System and method for controlling an implantable blood pump

Abstract

The present invention provides systems and methods for controlling an implantable pump. For example, an exemplary controller for controlling the implantable pump may rely solely on current measurements of the actuator. The controller is robust to pressure and flow changes inside the pump head and allows rapid changes in the operating point of the pump. For example, the controller includes a two-stage nonlinear position observer module based on a reduced-order model of an electromagnetic actuator. The controller includes an algorithm that estimates the position of a moving component of the implantable pump based on current measurements of the actuator and adjusts the operation of the pump accordingly. Alternatively, the controller may rely on position measurements and / or velocity estimates.

Description

[0001] CROSS REFERENCE TO RELATED APPLICATIONS

[0002] This application claims the benefit of priority to U.S. Provisional Patent Application No. 62 / 819,436, filed on March 15, 2019, the entire contents of which are incorporated herein by reference. Technical Field

[0003] The present invention relates generally to implantable cardiac pumps with an undulating membrane and improved hydraulic performance designed to reduce hemolysis and platelet activation, and more particularly to controlling implantable pumps. Background Art

[0004] The human heart consists of four main chambers with two ventricles and two atria. Generally speaking, the right side of the heart receives deoxygenated blood from the body into the right atrium and pumps it to the lungs via the right ventricle. The left side of the heart receives oxygen-rich blood from the lungs into the left atrium and pumps it to the aorta via the left ventricle for distribution throughout the body. The left ventricle may become less efficient and therefore unable to pump oxygenated blood to the body due to any of several diseases, including coronary artery disease, high blood pressure (hypertension), valvular regurgitation and calcification, myocardial damage due to infarction or ischemia, myocarditis, congenital heart defects, arrhythmias, or various infectious diseases.

[0005] The Centers for Disease Control and Prevention (CDC) estimates that approximately 5.1 million people in the United States suffer from some form of heart failure. Heart failure is generally divided into four different stages, the most severe of which is end-stage heart failure. End-stage heart failure is diagnosed when a patient continues to have symptoms of heart failure while at rest despite medical treatment. Patients at this stage may have systolic heart failure, which is characterized by a decreased ejection fraction. In patients with systolic heart failure, the walls of the ventricles, which are normally thick in healthy patients, become thin and fragile. As a result, the volume of oxygenated blood ejected into the circulation during systole is reduced, and this continues a downward spiral until death. Patients diagnosed with heart failure have a one-year mortality rate of approximately 50%.

[0006] For patients who have reached end-stage heart failure, treatment options are limited. In addition to continuing medical therapy typically prescribed during early heart failure, heart transplantation and implantation of a mechanical assist device are typically recommended. While heart transplantation can significantly extend a patient's lifespan with a mortality rate of more than one year, patients on the waiting list often die while waiting months and sometimes years for a suitable donor heart. Currently, the only alternative to heart transplantation is a mechanical implant. While the design of mechanical implants has improved in recent years, such implants typically extend a patient's lifespan by a few years at most and contain several recurring lesions.

[0007] One type of mechanical implant commonly used for patients with end-stage heart failure is a left ventricular assist device (LVAD). The LVAD is a surgically implanted pump that draws oxygenated blood from the left ventricle and pumps it directly to the aorta, thereby relieving (reducing) the pumping work of the left ventricle. LVADs are often used as "bridge-to-transplant therapy" or "destination therapy." When used as bridge-to-transplant therapy, an LVAD is used to extend the life of a patient awaiting a heart transplant. When a patient is not a candidate for a heart transplant, an LVAD can be used as a destination therapy to extend the patient's life, or improve quality of life, but this extension is generally only a few years.

[0008] In general, an LVAD comprises an inlet cannula, a pump, and an outlet cannula, and is coupled to an external battery and control unit. The inlet cannula is typically connected directly to the left ventricle, for example, at the apex of the heart, and delivers blood from the left ventricle to the pump. An outlet cannula, typically connected to the aorta distal to the aortic valve, delivers blood from the pump to the aorta. Typically, the outlet cannula of the pump is extended using a hose-type structure (e.g., a Dacron graft) to reach the appropriate delivery location on the aorta. Early LVAD designs were reciprocating types, but more recently rotary pumps and centrifugal pumps have been used.

[0009] Isaacson's U.S. Patent No. 4,277,706, entitled "Actuator for a Heart Pump," describes an LVAD with a reciprocating pump. The pump described in the Isaacson patent includes a housing having an inlet and an outlet; a chamber inside the pump connected to the inlet and the outlet; a flexible diaphragm extending across the chamber; a plate secured to the diaphragm; and a ball screw configured to reciprocate to drive the plate and the connected diaphragm from one end of the chamber to the other end to simulate contraction and relaxation. The ball screw is actuated by a DC motor. The Isaacson patent also describes a controller configured to manage the rotation of the ball screw to control starting, stopping, and reversing of direction to control blood flow into and out of the pump.

[0010] Previously known reciprocating pump LVADs have several disadvantages. Such pumps are often bulky, heavy, and may require the removal of bone and tissue in the chest cavity for implantation. They also require a large amount of energy to displace the blood by compressing the cavity. In addition, the pump subjects the blood to significant pressure fluctuations as it passes through the pump, resulting in high shear forces and risk of hemolysis. These pressure fluctuations may be amplified at higher blood flow rates. In addition, depending on the geometry of the pump, areas with little or no flow can cause flow stagnation, which can lead to thrombosis and potentially fatal medical conditions, such as stroke. Ultimately, a positive displacement pump like the pump described in the Exaxon patent is not able to achieve pulsatility similar to that of a natural heart, e.g., about 60 to 100 beats per minute, while maintaining a physiological pressure gradient.

[0011] LVADs utilizing rotational and centrifugal configurations are also known. For example, U.S. Pat. No. 3,608,088 to Reich, entitled "Implantable Blood Pump," describes a centrifugal pump for assisting a failing heart. The Reich patent describes a centrifugal pump having: an inlet connected to a rigid cannula coupled to the left ventricular cavity; and a Dacron graft extending from the pump diffuser to the aorta. The pump includes an impeller that rotates at high speed to accelerate the blood, and the pulsation of a natural heart is simulated by varying the rotational speed or introducing a fluid oscillator.

[0012] U.S. Patent No. 5,370,509 to Golding, entitled "Sealless Rotary Dynamic Pump with Fluid Bearings," describes an axial blood pump capable of being used as a heart pump. One embodiment described relates to an axial blood pump having impeller blades aligned with the axis of a blood inlet and a blood outlet. U.S. Patent No. 5,588,812 to Taylor, entitled "Implantable Electric Axial Blood Pump," describes an axial blood pump similar to the pump of the Golding patent. The pump described in the Taylor patent has a pump housing defining a cylindrical blood conduit through which blood is pumped from the inlet to the outlet, and rotor blades that rotate along the axis of the pump to accelerate blood flowing through the blood conduit.

[0013] Although improvements have been made to previously known LVAD devices, those pump designs are not without problems. Like reciprocating pumps, rotary and centrifugal pumps are often bulky and difficult to implant. Although rotary pumps are mechanically different from positive displacement pumps, they also exhibit undesirable characteristics. Like positive displacement pumps, rotary pumps apply significant shear forces to the blood, thereby creating the risk of hemolysis and platelet activation. The nature of the rotation of the disk or blades around the axis creates high-speed and low-speed areas as well as vibration and heating. Specifically, the area near the edge of the disk or blade farthest from the axis of rotation experiences higher angular velocities and therefore higher flow rates compared to the area closest to the axis of rotation. The radial velocity distribution generated along the rotating blades causes high shear forces to be applied to the blood. In addition, stagnation or low flow rates near the axis of rotation may lead to thrombosis.

[0014] Although centrifugal pumps may be able to produce pulsating flow by varying the rotational speed of the associated disks or blades, this only exacerbates the problems created by the steep radial velocity profile and high shear forces. Conventionally, the output of currently available rotary pumps, measured as flow rate and given head pressure, is controlled by varying the rotational speed of the pump. Given the mass of the rotating components, the angular velocity of the rotating components, and the resulting inertia, the change in rotational speed cannot be instantaneous, but must be gradual. Therefore, although a centrifugal pump can mimic pulsating flow with gradual speed changes, the pulsation produced is not "on demand" and does not resemble a typical physiological pulsation.

[0015] In addition, rotary pumps often result in non-physiological pressures being applied to the blood. Such high operating pressures have the unwanted effect of causing blood vessels to overstretch, which in the presence of continuous flow may lead to vascular fibrosis and inelasticity. This in turn may lead to a loss of elasticity in the circulatory system, thereby promoting calcification and plaque formation. Further, if the rotational speed of the pump is varied to simulate pulsatile flow or to increase the flow rate, the rotary pump is unlikely to operate at its optimal operating point, thereby reducing efficiency and increasing energy losses and heat generation.

[0016] LVADs can also be configured to increase blood flow to match the patient's needs. Numerous publications and patents describe methods for adjusting LVAD pump flow to match the flow required by the patient. For example, U.S. Patent No. 7,520,850 to Blockway, entitled "Feedback Control and Ventricular Assist Device," describes a system and method for controlling a ventricular assist device using pressure feedback. The system described in the Blockway patent attempts to maintain a constant filling of the ventricle by measuring ventricular pressure and / or ventricular volume. Although such systems can achieve flow rates of up to 8 or 9 liters / minute, these flow rates are generally beyond the efficient operating range of current rotary pumps, which are typically tuned to operate in the range of 4 to 6 liters / minute. Therefore, increasing the flow rate in the rotary pump to match the patient's needs will produce non-optimal pump performance.

[0017] Pumps other than rotary and positive displacement types are known in the art for displacing fluids. For example, U.S. Pat. Nos. 6,361,284 and 6,659,740 to Delevingne, entitled "Vibrating Membrane Fluid Circulator," both describe pumps in which a deformable membrane is vibrated to propel fluid through a pump housing. In these patents, the vibratory motion applied to the deformable membrane induces wave-like undulations in the membrane that propel the fluid along a channel. Different flow rates can be achieved by controlling the excitation applied to the membrane.

[0018] U.S. Patent No. 7,323,961 to Delevingne, entitled "Electromagnetic Machine with Deformable Membrane," describes a device in which a membrane is coupled in tension along its outer edge to an electromagnetic device arranged to rotate about the membrane. As the electromagnetic device rotates, the outer edge of the membrane deflects slightly in a direction normal to the plane of the membrane. These deflections induce wave-like undulations in the membrane that can be used to move a fluid in contact with the membrane.

[0019] Delevé's U.S. Patent No. 9,080,564, entitled "Diaphragm Circulator," describes a tensile deformable membrane in which undulation is formed by electromechanically moving a magnetized ring attached to the outer edge of the deformable membrane over a coil. Axial displacement of the magnetized ring causes the membrane to undulate. As in the '961 patent, the membrane undulation can be controlled by manipulating the magnetic attraction. Delevé's U.S. Patent No. 8,714,944, entitled "Diaphragm Pump with Corrugated Diaphragm of Improved Efficiency," and Delevé's U.S. Patent No. 8,834,136, entitled "Corrugated Diaphragm Pump," teach similar types of vibrating membrane pumps.

[0020] None of Delevingne's aforementioned patents describe a vibrating membrane pump suitable for use in a biological environment or capable of pumping blood for extended periods that present a low risk of flow stasis leading to thrombosis.

[0021] U.S. Patent Publication Nos. 2017 / 0290966 and 2017 / 0290967 to Barthes (each of which is incorporated herein by reference in its entirety) describe an implantable cardiovascular blood pump having a flexible membrane coupled to an electromagnetic actuator assembly that causes wave-like undulations to propagate along the flexible membrane to propel blood through the pump while avoiding thrombosis, hemolysis, and / or platelet activation. The Barthes pump generates hydraulic power (flow and pressure) by translating the linear motion of the electromagnetic actuator to the flexible membrane, which deforms through its interaction with the blood, thereby translating energy to the blood. The flexible membrane is oriented at a 90° angle to the motion of the linear actuator, so that the outer edge of the membrane is the first element to contact the blood. Therefore, there is a risk of energy loss at the entrance of the membrane, which affects the hydraulic power generation of the pump.

[0022] There is a need for a high performance implantable pump having light weight, small size, and fast start and stop response, which can operate efficiently and with improved hydraulic performance and minimal blood trauma over a wide range of flow rates.

[0023] The design of this high-performance implantable pump that meets all the requirements mentioned above poses many challenges in terms of mechanical design and manufacturing processes. It is also a challenge from a control perspective because, unlike a rotary pump, the operating point of a vibrating diaphragm pump is set by the frequency and amplitude of the film excitation. In fact, the higher the stroke of the frequency or fluctuation, the higher the pressure head of the implantable pump will be. The stroke needs to be accurately set at a sufficient speed to be able to switch the operating point of the pump quickly enough to re-form a sufficient pulse synchronized with the heartbeat. At the same time, the stroke must be limited to avoid damaging the film, blood or the internal spring assembly of the pump due to excessive pressure. This phenomenon may be caused by providing too much power to the actuator or by the disturbing force effect caused by the remaining activity of the left ventricle. Due to the specific medium (blood) in which the pump operates, it may be better to avoid adding position, velocity or acceleration sensors, which will significantly increase the complexity and size of the pump, which is already difficult to design.

[0024] Attempts to bypass the use of motion sensors include measuring the current ripple generated by the pulse width modulated (PWM) voltage input to estimate the equivalent circuit inductance related to the magnet position. (For example, see MF Rahman et al. Position Estimation in Solenoid Actuators, IEEE Transactions on Industrial Applications, Vol. 32, No. 3, pp. 552-559, June 1996). This method is only effective when the speed of the magnetic component is close to zero, which is not the case for vibrating membrane pumps operating at frequencies close to 100 Hz. Other methods calculate the back electromotive force (back EMF proportional to the speed) from the reverse equivalent circuit and directly integrate the estimated speed to obtain the position. (For example, see J. Zhang et al.'s study on self-sensors of piston stroke of linear moving magnet compressors, IEEE Sensor Journal, Vol. 9, No. 2, pp. 154-158, February 2009). This last method only requires knowledge of electrical parameters and does not require information about the mechanical subsystem of the actuator. However, the coil current derivative must be calculated, which is important in noisy environments.

[0025] For example, one approach proposes to estimate a velocity observer that does not rely on calculating the back EMF at any time. (See, for example, J. Latham et al., Parameter Estimation and a Series of Nonlinear Observers for System Dynamics of a Linear Steam Compressor, IEEE Transactions on Industrial Electronics, Vol. 63, No. 11, pp. 6736-6744, November 2016). The position derived from integrating the estimated velocity is sensitive to measurement biases that propagate into the velocity estimate, which causes drift when integrating. This effect can be limited by adding another stage to the observer. (See, for example, P. Mercorelli, Sensorless Control of Intake Valves in Internal Combustion Engines, IEEE Transactions on Industrial Electronics, Vol. 64, No. 4, pp. 3402-3412, April 2017). This additional stage adds partial knowledge about the mechanical subsystem of the actuator and is robust to unknown finite forces. However, these studies are limited to the linear domain of the actuator, where the parameters of the actuator's equivalent circuit can be approximated as constants, which is not valid for vibrating membrane pumps where the actuator is made as small as possible.

[0026] In view of the foregoing, there is a need for controlling a high performance implantable pump having light weight, small size, and fast start and stop response without relying on position, velocity, or acceleration sensors, for example.

[0027] It would further be desirable to provide an improved controller for controlling a high performance implantable pump that relies on position measurements. Summary of the invention

[0028] The present invention overcomes the shortcomings of previously known LVAD systems and methods by providing an implantable pump system having an undulating membrane capable of producing a wide range of physiological flow rates when low shear forces are applied to the blood, thereby reducing hemolysis and platelet activation relative to previously known systems.

[0029] According to one aspect of the present invention, the implantable blood pump system includes an implantable blood pump configured to be implanted at a patient's heart, and a controller operatively coupled to the implantable blood pump. The implantable blood pump includes: a housing having an inlet and an outlet; a deformable membrane disposed within the housing; and an actuator having a fixed component and a moving component coupled to the deformable membrane. The actuator is powered by an alternating current, which causes the moving component to reciprocate relative to the fixed component at a predetermined frequency and amplitude, thereby causing the deformable membrane to generate a predetermined blood flow that exits from the inlet through the outlet.

[0030] In addition, the controller is programmed to operate the actuator to cause the moving component to reciprocate at the predetermined frequency and amplitude relative to the fixed component, receive a signal indicative of the alternating current via a current sensor operatively coupled to the controller, determine the position of the moving component based on the signal indicative of the alternating current, and adjust the operation of the actuator based on the position of the moving component to cause the moving component to reciprocate at the adjusted predetermined frequency and amplitude relative to the fixed component, thereby causing the deformable membrane to generate an adjusted predetermined blood flow exiting from the inlet through the outlet. For example, the adjusted predetermined blood flow may be a pulse synchronized with the patient's heartbeat.

[0031] The controller may be programmed to determine the position of the moving component based on the signal indicative of the alternating current by estimating the speed of the moving component. For example, the controller may be programmed to estimate the speed of the moving component based on various positions of the moving component and an isotropic energy W value of a finite element model (FEM) of alternating current. Additionally, the controller may be programmed to determine the position of the moving component based on the estimated speed of the moving component by determining the speed of the moving component.

[0032] In addition, the controller can be programmed to adjust the operation of the actuator to cause the moving component to reciprocate at the adjusted predetermined frequency and amplitude relative to the fixed component while limiting overshoot. For example, the controller can include a proportional integral (PI) controller programmed to limit overshoot by canceling errors due to unmodeled dynamics of the implantable blood pump. The controller can be programmed to determine the position of the moving component based on the signal indicative of the alternating current and changes in inductance and back EMF coefficients.

[0033] According to one aspect of the present invention, the fixed component includes an electromagnetic assembly that generates a magnetic field. In addition, the moving component may include a magnetic ring that is concentrically suspended around the electromagnetic assembly and is designed to reciprocate over the electromagnetic assembly at the predetermined frequency and amplitude in response to the magnetic field. The electromagnetic assembly may include first and second electromagnetic coils, such that the magnetic ring is caused to move when at least one of the first or second electromagnetic coils is powered by the alternating current. In addition, the magnetic ring induces wave-like deformation in the deformable film by reciprocating over the electromagnetic assembly.

[0034] In addition, the implantable blood pump may include first and second suspension rings concentrically disposed around and coupled to the fixed component and the mobile component. Thus, the mobile component may be coupled to the deformable membrane and each of the first and second suspension rings via a plurality of rods, such that the first and second suspension rings permit the mobile component to reciprocate relative to the fixed component. The first and second suspension rings may apply a spring force to the mobile component when the mobile component reciprocates relative to the fixed component.

[0035] Additionally, the implantable blood pump may further include a rigid ring coupled to the moving assembly and to the deformable membrane. Furthermore, a bottom surface of the actuator and an interior portion of the housing adjacent to the outlet may form a flow channel within which the deformable membrane is suspended. Thus, the deformable membrane may have a central orifice adjacent to the outlet. Additionally, the actuator and an interior surface of the housing adjacent to the inlet may form a delivery channel extending from the inlet to the flow channel. The implantable blood pump system may further include a rechargeable battery for delivering the alternating current to power the implantable blood pump.

[0036] According to another aspect of the present invention, an alternative exemplary implantable blood pump system is provided. The system may include the implantable blood pump sized and shaped to be implanted at the heart of the patient described above and a controller operatively coupled to the implantable blood pump. For example, the controller may be programmed to: operate the actuator to cause the moving component to reciprocate at the predetermined frequency and amplitude relative to the fixed component; receive a signal indicating the strength of the magnetic field of a magnet coupled to the moving component via a sensor (e.g., a Hall effect sensor) operatively coupled to the controller, the sensor being fixed relative to the fixed component; determine the position of the moving component based on the signal indicating the strength of the magnetic field; and adjust the operation of the actuator based on the position of the moving component to cause the moving component to reciprocate at an adjusted predetermined frequency and amplitude relative to the fixed component, thereby causing the deformable membrane to produce an adjusted predetermined blood flow that exits from the inlet through the outlet.

[0037] For example, the sensor may be coupled to the fixed component or the housing. The controller may be further programmed to estimate blood flow exiting from the inlet through the outlet based on the position of the mobile component. Additionally, the controller may be further programmed to detect a fault by comparing an average residual value based on the position of the mobile component to a predetermined threshold.

[0038] According to another aspect of the present invention, the controller may be programmed to: operate the actuator to cause the moving component to reciprocate at the predetermined frequency and amplitude relative to the fixed component; receive a signal indicating the strength of the magnetic field of a magnet coupled to the moving component via a sensor operatively coupled to the controller, the sensor being fixed relative to the fixed component; estimate the speed of the moving component based on the signal indicating the strength of the magnetic field; and adjust the operation of the actuator based on the speed of the moving component to cause the moving component to reciprocate at an adjusted predetermined frequency and amplitude relative to the fixed component, thereby causing the deformable film to produce an adjusted predetermined blood flow that leaves the inlet and passes through the outlet. BRIEF DESCRIPTION OF THE DRAWINGS

[0039] Figure 1 An exemplary embodiment of a pump system of the present invention comprising an implantable pump, a controller, a battery, a programmer, and a mobile device is depicted.

[0040] Figure 2 yes Figure 1 Perspective view of an implantable pump.

[0041] Figure 3A and 3B 1 and 2 are perspective and schematic diagrams, respectively, of electronic components of an exemplary embodiment of a controller of the present invention.

[0042] Figure 4 is a plan view of an external battery for use in the pump system of the present invention.

[0043] Figure 5A and 5B 1 and 2 are respectively a perspective view and a schematic diagram of electronic components of an exemplary embodiment of a programmer of the present invention.

[0044] Figure 6 is a perspective view of a pump assembly of the present invention.

[0045] Figure 7 is a perspective cross-sectional view of an implantable pump of the present invention.

[0046] Figure 8 is an exploded view of the implantable pump of the present invention.

[0047] Fig. 9 is a perspective cross-sectional view of a pump assembly of the present invention.

[0048] Fig.10 is a perspective cross-sectional view of a film assembly of the present invention.

[0049] Fig.11 is a perspective cross-sectional view of a moving component of a pump assembly according to a first embodiment of the present invention.

[0050] Fig.12 is a cross-sectional view of an implantable pump of the present invention.

[0051] Fig.13 is a cross-sectional view of the lower portion of the implantable pump depicting the flow channel and the membrane assembly in the resting position.

[0052] Fig.14 is a cross-sectional view of the lower portion of an implantable pump depicting the flow channel and the membrane assembly with membrane undulations.

[0053] Fig.15A It is for Figure 1 A cross-sectional view of an alternative exemplary embodiment of an implantable pump of the present invention having improved hydraulic performance for use in a pump system of FIG.

[0054] Fig. 15B yes Fig.15A Perspective view of an implantable pump.

[0055] Fig.16A Diagram illustrating blood flow across a planar annular membrane support, and Fig. 16B Illustrated is blood flow using a pump assembly having a skirt according to one aspect of the present invention.

[0056] Fig.17 A graph illustrating the relationship between maximum hydraulic power and height of the skirt is shown.

[0057] Fig.18 is a cross-sectional view of yet another alternative exemplary embodiment of an implantable pump of the present invention having improved hydraulic performance, wherein the outflow cannula is coaxially disposed within the inflow cannula.

[0058] Fig.19 It is for Figure 1 A cross-sectional view of yet another alternative exemplary embodiment of an implantable pump of the present invention having a ring and a skirt with improved hydraulic performance for use in a pump system of FIG.

[0059] Fig. 20 It is for Figure 1 A cross-sectional view of yet another alternative exemplary embodiment of an implantable pump of the present invention having a ring, a skirt, and an expandable portion with improved hydraulic performance for use in a pump system of FIG.

[0060] Fig.21A H illustrate various configurations for coupling a battery to the controller of the present invention, and Fig.21I The diagram illustrates a controller coupled to a power supply.

[0061] Fig. 22is a flow chart illustrating the steps of an exemplary method for controlling an implantable pump constructed according to the principles of the present invention.

[0062] Fig.23 is a FEM model of a subset of the actuator assembly of an implantable pump constructed in accordance with the principles of the present invention.

[0063] Fig.24A Displays the energy changes with the position of the magnetic ring and the coil current. Fig. 24B The displayed force varies with the position of the magnetic ring and the coil current. Fig.24C The circuit inductance of the equivalent circuit is shown to vary with the position of the magnetic ring and the coil current, and Fig.24D The circuit EMF coefficient of the equivalent circuit is shown as a function of the magnetic ring position and the coil current.

[0064] Fig.25 is a graph illustrating spring reaction force as a function of position of the implantable pump.

[0065] Fig.26 is a schematic diagram of parameters of an equivalent circuit according to the principle of the present invention.

[0066] Fig.27A 1 to 4 are schematic diagrams of power electronic devices constructed according to the principles of the present invention.

[0067] Fig.28 The diagram illustrates ADC sampling for measuring current in accordance with the principles of the present invention.

[0068] Fig.29 is a diagram illustrating multiple levels of control of a controller constructed according to the principles of the present invention.

[0069] Fig. 30A to C respectively illustrate the changes of resistance, inductance and EMF coefficient with the position of the magnetic ring and the coil current according to the principle of the present invention.

[0070] Fig.31A and 31B The system response to the desired stroke is illustrated with respect to coil current and magnetic ring position, respectively.

[0071] Fig.32A and 32B The system response to operating point changes is illustrated with respect to amplitude and frequency, respectively.

[0072] Fig.33A 1 to 2 D illustrate stroke output error maps of a system in accordance with the principles of the present invention.

[0073] Fig.34 is a diagram illustrating multiple levels of control of a position measurement dependent controller constructed in accordance with the principles of the present invention.

[0074] Fig.35 is a diagram illustrating an alternative multi-level sensorless control of a controller constructed in accordance with the principles of the present invention.

[0075] Fig.36 is a diagram illustrating yet another alternative multi-level sensorless control of a controller constructed in accordance with the principles of the present invention. DETAILED DESCRIPTION

[0076] The implantable pump system of the present invention is particularly suitable for use as a left ventricular assist device (LVAD) and includes an undulating membrane pump suitable for long-term implantation in patients with end-stage heart failure. The implantable pump system constructed according to the principles of the present invention includes an implantable pump and an external battery, a controller and a programmer. The implantable pump includes: a housing having an inlet and an outlet; a flexible membrane; and an actuator assembly. When configured as an LVAD, the housing includes an inlet cannula inserted into the patient's left ventricle near the apex of the heart and an outlet cannula surgically placed to communicate with the patient's aortic fluid. By activating the actuator assembly within the implantable pump, the membrane is induced to undulate, thereby causing blood to be drawn into the pump through the inlet cannula and discharged into the aorta through the outlet cannula. The flow rate and pulsatility can be manipulated by changing one or more of the frequency, amplitude and duty cycle of the actuator assembly.

[0077] To improve hydraulic performance, the implantable pump may include a skirt disposed within the housing to direct blood flow from the inlet toward the outlet of the pump. The skirt may be positioned within the housing so that blood flows across opposite sides of the skirt and toward the undulating membrane immediately after activation of the pump.

[0078] Reference now Figure 1 , describes a pump system 10 constructed according to the principles of the present invention. The pump system 10 includes an implantable pump 20, a controller 30, a battery 40, a programmer 50, and a software module optionally programmed to run on a mobile device 60. The implantable pump 20 is configured to be implanted in the chest of a patient such that an inlet cannula 21 is coupled to the left ventricle LV of the heart H. The outlet cannula 22 of the pump 20 is configured to be coupled to the aorta A. The inlet cannula 21 is preferably coupled to the apex of the left ventricle LV, while the outlet cannula 22 is coupled to the aorta above the level of the cardiac arteries near the ascending aorta. The implantable pump 20 can be attached to the chest of the patient using ring suturing or other conventional techniques. The outlet cannula 22, which can include a Dacron graft or other synthetic material, is coupled to an outlet 23 of the implantable pump 20.

[0079] Now also refer to Figure 2In a preferred embodiment, implantable pump 20 is composed of an upper housing portion 24 joined to a lower housing portion 25 along an interface 26, for example, by threads or welding, to form a fluid-tight pump housing 27 that may have a cylindrical shape. Upper housing portion 24 includes an inlet sleeve 21 and an electrical conduit 28 for receiving electrical wires from controller 30 and battery 40. Lower housing portion 25 includes an outlet 23 coupled to an outlet sleeve 22, such as Figure 1 The pump housing 27 is made of a biocompatible material such as stainless steel and is sized to be implanted within the chest cavity of a patient.

[0080] Reference again Figure 1 In one embodiment, the controller 30 and battery 40 are external and sized to be placed on a belt or garment worn by the patient. Both the controller 30 and battery 40 are electrically coupled to the implantable pump 20, for example, via a cable 29 that extends through a percutaneous opening in the patient's skin and into an electrical conduit 28 of a pump housing 27. Illustratively, the battery 40 is electrically coupled to the controller 30 via a cable 41 integrated into a belt 42. In an alternative embodiment, the controller 30 may be enclosed within a biocompatible housing and sized to be implanted subcutaneously in the patient's abdomen. In this alternative embodiment, the controller 30 may include a wireless transceiver for bidirectional communication with an external programming device and also include a battery that is continuously and inductively charged via an external battery 40 and an external charging circuit. As will be appreciated, the aforementioned alternative embodiment avoids the use of a percutaneous cable 29 and thus eliminates a common source of infection for conventional LVAD devices.

[0081] Battery 40 preferably comprises a rechargeable battery capable of powering implantable pump 20 and controller 30 for a period of several days (e.g., 3 to 5 days) before requiring recharging. As is conventional with rechargeable batteries, battery 40 may include a separate charging circuit that is not shown. Battery 40 is preferably disposed within a housing that is suitable for carrying on a belt or holster so as not to interfere with the patient's daily activities.

[0082] Programmer 50 may consist of a conventional laptop computer that is programmed to execute a programmed software routine for use by a clinician or medical professional for configuring operating parameters and providing the operating parameters to controller 30. Configuration and operating parameter data is stored in a memory associated with controller 30 and is used by the controller to control the operation of implantable pump 20. As described in further detail below, controller 30 directs implantable pump 20 to operate with specific parameters determined by programmer 50. Programmer 50 is preferably coupled to controller 30 via cable 51 only when initially setting or periodically adjusting the operating parameters of the implantable pump, for example, when a patient visits a clinician.

[0083] According to another aspect of the present invention, mobile device 60, which may be a conventional smartphone, may include an application for bidirectionally and wirelessly communicating with controller 30, for example, via WiFi or Bluetooth communications. The application on mobile device 60 may be programmed to permit the patient to send instructions to the controller to modify or adjust a limited number of operating parameters of implantable pump 20 stored in controller 30. Alternatively or in addition, mobile device 60 may be programmed to receive from controller 30 and display on screen 61 of mobile device 60 data relating to the operation of implantable pump 20 or alarms or status messages generated by controller 30.

[0084] about Figure 3A and 3B , the controller 30 is described in more detail. Figure 1 , the controller 30 may be sized and configured to be worn on the outside of the patient's body and may be incorporated into clothing such as a belt or vest. The controller 30 includes an input port 31, a battery port 32, an output port 33, an indicator light 34, a display 35, a status light 36, and a button 37.

[0085] Input port 31 is configured to periodically and removably accept cable 51 to establish an electrical connection between programmer 50 and controller 30, for example, via a USB connection. In this manner, a clinician may be in contact with controller 30 to set or adjust operating parameters stored in controller 30 for controlling the operation of the implantable pump. Additionally, when programmer 50 is coupled to controller 30, the clinician may also download data related to the operation of the implantable pump from controller 30, such as actuation statistics for processing and presentation on display 55 of programmer 50. Alternatively or in addition, controller 30 may include a wireless transceiver for wirelessly communicating this information with programmer 50. In this alternative embodiment, wireless communication between controller 30 and programmer 50 may be encrypted with an encryption key associated with a unique identification code (e.g., a serial number) of the controller.

[0086] The battery port 32 is configured to removably receive a Figure 14, and the like. The battery port 32 is illustratively shown as a cable 41 integrated with the waist belt 42, such that the cable 41 is routed through the waist belt and extends around the patient's back until it is coupled to the controller 30. In this manner, the battery 40 can be removed from the waist belt 42 and disconnected from the controller 30 to enable the patient to periodically replace the battery with a fully charged battery. It is expected that the patient will have access to at least two batteries, such that while one battery is coupled to the controller 30 to power the controller and the implantable pump, the other battery can be connected to a charging station. Alternatively or in addition, the battery port 32 can be configured to accept a cable that is directly coupled to a power supply (permitting the battery 40 to be removed while the patient is lying supine in bed (e.g., sleeping), such as a much larger battery / charger combination).

[0087] Output port 33 is electrically coupled to cable 29, which in turn is coupled to implantable pump 20 through electrical conduit 28 of pump housing 27. Cable 29 both provides energy to energize implantable pump 20 according to configuration settings and operating parameters stored in controller 30 and receives data from sensors disposed in implantable pump 20. In one embodiment, cable 29 may include an electrical cable with a biocompatible coating and be designed to extend percutaneously. Cable 29 may be impregnated with a drug to reduce the risk of infection, the transmission of potentially hazardous substances, or to promote healing if it extends through the patient's skin.

[0088] As mentioned above, controller 30 may include indicator lights 34, display 35, status lights 36, and buttons 37. Indicator lights 34 may visually display information related to the operation of the system, such as the remaining life of battery 40. Display 35 may be a digital liquid crystal display that displays real-time pump performance data, physiological data of the patient (e.g., heart rate), or operating parameters of the implantable pump (e.g., target pump pressure or flow rate), etc. When a particular parameter condition is determined to exceed a preprogrammed threshold, an alarm may be audible and the alarm may be displayed on display 35. Status lights 36 may include light emitting diodes (LEDs) that are turned on or off to indicate whether a particular functionality of the controller or implantable pump is active. Button 37 may be used to wake up display 35, to set or silence an alarm, etc.

[0089] about Figure 3B ,describe Figure 3A The components of the illustrative embodiment of the controller 30. Figure 3AIn addition to the components of controller 30 described, controller 30 further includes a microprocessor 38, memory 39, battery 43, optional transceiver 44, and amplifier circuitry 45. The microprocessor may be a general purpose microprocessor in which programming for controlling the operation of implantable pump 20 is stored in memory 39. Memory 39 may also store configuration settings and operating parameters for implantable pump 20. Battery 40 supplies power to controller 30 to provide operational continuity when battery 40 is periodically replaced. Optional transceiver 44 facilitates wireless communication with programmer 50 and / or mobile device 60 via any of a number of well-known communication standards, including BLUETOOTH™, ZigBee, and / or any IEEE 802.11 wireless standards, such as Wi-Fi or Wi-Fi Direct. Controller 30 may further include amplifier circuitry 45 for amplifying electrical signals transmitted between controller 30 and implantable pump 20.

[0090] Reference now Figure 4 , depicting battery 40. Battery 40 provides power to implantable pump 20 and may also provide power to controller 30. Battery 40 may be comprised of a single battery or multiple batteries disposed within a housing, and is preferably sized and configured to be worn on the exterior of the patient's body, such as on a belt 42. A battery life indicator 46 may be provided on the exterior of battery 40 to cause an indication of the extent of the remaining charge of the battery. Cable 41 may have one end removably coupled to battery 40 and another end removably coupled to a battery port of controller 30 to supply power to energize implantable pump 20. In one embodiment, battery 40 may be recharged using a separate charging station, as is known in the art of rechargeable batteries. Alternatively or in addition, battery 40 may include a port 47 that may be removably coupled to a transformer and cable to allow the battery to be recharged using a conventional household electrical outlet (e.g., 120V, 60Hz AC power).

[0091] Reference now Figures 5A to 5B, describing a programmer 50. The programmer 50 may be a conventional laptop computer loaded with a programmed software routine for configuring the controller 30 and setting the operating parameters that are used by the controller 30 to control the operation of the implantable pump 20. As discussed above, the programmer 50 is typically located in a clinician's office or hospital and is coupled to the controller 30 via a cable 51 or wirelessly to initially set up the controller 30, and then periodically adjust the operating parameters as needed as required. The operating parameters of the controller 30 set using the programmed routine of the programmer 50 may include, but are not limited to, applied voltage, pump frequency, pump amplitude, target flow rate, pulsatility, etc. When first implanted, the surgeon or clinician may use the programmer 50 to transfer the initial operating parameters to the controller 30. After implantation, the patient may periodically return to the clinician's office to adjust the operating parameters, which may again be made using the programmer 50.

[0092] Programmer 50 may be any type of conventional personal computer device, such as a laptop or tablet computer with touch screen capability. Figure 5B As illustrated in FIG, the programmer 50 preferably includes a processor 52, a memory 53, an input / output device 54, a display 55, a battery 56, and a communication unit 57. The memory 53 may include an operating system for the programmer and programmed routines required to communicate with the controller 30. The communication unit 57 may include any of a number of well-known communication protocols such as BLUETOOTH™, ZigBee, and / or any IEEE 802.11 wireless standard such as Wi-Fi or Wi-Fi Direct. Figure 5A , the programmed routine for programming and communicating with controller 30 may also provide data identifying the operating parameters by which controller 30 controls implantable pump 20 for display on the screen of programmer 50. When programmer 50 is coupled to controller 30 via a wired or wireless connection, the programmed routine may also enable the programmer to download operational data or physiological data communicated by the implantable pump from the controller and display the information in real time. The transmitted data may then be processed and displayed on the screen of programmer 50.

[0093] Reference now Figure 6 and 7 , illustrates a preferred embodiment of a pump assembly 70 and an implantable pump 20. However, it should be understood that the pump assembly and the implantable pump and the components included therein may have other configurations without departing from the invention described herein. Figure 6 and 7 Shapes and sizes that are different from those illustrated in FIG. Figure 7, the pump assembly 70 is configured to fit within the pump housing 27. To secure the pump assembly 70 within the pump housing 27, the pump assembly 70 may include a retaining ring 71 that may extend from and around the stator assembly 72 and may be captured between the upper housing portion 24 and the lower housing portion 25 when the housing portions are assembled, as shown in FIG. Figure 7 . In this way, the stator assembly 72 may be suspended within the pump housing in a close-fitting relationship with the inner wall of the pump housing. The fixing ring 71 is preferably a rigid annular structure concentrically disposed around the stator assembly 72, the rigid annular structure having a diameter greater than the stator assembly 72. The fixing ring 71 may be rigidly coupled to the stator assembly 72 via struts 73. The struts 73 may form a gap 74 between the fixing ring 71 and the stator assembly 72, the gap 74 preferably being approximately 0.05 mm at its most restricted point.

[0094] like Figure 7 , the pump assembly 70 can be disposed in the pump housing 27 such that the retaining ring 71 is captured on a step 75 formed between the upper housing portion 24 and the lower housing portion 25. In this way, the stator assembly 72 can be suspended within the pump housing 27, and the stator assembly 72 is prevented from moving within the pump housing 27. The pump housing 27 is preferably sized and configured to conform to the pump assembly 70 so that the stator assembly 72 does not contact the interior of the pump housing at any location other than the retaining ring 71.

[0095] Figure 8 2 is an exploded view of the implantable pump 20, which depicts the arrangement of the internal components of the pump assembly 70 arranged between the upper housing portion 24 and the lower housing portion 25. In particular, the pump assembly 70 may include a stator assembly 72, a magnetic ring assembly 76, a first electromagnetic coil 77, a second electromagnetic coil 78, a fixing ring 71, a first suspension ring 79, a second suspension ring 80, a rod 81, and a membrane assembly 82. The stator assembly 72 may include a tapered section 83, electromagnetic coil holder portions 84, 85, and 86, and a flanged portion 87. The magnetic ring assembly 76 may include a magnetic ring 88 and magnetic ring holder portions 89 and 90. The first electromagnetic coil 77 and the second electromagnetic coil 78 together with the electromagnetic coil holder portions 84, 85, and 86 may form an electromagnet assembly 91. The electromagnet assembly 91 together with the stator assembly 72 forms an actuator assembly. The actuator assembly together with the magnetic ring assembly 76 then forms the actuator system of the implantable pump 20 .

[0096] The first electromagnetic coil 77 and the second electromagnetic coil 78 may be concentrically sandwiched between the electromagnetic coil holder portions 84, 85 and 86 to form an electromagnet assembly 91. The tapered section 83, which may be coupled to the fixed ring 71 and the first suspension spring 79, may be concentrically located on top of the electromagnet assembly 91. The magnetic ring 88 may be disposed together with the magnetic ring holder portions 89 and 90 to form the magnetic ring assembly 76, which may be concentrically disposed to reciprocate above the electromagnet assembly 91. The second suspension ring 80 may be concentrically disposed under the electromagnet assembly 91. The flanged portion 87 may be concentrically disposed under the second suspension ring 80. The rod 81 may engage the first suspension ring 79, the magnetic ring assembly 76 and the second suspension ring 80 at equidistant positions around the actuator assembly. The membrane assembly 82 may be concentrically positioned under the flanged portion 87 and engaged with the rod 81.

[0097] about Fig. 9 Additional details of the pump assembly 70 are provided. Specifically, the actuator assembly 95 includes the stator assembly 72 and the electromagnet assembly 91, which includes the first electromagnetic coil 77 and the second electromagnetic coil 78. During use of the implantable pump 20, the actuator assembly 95 remains fixed relative to the pump housing 27. The first electromagnetic coil 77 and the second electromagnetic coil 78 can be separated by the electromagnetic retainer portion 85. The controller 30 and the battery 40 are electrically coupled to the electromagnetic coils 77 and 78 via the cable 29 extending through the electrical conduit 28 of the pump housing 27 to supply current to the electromagnetic coils 77 and 78. The first electromagnetic coil 77 and the second electromagnetic coil 78 can be in electrical communication with each other or can be configured to operate independently and make separate wired connections with the controller 30 and the battery 40 via the cable 29.

[0098] The electromagnetic coils 77 and 78 may be made of any conductive metal material, such as copper, and may further be composed of one or more smaller metal wires wound into a coil. The wires of the electromagnetic coils are insulated to prevent shorting to adjacent conductive materials. Other components of the pump assembly 70, such as the stator assembly 72, are also preferably insulated and / or made of non-conductive materials to reduce unwanted transmission of electrical signals.

[0099] The actuator assembly 95 may be surrounded by a first suspension ring 79 and a second suspension ring 80. The suspension rings 79 and 80 may be annular in shape and fit concentrically around the actuator assembly 95. The first suspension ring 79 is preferably rigidly attached to the tapered section 83 near the top portion of the stator assembly 72 via a post 73 extending from the suspension ring to the stator assembly. As discussed above, the post 73 may also attach the fixing ring 71 to the stator assembly 72. The fixing ring 71 and the first suspension spring 79 may be sized and positioned so that a gap of no less than 0.5 mm exists between the first suspension ring 79 and the fixing ring 71. The second suspension ring 80 may similarly be rigidly attached below the electromagnet assembly 91 near the bottom of the stator assembly 72 via a post. The suspension rings 79 and 80 are preferably sized and shaped so that when the suspension rings 79 and 80 are positioned around the actuator assembly 95 , a gap of no less than 0.5 mm exists between the actuator assembly 95 and the suspension rings 79 and 80 .

[0100] The first and second suspension rings 79, 80 may comprise stainless steel having elastic properties and exhibiting a spring force when deflected in a direction normal to the plane of the spring. The first and second suspension rings 79, 80 may be substantially rigid relative to forces applied tangentially to the rings. In this manner, the first and second suspension rings 79, 80 may exhibit spring tension when deformed upward and downward relative to the vertical axis of the actuator assembly, but may be rigid against movement along any other axis, such as tilting or torsional movement.

[0101] The magnetic ring assembly 76 can be annular in shape and concentrically surround the actuator assembly 95. The magnetic ring 88 can include one or more materials that exhibit magnetic properties, such as iron, nickel, cobalt, or various alloys. The magnetic ring 88 can be made of a single integral component or include several magnetic components coupled together. The magnetic ring assembly 76 can be sized and shaped so that a gap of not less than 0.5 mm exists between the outer lateral surface of the actuator assembly 95 and the inner surface of the magnetic ring assembly 76 when it is concentrically positioned above the actuator assembly 95.

[0102] The magnetic ring assembly 76 may be positioned concentrically about the actuator assembly 95 between the first and second suspension rings 79, 80 and may be rigidly coupled to the first and second suspension rings 79, 80. The magnetic ring assembly 76 may be rigidly coupled to the suspension rings by more than one rod 81 evenly spaced about the actuator assembly 95 and configured to extend parallel to the central axis of the pump assembly 70. The suspension rings 79 and 80 and the magnetic ring assembly 76 may be engaged such that the magnetic ring assembly 76 is suspended equidistantly between the first and second electromagnetic coils 77, 78 when the suspension rings are in their non-deflected shapes. Each of the suspension rings 79 and 80 and the magnetic ring holder portions 89 and 90 may include a rod receiving area for engaging with the rod 81 or may be attached to the rod 81 in any suitable manner that causes the suspension rings 79 and 80 and the magnetic ring assembly 76 to be rigidly attached to the rod 81. The rod 81 may extend beyond the bail 79 and 80 to engage other components, such as the flanged portion 87 and the membrane assembly 82 .

[0103] The first electromagnetic coil 77 may be activated by a controller that applies an electrical signal from the battery 40 to the first electromagnetic coil 77, thereby inducing a current in the electromagnetic coil and generating a magnetic field surrounding the electromagnetic coil 77. The direction of the current in the electromagnetic coil 77 and the polarity of the magnetic ring assembly 76 closest to the electromagnetic coil 77 may be configured so that the first electromagnetic coil magnetically attracts or repels the magnetic ring assembly 76 as desired. Similarly, a magnetic field may be formed in the second electromagnetic coil 78 by introducing current in the second electromagnetic coil. The direction of the current in the second electromagnetic coil 78 and the polarity of the magnetic ring assembly 76 closest to the second electromagnetic coil may also be similarly configured so that the first electromagnetic coil 77 magnetically attracts or repels the magnetic ring assembly 76 when an appropriate current is induced in the second electromagnetic coil 78.

[0104] Since the magnetic ring assembly 76 can be rigidly attached to the rod 81 (which can in turn be rigidly attached to the first and second suspension rings 79 and 80), the elastic properties of the suspension rings permit the magnetic ring assembly 76 to move upward toward the first electromagnetic coil 77 or downward toward the second electromagnetic coil 78, depending on the polarity of the magnetic field generated by the electromagnetic ring. In this way, when the magnetic ring assembly 76 experiences an upward magnetic force, the magnetic ring assembly 76 deflects upward toward the first electromagnetic coil 77. When the rod 81 moves upward with the magnetic ring assembly 76, the rod 81 causes the suspension rings 79 and 80 to elastically deform, which forms a spring force opposite to the direction of movement. When the magnetic field generated by the first electromagnetic coil collapses when the current stops, this downward spring force causes the magnetic ring assembly to return to its neutral position. Similarly, when the magnetic ring assembly 76 is magnetically attracted downward, the magnetic ring assembly 76 deflects downward toward the second electromagnetic ring 78. When rod 81 moves downward with magnetic ring assembly 76, rod 81 imposes elastic deformation of the first and second suspension rings, thus generating spring forces in opposite directions. When the magnetic field generated by the second electromagnetic ring collapses when the current stops, this upward spring force causes the magnetic ring assembly to return to its neutral position again.

[0105] The electromagnetic coils 77 and 78 may be energized individually, or alternatively, the electromagnetic coils 77 and 78 may be connected in series so that the electromagnetic coils are activated simultaneously. In this configuration, the first magnetic coil may be configured to experience a current direction opposite to the current direction of the second electromagnetic coil. Thus, when current is induced to the first electromagnetic coil 77 to attract the magnetic ring assembly 76, the same current is applied to the second electromagnetic coil 78 to induce a current that causes the second electromagnetic coil 78 to repel the magnetic ring assembly 76. Similarly, when current is induced to the second electromagnetic coil 78 to attract the magnetic ring assembly 76, the current applied to the first electromagnetic coil 77 causes the first electromagnetic coil to repel the magnetic ring assembly 76. In this way, the electromagnetic coils 77 and 78 work together to cause deflection of the magnetic ring assembly 76.

[0106] By manipulating the timing and strength of the electrical signals applied to the electromagnetic coils, the frequency at which the magnetic ring assembly 76 deflects toward the first and second electromagnetic coils can be altered. For example, by alternating the current induced in the electromagnetic coils more frequently, the magnetic ring assembly can be caused to cycle up and down more times in a given cycle. By increasing the amount of current, the magnetic ring assembly can be deflected at a faster rate and caused to travel a longer distance.

[0107] Alternatively, the first electromagnetic coil 77 and the second electromagnetic coil 78 may be energized independently. For example, the first electromagnetic coil 77 and the second electromagnetic coil 78 may be energized at varying intensities; one electromagnetic coil may be coordinated to decrease in intensity while the other electromagnetic coil increases in intensity. In this way, the intensity of the signal applied to the second electromagnetic coil 78 to cause downward magnetic attraction may be increased while the intensity of the signal applied to the first electromagnetic coil 77 causes the upward magnetic attraction to decrease.

[0108] According to one aspect of the present invention, the movement of the magnetic ring assembly 76 can translate to a film assembly 82 which can be disposed concentrically below the stator assembly 72. The film assembly 82 is preferably rigidly attached to the magnetic ring assembly 76 by a rod 81. Fig. 9 In the embodiment depicted in , the rod 81 may extend beyond the second suspension ring 80 and couple to the film assembly 82 .

[0109] Reference now Fig.10 , an embodiment of the film assembly 82 is described in more detail. The film assembly 82 may include a rigid film ring 96 and a film 97. The rigid film ring 96 exhibits a rigid nature under the typical forces experienced during the entire operating range of the present invention. A rod receiving site 98 may be formed into the rigid film ring 96 to engage the film assembly 82 with the rod 81. Alternatively, the rod 81 can be attached to the rigid film ring 96 in any other way that the movement of the magnetic ring assembly 76 is directly translated to the rigid film ring 96. The rigid film ring 96 can be attached to the film 97 and the film is held in tension. The film 97 can be molded directly onto the rigid film ring 96 or can be attached to the rigid film ring 96 in any way that the film 97 is uniformly held in tension along its circumference. The film 97 may alternatively include a flexible pleated structure, and the film 97 is attached to the rigid film ring 96 at the pleated structure to increase the ability of the film to move, wherein the film is attached to the rigid film ring 96. The film 97 may further include a circular orifice 99 disposed in the center of the film.

[0110] In a preferred embodiment, the film 97 has a thin planar shape and is made of an elastomer having elastic properties and good durability. Alternatively, the film 97 may have a uniform thickness from the film ring to the circular orifice. As yet another alternative, the thickness of the film 97 may vary and exhibit a more complex geometry. For example, Fig.10 As shown in FIG. 1 , the membrane 97 may have a reduced thickness as it extends from the rigid membrane ring 96 to the circular orifice 99. Alternatively or in addition, the membrane 97 may incorporate a metal element such as a helical spring to enhance the spring force of the membrane in a direction normal to the plane of the membrane, and this spring force may vary radially along the membrane. In yet another embodiment, the membrane 97 may be preformed in an undulating shape.

[0111] Fig.11 Will Figures 6 to 9 The moving portion of the embodiment of the pump assembly 70 shown in FIG. 7 is depicted as a non-ashing element. The non-moving portion of the pump assembly, including the actuator assembly 95 and the electromagnet assembly 91 (partially shown), can be fixed to the pump housing 27 by a fixed ring 71. The moving portion of the pump assembly 70 may include a rod 81, a first suspension spring 79, a magnetic ring assembly 76, a second suspension spring 80, and a film assembly 82. When the magnetic ring assembly 76 moves up and down, the rod 81 will move rigidly to the film assembly 82. Given the rigidity of the rod, when the magnetic ring assembly 76 moves a specific distance up or down, the film assembly 82 can travel the same distance. For example, when the magnetic ring assembly 76 moves 2mm from a position near the first electromagnetic coil 77 to a position near the second electromagnetic coil 78, the film assembly 82 may also travel 2mm in the same direction. Similarly, the frequency at which the magnetic ring assembly 76 passes through the space between the first electromagnetic coil and the second electromagnetic coil can be the same frequency at which the film assembly 82 travels the same distance.

[0112] Reference now Fig.12 ,exist Figures 6 to 9 In the embodiment of the implantable pump 20 described in , blood can enter the implantable pump 20 from the left ventricle through the inlet cannula 21 and flow down the pump assembly 70 into the delivery channel 100 defined by the interior surface of the pump housing 27 and the exterior of the pump assembly 70. The delivery channel 100 begins at the top of the stator assembly 72 and extends between the tapered section 83 and the interior of the pump housing 27. As the blood moves down along the tapered section 83, it is directed through the gap 74 and into the vertical portion of the delivery channel 100 in the area between the pump housing 27 and the actuator assembly 95 and included in the gap between the magnetic ring assembly 76 and the electromagnet assembly 91. The delivery channel 100 extends downward to the flanged portion 87 of the stator assembly 72, which delivers the blood into the flow channel 101 within which the membrane assembly 82 is suspended. The delivery channel 100 delivers the blood to the membrane assembly 82 by directing the blood from the inlet cannula 21 to the flow channel 101 through the delivery channel 100. By actuating the electromagnetic coils 77 and 78, the membrane 97 can be caused to undulate within the flow channel 101 to induce a wave-like configuration in the membrane 97 that moves from the edge of the membrane toward the circular orifice 99. Thus, when blood is delivered from the delivery channel 100 to the membrane assembly 82, it can be radially propelled along both the top and bottom of the membrane 97 toward the circular orifice 99 and out of the outlet 23 therefrom.

[0113] According to one aspect of the invention, the heaving membrane pump described herein avoids thrombosis by placing all moving parts directly within the main flow path, thereby reducing the risk of flow stagnation. Specifically, Fig.11The moving components depicted in FIG, including magnetic ring assembly 76, suspension rings 79 and 80, rod 81, and membrane assembly 82, are all located within delivery channel 100 and flow channel 101. Flow stagnation can be further avoided by eliminating secondary flow paths that may experience significantly slower flow rates.

[0114] Now turn to Fig.13 and 14 , showing the lower portion of the implantable pump 20 including the flanged portion 87, the membrane assembly 82, and the lower housing portion 23. The delivery channel 100 can be in fluid communication with the membrane assembly 82 and a flow channel 101, which is defined by the bottom surface of the flanged portion 87 and the interior surface of the lower housing portion 25. The flanged portion 87 can include a feature 102 that extends downward as the bottom of the flanged portion 87 moves radially inward. The interior surface of the lower housing portion 25 can also be inclined upward as it extends radially inward. The upward inclination of the interior surface of the lower housing portion 25 combined with the bottom surface of the flanged portion 87 moving downward causes the flow channel 101 to narrow as the channel moves radially inward from the delivery channel 100 to the circular orifice 99 of the membrane 97 (which is disposed around the pump outlet 23).

[0115] As explained above, the membrane assembly 82 can be suspended by the rod 81 in the flow channel 101 below the bottom surface of the flanged portion 87 and above the inner surface of the lower housing portion 25. The membrane assembly 82 can freely move up and down in the vertical direction in the flow channel 101, and the movement is limited only by the suspension rings 79 and 80. The membrane assembly 82 can be prevented from twisting, tilting, or moving in any direction other than upward and downward in the flow channel 101 by the rigid rod 81 and by the suspension ring.

[0116] The flow channel 101 is divided into an upper flow channel and a lower flow channel by the membrane 97. The geometry of the membrane 97 may be angled so that when the membrane assembly 82 is at rest, the top surface of the membrane 97 is parallel to the bottom surface of the flanged portion 87 and the bottom surface of the membrane 97 is parallel to the opposing surface of the lower housing portion 25. Alternatively, the membrane 97 may be sized and shaped so that when the membrane assembly 82 is at rest, the upper and lower flow channels narrow as they move radially inward from the delivery channel 100 to the circular orifice 99 in the membrane 97.

[0117] Now also refer to Fig.14, when the rod 81 causes the rigid membrane ring 96 to move up and down in the flow channel 101, the outermost portion of the membrane 97 closest to the rigid membrane ring 96 moves up and down with the rigid membrane ring 96. The membrane 97, which is flexible and has elastic properties, gradually translates the up and down movement of the membrane portion closest to the rigid membrane ring 96 along the membrane 97 toward the circular orifice 99. This movement across the flexible membrane 97 causes a wave-like deformation in the membrane, which can propagate inward from the rigid membrane ring 96 toward the orifice 99.

[0118] The waves formed in the undulating film can be manipulated by changing the speed at which the rigid film ring 96 moves up and down and the distance that the rigid film ring 96 moves up and down. As explained above, the amplitude and frequency of the up and down movement of the rigid film ring 96 are determined by the amplitude and frequency of the reciprocating movement of the magnetic ring assembly 76 over the electromagnet assembly 91. Therefore, the waves formed in the undulating film can be adjusted by changing the frequency and amplitude of the reciprocating movement of the magnetic ring assembly 76.

[0119] When blood is introduced from delivery channel 100 into flow channel 101, undulations in membrane 97 cause the blood to be propelled toward circular orifice 99 and out of pump housing 27 via outlet 23. Energy transferred from the membrane to the blood is directed radially inward along the length of the membrane toward orifice 99 and along both sides of membrane 97 propelling blood along the flow channel toward outlet 23.

[0120] For example, when the rigid film ring 96 and the magnetic ring assembly 76 move downward in conjunction, the upper portion of the flow channel 101 near the delivery channel 100 expands, causing blood from the delivery channel 100 to fill the upper portion of the flow channel near the outer region of the film 97. As the rigid film ring 96 moves upward, the upper portion of the flow channel 101 begins to narrow near the rigid film ring 96, causing the wave-like deformation to translate across the film. As the wave propagates across the film 97, the blood in the upper portion of the flow channel 101 is pushed toward the circular orifice and eventually leaves the pump housing 27 through the outlet 23. At the same time, when the rigid film ring 96 moves upward, the lower portion of the flow channel 101 closest to the outer portion of the film 97 begins to enlarge, allowing blood from the delivery channel 100 to flow into this area. Subsequently, when the rigid film ring 96 is pushed downward again, the area of ​​the lower portion of the flow channel 101 closest to the outer portion of the film 97 begins to narrow, causing the wave-like deformation that pushes the blood toward the outlet 23 to translate across the film.

[0121] By manipulating the waves formed in the undulating membrane (by varying the frequency and amplitude of the up and down movement of the magnetic ring assembly 76), the pressure gradient within the flow channel 101 can be adjusted and ultimately the flow rate of blood moving through the flow channel 101 can be adjusted. Properly controlling the movement of the magnetic ring assembly 76 permits oxygen-rich blood to be efficiently and safely pumped from the left ventricle to the aorta and to the body as needed.

[0122] In addition to simply pumping blood from the left ventricle to the aorta, the implantable pump 20 of the present invention can also be operated to closely mimic physiological pulsatility without losing pump efficiency. In the embodiments detailed above, pulsatility can be achieved almost immediately by changing the frequency and amplitude of the movement of the magnetic ring assembly 76 to form the desired flow output, or stopping the movement of the magnetic ring assembly for a period of time to form a low flow output or no flow output for a period of time. Unlike typical rotary pumps (which require a specific time period to obtain a set number of rotations per minute to achieve the desired fluid displacement and pulsatility), the implantable pump 20 can achieve the desired flow output almost immediately and similarly can stop the output almost immediately due to the very low inertia generated by the small moving mass of the moving components of the pump assembly. The ability to start and stop on demand allows rapid changes in pressure and flow. Together with the frequency and amplitude, the duty cycle defined by the percentage of time that the film 97 is excited within a set time period can be adjusted to achieve the desired flow output and pulsatility without losing pump efficiency. Even if the frequency and amplitude are held constant, the flow rate can be varied by manipulating the duty cycle between 0 and 100%.

[0123] According to another aspect of the present invention, the controller 30 can be programmed by the programmer 50 to operate at a selected frequency, amplitude, and duty cycle to achieve a wide range of physiological flow rates and physiological pulsatility. For example, the programmer 50 can instruct the controller 30 to operate the implantable pump 20 at a given frequency, amplitude, and / or duty cycle during the time period when the patient is typically asleep and can instruct the controller 30 to operate the implantable pump 20 at a different frequency, amplitude, and / or duty cycle during the time period when the patient is typically awake. The controller 30 or the implantable pump can also include an accelerometer or position indicator to determine whether the patient is supine or ambulatory, the output of which can be used to move from one set of pump operating parameters to another set of pump operating parameters. When the patient feels some discomfort or the physician determines that the parameters are not optimized, the physician can change at least one or more of the frequency, amplitude, and duty cycle to achieve the desired functionality. Alternatively, the controller 30 or the mobile device 60 can be configured to change one or more of the frequency, amplitude, and duty cycle to meet the needs of the patient.

[0124] The implantable pump 20 may further include one or more additional sensors for adjusting the flow output and pulsatility according to the patient's needs. The sensors may be incorporated into the implantable pump 20, or alternatively or additionally may be implanted in or elsewhere on the patient. The sensors are preferably in electrical communication with the controller 30, and may monitor operating parameters that measure the performance of the implantable pump 20 or physiological sensors that measure a physiological parameter of the patient, such as heart rate or blood pressure. By using one or more physiological sensors, pulsatile flow may be synchronized with the patient's cardiac cycle by monitoring blood pressure or muscle contraction, for example, and synchronizing the duty cycle according to the sensed output.

[0125] The controller 30 may compare the physiological sensor measurements to the current implantable pump output. If it is determined by analyzing the sensor measurements that the demand exceeds the current output, the frequency, amplitude and / or duty cycle may be automatically adjusted to meet the current demand. Similarly, the controller may determine that the current output exceeds the demand and therefore modify the output by changing the frequency, amplitude and / or duty cycle. Alternatively or in addition, the controller 30 may issue an alarm when it is determined that the demand exceeds the current output. Similarly, the operational measurements from the operational sensors may be compared to a predetermined threshold and the controller 30 may issue an alarm when the measurements exceed the predetermined threshold or a fault is detected.

[0126] The implantable pump 20 is sized and shaped to produce physiological flow rates, pressure gradients, and pulsatility at an operating point that achieves maximum efficiency. Specifically, the implantable pump 20 can be sized and shaped to produce a physiological flow rate ranging from 4 to 6 liters / minute at a pressure gradient lower than the threshold associated with hemolysis. Moreover, in order to mimic a typical physiological pulsation of 60 times / minute, the implantable pump 20 can be pulsated approximately once per second. To achieve this pulsatility, a 50% duty cycle can be utilized, with an "on" period of 0.5 seconds and an "off" period of 0.5 seconds. For a given system, maximum efficiency at a specific operating frequency, amplitude, and voltage can be achieved by manipulating the shape and size of the blood flow channel, the elastic properties of the suspension ring, the mass of the moving parts, the film geometry, and one or more of the elastic properties and friction properties of the film while producing a flow rate of 4 to 6 liters / minute at a 50% duty cycle. In this way, the implantable pump 20 can be designed to produce a desirable physiological output while continuing to function at optimal operating parameters.

[0127] By adjusting the duty cycle, implantable pump 20 can be configured to produce a wide range of output flow rates under physiological pressure gradients. For example, for an exemplary LVAD system configured to produce 4 to 6 liters / minute at a 50% duty cycle, the optimal operating frequency may be 120 Hz. For this system, the flow output may be increased to 10 liters / minute or decreased to 4 liters / minute, for example, by changing only the duty cycle. When the duty cycle and frequency are operated independently of each other, the duty cycle may be manipulated between 0% and 100% while leaving the frequency of 120 Hz unaffected.

[0128] The implantable pump system described herein, tuned to achieve physiological flow rates, pressure gradients, and pulsatility, also avoids hemolysis and platelet activation by applying low to moderate shear forces to the blood, similar to the shear forces applied by a healthy heart. The moving components are rigidly attached to each other and do not incorporate any components that would cause friction, such as mechanical bearings or gears. In the embodiments detailed above, the delivery channel 100 can be sized and configured to also avoid friction between the moving magnetic ring assembly 76, the suspension rings 79 and 80, the rod 81, and the lower housing portion 25 by sizing the channel so that a gap of at least 0.5 mm is maintained between all moving components. Similarly, the magnetic ring assembly 76, the suspension rings 79 and 80, and the rod 81 can all be offset from the stator assembly 72 by at least 0.5 mm to avoid friction between the stator assembly and the moving components.

[0129] Reference now Fig.15A and 15B , describes an alternative exemplary embodiment of the pump assembly of the present invention. The implantable pump 20' is similar to the one described in Figure 7 , 812 , wherein similar components are identified by similar quotation marks. Implantable pump 20′ differs from implantable pump 20 in that membrane assembly 82′ includes a skirt 115 coupled to membrane 97′. The skirt illustratively includes a first portion 115a and a second portion 115b. The first portion 115a of skirt 115 extends upwardly toward inlet 21′ in a first direction (e.g., parallel to the longitudinal axis of stator assembly 72′ and / or the central axis of pump housing 27′) within delivery channel 100′. The second portion 115b of skirt 115 is bent toward outlet 23′ so that the second portion 115b is coupled to membrane 97′ so that membrane 97′ is oriented in a second direction, e.g., perpendicular to the first portion 115a of skirt 115. For example, the skirt 115 may have a J-shaped cross-section such that the first portion 115a forms a cylindrical ring around the stator assembly 72' and the second portion 115b has a predetermined radius of curvature that allows blood to flow smoothly from the delivery channel 100' across the skirt 115 to the outer edge of the membrane 97' and into the flow channel 101' while reducing blood flow stagnation. The skirt 115 interrupts the flow recirculation of blood within the delivery channel 100' and improves the hydraulic power generated for a given frequency while minimizing blood damage. In addition, the J-shape of the skirt 115 around the stator assembly 72' is significantly stiffer than a flat rigid membrane ring, thereby reducing buckling and fatigue as well as resistance to blood moving across the membrane 97'.

[0130] The skirt 115 exhibits a rigid nature under typical forces experienced during the entire operating range of the present invention and can be made of a biocompatible metal (e.g., titanium). The skirt 115 is preferably impermeable so that blood cannot flow through the skirt 115. A rod receiving site 98' can be formed into the skirt 115 to engage the film assembly 82' with the rod 81'. Alternatively, the rod 81' can be attached to the skirt 115 in any other manner that directly translates the movement of the magnetic ring assembly 76' to the skirt 115.

[0131] As the magnetic ring assembly 76' moves up and down, the rod 81' will move the J-shape of the skirt 115 that rigidly translates to the film assembly 82'. Given the rigidity of the rod, when the magnetic ring assembly 76' travels a certain distance up or down, the film assembly 82' can travel the same distance. For example, when the magnetic ring assembly 76' travels 2mm from a position near the first electromagnetic coil 77' to a position near the second electromagnetic coil 78', the film assembly 82' can also travel 2mm in the same direction. Similarly, the frequency at which the magnetic ring assembly 76' traverses the space between the first electromagnetic coil and the second electromagnetic coil can be the same as the frequency at which the film assembly 82' travels the same distance.

[0132] The skirt 115 may be attached to the membrane 97' and hold the membrane 97' in tension. The membrane 97' may be molded directly onto the skirt 115 or attached to the skirt 115 in any manner that may hold the membrane 97' in tension evenly along its circumference. For example, the skirt 115 may be coated with the same material used to form the membrane 97' and the coating on the skirt 115 may be integral with the membrane 97'.

[0133] Blood can enter the implantable pump 20' from the left ventricle through the inlet cannula 21' and flow down the pump assembly into the delivery channel 100'. As the blood moves down the tapered section 83', it is directed through the gap 74' and into the vertical portion of the delivery channel 100' in the area between the pump housing 27' and the actuator assembly 95'. Fig.15A , the skirt 115 divides the delivery channel 100′ into an upper delivery channel 100a and a lower delivery channel 100b, so that the blood flow through the delivery channel 100′ is divided into a flow channel 101a through the upper delivery channel 100a and a flow channel 101b through the lower delivery channel 100b, wherein the flow channels 101a and 101b are separated by the film 97′. As will be understood by those skilled in the art, the volume of blood flow through each of the delivery channels 100a and 100b may depend on the diameter of the first portion 115a of the skirt 115. For example, the larger the diameter of the first portion 115a of the skirt 115, the larger the volume of the delivery channel 100a and the smaller the volume of the delivery channel 100b. The ratio of the volume of delivery channel 100a to the volume of delivery channel 100b may be, for example, 1:1, 1:2, 1:3, 1:4, 2:1, 3:1, 4:1, etc., depending on the desired blood flow on each surface of membrane 97'.

[0134] By directing blood from the inlet cannula 21' across the skirt 115 in the delivery channel 100', the blood flow is divided into the delivery channels 100a and 100b and the flow channels 101a and 101b, respectively, so that the blood flows across the upper and lower surfaces of the membrane 97' of the membrane assembly 82'. Fig.16A As shown in , blood flow through a pump having a planar rigid membrane ring spaced a relatively small distance from the pump housing can allow unrestricted blood flow across the upper surface of the flexible membrane while restricting blood flow across the lower surface of the flexible membrane. Fig. 16B As depicted in , blood flow through a pump having a J-shaped skirt can be distributed in a desired ratio across both the upper and lower sides of the flexible membrane.

[0135] Return to reference Fig.15ABy actuating the electromagnetic coils 77' and 78', the membrane 97' can be caused to undulate within the flow channels 101a and 101b to induce a wave-like configuration in the membrane 97', which moves from the edge of the membrane 97' toward the circular orifice 99'. Therefore, when blood is delivered from the delivery channel 100' to the membrane assembly 82', it can be radially urged along both the upper and lower surfaces of the membrane 97' toward the circular orifice 99' and from there out of the outlet 23'. The distribution of blood flow across the upper and lower surfaces of the membrane 97' reduces blood recirculation within the delivery channel 101' and reduces repeated exposure of blood to high shear stress areas, which results in a significant improvement in the hydraulic performance of the implantable pump 20'.

[0136] Reference now Fig.17 , describes the relationship between the maximum hydraulic power of the pump system and the height of the J-shaped skirt. As the height of the vertical portion of the skirt increases, the maximum hydraulic power of the pump increases at a nonlinear rate. For example, Fig.17 As shown in FIG, operating a pump with a flat rigid membrane ring at 60 Hz generates a maximum hydraulic power of 0.15 W, operating at 90 Hz generates a maximum hydraulic power of 0.47 W, and operating at 120 Hz generates a maximum hydraulic power of 1.42 W. Operating a pump with a skirt (having an extension height of 2 mm measured from the top surface of the membrane ring to the top of the J-shaped skirt) at 60 Hz generates a maximum hydraulic power of 0.16 W, operating at 90 Hz generates a maximum hydraulic power of 0.85 W, and operating at 120 Hz generates a maximum hydraulic power of 1.54 W. Operating a pump with a skirt (having an extension height of 4 mm) at 60 Hz generates a maximum hydraulic power of 0.43 W, operating at 90 Hz generates a maximum hydraulic power of 1.06 W, and operating at 120 Hz generates a maximum hydraulic power of 2.44 W. A pump with a skirt (having an extended height of 10 mm) operating at 60 Hz will produce a maximum hydraulic power of 0.75 W, a maximum hydraulic power of 1.89 W at 90 Hz, and a maximum hydraulic power of 4.03 W at 120 Hz. A pump with a skirt (having an extended height of 18 mm) operating at 60 Hz will produce a maximum hydraulic power of 1.16 W, a maximum hydraulic power of 3.08 W at 90 Hz, and a maximum hydraulic power of 9.13 W at 120 Hz. Thus, the height of the skirt 115 is preferably at least 2 mm, and more preferably at least 4 mm, at least 10 mm, and / or at least 18 mm. Thus, the implantable pump 20' can be operated at significantly lower frequencies to achieve the same hydraulic output as a pump with a planar rigid membrane ring operating at a higher frequency, while reducing blood damage and increasing fatigue life of the membrane 97' and spring.

[0137] Reference now Fig.18, depicting an alternative exemplary embodiment of a pump assembly of the present invention having a J-shaped skirt. Implantable pump 20″ is similar to Fig.15A The implantable pump 20' is constructed as described in the foregoing, wherein similar components are identified by similar double quotation mark numbers. In addition, the implantable pump 20" includes a similar Fig.15A The implantable pump 20″ differs from the implantable pump 20′ in that the inlet 21″ is coupled to the inflow cannula 116, and the outlet 23″ is coupled to the outflow cannula 117, so that the outflow cannula 117 is coaxially disposed within the inflow cannula 116, as described in U.S. Patent Publication No. 2017 / 0290967 to Barthes, the entire contents of which are incorporated herein by reference. Therefore, during operation, blood flows into the inlet 21″ via the inflow cannula 116, passes through the delivery channel 100″ into the flow channel 101″ across the membrane 97″, and exits through the outlet cannula 117 via the outlet 23″.

[0138] Reference now Fig.19 , describes another alternative exemplary embodiment of the pump assembly of the present invention. The implantable pump 20' is similar to Fig.15A and 15B ' is constructed to resemble the implantable pump 20' described in , wherein similar components are identified with similar double quotation mark numerals and similar triple quotation mark numerals. Implantable pump 20'" differs from implantable pump 20'" in that implantable pump 20'" includes a rigid ring 118 fixed around stator assembly 72'". Ring 118 extends longitudinally within delivery passage 100'" parallel to the longitudinal axis of stator assembly 72'" such that ring 118 forms a cylindrical ring around stator assembly 72'".

[0139] In addition, the membrane assembly 82'' of the implantable pump 20'' includes a skirt 119 coupled to the membrane 97''. The upper portion of the skirt 119 is substantially parallel to the ring 118, and the lower portion of the skirt 119 is bent toward the outlet 23'' so that the skirt 119 is coupled to the membrane 97'' perpendicular to the ring 118. For example, the skirt 119 may have a J-shaped cross-section having a predetermined radius of curvature that allows blood to flow smoothly from the delivery channels 100a'' and 100b'' across the skirt 119 to the outer edge of the membrane 97'' within the flow channel 101'' while reducing blood flow stagnation. In summary, the ring 118 and the skirt 119 interrupt the blood flow recirculation within the delivery channel 100'' and improve the hydraulic power generated for a given frequency while minimizing blood damage. The distance between the ring 118 and the skirt 119 is minimized to reduce blood leakage between the delivery channels 100a″ and 100b″ and reduce blood damage when the skirt 119 reciprocates in response to the magnetic field generated by the magnetic ring assembly 76″′ (as described in further detail below). In addition, the J-shape of the skirt 119 is significantly stiffer than a flat rigid film ring, thereby reducing flexure and fatigue as well as resistance to blood moving across the film 97″′.

[0140] The skirt 119 is preferably impermeable so that blood cannot flow through the skirt 119, and exhibits rigid properties under typical forces experienced during the entire operating range of the present invention, and can be made of a biocompatible metal (e.g., titanium). A rod receiving site can be formed into the skirt 119 to engage the film assembly 82'" with the rod. Alternatively, the rod can be attached to the skirt 119 in any other manner that directly translates the movement of the magnetic ring assembly 76'" to the skirt 119.

[0141] As the magnetic ring assembly 76'' moves up and down, the movement is rigidly translated by the rod to the skirt 119 of the film assembly 82'''. Given the rigidity of the rod, when the magnetic ring assembly 76''' moves a particular distance up or down, the film assembly 82''' can travel the same distance. For example, when the magnetic ring assembly 76''' travels 2mm from a position near the first electromagnetic coil 77''' to a position near the second electromagnetic coil 78''', the film assembly 82''' can also travel 2mm in the same direction. Similarly, the frequency at which the magnetic ring assembly 76''' crosses the space between the first electromagnetic coil and the second electromagnetic coil can be the same as the frequency at which the film assembly 82''' travels the same distance.

[0142] Skirt 119 may be attached to membrane 97'" and hold membrane 97'" in tension. Membrane 97'" may be molded directly onto skirt 119 or may be attached to skirt 119 in any manner that will hold membrane 97'" uniformly in tension along its circumference. For example, skirt 119 may be coated with the same material used to form membrane 97'" and the coating on skirt 119 may be integral with membrane 97'".

[0143] Blood can enter the implantable pump 20"' from the left ventricle through the inlet 21"' and flow down the pump assembly into the delivery channel 100"'. As the blood moves down the tapered section 83"', it is directed through the gap 74"' and into the vertical portion of the delivery channel 100"' in the area between the pump housing 27"' and the actuator assembly 95"'. Fig.19 As shown in the figure, the ring 118 divides the delivery channel 100"' into an upper delivery channel 100a" and a lower delivery channel 100b", so that the blood flow through the delivery channel 100"' is divided into a flow channel 101a" through the upper delivery channel 100a" and a flow channel 101b" through the lower delivery channel 100b" and crosses the skirt 119 with minimal leakage between the delivery channel 100a" and the delivery channel 100b", wherein the flow channels 101a" and 101b" are separated by a film 97"'.

[0144] As will be appreciated by one skilled in the art, the volume of blood flow through each of the delivery channels 100a″ and 100b″ may depend on the diameter of the ring 118 and the radius of curvature of the skirt 119. For example, the larger the diameter of the ring 118, the larger the volume of the delivery channel 100a″ and the smaller the volume of the delivery channel 100b″. Depending on the desired blood flow on each surface of the membrane 97″′, the ratio of the volume of the delivery channel 100a″ to the volume of the delivery channel 100b″ may be, for example, 1:1, 1:2, 1:3, 1:4, 2:1, 3:1, 4:1, etc. By directing blood from the inlet cannula 21″ across the ring 118 within the delivery channel 100″′, the blood flow is divided into the delivery channels 100a″ and 100b″ and across the skirt 119 to reach the flow channels 101a″ and 101b″, respectively, so that the blood flows across the upper and lower surfaces of the membrane 97″′ of the membrane assembly 82″′.

[0145] By actuating the electromagnetic coils 77″′ and 78″′, the membrane 97″′ can be caused to undulate within the flow channels 101a″ and 101b″ to induce a wavy configuration in the membrane 97″′, which moves from the edge of the membrane 97″′ toward the circular orifice 99″′. Therefore, when blood is delivered from the delivery channel 100″′ to the membrane assembly 82″′, it can be radially pushed along both the upper and lower surfaces of the membrane 97″′ toward the circular orifice 99″′ and from there exit the outlet 23″′. The distribution of blood flow across the upper and lower surfaces of the membrane 97″′ reduces blood recirculation within the delivery channel 101″′ and reduces repeated exposure of blood to areas of high shear stress, which results in a significant improvement in the hydraulic performance of the implantable pump 20″′.

[0146] Reference now Fig. 20 , describes yet another alternative exemplary embodiment of the pump assembly of the present invention. Implantable pump 20"" is similar to Fig.19 The implantable pump 20'" is constructed from the implantable pump 20'" described in , wherein similar components are identified with similar double-quoted, similar triple-quoted, and similar quadruple-quoted numerals. The implantable pump 20'" differs from the implantable pump 20'" in that the implantable pump 20'" includes an expandable portion 120 coupled between the ring 118' and the upper portion of the skirt 119'. The expandable portion 120 is impermeable and prevents leakage between the delivery channels 100a'"' and 100b'"'. Preferably, the expandable portion 120 has a pleated configuration that can expand and contract to permit efficient reciprocating movement of the skirt 119' relative to the ring 118'. For example, the expandable portion 120 may include a plurality of bellows having a first end coupled to the ring 118' and a second end coupled to the skirt 119'.

[0147] The expandable portion 120 may be molded directly onto the skirt 119′ or affixed to the skirt 119′ in any manner that will hold the expandable portion 120 evenly along its circumference. Similarly, the expandable portion 120 may be molded directly onto the ring 118′ or affixed to the ring 118′ in any manner that will hold the expandable portion 120 evenly along its circumference. The skirt 119′ may be coated with the same material used to form the membrane 97′″ and / or the expandable portion 120 and the coating on the skirt 119′ may be integral with the membrane 97′″ and / or the expandable portion 120.

[0148] like Fig. 20As shown in , the expandable portion 120 extends longitudinally within the delivery channel 100"" parallel to the longitudinal axis of the stator assembly 72"". Therefore, during operation, blood is guided within the delivery channel 100"" from the inlet cannula 21"" across the ring 118' and the expandable portion 120, and is divided into delivery channels 100a"' and 100b"' and across the skirt 119' to reach the flow channels 101a"' and 101b"' respectively, so that the blood flows across the upper and lower surfaces of the membrane 97"" of the membrane assembly 82"".

[0149] As the magnetic ring assembly 76"" moves up and down, the movement is rigidly translated by the rod to the skirt 119' of the film assembly 82"" and is therefore translated to the expandable portion 120. For example, when the magnetic ring assembly 76"" travels a particular distance up or down, the film assembly 82"" travels the same distance, causing the expandable portion 120 to expand and contract the same distance within the delivery channel 100"" parallel to the longitudinal axis of the stator assembly 72"". Similarly, the frequency at which the magnetic ring assembly 76"" traverses the space between the first electromagnetic coil and the second electromagnetic coil can be the same as the frequency at which the film assembly 82"" travels the same distance.

[0150] Reference now Figures 21A to 21H , various configurations are provided for energizing the implantable pumps of the present invention described above (e.g., implantable pumps 20, 20", 20"', and 20"". Fig.21A , the controller 30 includes an output port 33 electrically coupled to the cable 29 as described above, which in turn is coupled to the implantable pump. The controller 30 also includes a power connector 103, which can be electrically coupled to a battery, an expansion port of a battery, or an AC / DC power supply. For example, the power connector 103 can be male, while the connector of the corresponding battery or the expansion port is female.

[0151] In one embodiment, Fig.21B, the controller 30 includes two power connectors, for example, a first power connector 103 and a second power connector 104. As described above, the first power connector 103 can be electrically coupled to a first battery, to a first expansion port of the first battery, or to a first AC / DC power supply, and the second power connector 103 can be electrically coupled to a second battery, to a second expansion port of the second battery, or to a second AC / DC power supply. In this embodiment, the first power connector 103 and the second power connector 104 can both be convex. In addition, the controller 30 includes a circuit system for switching between power sources so that energy is selectively transmitted to the controller 30 from at least one of the first or second battery / power supply. For example, the circuit system can switch between the first battery and the second battery intermittently or after the remaining power of one of the batteries reaches a predetermined threshold.

[0152] Reference now Fig. 21C 4 to 5, illustrate several configurations in which the controller 30 is directly electrically coupled to the battery 40, so that the patient can wear the controller 30 and the battery 40 together, for example, via a handbag, backpack, or holster. Fig. 21C As shown in Fig.21A The controller 30 can be electrically coupled to the battery 40 via the power connector 103, where the power connector 103 is male and the battery 40 has a corresponding female connector. Fig.21D The diagram illustrates the controller 30 electrically coupled to a battery 40, wherein the battery 40 is of a smaller size, and therefore has a lower capacity, and Fig.21E The controller 30 is illustrated electrically coupled to a battery 40, wherein the battery 40 is of a larger size, and therefore has a higher capacity. As will be appreciated by those skilled in the art, the battery 40 may be of various sizes depending on the needs of the patient.

[0153] Reference now Fig.21F H, several configurations are illustrated in which the controller 30 is remotely electrically coupled to the battery 40 so that the weight and volume of the controller 30 and the battery 40 are distributed and can be worn solely by the patient, for example, via a belt or vest. Fig.21F , the cable 41 that electrically couples the controller 30 to the battery 40 is electrically coupled to the first power connector port 105 via a spill 106, which is a hardwired junction between the cable 41 and the first power connector port 105. The power connector port 105 includes a power connector 107 that can be electrically coupled to the battery. For example, the power connector 107 can be male, while the connector for the corresponding battery is female.

[0154] like Figure 21G, the controller 30 can be remotely electrically coupled to the battery 40 via a cable 41. The cable 41 is electrically coupled to the controller 30 at one end via a second power connector port 108 and a spill 114 (which is a hardwired junction between the cable 41 and the second power connector port 108), and is electrically connected to the battery 40 at the other end via a first connector port 105 and a spill 106. For example, the power connector 103 of the controller 30 can be male, while the connector corresponding to the second power connector port 108 is female, and the power connector 107 of the first power connector port 105 can be male, while the connector corresponding to the battery 40 is female.

[0155] In one embodiment, Fig.21H As shown in FIG. 1 , the controller 30 can be remotely electrically coupled to multiple batteries, such as battery 40A and battery 40B, via a single second power connector port 108. Fig.21H , the second power connector port 108 includes a first spill 114A and a second spill 114B, so that the controller 30 is remotely electrically coupled to the battery 40A via the cable 41A and remotely electrically coupled to the battery 40B via the cable 41B. Specifically, the cable 41A is electrically coupled to the controller 30 via the second power connector port 108 and the first spill 114A at one end, and is electrically coupled to the battery 40A via the first connector port 105A and the spill 106A at the other end, and the cable 41B is electrically coupled to the controller 30 via the second power connector port 108 and the second spill 114B at one end, and is electrically coupled to the battery 40B via the first connector port 105B and the spill 106B at the other end. In this embodiment, the controller 30 may include circuitry for switching between the battery 40A and the battery 40B, so that energy is selectively transmitted to the controller 30 from at least one of the battery 40A and the battery 40B. For example, the circuit system may switch between battery 40A and battery 40B intermittently or after the remaining charge of one of the batteries reaches a predetermined threshold. Alternatively, controller 30 may receive energy from battery 40A and battery 40B simultaneously.

[0156] In another embodiment, if Fig.21I, for example, when the patient is resting at the bedside, the controller 30 is electrically coupled to the AC / DC power supply 109, which can be plugged into an electrical outlet via an AC plug 113. Specifically, the AC / DC power supply 109 is electrically coupled to the controller 30 via a cable 41, such that the cable 41 is electrically coupled to the controller 30 at one end via the second power connector port 108 and the spill port 114, and is electrically coupled to the AC / DC power supply 109 at the other end via the first power supply port 110. In addition, the AC / DC power supply 109 is electrically coupled to the plug 113 via the cable 112 and the second power supply port 111.

[0157] Controller 30 may include an internal battery such that the internal battery powers controller 30 and the implantable pump during the time required to replace and / or recharge battery 40. Thus, controller 30 may include circuitry for switching between power sources such that energy is transferred from the internal battery to controller 30 when battery 40 is disconnected from controller 30, and from battery 40 to controller 30 when battery 40 is electrically coupled to controller 30. Additionally, the circuitry may allow battery 40 to charge the internal battery while also supplying energy to the implantable pump until the internal battery is recharged to a desired amount, at which point the circuitry allows battery 40 to supply energy to the implantable pump alone. Similarly, when controller 40 is electrically coupled to AC / DC power supply 109, the circuitry may allow AC / DC power supply 109 to charge the internal battery while also supplying energy to the implantable pump until the internal battery is recharged to a desired amount, at which point the circuitry allows AC / DC power supply 109 to supply energy to the implantable pump alone.

[0158] According to some aspects of the present invention, systems and methods are provided for controlling an implantable pump (e.g., implantable pumps 20, 20", 20"' and 20"") constructed according to the principles of the present invention without the need for position, velocity or acceleration sensors. Specifically, an exemplary controller for an implantable pump (e.g., controller 30) can rely solely on current measurements of the actuator. The controller is robust to pressure and flow changes within the pump head and allows the operating point of the pump to change rapidly. For example, the controller includes a two-stage nonlinear position observer module based on a reduced order model of the electromagnetic actuator. Since the performance requirements of the actuator are very few, a linear approximation of the equivalent circuit is not sufficient. In order to meet the required operating range of the controller, the controller includes parameter changes of the relevant state variables. The means for identifying the model of the actuator is given by recursive least squares (RLS), so the means can be incorporated into the position observer module of the controller in a sensible manner. A forgetting factor is further included in the RLS to capture the model parameter changes of the relevant state variables.

[0159] Reference now Fig. 22 , a flow chart illustrates the steps of an exemplary method 2200 for controlling an implantable pump (e.g., implantable pumps 20, 20", 20"', and 20') constructed according to the principles of the present invention. First, a finite element method (FEM) model of an electromagnetic actuator (e.g., electromagnet assembly 91 and magnetic ring assembly 76) is transformed into a lumped parameter model represented by a system of ordinary differential equations (ODEs). The FEM model is set up by creating a subset of the actuator's geometry to save computational time, such as Fig.23 At step 2202, the same energy W of the implantable blood pump system is calculated for various magnetic ring positions and coil currents, such as Fig.24A For example, the isoenergy W can be approximated by a lookup table storing the output isoenergy W values ​​simulated by the FEM model.

[0160] At step 2204, partial derivatives of the same energy W are calculated and identified as parameters of the equivalent circuit, which is expressed as:

[0161]

[0162] in:

[0163]

[0164]

[0165] The equation of motion for one degree of freedom of the magnetic ring of the implantable pump is given by:

[0166]

[0167] in:

[0168]

[0169] F springs =ax 3 +bx

[0170] V in =Input voltage

[0171] x = magnetic ring position

[0172] I = coil current

[0173] R = coil resistance

[0174] L = coil inductance

[0175] E = Anti-EMF Factor

[0176] Fig. 24B , 24Cand 24D respectively illustrate that the force derived from the same energy W varies with the position of the magnetic ring and the coil current, the inductance of the equivalent circuit derived from the same energy W varies with the position of the magnetic ring and the coil current, and the EMF coefficient of the equivalent circuit derived from the same energy W varies with the position of the magnetic ring and the coil current. Fig.25 is a graph illustrating the relationship between spring reaction force and the position of a magnetic ring of an implantable pump. Fig.26 Schematic diagram of the parameters of the equivalent circuit and the equation of motion for one degree of freedom of the magnetic ring of the implantable pump described above.

[0177] Taking into account the design-induced nonlinearity measured by the manufacturer of the electromagnetic actuator, the spring reaction force F springs Identify as a cubic polynomial. Assume that the membrane force F membrane is finite and piecewise continuous. This ambiguous description of the membrane force is due to the lack of a full understanding of the fluid-structure interaction occurring between the membrane of the pump and the fluid and the possibility of synthesizing a controller that would not require more assumptions about this force than are given.

[0178] At step 2206, the controller operates the electromagnetic actuator of the implantable pump to cause the moving component (e.g., magnetic ring assembly 76) to reciprocate at a predetermined stroke (e.g., frequency and amplitude). At step 2208, the controller receives a signal indicative of the system's alternating current (e.g., coil current) from, for example, a current sensor positioned inside the power electronics of the implantable pump system.

[0179] For example, if Fig.27A As illustrated in FIG. , an implantable blood pump, which can be considered as an inductive load, can be driven with the help of an H-bridge configuration. Fig.27A As illustrated in the schematic diagram of the power electronics of , the power electronics may include an H-bridge 130 , a shunt resistor 132 , a current sensor 140 , and an optional voltage sensor 150 . Fig.27B The H-bridge illustrated in FIG. 1 is driven to produce a specific voltage waveform when powered directly or through a DC / DC voltage converter from a power supply, such as a battery 40. As will be appreciated by those skilled in the art, the power electronics device may include a single H-bridge with two coils connected in series or in parallel or two H-bridges with one H-bridge per coil.

[0180] Fig.27C is a schematic diagram of a current sensor for measuring the current of an actuator. Fig.27D is a schematic diagram of an optional voltage sensor used to measure the voltage of the actuator. The use of voltage feedback is optional, given that the algorithm controls the H-bridge and therefore knows the applied voltage. Fig.28As illustrated in FIG, the analog-to-digital converter (ADC) sampling is synchronized with the midpoint of the transistor pulse width modulation (PWM) signal to remove the transistor switching glitch noise from the current and voltage measurements. Thus, the current and / or pump voltage measurements are sent to an algorithm running on the controller, and the algorithm estimates the position of the actuator and determines the H-bridge voltage required to impose a specific position oscillation.

[0181] Specifically, based on the current measurement, the controller can control the excitation of the deformable membrane (e.g., membrane 97) while controlling the unmodeled force F for the deformable membrane. menbrane is robust. Therefore, the implantable pump system may not require position, velocity, or acceleration sensors. For example, the controller includes a position observer module with two stages. During the first stage (step 2210), the position observer module estimates the velocity of the magnetic ring based on the AC measurement and the parameters of the equivalent circuit using the equation described above:

[0182]

[0183] For example, the estimated speed can be expressed as:

[0184]

[0185] If left as is, the derivative in the above equation will make the estimate extremely sensitive to measurement noise. To handle this estimation problem, a derivative estimator was developed:

[0186]

[0187] where T is the length of the integration window. This estimate is directly implemented as a discrete finite impulse response (FIR) filter using the trapezoidal method:

[0188]

[0189] Where N is selected so that T = NT s , And w i =T s , i=1,...N-1 integers.

[0190] Next, the second stage of the position observer module is implemented (step 2212), where the position observer module determines the speed of the magnetic ring based on the estimated speed calculated during step 2210. For example, it is inferred that: if and is the observed position and speed, then the position observer module can be expressed as:

[0191]

[0192] Where A is a constant matrix that regroups the linear terms of the estimated velocity and F(t) is a function that regroups the nonlinear elements, and k1 and k2 are two gains selected to ensure the following equation:

[0193]

[0194] At step 2214, the position observer module of the controller determines the position of the magnetic ring based on the determined velocity described above. Thus, based on the observed position of the magnetic ring, the stroke controller will be able to set the excitation of the deformable membrane via the electromagnetic actuator to the desired frequency and amplitude while limiting overshoot. Thus, at step 2216, the controller cancels the error due to the unmodeled dynamics of the implantable pump to limit overshoot. For example, if Fig.29 As illustrated in FIG. 1 , the controller may include a feedforward module and a PI controller module. The feedforward module converts the desired position x at each time step into d is considered as input to calculate the input voltage as:

[0195]

[0196]

[0197] Among them I d can be calculated as:

[0198] I d =Φ(x d , F mag )

[0199] Generate reference signal x d As:

[0200]

[0201]

[0202]

[0203] where k f is a positive real number that guarantees the stability of H(s).

[0204] The PI controller module then cancels the remaining error due to the unmodeled dynamics by adjusting the excitation signal. This can be implemented using various methods. For example, its instantaneous value can be modified directly, or alternatively, another method is to modify its amplitude, or both its amplitude and its instantaneous value, on a different feedback loop. If amplitude modification is used, one way to estimate it is to define an amplitude estimator that is valid if x(t) is sufficiently close to a sine function

[0205]

[0206] At step 2218, the controller adjusts the operation of the electromagnetic actuator to cause the magnetic ring to reciprocate at the adjusted frequency and / or amplitude, thereby causing the deformable membrane to produce an adjusted predetermined blood flow across the implantable pump.

[0207] To capture the variations in inductance and back EMF coefficients with the position of the magnetic ring and the coil current, the controller uses a recursive least squares estimator.

[0208] The parameters R, L and E described above are unknown and vary slowly over time. in , I and x are piecewise continuous and finite, and are all equal to zero at t = 0. The problem is set by integrating (1) over t:

[0209]

[0210] It can be expressed as:

[0211] y=Ψ T θ

[0212] Ψ T =[∫V in ∫I x]

[0213]

[0214] For each sample, n>0:

[0215]

[0216]

[0217] K n =Ψ n Q n

[0218]

[0219]

[0220] Where λ is the chosen forgetting factor, so λ<1, P o is the initial covariance matrix, and is the initial estimate of the parameters.

[0221] The resulting estimated data is then fit with a polynomial of appropriate degree and stored into a lookup table that associates the corresponding inductance with the emf factor for each (x, I) combination. The lookup table is used in the speed estimator:

[0222]

[0223] Experimental Results

[0224] A numerical model of the implantable pump and controller was built under Matlab / Simulink to test the implementation of the controller and the identification of the model parameters. The actuator model was compared with the measurements and adjusted accordingly. The reaction force of the spring was measured by using a tension tester, which was also used to measure the magnetic force of the actuator by applying an arbitrary constant current to the electromagnetic coil of the actuator while measuring the force. The back EMF coefficient was derived based on the force measured at different currents and magnetic ring positions. When the movement of the magnetic ring is blocked to cancel the effect of the back EMF, the inductance and resistance can be estimated with the help of an LRC meter. When the input current of the LRC meter is limited (<20mA), the inductance can be estimated only in this limited area. In general, the performance of the actual actuator is reduced compared to the model (lower inductance, magnetic force and EMF). This can be attributed to imperfect manufacturing processes, such as the winding of the coil. The film force can be simulated by the viscous friction term, which is a reasonable first approximation:

[0225]

[0226] With these validations, variations of parameters are identified and controller structures are tested. In particular, different position observer implementations are compared to demonstrate the interest in using varying electrical parameters rather than linear approximations.

[0227] exist Figures 30A to 30C The results of the identification are shown in Figure 1. To ensure fast convergence, two excitation signals are applied to the actuator. The voltage excitation consists of a high frequency (500 Hz) square wave voltage so that the voltage of the inductor is never close to zero, and the low frequency (0.1 Hz) sinusoidal component of the voltage of the resistor arrives at each position. To ensure that the back EMF is represented in the response, the external sinusoidal force is simulated at 50 Hz. In order to filter out high frequency variations and eventually noise while capturing low frequency variations of the parameters, λ is set to 0.999 via trial and error. Recursive least squares identification is run with different initial conditions. Measurement errors (noise, bias, gain) are simulated to verify that their effects do not hinder convergence and are helpful for diagnosing problems in future experiments. The RLS algorithm filters out high frequency noise very easily, but gain errors lead to overestimation or underestimation of the parameters, while bias and low frequency noise increase the estimation error over time.

[0228] A discrete version of the controller is implemented on Simulink to mimic the operation that would be accomplished by compiling it on the hardware target. Since the frequency response of the derivative estimator described above depends on the length of the integration window and the sampling rate, the signal to be derived can have a frequency of up to 100 Hz. We set T s =2.10-5 s and N = 6 and (ie, 1, 2.10 -4 s), which is a good compromise between noise attenuation and performance. Fig.31A and 31B The actuator's response to start-up at t=0s is shown to a nominal constant operating point (ie, constant amplitude and frequency). Both current and magnetic ring position are reasonably sinusoidal, and after a transition period, the amplitude of position reaches the desired amplitude, and the position observer module output tracks the position change.

[0229] Fig.32A and 32B Two examples of operating point changes are shown: frequency change and stroke change, which are two ways to increase or decrease blood flow through the pump. Fig.32A and 32B As shown in , overshoot occurs during stroke changes. If not kept below a safe level, the overshoot may create overstress that may damage the membrane and spring. Overshoot can be avoided by making the desired stroke signal change smoother.

[0230] Figures 33A to 33D is a comparison between the 4 different position observer modules in its first stage. In each case, the inductance and back EMF are implemented as constant approximations and as functions. When the controller must maintain the stroke over a wide range of strokes and frequencies, and when the different reaction forces of the membrane are unknown, the combination of these three parameters must be tested to evaluate the performance of the controller. In this way, the variation of the membrane force is simulated according to the flow and pressure inside the pump head by varying μ, and an error variable e evaluated over a range of strokes, frequencies and μ is formed:

[0231]

[0232] Here max x & min x are calculated based on one oscillation period. This formula for e can be compared with the maximum permissible error ε: Each operating point [S d , f d ], which exhibits e<ε (the stroke will be maintained to deliver the required flow without risk of damaging the device by overshoot). With this performance indicator it is observed that accounting for changes in inductance and back EMF in the speed estimator results in an increase in the operating range of the stroke controller.

[0233] Reference now Fig.34, providing an exemplary multi-stage control of a controller that relies on position measurement and is constructed according to the principles of the present invention. Specifically, the magnetic ring position can be measured via one or more sensors and used to control the operation of the pump. For example, a sensor such as a Hall effect sensor can be coupled to a component that remains fixed relative to the stator assembly (e.g., the stator assembly itself or the housing of the pump). In addition, a small permanent magnet can be coupled to any moving component of the pump that moves relative to the stator assembly, such as a magnetic ring assembly or a rigid film ring. Therefore, the strength of the magnetic field generated by the permanent magnet can be measured via the Hall sensor during the operation of the pump. For example, as the permanent magnet moves closer to the Hall sensor, the strength of the magnetic field generated by the permanent magnet and measured by the sensor increases.

[0234] The above-described magnetic ring position measurement and the current of the magnetic coil can be used to form Fig.34 A control scheme based on pump dynamics is illustrated in . First, the appropriate trajectory can be based on the inverse dynamics of the pump system.

[0235] A suitable trajectory can be written as:

[0236]

[0237] The dynamics of the pump can be written as:

[0238]

[0239] where F a It is the magnetic force generated by the actuator plus the reaction force of the suspension spring. Given the desired stroke amplitude S d and frequency f d , it is possible to generate a feasible trajectory for the pump. If the desired position x can be derived twice with respect to time d And those derivatives v d and is continuous, then this trajectory exists. a When (x, I) is a diffeomorphism, there exists a mapping φ a (x, Fd), for example:

[0240]

[0241]

[0242] exist and In the case of m, the Kalman filter block can be used to calculate m f and two estimates of μ. In addition, the feedforward voltage V required for the pump to follow the desired trajectory ff For each k>0 (each number is related to time, such as t=kT s , T sis the sampling period of the controller) is calculated as:

[0243]

[0244] Since any model has errors, the feedback voltage alone is not sufficient to follow the trajectory accurately enough. Therefore, to complete Fig.34 The control scheme adds the feedback voltage, which can be written as:

[0245]

[0246] Where K = [k1 k2 k3 k4] and tracking error e k ,For example and z k =z k-1 +T s (x d -x k ). and is x calculated by the extended Kalman filter and An appropriate choice for the matrix K may be selected, for example, by using a linear quadratic integration method applied to a linear approximation of the pump dynamics.

[0247] Given the measurements of x and I corrupted by noise, the goal of the extended Kalman filter is to compute the variables x and And the time-varying parameters μ(t) and m f An accurate estimate of (t).

[0248] Using the dynamics of the pump discretized using the Euler method, the estimated variables and parameters can be written in vector form as:

[0249]

[0250] When measuring position x k When x k =CX k , where C =

[1000] . For each k>0, we have:

[0251]

[0252]

[0253] in is a vector The predicted estimate and P k+1|k is the forecast covariance matrix. J k is the Jacob matrix:

[0254]

[0255] And Q is composed of 4 diagonal terms q 1，...， The process covariance matrix composed of 4>0 is:

[0256]

[0257] In the case of measurement covariance R>0, the correction gain L k+1 , Corrected Estimate And the covariance matrix P k+1|k+1 is calculated as:

[0258] L k+1 =P k+1|k C T (CP k+1|k C T +R) -1

[0259]

[0260]

[0261] Additionally, blood flow out of the inlet through the outlet of the pump may be estimated based on the Kalman filter estimate of μ described above. For example, when it has been empirically demonstrated that the change in μ is proportional to the pump flow rate q given the desired stroke and frequency f When there is a strong correlation between the changes in f can be written as a function such as:

[0262]

[0263] From this function, an estimate of the pump flow rate can be calculated, which does not require a flow sensor and can be used to set the operating point of the pump.

[0264] In addition, the Kalman filter estimation described above can be used to detect faults in the pump system. Specifically, the estimated residual of the Kalman filter can be used to always monitor the operating conditions of the pump and any changes that must be caused by a fault can be detected almost immediately. Therefore, for each k>0, the residual εk can be calculated as:

[0265]

[0266] Therefore, we can run k Statistical analysis of ε can be performed to detect anomalies. For example, ε can be calculated for any number of samples. k In nominal operating conditions, the mean of the residuals should be close to zero. Therefore, any time the value is determined to be above or below a threshold, a mechanical fault can be detected. In addition, abnormal operating conditions can be detected by comparing the derivative of the estimated value with the parameter having its expected nominal value.

[0267] Reference now Fig.35 , another exemplary sensorless multi-level control of a controller constructed according to the principles of the present invention is provided. Fig.35 The control scheme is similar to Fig.34 The control scheme is constructed by Fig.35 The Kalman filter does not rely on magnetic ring position measurements. Instead, Fig.35 The Kalman filter relies on velocity estimation

[0268] Using the dynamics of the pump discretized using the Euler method, the estimated variables and parameters can be written as the following vector

[0269]

[0270] When the speed observer module estimates the speed hour, Where C =

[1000] . For each k>0:

[0271]

[0272]

[0273] in is a vector The predicted estimate and P k+1|k is the forecast covariance matrix. J k is the Jacobian matrix:

[0274]

[0275] And Q is composed of 4 diagonal terms q 1，...， The process covariance matrix composed of 4>0 is:

[0276]

[0277] In the case of measurement covariance R>0, the correction gain L k+1 , Corrected Estimate And the covariance matrix P k+1|k+1 can be calculated as:

[0278] L k+1 =P k+1|k C T (CP k+1 | k C T +R) -1

[0279]

[0280]

[0281] The speed observer block can be calculated in various ways. For example, the speed observer block can be calculated as:

[0282]

[0283] It can be approximated by a first-order finite difference method:

[0284]

[0285] Alternatively, can be calculated as:

[0286]

[0287] N is selected so that T = NT s , And w i =T s , i=1,...N-1 integers.

[0288] Alternatively, the speed observer block can be calculated using a set of nonlinear equations, such as:

[0289]

[0290]

[0291]

[0292] Reference now Fig.36 , provides another exemplary sensorless multi-level control of a controller constructed according to the principles of the present invention. Fig.36 As shown in Fig.34 The estimated position of the position observer is used in the extended Kalman filter. Fig.36 The estimated speed, position and parameters are used in the control scheme.

[0293] Although various illustrative embodiments of the present invention are described above, it will be apparent to those skilled in the art that various changes and modifications may be made herein without departing from the present invention. Fig. 9 The pump assembly 70 shown in FIG. 1 may be customized differently and may include additional or fewer components of various sizes and compositions. The appended claims are intended to cover all such changes and modifications as come within the true spirit and scope of the present invention.

Claims

1. An implantable blood pump system, comprising: An implantable blood pump configured to be implanted in a patient's heart, the implantable blood pump comprising: a housing having an inlet and an outlet; a deformable membrane disposed in the housing; and an actuator comprising a fixed component and a moving component coupled to the deformable membrane, the actuator being configured to be powered by an alternating current that causes the moving component to reciprocate relative to the fixed component at a predetermined frequency and amplitude, thereby causing the deformable membrane to generate a predetermined blood flow from the inlet outward through the outlet; and a controller operatively coupled to the implantable blood pump, the controller being configured to be programmed to: operating the actuator to cause the moving component to reciprocate relative to the fixed component at the predetermined frequency and amplitude; receiving a signal indicative of the alternating current via a current sensor operatively coupled to the controller; determining a position of the moving component by estimating a speed of the moving component based on the signal indicative of the alternating current; and adjusting the operation of the actuator based on the position of the moving component to cause the moving component to reciprocate relative to the fixed component at an adjusted predetermined frequency and amplitude, thereby causing the deformable membrane to generate an adjusted predetermined blood flow from the inlet outward through the outlet, The controller is programmed to estimate the speed of the moving component based on various positions of the moving component and an isotropic energy W value of a finite element model (FEM) of alternating current. 2 . The system of claim 1 , wherein the controller is programmed to determine the position of the moving component by determining the speed of the moving component based on an estimated speed of the moving component.

3. The system of claim 1, wherein the controller is programmed to adjust operation of the actuator to cause the moving component to reciprocate relative to the fixed component at the adjusted predetermined frequency and amplitude while limiting overshoot.

4. The system of claim 3, wherein the controller comprises a PI controller configured to be programmed to limit overshoot by canceling errors due to unmodeled dynamics of the implantable blood pump.

5. The system of claim 1, wherein the controller is programmed to determine the position of the moving component based on the signal indicative of the alternating current and changes in inductance and back EMF coefficients.

6. The system of claim 1, wherein the adjusted predetermined blood flow comprises pulses synchronized with the patient's heartbeat.

7. The system of claim 1 , wherein the fixed component comprises an electromagnetic assembly configured to generate a magnetic field, and wherein the moving component comprises a magnetic ring suspended concentrically about the electromagnetic assembly and configured to reciprocate over the electromagnetic assembly at the predetermined frequency and amplitude in response to the magnetic field.

8. The system of claim 7, wherein the electromagnetic assembly comprises first and second electromagnetic coils, and wherein the magnetic ring is caused to move when at least one of the first or second electromagnetic coils is powered by the alternating current.

9. The system of claim 7, wherein the magnetic ring is configured to induce wave-like deformation in the deformable membrane by reciprocating over the electromagnetic assembly.

10. The system of claim 1, further comprising first and second suspension rings concentrically disposed about and coupled to the fixed component and the moving component.

11. The system of claim 10, wherein the moving assembly is coupled to the deformable membrane and each of the first and second suspension rings via a plurality of rods.

12. The system of claim 10, wherein the first and second suspension rings permit reciprocating movement of the moving component relative to the fixed component.

13. The system of claim 10, wherein the first and second suspension rings apply a spring force to the moving assembly when the moving assembly reciprocates relative to the fixed assembly.

14. The system of claim 1, further comprising a rigid ring coupled to the moving assembly, the rigid ring further coupled to the deformable membrane.

15. The system of claim 1, wherein a bottom surface of the actuator and an interior portion of the housing adjacent the outlet form a flow channel, the deformable membrane being suspended within the flow channel.

16. The system of claim 15, wherein the deformable membrane includes a central orifice adjacent the outlet.

17. The system of claim 15, wherein the actuator and an interior surface of the housing adjacent the inlet form a delivery channel extending from the inlet to the flow channel.

18. The system of claim 1, further comprising a rechargeable battery configured to deliver the alternating current to power the implantable blood pump.

Images