 582 is greater than the cross- sectional area of clearance 580 (see FIG. 4D).
[00105] In the event that any portion of rotor 540 moves closer to inner surface 582 of housing 502 (i.e., such that the clearance 580 becomes smaller at that location), the variable pressure surfaces 520 compress the blood flowing past them at the circumferential location that is closest to inner surface 582 of housing 502 to generate a force opposing this motion. This radial force resists the radial force or motion that is moving the rotor towards the housing and would otherwise destabilize the radial position of rotor 540 within housing 520.
[00106] Referring now to FIG. 7, another embodiment of an axial flow pump 500a will now be described. As in the previous embodiment, pump 500a includes an outer housing or casing 502a and a rotor 540a coupled to an impeller 550a and a stator or diffuser 560a. Rotor 540a and impeller 550a may be coupled together by any suitable means, such as a threaded screw type connection 509a that allows rotor 540a to rotate impeller 550a. Impeller 550a and stator 560a are rotatably coupled to each other with a rotational linkage (not shown) such that stator 560a remains stationary within housing 502a as impeller 550a rotates. In this embodiment, stator 560a has a downstream end that includes a cap 590a having convex outer surface 592a near the outlet 512a of the primary blood flow path 520a.
[00107] Similar to the previous embodiment, rotor 540a includes one or more grooves, channels or the like 570a extending around an outer surface 572a of rotor 540a. Grooves 570a preferably extend around outer surface 572a in a spiral or helical direction similar to ribs 270 and function in the same manner to draw blood into a secondary flow path 530a that passes through a clearance 580a between rotor 540a and an inner surface 582a of housing 202a. [00108] Referring now to FIG. 8, an alternative embodiment of an intracardiac device 300 includes an outer casing 302 having a substantially cylindrical main body 304 with first and second ends 306, 308. Main body 304 may have a substantially uniform outer diameter to facilitate insertion of device 300 into an artery or specific delivery device. Main body 304 includes an inlet 310 located between first and second ends 306, 308, a first outlet 312 at, or near, second end 308 and a second outlet 314 at, or near, first end 306. Inlet 310 and first outlet 312 are fluidly coupled to each other to define a primary blood flow path 320 through an internal lumen in casing 302. Inlet 310 is also fluidly coupled to second outlet 314 to define a secondary blood flow path 330 through an internal lumen of casing 302.
[00109] Device 300 further includes a motor stator (not shown) that is preferably integral with outer casing 302 and may include stator windings and a back iron. A tubular rotatable element 340 is positioned within casing 302 between inlet 310 and second outlet 314. Rotatable element or rotor 340 is configured to be rotated (i.e., driven) by the motor stator. In one embodiment, the motor stator includes one or more permanent magnets and rotatable element 340 includes one or more magnets such that rotatable element 340 may be rotated by a suitable magnetic field, as is known in the art.
[00110] Rotor 340 is coupled to an impeller 350 that comprises a hub 352 and one or more rotating blades 354 for drawing blood through inlet 310. Device 300 may further include a diffuser or stator (not shown) that is configured to redirect blood flow from the radial direction to the longitudinal direction and to reduce turbulence of the blood flow passing through blades 354 and into outlet 312. In one embodiment, blades 354 have a substantially helical shape such that the blades 234 spiral around hub 352 from the upstream end to the downstream end.
[00111] Rotor 340 preferably includes one or more ribs 370 (or channels) extending from an outer surface 372 thereof. Ribs 370 may comprise blades, vanes, fins or other projections that extend around outer surface 372 and are configured to draw fluid into casing from inlet 310 as element 340 rotates around its longitudinal axis. Ribs 370 preferably have a substantially helical shape with generally the opposite orientation as impeller blades 354 such that the flow of blood in secondary blood flow path 330 is in substantially the opposite direction as primary blood flow path 320. Similar to the device shown in FIGS. 7-10, rotor 240 defines a clearance (not shown) between its outer surface and the inner surface of casing 202.
[00112] The blood flowing through secondary flow path 330 creates a force against device 300 that is in the opposite direction as the force created by the blood flowing through impeller 354 in the primary blood path 320. The mass flow rate of the blood in secondary flow path 330 is significantly less than the mass flow rate of the blood in primary flow path 320 in order to ensure that the majority of the power applied to pump 300 is consumed with the primary goal of propelling blood through the primary flow path and into the aorta to support function of the left ventricle. In certain embodiments, mass flow rate of the blood in secondary flow path is about 1% to about 20%, preferably about 5% to about 10% of the mass flow rate of the blood in primary flow path 320.
[00113] Since the mass flow rate of the secondary flow path is less than the primary flow path, additional forces must be applied to maintain axial stability of the rotor and impeller. To that end, device 300 includes one or more fluid pressure elements that provides resistance to the flow of blood along secondary flow path 330. This resistance at least partially offsets these axial forces and serves to arrest the forward translation of rotor 340 and impeller 350 within casing 302.
[00114] In one embodiment, the fluid pressure elements comprises an enlarged bulb 392 coupled to, or integral with, rotor 340 and having an outer diameter larger than the outer diameter of rotor 340. Bulb 392 includes an outer surface 396, a first inclined surface 394 adjacent the outer surface of rotor 340 that is transverse to the flow of blood in secondary flow path 330 and a second inclined surface 398 adjacent inlet 310. Outer casing 302 includes a substantially cylindrical inner surface 382 that surrounds rotor 340 to provide clearance 380. This inner surface 382 includes an inclined portion 399 that extends alongside inclined surface 394 of bulb 392 to form a clearance 395 therebetween.
[00115] Clearance 395 has a smaller cross-sectional area than clearance 380. Thus, fluid flowing through clearance 395 is compressed creating a higher fluid pressure within this area. This higher fluid pressure applies a force against inclined surface 394 of bulb 392. The force applied against bulb 392 is in the opposite direction of forces applied by impeller 350 and rotor 340 and therefore at least partially resists these axial forces to maintain the axial position of impeller 350 and rotor 340 relative to housing.
[00116] Referring now to FIG. 9, another embodiment of an intracardiac device 400 comprises an outer casing 402 having a substantially cylindrical main body 404 with first and second ends 406, 408. Main body 404 includes a first inlet 410 located between first and second ends 406, 408, an outlet 412 at, or near, second end 408 and a second inlet (or outlet) 414 at, or near, first end 406. First inlet 410 and outlet 412 are fluidly coupled to each other to define a primary blood flow path 420 through an internal lumen in casing 402. Second inlet (or outlet) 414 is fluidly coupled to either or both of first inlet 410 and outlet 412 to define a secondary blood flow path 430 through an internal lumen of casing 402.
[00117] Similar to previous embodiments, device 400 also includes a motor stator (not shown) and a rotor 440 positioned within casing 402 between first and second inlets 410, 414. Rotor 440 is coupled to an impeller 450 that comprises a hub 452 and one or more rotating blades 454 for drawing blood through inlet 410. Rotor 440 preferably includes one or more ribs 470 extending from an outer surface 472 of rotatable element 440. Ribs 470 may comprise blades or other projections that extend around outer surface 472 and are configured to draw fluid into casing from inlet 410 as rotor 440 rotates around its longitudinal axis. Alternatively, ribs 470 may be oriented to draw blood from inlet 414. Ribs 470 preferably have a substantially helical shape and may be oriented in the same or the opposite direction as impeller blades 354, as described in the embodiments of FIGS. 3-6.
[00118] Rotor 440 defines a clearance 480 between its outer surface 472 and the inner surface 482 of casing 402. This clearance 480 provides the space for secondary blood flow path 430. Blood flowing through secondary flow path 430 ensures that rotatable element 440 does not contact casing 402.
[00119] In this embodiment, the fluid pressure element comprises an enlarged bulb 492 coupled to rotor 440 having an outer diameter larger than the outer diameter of rotatable element 440. Bulb 492 is located near second inlet 414 on the opposite side of rotor 440 from impeller 454. An axial magnetic bearing 481 is located on the side of rotatable element 440 adjacent to or near impeller 454. Locating axial magnetic bearing 481 closer to impeller reduces the distance of the magnetic field, thereby making it more efficient and requiring less power consumption to provide axial stability to the device.
[00120] Systems and methods for transferring energy and/or power to an implanted medical device will now be described. In the representative embodiments, the devices are implantable intracardiac devices, such the ventricular assist devices (VADs) described above. However, it will be recognized that these systems may be used to power mechanical circulatory support devices (MCS), artificial hearts, ECMO devices, implantable heart monitors and defibrillators, pacemakers, or other intracardiac devices as well as patient and disease management devices such implantable pulmonary artery pressure monitors.
[00121] FIGS. 10-12 are schematic diagrams of a ventricular assist system 700 that may include one or more pumps 702, such as the axial pumps described above. System 700 includes an external device 706 that comprises a wireless power transmitter 720, a magnetic coil (not shown), a power source 722, such as a rechargeable battery or the like, an antenna 712 and the associated electronics 726 for transferring energy or power from antenna 712 to an internal controller 708. In some embodiments, external device 706 may include a user display 728 for providing information to the patient related to various parameters of the system, such as the power delivered to the motor 702, the speed of the pump and the like. User display 728 may also include a user interface that provides input controls for the patient to directly modulate certain parameters of the system (discussed below). Device 706 may also include a suitable coupler for removably coupling the power source to an external charger 710. Alternatively, the power source may be situated remotely from device 706 and may be coupled to device 706 wirelessly or via a direct wired connection.
[00122] In certain embodiments, the wireless power transmitter 720 within device 706 includes an amplifier or controller AC power supply that is operably coupled to a drive loop, to provide RF energy to the drive loop. A sensor, such as a directional coupler, vector network analyzer or the like, provides information from the drive loop. The drive loop may comprise a single-turn or multi-turn drive loop.
[00123] Device 706 may comprise a wearable device that can be attached to, or worn by, the patient. Preferably, the wearable device is secured near a portion of the patient’s body directly over, or near, antenna 714 or internal controller 708. In some embodiments, device 706 also includes an attachment element (not shown) for attaching device 706 to a patient. The attachment element may comprise any suitable releasable coupling element, such as fasteners, snaps, interference fit structures, Velcro and the like. Wearable device 706 may be configured for direct attachment to the patient’ s outer skin surface or for attachment to a variety of different wearable garments, such as pants, belts, chest straps, pendants, sashes, hats, jackets, shirts, vests, shorts, skirts, bibs, coveralls. The wearable garment may include additional features, such as multiple hardpoints, straps or the like, for ensuring that the antenna contacts the patient’s skin surface and engages this surface sufficiently to transmit the power therethrough with minimal losses. The wearable garment may also include a waterproof outer shell around to insulate the antenna, transmitter and associated electronic circuits from water or other fluids that may contact the garment.
[00124] System 700 may further include one or more relay resonators (not shown) positioned between external device 706 and internal controller 708. In this embodiment, one or more of the relay resonators may be provided in a wearable device, while device 706 remains in a position remote to the patient. Alternatively, the relay resonator may be disposed in or on a different wearable device. In certain embodiments, the relay resonator may be larger than implanted resonator in controller 708 and is operable to increase the range of the wireless energy transfer.
[00125] The internal controller 708 may be implanted in a suitable location within the patient. Controller 708 may be implanted subcutaneously within the patient, or it may be implanted at another suitable location exterior to the thoracic cavity. Controller 708 comprises an antenna 714 for receiving power from transmitter 706, a power source 730, such as a rechargeable battery, a motor driver 732 for transferring the power to pump 702 and associated electronics 734, such as memory, power recovery, telemetry and the like. Controller 708 may further include one or more sensors 736 that detect a variety of operational parameters for the pump 702, such as the power transmitted to the pump, the pump speed, the maximum output pressure, the negative intake pressure and the like.
[00126] Motor driver 732 communicates with pump 702 to drive the pump motor and control blow flow through the pump. In some embodiments, controller 702 wirelessly communicates with the heart pump. In other embodiments, controller 702 is connected to pump 702 via direct wire connections. In other embodiments, controller 708 is integrated into the housing of the heart pump.
[00127] Controller 708 may have the ability to monitor the function of the heart pump and/or the cardiac function of the patient. In certain embodiments, controller 708 includes one or more sensing electrodes (not shown) to receive, filter, amplify and analyze an EKG signal. The controller may measure real time function and power consumption of the heart. These measures can then be used to derive many variables of pump function, including speed, flow, suction, pressure head of the pump and an occlusion event. Controller 708 may also have multiple modes, such as a continuous flow mode and/or a pulsatile flow mode, wherein the pump speed is attuned to the systole and diastole periods of the cardiac cycle of the patient. The system may further include an external control unit with a user interface for controlling the specific mode of operation of controller 708, which may include fixed speed (RPM) operation, fixed flow rate operation as well as fixed power operation. A more complete description of one representative controller for use with the system described herein can be found in U.S. Patent No. 9,919,088, the complete disclosure of which is incorporated herein by reference in its entirely for all purposes.
[00128] In some embodiments, internal controller 708 may include a load loop operably connected to provide energy to the pump 702, and a receiver resonator that is inductively coupled to the load loop. During operation, the transmitter resonator and the receiver resonator may form a magnetically coupled resonator (MCR), such that the pump 702 is energized from RF energy from the amplifier that is inductively transmitted from the drive loop, to the MCR, and is inductively transmitted from the MCR to the load loop. MCRs induce power transfer between two components through a matching of the resonance frequency between a source resonator and a receiver resonator. A controller may be operable to receive data from the sensor, and to control the operating parameters to optimize the energy transfer efficiency in the MCR. A more complete description of suitable wireless power transmitters can be found in U.S. Patent Nos. 8,299,652, 8,827,889 and 9,415,149, the complete disclosures of which are incorporated herein by reference in their entirety for all purposes.
[00129] Transmitter 706 is also configured to transmit various control signals to internal controller 708. Likewise, controller 708 is operable to control operation of pump 702 and to transmit data back to transmitter 706. The control signals provide feedback control to the pump based on physiological requirements of the patient. In some embodiments, the control signals are based on the power transferred to the receiver 708. These control signals may, for example monitor the dynamic power coupling between the transmitter and the receiver to ensure the efficient transfer of power therebetween.
[00130] Power may be transferred from wearable device 706 through the air 712 and the patient’s tissue 714 to internal controller 708. In some embodiments, wearable device 706 may be in direct contact with the patient’s tissue, which reduces or eliminates the amount of air 712 in the power transmission pathway. Internal controller 708 then transfers the power to the motor in pump 702, which drives the impeller and provides work to the blood 704 to propel the blood through pump 702. Power may be lost between all of these components due to various inefficiencies. For example, power may be lost between the receiver and transmitter due to a number of factors, including the distance between the coils, the offset between the center of the coils, the substance between the coils and the angle between the coils. The position and orientation of wearable device 706 may, therefore, change the efficiency of this power transfer, which may in turn effect the operation of pump 702. In certain embodiments, the wearable device 706 and/or the internal controller 708 include sensors (not shown) that detect the power transferred from wearable device 706 and the power received by internal controller 708. A controller (not shown) housed within, or coupled to, wearable device 706 calculates the difference between these two power values to ensure that the power loss remains within an acceptable range to operate pump 702. [00131] In certain embodiments, the wearable device 706 and/or the internal controller 708 may also include sensors indicating the position and/or orientation of the wearable device 706 relative to the internal controller 708. In certain embodiments, the sensors indicate the absolute position of the transmitter. In other embodiments, the sensors may indicate the position of the transmitter relative to the receiver. Suitable sensors may include capacitive displacement sensors, eddy-current sensors, Hall effect type position sensors, inductive sensors, laser doppler sensors, linear variable differential transformers (LVDTs), photodiode arrays, piezo-electric transducers, position encoders, potentiometers, optical proximity sensors, magnetic angle sensors, TMR, GMR or AMR angle sensors, orientation sensors, RF interferometry based sensors and the like. The sensors may be coupled to external device 706, internal controller 710 or both. The controller is configured to compare the position and orientation of the transmitter and receiver with the power delivered to the motor within the pump to, for example, determine if the wearable device 706 is positioned correctly on the patient (i.e., at the optimal distance, angle and/or coil center offset to achieve an acceptable power transfer therebetween). [00132] In one such embodiment, system 700 includes sensors that detect the physical distance between the antenna coils in wearable device 706 and internal controller 708. The sensors are coupled to the controller and configured to transmit this distance to the controller, either wirelessly, or through wearable device 706. The controller is configured to compare this distance with the power loss detected between the receiver and the transmitter and/or the absolute power delivered to the motor within the implanted device to determine if the coils are, for example, positioned close enough to each other to provide sufficient power transfer to operate pump 702. In some embodiments, controller uses the extra power received to charge an implanted battery and drain energy from it when the received power is insufficient.
[00133] In another embodiment, system 700 includes one or more sensors that detect the relative angle of the coils in transmitter 706 and receiver 708. The sensors are coupled to the controller and configured to transmit this angle data to the controller. The controller is configured to compare this angle data with the power loss detected between the receiver and the transmitter and/or the absolute power delivered to the motor within the implanted device to determine if the coils are, for example, oriented at an angle close enough to parallel to provide sufficient power transfer to operate pump 702.
[00134] In yet another embodiment, system 700 includes one or more sensors that detect the offset (if any) between the centers of the coils on the transmitter and the receiver. The sensors are coupled to the controller and configured to transmit this data to the controller. The controller is configured to compare this data with the power loss detected between the receiver and the transmitter and/or the absolute power delivered to the motor within the implanted device to determine if the coils are, for example, centered relative to each other to provide sufficient power transfer to operate pump 702.
[00135] In certain embodiments, one or more of the controllers is configured to automatically adjust parameters of the wireless power based on either detecting the power delivered to the motor within the pump (i.e., if this power drops below a threshold level), or detecting changes in the relative position and/or orientation of the magnetic coils in the transmitter and receiver. In certain embodiments, the power delivered to transmitter 706 may be adjusted directly to account for these changes. In other embodiments, the frequency of the amplifier is adjusted to adapt to changes in the position and/or orientation of the magnetic coils in the transmitter and receiver. In yet another embodiment, the coupling between the magnetic coils is actively controlled with one or more matching networks that are operable to adjust the impedance in the system, such that the power level delivered to the motor remains at or above a threshold level.
[00136] System 700 may further comprise a user interface (not shown) within user display 728 or wirelessly coupled to external device 706 and including one or more alert indicators that indicate whether the wearable device is positioned at the optimal distance and/or orientation relative to the receiver 708. The alert indicators may be visual, audible, tactile (e.g., vibration) or the like, and they may be housed on, or within, wearable device 706 or wirelessly coupled to wearable device 706, for example, on a separate mobile device or the like. The user interface provides immediate feedback to the patient and/or the healthcare professional that the wearable device 706 should be repositioned to establish sufficient power transfer to pump 702.
[00137] In one such embodiment, the user interface includes one or more position indicators that indicate: (1) a distance between the wearable device 706 and the receiver 708; and/or (2) the positional offset between the centers of the coils in these two devices. The position indicator alerts the patient if the wearable device 706 is not positioned properly to achieve an efficient power coupling with the receiver 708.
[00138] In another embodiment, the wearable device 706 includes an angle indicator that alerts the patient of an unsuitable angle between the coils. Generally, the closer these two coils are to a parallel angle relative to each other, the less power will be lost during transfer. This angle indicator provides an alert to the patient if the wearable device 706 needs to be repositioned to reestablish this angle. [00139] In certain embodiments, system 700 is configured to automatically provide a constant level of power to pump 702 and/or to blood 704. This ensures that pump 702 will continuously pump blood flow through the heart at a sufficient rate regardless of any changes in the system that would otherwise reduce this level of power, such as power loss due to inefficiencies and/or changes in the relative positions of the coils within the transmitter and receiver, including the distance between the coils, the offset between the center of the coils, changes in the substance(s) or material(s) located between the coils and the angle between the coils.
[00140] In this embodiment, the implant will include a sensor (not shown) for continuously measuring the power received by the pump. This sensor may be, for example, housed within the implant housing and coupled to the power electronics provided to the pump. The sensor is coupled to internal controller 708 (either directly through wired connections or wirelessly). Internal controller 708 receives this data related to the power received by the pump and transmits it to an external controller that is coupled to, or disposed within, external device 706. External device 706 is configured to modulate one or more of the parameters of the wireless power transmission based on this data, such that the power received by the pump remains substantially constant, e.g., about 5 Watts to about 20 Watts, or about 7 Watts to about 17 Watts. In an exemplary embodiment, the second lever of power is about 15 Watts.
[00141] The system is preferably configured to deliver a constant amount of power to the pump to ensure that the pump draws flows through the heart at a target rate. In an exemplary embodiment, the target power delivered to the blood is about 1.5 W to about 2.0 W or about 1.73 W, which is sufficient power to pump at least about 6 Liters of blood per minute or 100 mL/second (i.e., based on a mean arterial pressure (MAP) of about 130 mm Hg). In this regard, the implant may include an additional sensor located, for example, within the implant housing near its outlet to measure the power delivered by the pump to the blood. This additional sensor may be coupled to internal controller 708 such that the system can modify parameters in the event that the pump is not transferring the target power to the blood.
[00142] In certain embodiments, the controller is programmed to direct a continuous level of power (and thus a continuous level of blood flow through the pump). In other embodiments, the controller is programmed to direct pulsatile flow that may be, for example, synchronized with physiological blood flow through the patient’s heart. For example, the power level (or the speed of the pump) may be increased or decreased during a systole or diastole state in the heart. [00143] FIG. 9 illustrates a more simplified version of a system 900 that includes a power source 910, a transmitting resonator 906, a receiving resonator 914 and an implanted device 902 that may include a motor and a pump, as described above. As shown, system 900 will provide a constant power level of about 5 W to about 20 W to the pump, preferably about 13 W to about 17 W, or about 15 W. This power level, however, is lost as it is transmitted from power source 910, through resonator 906, the air 920, the patient’s skin 930, the patient’s tissue 940 and the implanted device 902. Thus, system 900 is designed such that power source 910 generates about 20 W to about 40 W, preferably about 30 W, of power. Transmitting resonator 906 is designed to delivery this power through air 920, skin 930 and tissue 940 to receiving resonator 914, which delivers about 10 to 20 W, or about 15 W, to implanted device 902. Device 902 may include one of the embodiments described and delivers about 1 W to 3 W, or about 1.5 to 2 W, preferably about 1.73 W, to the blood flowing through device 902 at a rate of approximately 6 Liters / minute. Power usage would be less at lower flow rates, such as 4 Liters / minute.
[00144] The overall efficiency of the system ensures that an appropriate amount of power is transmitted to the device 902 regardless of any other inefficiencies in the pathway between the energy source 910 and the receiving resonator 914. In certain embodiments, the power level transmitted to the implanted device is at least about 40% of the power level transmitted from the energy source 910, preferably at least about 50%.
[00145] The system also is configured to minimize the specific absorption rate (SAR) of the tissue 940 between the transmitting and receiving resonators. The SAR is generally defined as the measure of the rate of radiofrequency (RF) energy absorption within the body tissue. In certain embodiments, the system is configured to maintain the SAR at or below about 1.5 W/Kg.
[00146] FIG. 14 illustrates another ventricular assist system 800 that may include one or more implantable pumps 801, such as the axial pumps described above. The system generally comprises five components: (1) system-specific implant tools; (2) an externally worn power storage and transmission pack; (3) an implanted transmitted power receiver, battery and blood pump controller; (4) an implanted blood pump; (5) hand-held control devices used by the patient and clinician to monitor and adapt the blood pump operation and to recharge the battery. [00147] The patient sourced treatment and monitoring unit may be wirelessly coupled, e.g., via Bluetooth, to a patient’s smartphone or tablet to report clinical parameters and diagnostic parameters of interest to the patient. The portable patient power source may comprise a charging apparatus that can be worn by the patient to wirelessly charge the implanted pump and pump controller as well as diagnostic and monitoring capabilities via device telemetry to ensure pump and controller charging and operation. The therapy control and monitoring system may contain an expanded set of diagnostic and monitoring capabilities for the physician to be able to reprogram the pump and controller. All of these systems may communicate to a patient safety and monitoring network similar to cloud-based networks currently in use by cardiac rhythm management devices.
[00148] The small size and unique dimensions of the blood pump and blood conduiting allow the pump mechanism to be implanted using minimally invasive surgical techniques, an important advantage when seeking to reduce trauma, improve adverse events and implant recovery times associated with the system. The system includes pre-operative precure planning tools, and specialized implant tools used to place and secure the blood pump and blood conduiting to the desired anatomical structures.
[00149] The continuous power needed to operate the blood pump and system controller is provided wirelessly, there are no physical, mechanical, or electrical connections between the implanted system and the patient carried external power source. Infection is the second most frequent adverse event after bleeding in the first 3 months following LVAD implantation. The use of wireless power technology allows this system to avoid the morbidity and mortality associated with infection due to the percutaneous connections between implanted system and external mechanical or electrical power source. Tablet, phone, and watch-based devices provide patients, caregivers, and clinicians access to information regarding the operating status of the external and implanted systems. Additionally, these devices allow authorized users the ability to modify pump operation, such as the pump flow, operating conditions, algorithms and how the system adjusts pump blood output to patient activity.
[00150] As shown, system 800 includes a first module 802, a second module 804 and a third module 806 that are located external of outer skin surface 840 of the patient and a fourth module 808 and a fifth module 810 that are located within the patient. Module 802 generally functions as an external battery charger, such as a DC or AC battery charger, for charging module 804. In an exemplary embodiment, module 802 operates with an AC supply in the range of about 110V to about 240 V and a frequency of about 50 Hz to about 60 Hz. It can also be a DC source, such as USB or 12 Volt cigarette lighter adapter. Module 802 is capable of charging a battery of at least about 800 kJ within about 2 hours (average 100 Watts, peak 200 Watts). [00151] Module 804 comprises a housing 811 and a power source 812, such as a rechargeable battery, within housing 804 for providing power to module 806. In an exemplary embodiment, the rechargeable battery provides usable power of at least about 500 kJ to about 1,000 kJ, or about 750 kJ to about 850k J, or about 800 kJ and is configured to inhibit or prevent deep discharge of the battery within to prolong its life. Module 804 may also include a suitable coupler for removably coupling the power source to module 802. Alternatively, the power source may be situated remotely from module 802 and may be coupled to module 802 via a direct wired or a wireless connection.
[00152] Module 804 may contain suitable electronics 813 for controlling the parameters of the power delivered to module 806. Module 804 may also include an electrical connection 815 to a computer or other processing device 814. Processing device 814 may provide a variety of software programs and memory for transferring data to and from system 800. Module 804 may be coupled to processing device 814 via wired connections, such as Ethernet, USB, RS-232 or the like, or wirelessly, such as WIFI, Bluetooth or the like.
[00153] Module 806 comprises a transmitting antenna 807 and the associated electronics for transferring energy or power from the antenna to internal modules 808, 810. Antenna 807 includes a magnetic coil that preferably has a diameter of about 20 cm or less and may be flexible or rigid. Module 806 may further include one or more magnets 809 to assist with the positioning of a transmitting antenna 807 relative to an implanted antenna 813 within Module 808. Module 806 may also be configured to provide visual or other feedback related to the alignment of coils within antennas 807, 813 within module 806 and module 808. In an exemplary embodiment, module 806 is configured to perform all of its functions with a maximum power consumption of about 30 Watts. Module 806 is further configured to transfer power in the range of about 20 to about 28 Watts, preferably about 25 Watts, to module 808.
[00154] In an exemplary embodiment, module 808 comprises a housing 811 that is implanted within the patient, e.g., subcutaneously. Housing 811 preferably comprises a watertight material, such as ceramic or the like, that hermetically seals housing 811 to ensure that the components within remain insulated from bodily fluids. Module 808 comprises a receiving antenna 813 and power recover circuitry, e.g., rectifiers or the like. Receiving antenna 813 includes a magnetic coil with a diameter that is preferably about 7 to 10 cm and may be flexible or rigid. Module 808 will also include one or more magnets 815 that cooperate with the magnet(s) 809 in module 806 to improve the relative positioning of module 806 and ensure that a constant level of power is delivered to the implant. Module 808 recovers RF power and is preferably configured to produce at least about 15 Watts of power to the implanted medical device 801. Module 808 operates with 24 Volts DC and may transmit data using in band or out of band modulation.
[00155] In one embodiment, RF electronics are disposed within module 806 and configured to provide a DC connection between modules 806 and 808. In another embodiment, RF circuitry is housed within module 806 and configured to provide an RF connection with module 808. In yet another embodiment, modules 806 and 808 are combined into a single housing to product a more compact design. In this embodiment, modules 806, 808 may be located externally of the patient, or implanted within the patient subcutaneously or in a location closer to module 810.
[00156] Module 810 comprises an implantable housing 820 that includes a controller 822 for controlling a motor within the implanted medical device 801, which may, for example, comprise an implantable pump for assisting with cardiac function as described above. Housing 820 preferably has a volume of 40 cc or less. Module 810 may further include a rechargeable battery 832 for providing short-term power to the pump within device 830 when, for example, external power is not available. In an exemplary embodiment, battery 832 has at least about 20 kJ power. System 800 may further include a wired or wireless connection between module 808 and module 810 for transferring power and/or data therebetween. In an exemplary embodiment, this connection is permanent. In certain embodiments, system 800 includes another wired or wireless connection between module 810 and device 801 to transfer power and data between module 810 and the motor 824. In an exemplary embodiment, this electrical connection is industry standard IS-1.
[00157] In certain embodiments, the transmitting and resonating coils are step up/step down resonant coil pairs. Preferably, each of the transmitting and receiving resonators comprise at least two coils.
[00158] Generally speaking, the resonators will comprise a housing having an outer surface and an interior. The magnetic coil is located on the outer surface and the electronics to drive the coil in the interior of the housing. The housing may be any suitable shape. In one embodiment, the housing for the transmitting coil generally has a circular shape with a domed outer surface, the housing for the receiving coil may be the same shape as the transmitting coil, or a different shape. In one embodiment, the receiving coil housing is substantially rectangular. Preferred coil designs include substantially spiral windings that extend around the outer surface from a starting position near the outer edge of the outer surface to an end position near the center of the outer surface. In preferred embodiments, the spiral winding will extend around a central portion of the outer surface (i.e., the winding will leave a substantially open area in the center of the outer surface of the housing). In certain embodiments, the spiral windings will each have between about 2 to about 10 turns, preferable between about 4 to about 7 turns.
[00159] The transmitting coil(s) will typically be placed a distance of about 1.0 cm to about 6.0 cm, preferably about 2.0 cm, from the receiving coil(s). This distance will be occupied by the subcutaneous fat and skin tissue in the patient’s outer skin surface, garment of the patient, fabric used for the construction of the vest worn by the subject as well as the material used for the encapsulation of the coils. During operation, the coils may be displaced from each other (i.e., shifted horizontally) and/or distanced from each other (i.e., moved toward and away vertically). This displacement may occur inadvertently (i.e., suboptimal placement of the coil, or movement of the patient after the transmitting coil has been), or it may be controlled by the system to maintain power levels, power efficiency and other parameters of the system. As discussed previously, the system ensures that the overall power efficiency and the power transmitted by the receiving resonator to the implant remains substantially constant despite such displacements between the receiving and transmitting coils, which could be accomplished by increasing the transmitter power or adjusting the coil and/or circuit parameters dynamically. A more complete description of suitable antennas for use herein is described in U.S. NonProvisional Patent Application Serial No. 18/177,290, filed March 2, 2023, the complete disclosure of which has been previously incorporated herein by reference.
EXAMPLES
[00160] Applicant tested the pumps shown in FIGS. 6 and 7. The housing diameter of the pumps are 10mm OD and, accounting for a nominal wall thickness, an ID of 9.2mm. The length of each pump was about 70 mm. To account for manufacturing tolerance and to minimize wall shear stress, a radial gap of 0.15mm was selected giving an impeller tip diameter of 8.9 mm. Three impeller blades with a nominal thickness of 0.5mm were chosen for rotor stability. To avoid harmonic resonance and to balance the cross-sectional area with flow separation, the stator was designed with five blades. To pre-emptively manage thermal concerns produced by the DC motor, an active cooling system was needed. This design incorporates non contacting hydrodynamic bearings in an axial configuration to aid in pumping fluid through the motor and provides flow over our designed operating conditions while to cooling the motor. As the development of the blood pump continued, a second design was developed to improve the efficiency of the initial design (see FIG. 7). [00161] A test apparatus was designed, built, and fabricated to facilitate the rapid evaluation of the different impeller/stator configurations. The test apparatus was built out of electropolished stainless steel, and a cavity was machined in the apparatus to accommodate the impeller and stator. A shaft was attached to the impeller and motor which was controlled by a 20V motor controller with a potentiometer to control motor speed. Tubing connectors were created for the inflow and outflow to test the impeller prototypes in mock loops. To ensure that the flow paths to the leading edge of the impeller was unimpeded and resulted in clean evaluation of the impeller and stator performance the test figure was CFD modeled. Two different impellers and stators pairs were fabricated from titanium and electropolished. The two impeller designs were then evaluated in the test apparatus.
[00162] All simulations were completed using Ansys CFX 2021 R2 system. A finite volume methods solver that specializes in analyzing rotor dynamics. All simulations were completed with the following boundary conditions: 25,000 rpms, assuming a head pressure rise of lOOmmHg. Output parameters included flow, thrust, torque and efficiency. The pressure selected for the modeling was chosen to account for anatomical pressures and to overcome and tubing losses that would occur during invitro testing. Simulations were also interrogated based on shear stress, uniformity of the flow, and maintaining boundary layers on working surfaces. [00163] Static mock flow loop experiments were performed as described previously (Monreal Inspired, Monreal RTCS) to evaluate hydrodynamic performance of the prototype. CoRISMA pumps over a range of test conditions to produce H-Q curves by measuring flow (Q) as a function of head pressure (H) over a wide range of afterloads (0-300 mmHg) and flows (0-5 L/min).
[00164] Four different prototype CoRISMA pump configurations were tested: 1) Housing #2 with impeller 1 (“2mvl”); 2) Housing #3 with impeller 2 (“3mv2”); 3) Housing #2 with impeller 2 (“2mv2”); 4) Housing #3 with impeller 1 (“3mvl”).
[00165] Each pump configuration was integrated into the static mock loop circuit (3/8” Tygon 3350) which contained lOOOmL of 3.5cP glycerol-distilled water at 37°C (viscosity validated using Cannon-Fenske calibrated viscometers (ACE Glass, Vineland, NJ) following ASTM Standard D445-21E2). The mock loop reservoir was placed in a 37°C water bath. A Hoffman clamp (restrictor) was used to increase the outflow resistance increasing pressure and decreasing flow. Two fluid-filled transducers (MLT0380/D, ADInstruments, Colorado Springs, CO) were used to acquire pressure data (pump inlet and outlet). A 9mm flow probe (PXL, Transonic Systems, Ithaca, NY) was placed on the pump outflow tubing. Motor shaft and room temperatures were recorded using temperature probes (MLT422/D, ADInstruments). Calibrated output voltages for pump speed (IV = 10,000 rpm) and current (IV = 1A) were recorded. Hemodynamic data were collected using a PowerLab 16/35 acquisition system (ADInstruments) and recorded using LabChart v.8.1.25 (ADInstruments).
[00166] Once the static mock loop was primed, the pump’s saline infusion system (set to 4mL/hr.) was unclamped and pump speed was set to 15,000 rpm (and the loop to no resistance). Flow was decreased in 0.25 L/min increments by increasing the outflow resistance (Hoffman clamp) and data were collected at each increment. Pump speed was then increased by 2,500 rpm increments to 25,000 rpm, and stepwise data were reacquired at each rpm increment. Pump delta pressure, pump current, motor shaft temperature, and room temperature were documented at each increment. The experimental design was repeated for each pump configuration. Mean values for each parameter at each test condition were calculated and H- Q curves were plotted using Prism v.10.1.1 (323) (GraphPad, San Diego, CA).
[00167] Dynamic mock loop experiments were performed using a pneumatic ventricle to quantify in vitro hemodynamic performance of the prototype pumps. Four different prototype pump configurations (described above) were tested.
[00168] Each pump configuration was integrated into the loop (3/8” and 3/4” Tygon 3350) in a left atrial to aortic circuit (inflow connected to a silicone atrium, outflow to the aortic position). The silicone ventricle (Derby City Supply, Taylorsville KY) was powered by a pneumatic driver (Ventricular Assist Device Pneumatic Drive System, Thoratec, Pleasanton, CA (now Abbott)). The dynamic loop was primed with 3600mL of a 3.5cP glycerol -distilled water solution at 37°C (viscosity validated using Cannon-Fenske calibrated viscometers (ACE Glass, Vineland, NJ) following ASTM Standard D445-21E2)) and the loop reservoir was maintained in a 37°C water bath. A compliance chamber was integrated into the arterial side of the loop. A Hoffman clamp was used to adjust aortic resistance. Five fluid-filled transducers (MLT0380/D, ADInstruments) were used to acquire pressure data from the silicone ventricle, aorta, venous position, pump inlet, and pump outlet. Flow probes (PXL, Transonic Systems) were placed on the proximal aorta (LV flow, 16mm), distal aorta (total flow, 16mm) and pump outflow (9mm). Motor shaft and room temperature were recorded using temperature probes (MLT422/D, ADInstruments). Calibrated output voltages for pump speed (1 V = 10,000 rpm) and current (1 V = 1 A) were recorded.
[00169] The mock loop was tuned to a heart rate of 80 bpm, mean aortic pressure of 50 ± 10 mmHg, mean aortic flow of 3 ± 0.5 L/min, and venous pressure of 15 ± 5 mmHg. The pump’s saline infusion system (set to 4 mL/hr) was unclamped. Baseline data (pressures, flows, current, temperatures) was collected with the pump at 15,000 rpm and the dynamic mock ventricle off. The dynamic ventricle was then turned on and data reacquired at 15,000 rpm to 25,000 rpm in 2,500 rpm increments. Data were collected using a PowerLab 16/35 acquisition system, recorded using LabChart (AD Instruments), and plotted using Prism (GraphPad).
[00170] At a maximum pressure of lOOmmHg the REV 1 impeller system (see FIG. 6), the CFD simulation revealed a pump flow of 4.56 L/min. In the modified REV 2 design (see FIG. 7), the pump performance improved to flow to 4.82 L/min under the same conditions. REV. 2 of the design showed less torque required to generate a 6% increase in fluidic flow. REV. 2 also shows that the stator had a larger area of high pressure indicating a decrease in friction suggesting increased efficiency. Additionally, the pressure contours on the REV. 2 impeller system appeared broader indicating a more consistent amount of work done by the diffuser fins with respect to radial and axial engagement relative to REV. 1. TABLE 1 shows the Pump performance 25K RPM @ lOOmmHg delta pressure.
TABLE 1
[00171] Experiments were performed to quantify in vitro hydrodynamic performance (H-Q curves) of the prototype pump in four different configurations of housing + impeller combinations. As shown in FIG. 16, H-Q curves generated were comparable across all four housing + impeller configurations tested. At 25,000 rpm and no resistance, configuration 2mvl achieved a maximum flow of 5.8 L/min, configuration 3mv2 achieved a maximum flow of 5.0 L/min, configuration 2mv2 achieved a maximum flow of 4.8 L/min, and configuration 3mvl achieved a maximum flow of 5.5 L/min. Pump motor shaft temperature reached a high of 50.9°C in the 3mvl housing + impeller configuration at 25,000 rpm (mean room temperature 25.1°C). Air bubbles were observed exiting from the outlet of pump configurations 3mv2 and 2mv2. The original motor of pump configuration 3mvl failed during static mock loop testing at the 17,500 rpm test point and was replaced by a new motor. FIG. 16 shows the relationship of head pressure (H) to flow (Q) for the four different prototype pump housing + impeller configurations tested (2mvl, 3mv2, 2mv2, and 3mvl) in a static mock flow loop model (3.5cP viscosity).
[00172] Experiments were performed to quantify in vitro hemodynamic performance of the prototype pump in four different configurations of housing + impeller combinations. As shown in FIGS. 15A and 15B, hemodynamic performances were comparable across all four pump configurations tested. At 25,000 rpm, configuration 2mvl achieved a maximum flow of 4.5 L/min, configuration 3mv2 achieved a maximum flow of 4.2 L/min, configuration 2mv2 achieved a maximum flow of 4.0 L/min, and configuration 3mvl achieved a maximum flow of 4.5 L/min. Pump motor shaft temperature reached a high of 44.3°C in the 2mv2 housing + impeller configuration at 25,000 rpm (mean room temperature 22.5°C). Pump input power averaged 14.7 watts at baseline (15,000 rpm) and 29.9 watts at 25,000 rpm across all four pump configurations. Air bubbles were observed exiting from the outlet of pump configurations 3mv2 and 2mv2; this may have resulted in an aberrant flow probe readings observed at the 22,500 and 25,000 rpm test points, affecting the total flow values measured for 2mv2. The replacement motor used during the static testing of 3mvl failed at startup during the dynamic testing. The failed motor was removed with glycerol solution in the motor housing observed. The motor from pump configuration 2mv2 was used as the replacement motor for 3mvl to complete dynamic testing.
[00173] FIGS. 15A and 15B illustrate pump flow and pressure data, respectively, from the four different prototype pump housing + impeller configurations tested (2mvl, 3mv2, 2mv2, and 3mvl) in the dynamic mock loop. BL, baseline; LV, left ventricle (pneumatic); AP, delta pressure across the pump; MAP, mean aortic pressure (dynamic loop); LAP, left atrial pressure (dynamic loop). Note in the flow data (left) for pump configuration 2mv2 (green), air was observed exiting the pump outflow at pump speeds of 22,500 and 25,000 rpm, which may have contributed to aberrant measurement in the flow probe readings, affecting the total flow (= aorta + pump) values at these pump speeds.
[00174] Persons skilled in the art will understand that the devices and methods specifically described herein and illustrated in the accompanying drawings are non-limiting exemplary embodiments. The features illustrated or described in connection with one exemplary embodiment may be combined with the features of other embodiments. Various alternatives and modifications can be devised by those skilled in the art without departing from the disclosure. Accordingly, the present disclosure is intended to embrace all such alternatives, modifications, and variances. As well, one skilled in the art will appreciate further features and advantages of the present disclosure based on the above-described embodiments. Accordingly, the present disclosure is not to be limited by what has been particularly shown and described, except as indicated by the appended claims.
[00175] For example, in one aspect, a first embodiment is a system for supporting cardiac function in a patient. The system comprises an elongate housing configured for implantation into the patient exterior to the pericardial cavity, the housing having an inlet and an outlet spaced longitudinally from the inlet, the inlet and the outlet defining a primary blood flow path from a left atrium through at least a portion of the housing to an aorta; a motor coupled to the housing; and an impeller coupled to the motor and configured to propel blood through the primary blood flow path.
[00176] A second embodiment is the first embodiment wherein the housing is configured for implantation exterior to the thoracic cavity.
[00177] A third embodiment is any combination of above embodiments, wherein the housing is configured for implantation in an intercostal space.
[00178] A 4th embodiment is any combination of the above embodiments, further comprising an anchor for attaching the housing to one or more rib bones within the patient.
[00179] A 5th embodiment is any combination of the above embodiments, further comprising an introducer for creating a space in an intercostal muscle.
[00180] A 6th embodiment is any combination of the above embodiments, further comprising an anchor for attaching the housing to one or an intercostal muscle of the patient.
[00181] A 7th embodiment is any combination of the above embodiments, further comprising: a tube fluidly coupling the outlet with the aorta; and an anchor coupled to the tube and configured for securing the tube across a wall of the aorta.
[00182] An 8th embodiment is any combination of the above embodiments, wherein the wall is within a superior vena cava.
[00183] A 9th embodiment is any combination of the above embodiments, further comprising: a tube fluidly coupling the inlet with the left atrium; and an anchor coupled to the tube and configured for securing the tube across a wall between the right atrium and the left atrium.
[00184] A 10th embodiment is any combination of the above embodiments, wherein the wall is the atrial septum. [00185] An 11th embodiment is any combination of the above embodiments, further comprising: a transmitting resonator comprising at least one magnetic coil and configured to transmit power through an outer skin surface of the patient; and a receiving resonator configured for implantation within the patient, the receiving resonator comprising at least one magnetic coil.
[00186] A 12th embodiment is any combination of the above embodiments, further comprising a controller coupled to the transmitting resonator and configured to control the transmitting and receiving resonators such that a level of power remains at or above a threshold level.
[00187] A 13th embodiment is any combination of the above embodiments, wherein the receiving resonator is disposed within the housing.
[00188] A 14th embodiment is any combination of the above embodiments, further comprising a wearable device configured to be attached to, or worn by, the patient, wherein the transmitting resonator is housed within the wearable device.
[00189] A 15th embodiment is any combination of the above embodiments, wherein the motor comprises a stator coupled to the housing and a rotor disposed within the housing, wherein the impeller is coupled to the rotor.
[00190] A 16th embodiment is any combination of the above embodiments further comprising a second inlet within the housing defining a secondary blood flow path through at least a portion of the housing.
[00191] A 17th embodiment is any combination of the above embodiments, wherein the rotor comprises an external surface and one or more rotational elements extending from the external surface for drawing the blood through the secondary flow path.
[00192] An 18th embodiment is any combination of the above embodiments, wherein the rotor and the impeller are spaced from the internal surface of the housing and at least partially supported with hydrodynamic forces within the housing.
[00193] A 19th embodiment is any combination of the above embodiments, further comprising a fluid bearing suspending the impeller and the rotor within the housing.
[00194] A 20th embodiment is any combination of the above embodiments, wherein the fluid bearing comprises one or more surfaces that create a fluid pressure that resists axial forces generated by the impeller or the rotor. [00195] A 21st embodiment is any combination of the above embodiments, wherein the fluid bearing comprises one or more surfaces that create a fluid pressure that resists radial forces generated by the impeller or the rotor.
[00196] A 22nd embodiment is a system for supporting cardiac function in a patient, the system comprising: an elongate housing configured for implantation into the patient and having an inlet and an outlet spaced longitudinally from the inlet, the inlet and the outlet defining a primary blood flow path from a left atrium through at least a portion of the housing to an aorta; a motor coupled to the housing; an impeller coupled to the motor and configured to propel blood through the primary blood flow path; and a rechargeable power source configured for subcutaneous implantation into the patient and coupled to the housing.
[00197] A 23rd is embodiment is the 22nd embodiment and any combination of the above embodiments.
[00198] A 24th embodiment is any combination of the above embodiments, wherein the power source comprises a battery.
[00199] A 25th embodiment is any combination of the above embodiments, further comprising a power transmitter configured to transmit power through an outer skin surface of the patient to the rechargeable power source.
[00200] A 26th embodiment is any combination of the above embodiments, wherein the power source comprises a receiver for receiving the power transmitted by the power transmitter.
[00201] A 27th embodiment is any combination of the above embodiments, wherein the power source is wirelessly coupled to the housing.
[00202] A 28th embodiment is any combination of the above embodiments, further comprising an electrical connector coupling the power source to the housing.
[00203] A 29th embodiment is any combination of the above embodiments, further comprising: a transmitting resonator comprising at least one magnetic coil and configured to transmit a first level of power through an outer skin surface of the patient; and a receiving resonator configured for implantation within the patient, the receiving resonator comprising at least one magnetic coil.
[00204] A 30th embodiment is any combination of the above embodiments, further comprising a controller coupled to the receiving resonator and configured to control the rechargeable power source to delivery power to the housing when the first level of power falls below a threshold level. [00205] A 31st embodiment is any combination of the above embodiments, wherein the housing is configured for implantation exterior to the thoracic cavity.
[00206] A 32nd embodiment is any combination of the above embodiments, wherein the housing is configured for implantation in an intercostal space.
[00207] A 33rd embodiment is any combination of the above embodiments, further comprising an anchor for attaching the housing to one or more rib bones within the patient.
[00208] A 34th embodiment is any combination of the above embodiments, further comprising: a tube fluidly coupling the outlet with the aorta; and an anchor coupled to the tube and configured for securing the tube across a wall of the aorta.
[00209] A 35th embodiment is any combination of the above embodiments, further comprising: a tube fluidly coupling the inlet with the left atrium; and an anchor coupled to the tube and configured for securing the tube across a wall between the right atrium and the left atrium.
[00210] A 36th embodiment is a method for supporting cardiac function in a patient, the method comprising: implanting a pump at a target site within the patient exterior to the thoracic cavity; generating a fluid path from a left atrium to the pump; generating a fluid path from the pump to an aorta; and drawing blood from the left atrium into an inlet of the pump through a flow path such that the blood flows through an outlet of the pump and into the aorta.
[00211] A 37th embodiment is the 36th embodiment and any combination of the above embodiments.
[00212] A 38th embodiment is any combination of the above embodiments, further comprising implanting the pump within an intercostal space in the patient.
[00213] A 39th embodiment is any combination of the above embodiments, further comprising anchoring the pump within the intercostal space.
[00214] A 40th embodiment is any combination of the above embodiments, further comprising anchoring a tube through a wall between the left atrium and the right atrium and fluidly coupling the tube to the inlet of the pump.
[00215] A 41st embodiment is any combination of the above embodiments, further comprising implanting an anchor within the wall and securing the tube to the anchor.
[00216] A 42nd embodiment is any combination of the above embodiments, wherein the wall is an atrial septum. [00217] A 43rd embodiment is any combination of the above embodiments, further comprising anchoring a tube through a wall of an aorta and fluidly coupling the tube to the outlet of the pump.
[00218] A 44th embodiment is any combination of the above embodiments, further comprising implanting an anchor within the wall and securing the tube to the anchor.
[00219] A 45th embodiment is any combination of the above embodiments, further comprising implanting a battery within the patient and coupling the power source to the pump. [00220] A 46th embodiment is any combination of the above embodiments, further comprising implanting the battery subcutaneously within the patient.
[00221] A 47th embodiment is any combination of the above embodiments, further comprising recharging the battery by transmitting power to the power source through an outer skin surface of the patient.
[00222] A 48th embodiment is any combination of the above embodiments, further comprising wirelessly transmitting a level of power to the pump from a power source located external to the patient.
[00223] A 49th embodiment is any combination of the above embodiments, further comprising transmitting power from the battery to the pump when the level of power falls below a threshold level.
[00224] A 50th embodiment is any combination of the above embodiments, further comprising rotating an impeller within the pump to draw the blood through the primary blood flow path.
[00225] A 51st embodiment is any combination of the above embodiments, further comprising rotating a rotor coupled to the impeller to draw the blood through the secondary blood flow path.