Centrifugal LVAD with wireless power transmission and antithrombotic SLIC coating

Abstract

An exemplary left ventricular assist device (LVAD) and method employing design and material design to reduce the risk of thrombosis of LVAD implants in patients. One feature of the design includes a stent-implanted inlet member for the device that can reduce or completely eliminate flow stagnation that may induce blood protein adsorption, which may then trigger a chain reaction leading to thrombosis. Other features of the design include (i) a flexible rotor and / or pump housing that can reduce blood damage, (ii) a hydrophilic slippery coating on high-shear components such as the rotor to further reduce the risk of protein adsorption and thus thrombosis, (iii) a magnetic-based drive and bearing component that can improve the blood compatibility of the blood pump, and (iv) advanced control and charging.

Description

[Technical Field]

[0001] Related applications This PCT application claims priority and interest to U.S. Provisional Patent Application No. 63 / 505,873, filed on 2 June 2023, entitled "CENTRIFUGAL LVAD WITH WIRELESS POWER TRANSFER AND ANTITHROMBOTIC SLIC COATING," which is incorporated herein by reference in its entirety. [Background technology]

[0002] When cardiovascular disease and heart failure (HF) reach their final stages, there are no known drugs that are effective in treating them. Left ventricular assist devices (LVADs) are often used to help a failing heart maintain blood circulation to the body when a donor heart is unavailable. Common complications of LVAD devices are known to include post-implantation infections, thrombosis, stroke, and bleeding, any of which can significantly reduce the quality of life for LVAD patients.

[0003] Blood damage caused by mechanical forces from LVAD devices is known to cause side effects, including bleeding and thrombosis, requiring readmission and emergency surgery. Blood damage is often associated with hyperphysiological shear stress in LVAD devices from mechanical forces that destroy blood cells, cause hemolysis, activate platelets, and lead to thrombosis. There is currently one LVAD device on the market.

[0004] Improving left ventricular assist devices has advantages. [Overview of the Initiative]

[0005] Exemplary left ventricular assist devices (LVADs) and methods employing design and material designs to reduce the risk of thrombosis with LVAD implants in patients are disclosed. One feature of the design includes a stent-placed inlet member for the device that can reduce or completely eliminate flow stagnation that may induce blood protein adsorption, which can then trigger a chain reaction leading to thrombosis. Thrombosis is the formation of a thrombus in a blood vessel or heart that can effectively block or guide blood flow. Another feature of the design includes a flexible rotor and / or pump housing that can reduce blood damage. The design has been observed to improve pump efficiency hydrodynamically. Another feature of the design employs a hydrophilic, slippery coating on high-shear components such as the rotor to further reduce the risk of protein adsorption, and therefore thrombosis. Another feature of the design employs magnetically based drive and bearing components that can improve the blood compatibility of the blood pump. Each and the combined features considered can reduce the associated risks of LVADs, making them more accessible to more severely ill patients and altering treatment options for cardiovascular disease and heart failure (HF).

[0006] Exemplary LVADs may further include system improvements and control enhancements that leverage advanced control and electronics as well as advanced charging and energy management controls to improve device usability, reduce the likelihood of premature failure, and enhance device operation. In some embodiments, wireless power transmission and communication submodules may be implemented to eliminate percutaneous drive lines associated with serious infections. Various mechanical, electrical, material, and control system features may be used, in whole or in part, in new LVADs or in existing commercially available LVAD designs. LVADs using some or all of the mentioned features may (i) support patients for longer periods with less anticoagulation therapy, and (ii) operate with lower risk for emergency surgery to treat complications. In particular, an LVAD optimized using all of the mentioned features may potentially serve as a viable treatment comparable to existing heart transplants.

[0007] Exemplary AI-based algorithms and analyses are disclosed that employ state-of-the-art deep learning algorithms to analyze the relationship between different geometric shape parameters in design space and blood damage. AI-based algorithms can be used to assess patients for suitability for LVAD devices or treatments.

[0008] In addition to LVADs, various mechanical, electrical, material, and control system features can be used in other medical implants and devices. Examples of such devices include, but are not limited to, extracorporeal membrane oxygenation (ECMO) and cardiopulmonary bypass (CPB) machines. Cardiopulmonary bypass is a machine that temporarily takes over the functions of the heart and lungs during cardiac surgery by maintaining the circulation of blood and oxygen throughout the body. Extracorporeal membrane oxygenation is a form of extracorporeal life support device that provides long-term cardiac and respiratory support to people whose hearts and lungs cannot provide enough oxygen, gas exchange, or blood supply to sustain life.

[0009] Antithrombotic SLIC coatings can be used on blood-contact devices such as catheters, guidewires, dialyzers, oxygen dispensers, cardiac support systems, cardiac pacemakers, vascular grafts, stents, and heart valves to reduce the risk of thrombotic events. In some embodiments, less invasive treatments may be used. Exemplary SLIC coatings can reduce thrombosis in blood-contact devices and materials, as well as in fouling-free coated devices, for example, in contact lenses, chemical and biomedical sensors, food and beverage equipment, and marine equipment.

[0010] In one embodiment, an implantable left ventricular assist device (LVAD) is disclosed, comprising: a pump housing that forms a volume therein, the pump housing having an inlet and an outlet; a rotor disposed within the pump housing to rotate within the volume for pumping blood; a motor coupled to the rotor to drive the rotation of the rotor; and an inlet member coupled to the inlet of the housing, the inlet member having (i) an external structure configured to conform to and contact the outer wall at the base of the left ventricle; and (ii) an internal expandable body configured to move between a housing configuration and a deployable configuration that is located within and extends into the left ventricle, the internal expandable body being defined by (i) a first section having a first circumference sized to conform to the inlet of the pump; and (ii) a second section having a second circumference sized to substantially contact the inner wall of the left ventricle in the deployable configuration to eliminate flow stagnation within the left ventricle.

[0011] In some embodiments, the rotor comprises a rotor body and a plurality of curved flexible blades extending therefrom, the plurality of curved flexible blades being flexibly molded and formed from a flexible material (e.g., polyurethane rubber) to reduce shear stress on the components of the blood (and also improve pump efficiency, for example).

[0012] In some embodiments, the rotor includes an antithrombogenic coating (e.g., configured to reduce protein adsorption and thrombosis risk) that includes a hydrophilic and slippery polymer.

[0013] In some embodiments, the device further includes a drive system configured to drive the rotor to rotate within the volume of the pump housing, and an embedded control unit electrically coupled to the drive system to control the rotation of the drive system and the rotor. The embedded control unit includes a wireless power transmission circuit that includes one or more energy storage devices, a charging circuit, and electronics for controlling an inductive charging operation using an antenna coil and an external RF source (e.g., the transmission circuit does not allow a percutaneous drive line that could cause infection).

[0014] In some embodiments, the motor includes a magnetic levitation drive subassembly that includes one or more permanent magnets (e.g., radially passive magnets located in the inlet cannula portion of the rotor body) located within the rotor body, a stator disposed at a position in the pump housing proximate to the one or more permanent magnets (e.g., the stator includes permanent magnets, and the permanent magnets of the stator and the one or more permanent magnets within the rotor body are passively stabilized in the radial direction and inclination by the repulsive force between the permanent magnets of the rotor and the stator), and one or more magnetic bearings (passive magnetic levitation bearings located at the bottom of the VAD housing) disposed at one or more positions in the pump housing proximate to the plurality of curved flexible blades of the rotor.

[0015] In some embodiments, the inner expandable body of the inlet member includes a flexible mesh surrounded by or embedded in a fabric, and the flexible mesh is configured as self-expandable or balloon-expandable.

[0016] In some embodiments, the inner expandable body of the inlet member elutes an antithrombogenic and anti-restenosis drug.

[0017] In some embodiments, after deployment, the internally expandable body is defined by (i) a first section having a first circumference sized to couple to the inlet of the pump, and (ii) a second section having a second circumference sized to continuously contact and conform to the inner wall surface of the left ventricle and to conform to the shape of the left ventricle.

[0018] In some embodiments, the internally expandable body (and outer surface) is configured to deform with ventricular motion over a predetermined set of deformation cycles (e.g., to resist failure caused by fatigue).

[0019] In some embodiments, the fabric includes an absorbent material (e.g., absorbable over time to leave only new tissue) to promote endothelialization.

[0020] In some embodiments, the internally expandable body includes an absorbent material (e.g., absorbable over time to leave only new tissue) to promote endothelialization.

[0021] In some embodiments, the fabric includes polyester (e.g., PET such as Dacron), polytetrafluoroethylene (e.g., expanded polytetrafluoroethylene (ePTFE)), biological tissue (e.g., pericardial tissue), or combinations thereof.

[0022] In some embodiments, the internally expandable body, including the flexible mesh and fabric, is customized to be sized and shaped to match a scan of the patient's left ventricle (e.g., the scan is computationally modeled and a portion of the internally expandable body, e.g., the flexible mesh, is 3D printed to accommodate complex anatomical variations of the ventricle).

[0023] In some embodiments, the pump housing and / or inlet member are provided with a quick-connect connector for removably and reattachably connecting the pump housing to the inlet member (for example, the device can be installed and deployed with the pump housing disconnected from the inlet member, and the inlet member can be capped as needed).

[0024] In some embodiments, the pump housing comprises (i) an external structure and (ii) an internally deformable structure disposed within a portion of the surface defining the volume of the pump housing, wherein the internally deformable member includes a flexible material.

[0025] In some embodiments, the flexible material of the internally deformable member is the same as the flexible material of the rotor.

[0026] In some embodiments, the flexible material of the internally deformable member is different from the flexible material of the rotor.

[0027] In some embodiments, a portion of the surface defining the volume of the pump housing having a flexible material corresponds to the area of ​​contact with the multiple curved flexible blades during their rotation.

[0028] In some embodiments, a portion of the surface defining the volume of the pump housing having a flexible material corresponds to the area of ​​contact with the multiple curved flexible blades during (i) the rotation of the multiple curved flexible blades and (ii) the off-axis movement of the multiple curved flexible blades, which is configured to move off-axis via magnetic bearings and an electromagnetic drive subsystem.

[0029] In some embodiments, the external structure is made of a flexible material.

[0030] In some embodiments, the inner surface of the pump housing, which includes an internally deformable structure, includes an antithrombotic coating containing a hydrophilic and slippery polymer (for example, configured to reduce protein adsorption and thrombosis risk).

[0031] In some embodiments, the antithrombotic coating on the interface surface of the pump housing and / or the rotor is optimized for maximum antithrombotic response through systematic tailoring of hydroxylation parameters, molecular construction, and synthetic reaction coordinates.

[0032] In some embodiments, the antithrombotic coating on the interface surface of the pump housing and / or the antithrombotic coating of the rotor is formed by liquid-phase silanization of the flexible material on the interface surface of the pump housing and / or the antithrombotic coating of the rotor to form a silane (for example, by modifying the surface of the flexible material using oxygen plasma to achieve a high degree of substrate hydroxylation).

[0033] In some embodiments, the silane is formulated to (i) obtain a high graft density to the flexible material on the interface surface of the pump housing and / or to the antithrombotic coating of the rotor, and (ii) ensure antithrombotic properties.

[0034] In some embodiments, the coating is prepared by a synthesis reaction.

[0035] In some embodiments, the device further comprises an implantable subcutaneous coil, the wireless power transmission circuit of the implanted control unit being operably coupled to the implantable subcutaneous coil via a drive line conductor, and the implantable subcutaneous coil being configured to operate as a pair of coupling coils with an external coil configured to be mounted (i) in a patient's skin area and (ii) in close proximity to the implantable subcutaneous coil.

[0036] In some embodiments, one or more energy storage devices include (i) one or more rechargeable batteries and (ii) at least one of a supercapacitor and a hybrid supercapacitor, wherein one or more rechargeable batteries and at least one of the supercapacitor and hybrid supercapacitor have combined energy storage for at least 20 minutes (e.g., 30 minutes, 40 minutes, 50 minutes, 1 hour, 2 hours).

[0037] In some embodiments, the embedded control unit comprises (i) a processor and (ii) a memory having instructions stored thereon, wherein the execution of instructions by the processor causes the processor to execute a dynamic charging algorithm (for example, based on internal and external temperature sensors that pulse wireless energy transmission to control heating and charging levels).

[0038] In some embodiments, the embedded control unit includes a wireless communication interface (e.g., an edge device, cloud infrastructure, or a combination thereof) configured to operably connect with a remote controller.

[0039] In some embodiments, the remote controller includes (i) a processor and (ii) a memory having instructions stored thereon, wherein the execution of instructions by the processor causes the processor to execute a dynamic charging algorithm which includes a trained machine learning model or a model derived therefrom (for example, configured to monitor the current patient activity status to determine LVAD demand and immediate flow settings), and thereon it causes the output of the dynamic charging algorithm to be sent to an embedded control unit to coordinate the control operation.

[0040] In some embodiments, the embedded control unit is configured to execute a dynamic charging algorithm.

[0041] In some embodiments, the dynamic charging algorithm (e.g., locally within and / or externally within the LVAD) includes a dynamic motor control loop that can reduce the LVAD speed during times of less patient activity, and the dynamic motor control loop has an output for driving the motor's operation.

[0042] In some embodiments, the device further comprises a sensor configured to acquire electrical signals from the heart (e.g., ECG signals, waveforms, etc.), and the dynamic motor control loop includes one or more inputs, including a first input for receiving the acquired electrical signals.

[0043] In some embodiments, the pump housing includes a position sensor for the rotor (e.g., an eddy current sensor, e.g., configured to measure the gap between the rotor, e.g., the impeller tip and the pump housing), and one or more inputs to the dynamic motor control loop include a second input for receiving electrical signals acquired from the position sensor.

[0044] In some embodiments, trained machine learning algorithms are used to evaluate long-term LVAD performance data (e.g., to adjust settings).

[0045] In some embodiments, the device further comprises a magnetic levitation drive system including a magnetic levitation ventricular assist device (VAD), the VAD comprising an impeller attached to a rotor permanent magnet, a stator embedded in the VAD housing adjacent to the inlet cannula, and an active magnetic levitation bearing located at the bottom of the VAD housing with an eddy current sensor configured to monitor the gap between the impeller tip and the magnetic levitation device.

[0046] In some embodiments, an implantable left ventricular assist device is selected to be implanted in a patient identified via a trained ML algorithm used to evaluate candidate assessments.

[0047] In some embodiments, the trained ML algorithm is configured to estimate the likelihood of the presence of candidate criteria for LVAD implants (e.g., left fascicular branch block (LBBB), premature ventricular contraction (PVC), left ventricular hypertrophy (LVH), premyocardial infarction (AMI) and congestive heart failure (CHF), QRS duration, inferior myocardial infarction (IMI) and atrioventricular block (AVB), right ventricular hypertrophy (RVH), right fascicular branch block (RBBB), right atrial hypertrophy / overload (RAE), and combinations thereof).

[0048] In some embodiments, the stator forms a brushless DC motor (BLDC) with the rotor, and the stator and rotor diameters are optimized via numerical simulation (for example, to prevent flux saturation at operating speeds), and the stator length is optimized to generate motor torque that adequately satisfies hydraulic torque requirements.

[0049] In some embodiments, the system further comprises a magnetic levitation drive system including a magnetic levitation ventricular assist device (VAD), the VAD comprising an impeller attached to a rotor permanent magnet, a stator embedded in the VAD housing adjacent to the inlet cannula, and an active magnetic levitation bearing located at the bottom of the VAD housing with an eddy current sensor configured to monitor the gap between the impeller tip and the magnetic levitation device.

[0050] In some embodiments, the magnetic levitation ventricular assist device includes a magnetic bearing system (MBS) for controlling the rotation and translation of the rotor and impeller, with radially passive magnets fixed to the rotating impeller and permanent magnets mounted near the inlet cannula, and the magnets are configured such that the repulsive force between the rotor permanent magnets and the stator permanent magnets passively stabilizes the radial and inclined directions. [Brief explanation of the drawing]

[0051] [Figure 1]An exemplary left ventricular assist device, comprising a stent-inspired inlet member to reduce or completely eliminate flow stagnation, is shown according to an exemplary embodiment. [Figure 2A] An exemplary embodiment shows an example of an LVAD with a stent implanted or stent-like inlet member. [Figure 2B] An exemplary embodiment shows an example of an LVAD with a stent implanted or stent-like inlet member. [Figure 2C] An exemplary embodiment shows an example of an LVAD with a stent implanted or stent-like inlet member. [Figure 2D] An exemplary embodiment shows an example of an LVAD with a stent implanted or stent-like inlet member. [Figure 3A] Figure 1 shows an exemplary embodiment of a flexible rotor and / or pump housing for an exemplary LVAD, as well as design parameters. [Figure 3B] Figure 1 shows an exemplary embodiment of a flexible rotor and / or pump housing for an exemplary LVAD, as well as design parameters. [Figure 3C] Figure 1 shows an exemplary embodiment of a flexible rotor and / or pump housing for an exemplary LVAD, as well as design parameters. [Figure 3D] Figure 1 shows an exemplary embodiment of a flexible rotor and / or pump housing for an exemplary LVAD, as well as design parameters. [Figure 3E] Figure 1 shows an exemplary embodiment of a flexible rotor and / or pump housing for an exemplary LVAD, as well as design parameters. [Figure 3F] Figure 1 shows an exemplary embodiment of a flexible rotor and / or pump housing for an exemplary LVAD, as well as design parameters. [Figure 3G] Figure 1 shows an exemplary embodiment of a flexible rotor and / or pump housing for an exemplary LVAD, as well as design parameters. [Figure 4A] An exemplary embodiment shows a slippery hydrophilic (SLIC) surface in the exemplary LVAD of Figure 1 that can spread water droplets. [Figure 4B] An exemplary embodiment shows a slippery hydrophilic (SLIC) surface in the exemplary LVAD of Figure 1 that can spread water droplets. [Figure 4C] An exemplary embodiment shows a slippery hydrophilic (SLIC) surface in the exemplary LVAD of Figure 1 that can spread water droplets. [Figure 4D] An exemplary embodiment shows a slippery hydrophilic (SLIC) surface in the exemplary LVAD of Figure 1 that can spread water droplets. [Figure 5] An exemplary wireless energy transmission and communication system that can be implemented in the exemplary centrifugal LVAD of Figure 1 is shown according to an exemplary embodiment. [Figure 6] An exemplary LVAD comprising a drive system with magnetic levitation (also referred to as the “magnetic levitation system”) is shown according to an exemplary embodiment. [Figure 7A] This section outlines various characteristics of research related to heart rate synchronization velocity modulation. [Figure 7B] This section outlines various characteristics of research related to heart rate synchronization velocity modulation. [Figure 8A] This paper highlights various features of research on evaluating the hydraulic performance of flexible motors and impellers for LVADs. [Figure 8B] This paper highlights various features of research on evaluating the hydraulic performance of flexible motors and impellers for LVADs. [Figure 8C] This paper highlights various features of research on evaluating the hydraulic performance of flexible motors and impellers for LVADs. [Figure 8D] This paper highlights various features of research on evaluating the hydraulic performance of flexible motors and impellers for LVADs. [Figure 8E] This paper highlights various features of research on evaluating the hydraulic performance of flexible motors and impellers for LVADs. [Figure 9A]This paper highlights various aspects of research on the evaluation of slippery hydrophilic (SLIC) surfaces for LVADs. [Figure 9B] This paper highlights various aspects of research on the evaluation of slippery hydrophilic (SLIC) surfaces for LVADs. [Figure 10A] This paper highlights various features of research on the evaluation of wireless energy transmission and communication systems for LVADs. [Figure 10B] This paper highlights various features of research on the evaluation of wireless energy transmission and communication systems for LVADs. [Figure 10C] This paper highlights various features of research on the evaluation of wireless energy transmission and communication systems for LVADs. [Figure 10D] This paper highlights various features of research on the evaluation of wireless energy transmission and communication systems for LVADs. [Figure 11] This paper highlights various features of research on the evaluation of heart rate-synchronized velocity modulation control for LVADs. [Figure 12A] This document outlines various features of research on the evaluation of magnetic drive and bearing systems (MagLev) for LVADs. [Figure 12B] This document outlines various features of research on the evaluation of magnetic drive and bearing systems (MagLev) for LVADs. [Figure 12C] This document outlines various features of research on the evaluation of magnetic drive and bearing systems (MagLev) for LVADs. [Figure 12D] This document outlines various features of research on the evaluation of magnetic drive and bearing systems (MagLev) for LVADs. [Figure 13A] This paper highlights various features of research on the evaluation of trained AI models for assessing candidate recipients for LVADs. [Figure 13B] This paper highlights various features of research on the evaluation of trained AI models for assessing candidate recipients for LVADs. [Figure 13C]This paper highlights various features of research on the evaluation of trained AI models for assessing candidate recipients for LVADs. [Modes for carrying out the invention]

[0052] Several references, which may include various patents, patent applications, and publications, are cited in the bibliography and considered within the disclosures provided herein. Such citations and / or references are provided solely for the purpose of clarifying the description of the disclosed technology and do not constitute any acknowledgment that such references are “prior art” with respect to any aspect of the disclosed technology described herein. In terms of notation, "[n]" corresponds to the nth reference in the list. For example, [1] refers to the first reference in the list. All references cited and considered herein are incorporated herein by reference and to the same extent as if each reference were incorporated separately by reference.

[0053] Figures 1, 2A, and 2B show exemplary left ventricular assist devices 100 (shown as 100a and 100b), respectively, comprising stent-inspired inlet members 102 (shown as 102a and 102b) to reduce or completely eliminate flow stagnation, according to exemplary embodiments. The LVAD 100 may include some or all of the improved LVAD features described herein, including a stent-implanted inlet member 102 (referred to as “stent-implanted inlet” 102a, 102b, see Figures 2A–2B), a flexible rotor and / or pump housing 104 (Figures 3A–3B), and magnetic-based drive and bearing components (“maglev”) 106 (Figure 6) (collectively referred to as “flexible rotating components and magnetic components” 104, 106), a hydrophilic coating or surface 108 (see Figures 4A–4D), advanced control and electronics 110, and advanced charging and energy management controls 112 (collectively referred to as “advanced control and charging” 110, 112, see Figure 5), or a combination thereof.

[0054] In Figure 1, the LVAD100a includes a pump housing 114 that forms a volume 115 (not shown, see Figures 3A and 3B). The pump housing 114 includes an inlet 116 and an outlet 118. The LVAD100a includes a rotor 120 (e.g., a flexible impeller, not shown, see Figures 3A and 3B) disposed within the pump housing 110 to rotate within the volume 115 to pump blood. The LVAD100a includes a motor 122 (not shown, see Figures 4A and 4B) electrically coupled to the rotor 120 to drive the rotation of the rotor 120. The stented inlet member 102 is coupled to the inlet 116 of the housing 114, and the stented inlet member 102 has (i) an external structure 124 (e.g., fabric) configured to conform to and contact the outer wall at the base of the left ventricle, and (ii) an internal expandable body 126 configured to be positioned within the left ventricular chamber to expand and contact the tissue wall in the chamber (e.g., to reduce the risk of post-implant flow stagnation). The output of the LVAD 100 is connected to an outflow graft 128 that connects to the aorta. The external structure 124 may be made of fabric and sutured over the apex of the left ventricle.

[0055] The internally expandable body 126 is configured to move between a stowed configuration (located and extending within the left ventricle before being placed in the chamber) and a deployed configuration (deployed after being placed in the chamber). When deployed, the internally expandable body 126 forms a fan-shaped body that contacts the internal chamber / tissue of the left ventricle to avoid or reduce the possibility of flow stagnation within the chamber.

[0056] The structure and materials of the internally expandable body 126 are, in some embodiments, configured to be absorbable over time, for example, so that they are replaced by cardiac tissue after absorption, leaving only newly formed tissue behind. In some embodiments, other biocompatible and biodegradable (e.g., absorbable) materials may be used. The internally expandable body 126 may be terminated with a quick-release connector 129 to allow the pump housing 114 to be removable and attachable during implantation and any subsequent surgery. The quick-release connector allows the pump housing 114 to be connected to the stent-implanted inlet member 102, so that the LVAD 100 can be easily attached and replaced and the cap removed as needed.

[0057] The advanced charger 110 operates together with a subcutaneously implanted coil 132 that operates together with an external coil 134 coupled to an external charger 136. The charger 136 can communicate with the LVAD 100 via a low-power communication channel and / or via power line communication through the charging circuit.

[0058] Magnetic levitation bearing drive systems can reduce wear on the bearing system, extend the lifespan of the device, and improve the overall blood compatibility of the device. The system is configured to use active control in the axial direction, as well as passive control of radial position and tilt, to minimize the number of control axes. Fewer control axes can reduce the outer diameter and the number of sensors and electronic components. Magnetic levitation technology can also be applied to other direct blood contact medical devices and / or devices with rotating components to reduce induced hemolysis and wear on the bearing system.

[0059] Development.To deploy the stented inlet 102, an opening can be created at the apex of the LV by a coring tool. The crimped stent, with the fabric and suture ring, can then be inserted into the left ventricle. After removing the crimping tool, the suture ring can be attached to the left ventricle using sutures. Finally, the LVAD can be attached to the stented inlet using a quick-release connector.

[0060] Reduced risk of thrombosis. The location of the movable part of the LVAD, as well as the location of stagnant flow formed within the heart, can cause thrombus formation. Thromboembolic formation can increase the risk of ischemic stroke in LVAD patients, potentially blocking the entry point and causing pump failure. Points of blood stagnation can induce blood protein adsorption, leading to a chain reaction including thrombosis, ultimately resulting in sepsis and death. Studies have observed a high incidence of thrombosis near the entry cannula, at 96%. Conventional LVAD designs utilize sintered rough surfaces, among other things, to promote the formation of a biological neointimal layer at the cannula, and while the anti-vWF antibodies found in this layer were significantly reduced, the problem of thromboembolism was not eliminated.

[0061] The insertion angle and depth of the inflow cannula have been observed to affect the flow and stagnation areas around the cannula and within the pump, and have been associated with thrombosis.

[0062] Regarding damage to the moving parts, studies have shown that complications are associated with blood damage caused by the hyperphysiological shear stress of the LVAD on the moving parts of the device. When exposed to high mechanical forces, blood cells can be damaged or destroyed, leading to hemolysis and platelet activation, which can cause coagulation and thrombosis. Furthermore, high mechanical forces can degrade high molecular weight von Willebrand factor (vWF), a glycoprotein important in the blood coagulation process. Destruction of vWF is associated with gastrointestinal bleeding after LVAD implantation. Reducing blood damage improves blood compatibility and therefore improves outcomes of LVAD therapy.

[0063] Problems caused by moving parts and blood damage are not limited to LVADs but are prevalent in other blood circulation pumps, including extracorporeal membrane oxygenation (ECMO) and cardiopulmonary bypass (CPB) pumps. A promising method for treating uncontrolled bleeding that causes preventable death in the military is to induce hypothermia using CPB. Lower temperatures during CPB treatment put the injured soldier into a state of suspended animation, which can preserve organs and nerve function for extended periods to control bleeding or transfer to a competent medical facility. High flow rates similar to those required for LVADs are needed in the pump to rapidly lower body temperature to 20°C. For this purpose, improved mechanical and / or coating designs can provide ultra-low thrombosis surgery for CPB or other devices such as ECMO.

[0064] Thrombosis due to quasi-static blood substance interactions.Surface thrombosis due to the presence of whole blood can be elucidated as several events: protein adsorption, platelet adhesion / activation, leukocyte recruitment, and further activation of contact and complement pathways. Within seconds to minutes, serum proteins are adsorbed and conformation changes occur on the surface. This layer of adsorbed proteins can enable the subsequent adhesion and activation of nuclear platelets, which play a role in fibrin thrombus formation and the recruitment of leukocytes (including monocytes and neutrophils). The platelets then initiate an inflammatory immune response, promoting a cascade of events that leads to thrombosis within the device in contact with the blood. A surface that can eliminate unprecedented resistance to blood protein adsorption (the first step in the blood coagulation cascade) can hinder the blood coagulation cascade, thereby providing favorable quasi-static blood-molecular interactions.

[0065] In some cases, the affinity of blood proteins to hydrophilic (i.e., water-loving) surfaces is lower than that to hydrophobic (i.e., water-repellent) surfaces. Most physiological blood-contacting surfaces in the human body are hydrophilic, thereby reducing undesirable protein adsorption. The lower adsorption of blood proteins on hydrophilic surfaces compared to hydrophobic surfaces may be due to a combination of lower solid-liquid interface energy and the presence of a hydration layer. To reduce the solid-liquid interface energy (as nature prefers), blood proteins in aqueous media tend to adsorb at the high-energy solid-liquid interface of hydrophobic surfaces rather than the low-energy solid-liquid interface of hydrophilic surfaces. Thus, hydrophilic surfaces can represent lower blood protein adsorption than hydrophobic surfaces. Furthermore, unlike hydrophobic surfaces, hydrophilic surfaces have a hydration layer (i.e., water molecules bound to the surface) that must be replaced for protein adsorption. This presents an energy barrier, thereby preventing or reducing protein adsorption on a short timescale (seconds to minutes).

[0066] Exemplary stent-placed entrance member Referring to Figures 2A and 2B, two exemplary LVADs, 100a and 100b, are shown, stented or equipped with a stent-like inlet member. Figure 2A shows LVAD 100b. As used herein, the term “stent” refers to a small, expandable tube that is insertable into the left ventricle and expands to contact the tissue within the left ventricular chamber. The inlet member has, or may have, a structure similar to that of an arterial stent, but expands to contact the tissue within the left ventricle to minimize the risk of flow stagnation that may form around the interface between the LVAD and the left ventricle. In this example, there is no occlusion, as in the case of an artery or blood vessel.

[0067] To address potential flow stagnation and thrombosis and improve inflow hemodynamics, exemplary LVADs 100a, 100b include an internal expandable body 126 having a fan-shaped body that contacts the internal chamber / tissue of the left ventricle to avoid or reduce the possibility of flow stagnation pockets in the chamber. In some embodiments, the internal expandable body 126 is configured to move between a stowed configuration (located and extending within the left ventricle before being placed in the chamber) and an unfolded configuration (unfolded after being placed in the chamber). In the unfolded configuration, the internal expandable body 126 is defined by (i) a first section having a first perimeter sized to connect to the pump inlet, and (ii) a second section having a second perimeter sized to contact the inner wall / tissue of the left ventricle to eliminate left ventricular flow stagnation. In the storage configuration, the internal expandable body 126 is defined by (i) a first section having a first circumference sized to connect to the pump inlet, and (ii) a second section having a second circumference sized similarly to the first circumference.

[0068] As discussed above, in Figure 1, and here in Figure 2A, the LVAD 100a includes a pump housing 114 including an inlet 116 and an outlet 118. A stented inlet member 102a is coupled to the inlet 116 of the housing 114, and the stented inlet member 102 has an external structure 124 configured to contact and conform to to some extent the outer wall 202 at the base of the left ventricle (e.g., the LV apex). The stented inlet member 102 includes an internal expandable body 126 configured to be positioned in the left ventricular chamber 203 to expand and contact the tissue wall 205 in the chamber (e.g., to reduce the risk of flow stagnation after implantation). In Figure 2A, the tissue is shown to be partially transparent to provide a diagram of the internal expandable body 126 in an expandable configuration.

[0069] Inspired by self-expanding or balloon-expanding stents in transcatheter aortic valve technology, in some embodiments, the internally expandable body 126 includes a flexible mesh 204 that can be configured to be self-expanding or balloon-expanding. The self-expanding stent has interconnecting mesh elements, i.e., struts, configured to flexibly bend into a smaller configuration when placed under a force, allowing it to expand circumferentially to an expanded position when deployed. The balloon-expanding stent has interconnecting mesh elements, i.e., struts, configured to be in a smaller configuration when not subjected to force, but the struts are configured to expand relative to each other when subjected to force, i.e., via the balloon. The mesh 204 of the internally expandable body 126 can be used to function similarly to the self-expanding or balloon-expanding stents, but the expansion is (i) minimal in the proximal region 206 to the inlet 116 of the pump housing 114, and (ii) maximum in the distal region 208 of the device to form a fan-shaped expansion.

[0070] For this purpose, the internally expandable body 126 is defined in the deployed configuration by (i) a first section 210 having a first circumference sized to connect to the inlet 116 of the pump housing 114, and (ii) a second section 212 having a second circumference sized to contact the inner wall / tissue 205 of the left ventricle to eliminate left ventricular flow stagnation. In the retracted configuration, the internally expandable body 126 is defined by (i) a first section having a first circumference sized to connect to the inlet of the pump, and (ii) a second section having a second circumference sized similarly to the first circumference.

[0071] Flexible mesh patterns for expansion or compression. Figure 2C shows an exemplary pattern of mesh 204 from Figures 2A and 2B, including multiple changing sinusoidal struts 214 (or hourglass struts) including two adjacent struts 214a, 214b joined to each other at the peak portion of the sinusoidal curve. In Figure 2C, the changing sinusoidal struts have a constant periodicity, i.e., the distance between the peak and trough of the sinusoidal curve, and the longitudinal length 220 and amplitude / width 222 of the sinusoidal peak 216 increase from the proximal region 206 to the distal region 208. Figure 2D shows a mesh pattern for a transcatheter aortic valve, for example, mesh 204, having a set of generally triangular members 224. In the embodiment shown in Figure 2D, multiple triangular members 224 form a strut section 226a, joined to adjacent strut sections 226b at a connecting portion 228. Each generally triangular member may include notches to provide expansion and compression of the secondary structure. Mesh 204 can be manufactured, in one embodiment, by laser cutting a nitinol tube. Other manufacturing operations may be used.

[0072] material.Referring to Figure 2A, the internally expandable body 126 includes a fabric 230 that is sized and formed into a fan-shaped configuration. Examples of the fabric 208 include, but are not limited to, polyester or thermoplastic polymer resins (e.g., polyethylene terephthalate (PET) (e.g., Dacron®)). The fabric 208 may be stretchable (or non-stretchable) and may surround the mesh 204, or the mesh 204 may be embedded within the fabric 204. Other fabric materials that may be used include polytetrafluoroethylene (e.g., expanded polytetrafluoroethylene (ePTFE)), biological tissue (e.g., pericardial tissue transplanted from a human donor or animal), or combinations thereof.

[0073] For example, a self-expandable or balloon-expandable stent-like structure, enclosed in or embedded within a fabric (e.g., Dacron fabric), may be deployed in the left ventricle to replace a pump inlet cannula, as used in previous LVADs. The Dacron fabric or other fabrics described herein may promote endothelialization that can cover the entire inlet member 102. Similar fabrics have been successfully implemented to prevent paravalvular leakage in transcatheter aortic valves. Well-formed endothelium has been shown to provide a barrier against thrombus formation in relation to coronary stents, and the risk of thrombosis has been observed to be negatively correlated with the percentage of struts covered by endothelium. Inflow into the LVAD may be smooth with minimal occlusion to generally conform to the medial tissue or chamber shape of the left ventricle.

[0074] Patient-specific design. In some embodiments, the inlet member 102, including the fabric and mesh pattern, may be fabricated as a patient-specific design for a given patient, for example, to match a scan of the patient's left ventricle. The scan may be computationally modeled, and a portion of the internal expandable body, for example, a flexible mesh, may then be 3D printed or laser cut to form a structure that, once expanded, can adapt and / or conform to the complex anatomical deformation of the ventricle for the patient.

[0075] In some embodiments, computational modeling may include optimization via finite element simulation, for example, in Abaqus / CAE (Dassault Systèmes). Mesh 204 can be manufactured by laser cutting nitinol tubing.

[0076] Exemplary flexible rotor / casing Figures 3A–3G show the flexible rotor and / or pump housing and design parameters for the exemplary LVAD100a and 100b of Figure 1, according to exemplary embodiments. Specifically, Figures 3A and 3B show the exemplary flexible rotor and / or pump housing, respectively. Figures 3C–3G show the design parameters.

[0077] In the embodiment shown in Figure 3A, the LVAD 100 (indicated as 100c) includes a pump housing 114 (indicated as 114a) forming a volume 115. The pump housing 114a includes an inlet 116 and an outlet 118. The LVAD 100a includes a flexible rotor 120 (indicated as 120a) disposed within the pump housing 110 to rotate within the volume 115 for pumping blood. The flexible rotor 120a includes, for example, a plurality of curved blades 302 formed of a polymer material to form an impeller. Other parts of the rotor may also be formed from a polymer material. The curved blades 302 (indicated as 302a) may be coated or treated so that the surface forms a slippery hydrophilic antithrombotic coating or surface 304 (see Figure 3B). In the embodiment shown in Figure 3A, a portion of the surface may be coated with a slippery hydrophilic antithrombotic coating or treated as a slippery hydrophilic antithrombotic surface 304. In other embodiments, the entire inner surface (not shown) of the pump housing 114 may be coated with a slippery hydrophilic antithrombotic coating or treated as a slippery hydrophilic antithrombotic surface 304.

[0078] In Figure 3A, the flexible rotor 120a includes a projection section 306 for engaging with a bearing 308 located in the interior portion of the pump housing 114a. The LVAD 100c includes a motor 122, shown to have a stator 122a and permanent magnets 122b. The permanent magnets 122b are either fixedly coupled to or embedded in the rotor 120a to drive the rotation of the rotor 120a. A driver circuit 117 is embedded in the pump housing 114. The driver circuit 117 is electrically coupled to the stator 122a to drive the rotation of the motor 122 and to regulate the speed of the motor 122. The driver circuit 117 may be coupled to a sensor within the LVAD 100. The flexible rotor 120 and the pump housing 114 may constitute a magnetic levitation drive system (see Figure 6). Based on inviscid fluid dynamics theory, initial experiments and CFD tests indicate that the pump can provide pressures exceeding 90 mmHg at 5 L / min and 3000 RPM, which would be sufficient to assist patients with heart failure. Figure 3D shows the CFD results for a flexible rotor, which show good agreement with the experiments (Figure 3D).

[0079] Figure 3C shows an exemplary design 320 (based on a previous device, e.g., Abbott Labs' HeartMate3®), as well as the corresponding quantitative values ​​of pressure and flow rate for comparison. The assembly of device 320 is shown (subpanel A in Figure 3C) and evaluated for different rotational speeds (subpanel B in Figure 3C). Model 320 was designed based on a proven industrial design strategy. Device 320 was driven by an external shaft and evaluated for flow performance and flow characteristics.

[0080] The blades of the rotor 120 can be made of cast polyurethane rubber (for example, for durability and chemical stability). The rotor 120 can also be made of other flexible materials and by other methods, such as injection molding. The acrylic casing can be coated with a layer of the same material to make it flexible as well. The rotor blades have a wide range of hardness, for example, from Shore 10A to 100A. Flexible, colorless, transparent polyurethane (BJB Enterprise, Tustin, CA) 22 can be used for the rotor. The flexible rotor and casing can be coated with an antithrombotic coating.

[0081] Figure 3D shows the flow field and hemolysis index (HI) distribution for two rotor designs ("10-30" design and "70-70"). The "10-30" and "70-70" designs are optimized for blood flow but not yet optimized for blood damage. In-plane flow vectors are shown in subpanels A and C of Figure 3D, and hemolysis index contours are shown in subpanels B and D of Figure 3D. Figure 3D shows a comparison of the two rotor designs in terms of flow field to compare the hemolysis index of the two rotor designs under the same auxiliary flow rate and head rise for different RPMs.

[0082] The hemolysis index can be obtained by measuring the ratio of free hemoglobin to total hemoglobin and using the Euler hemolysis model described by Yu et al. The difference between the two rotors is the leading-edge blade angle and the trailing-edge blade angle, measured between the chordal and circumferential directions of the blades. The rotor "10-30" is an optimized design for flow, with a leading-edge angle of 10° and a trailing-edge angle of 30°. In fact, the flow vector follows the blades, unlike the large recirculation zone seen for the 70-70 rotor. However, when comparing the blood damage measured at the outlet of the LVAD model, the 70-70 rotor shows a 5% lower hemolysis index than the 10-30 rotor.

[0083] A flow field below the optimal level for the "70-70" rotor can provide less blood damage. With straighter blades, the pump can operate at a much lower RPM. 3163 RPM was used for the 10-30 rotor to produce the same head and flow, whereas the 70-70 only needed to operate at 2749 RPM. Lower RPMs can cause less shear to the blood, thus reducing blood damage. The shorter blades of the "70-70" rotor can also provide a smaller blood contact surface, which can reduce blood damage. In particular, the results in Figure 3D suggest that there is an optimal point for providing flow while minimizing blood damage. Figure 3E shows the hemolysis index as a function of blade number (subpanel A) and pump size (subpanel B).

[0084] Flexible rotor performance. Figure 3F shows the efficiency performance of a flexible rotor blade compared to a rigid rotor blade. The data shows that a flexible rotor with a flexible rotor blade can increase power production efficiency. Subpanel A of Figure 3F shows a four-piece mold for a rotor adapted from MacPhee and Beyene. Several hydrodynamic studies have shown that flexible wings or blades can provide higher performance to aircraft and turbomachines, especially outside of design conditions [13']~[15']. Due to the interaction of the wing and the fluid structure of the flow field, flexible rotor blades have been found to adapt to the surrounding flow, resulting in increased efficiency in turbomachines [16']. The results in Figure 3F show that flexible rotor blades can increase the flow efficiency of the LVAD100.

[0085] Figure 3G shows an exemplary design for an impeller and the corresponding geometric parameters for an exemplary LVAD (e.g., 100). In Figure 3G, the pump geometry is characterized by nine parameters: inlet diameter (D1), rotor diameter (D2), leading edge angle (β1), trailing edge angle (β2), number of blades (z), blade thickness (d), blade tip gap size (distance between the blade tip and the drain) (h), blade wrap angle (φ), and outlet diameter (D3).

[0086] Machine learning optimization. Machine learning can be used, for example, to optimize the LVAD design in Figure 3G. A program can be developed in MATLAB to generate an LVAD model based on these parameters. The 3D geometric shape can then be constructed in Solidworks. The hemolysis index (HI), a measure of blood damage, can be optimized for design conditions of 5 L / min and 90 mmHg. HI can be determined as a function of all parameters in the problem, where HI = function (RPM, D1, D2, β1, β2, z, d, h, φ, D3), and is evaluated using computational fluid dynamics (CFD).

[0087] The analysis can begin by simulating approximately 500 cases as a training dataset. 200 separate cases can be randomly generated and used to validate the model. The process can be automated using Matlab and Ansys Workbench. A deep neural network (DNN) architecture can be used to predict HI based on 10 input parameters. The DNN can be trained using the training dataset and evaluated using the validation set. After the model is accurately trained, stochastic gradient descent can be used to predict the parameter combination with the lowest possible HI. Efficiency can be optimized using a similar process.

[0088] Other studies have shown that flexible vanes can suppress laminar flow separation bubbles, which can cause performance degradation at low Reynolds numbers [17']. The concept of such flexible rotors has been used in the design of foldable transcatheter VADs [18']-[20'] and flexible pumps [21'], [22'] that provide superior flow efficiency performance to their rigid counterparts in certain flow regimes [22']. Shear stress reduction was not shown as a consideration in this study.

[0089] Considering the complex inflow conditions within a patient's heart, LVADs may never function under optimal flow design conditions. Blood damage can be significantly increased under complex, real-world pulsating conditions. Exemplary LVADs using flexible rotors (e.g., 100) can reduce blood damage by adapting to the complex inflow conditions of the heart.

[0090] Slippery hydrophilic antithrombotic coating for LVAD Figure 4A shows a slippery hydrophilic (SLIC) surface that allows water droplets to diffuse intuitively, rather than causing water to adhere as a conventional hydrophilic surface. While conventional hydrophilic surfaces can reduce blood protein adsorption on a short timescale, on a longer timescale (several minutes to one hour), blood proteins can adopt a hydrophilic fit, diffuse through the hydration layer, and then adsorb onto the surface. Because the water molecules within the hydration layer are disordered and loosely bind to conventional hydrophilic surfaces, blood proteins can permeate the hydration layer and its irreversible adsorption, as shown in Figure 4A.

[0091] On exemplary SLIC surfaces, water molecules within the hydration layer are highly regular and tightly bound to the surface, thereby forming an "ice-like" hydration layer, as shown in Figure 4B. Previous molecular dynamics (MD) simulations have shown that the ice-like hydration layer can reduce friction at the molecular scale, allowing bulk water molecules to slide or glide tangentially against surfaces with the ice-like hydration layer. The high tangential mobility of water molecules near the ice-like hydration layer can impart substantial lateral forces to adsorbed blood proteins, as indicated by the arrows in subpanel B of Figure 4B. These lateral forces are expected to remove adsorbed blood proteins and hinder subsequent steps in the blood coagulation cascade (e.g., platelet adhesion and activation, leukocyte adhesion, etc.), as shown in subpanel C of Figure 4C. Indeed, results from studies conducted show that blood proteins do not adsorb to exemplary SLIC coatings after one week. In contrast, state-of-the-art coatings can prevent protein adsorption for as little as minutes to hours, potentially making exemplary SLIC coatings a breakthrough in surface science.

[0092] Figure 4B shows the antithrombotic properties of the SLIC coating. Subpanel A of Figure 4B shows a schematic diagram of a disordered hydration layer on a conventional hydrophilic surface, resulting in high protein adsorption. Subpanel B shows a schematic diagram of an ice-like hydration layer on the SLIC coating, resulting in negligible protein adsorption. Subpanel C shows a water droplet that cannot slide on a conventional hydrophilic surface. Subpanel D shows a water droplet that slides on the SLIC coating at a low tilt angle.

[0093] Subpanels A and B of Figure 4C show schematic diagrams illustrating fully developed blood flow between a conventional hydrophilic surface wall and a SLIC coating, respectively. Subpanel C of Figure 4C shows a schematic diagram illustrating the antithrombotic properties of the SLIC coating on an LVAD. Additional examples and descriptions of hydrophilic coatings can be found in the published PCT application WO2019 / 074791A1, which is incorporated herein by reference in its entirety.

[0094] Thrombosis caused by dynamic interactions between blood substances. Thrombosis in LVADs due to dynamic flow conditions can result from hyperphysiological shear stress and flow stagnation in the device, potentially leading to hemolysis, platelet activation, and degradation of von Willebrand factor (vWF). When blood flows over a conventional LVAD surface, there is no slip between the blood and the conventional LVAD surface. In other words, non-slip boundary conditions can increase blood stagnation and shear stress when present at the blood-material interface. LVAD surfaces with a SLIC coating, which induces slipperiness at the blood-material interface due to an ice-like hydration layer, can reduce friction. When blood flow passes over such an LVAD surface with a SLIC coating, blood stagnation and shear stress are reduced, thereby reducing hemolysis, platelet activation, and degradation of vWF, and thus reducing thrombosis.

[0095] Consider stable, fully developed blood flow through conduits without and with SLIC coating. Without SLIC coating, conventional hydrophilic surfaces have non-slip boundary conditions at the wall, i.e., the velocity of blood in contact with the wall, U no SLIC This can result in a value of =0, leading to significant flow stagnation. Furthermore, non-slip boundary conditions then result in higher shear stress at the wall.

number

number

[0096] SLIC coating design. The SLIC coating with an ice-like hydration layer can improve both the static and dynamic interactions with blood, thereby reducing thrombosis in LVADs. The SLIC coating can have both hydrophilicity and slipperiness.

[0097] Hydrophilicity: The basic macroscopic measure of the wetting property of a surface is the equilibrium contact angle θ. On a non-textured (or smooth) surface, for any contacting liquid, the equilibrium contact angle θ is given by Young's equation as γ LV cosθ = γ SV −γ SL . Based on the equilibrium contact angle θ of a water droplet, the surface can be classified as hydrophilic when θ < 90° and hydrophobic when θ > 90°. Young's equation shows that a high solid surface energy γ SV results in a low equilibrium contact angle. Thus, surface chemistries with high solid surface energy (e.g., PEG-based, ionic, zwitterionic, etc.) are a logical choice for designing hydrophilic surfaces.

[0098] Slipperiness: The basic macroscopic measure of slipperiness is the contact angle hysteresis Δθ, which is the difference between the advancing θ adv and the receding θ rec contact angles on the solid surface (i.e., Δθ = θ adv −θ recIt is defined as the contact angle hysteresis Δθ. Contact angle hysteresis Δθ can arise from chemical and physical heterogeneity of the surface. Physically, contact angle hysteresis is a measure of the energy dissipated during the movement of a droplet along a solid surface. When the surface exhibits a very high contact angle hysteresis (regardless of the equilibrium contact angle), the droplet can remain stationary on the surface (subpanel C in Figure 4B). On the other hand, when the surface exhibits a very low contact angle hysteresis (regardless of the equilibrium contact angle), the droplet can slide on the surface. Thus, low contact angle hysteresis can lead to a low sliding angle ω (i.e., the minimum angle at which the surface must be tilted for the droplet to slide) (subpanel D in Figure 4B). Based on the balance between the work done by gravity and the energy dissipated by adhesion, the sliding angle on a smooth solid surface is given by ρVg sinω≈γ LV D TCL (cos θ rec -cos θ adv It can be given as ), where ρ and V are the density and volume of the droplet, respectively, g is the acceleration due to gravity, and D TCL ω is the width of the three-phase contact line perpendicular to the sliding direction. The sliding angle ω is the inverse measure of how easily a droplet slides on a solid surface. Slippery surfaces with low sliding angles and low contact angle hysteresis can be designed with surfaces having low physical heterogeneity (i.e., smooth surfaces) and low chemical heterogeneity (i.e., high graft density). Smooth surfaces with ultra-high graft density are ideal for designing slippery surfaces.

[0099] Conventional hydrophilic surfaces have high solid surface energy γ SVWhile low water contact angles can be represented, they can also represent high adhesion to the surface due to high contact angle hysteresis. On a macroscopic scale, high adhesion can manifest as droplets that cannot slide through the surface, even on vertical surfaces, making the surface slippery. Exemplary SLIC surfaces are counterintuitive because they are not only hydrophilic but also exhibit high slipperiness, possessing very low sliding angles and contact angle hysteresis. Recent molecular dynamics (MD) simulations have shown that slipperiness and low adhesion can manifest on hydrophilic surfaces with an ice-like hydration layer having strong hydrogen bonds that may not be hindered by any protein fitting. Such ice-like hydration can be achieved on smooth surfaces with a high graft density of hydrophilic molecules on the surface. The ice-like hydration layer results in high liquid mobility near the solid-liquid interface.

[0100] Exemplary SLIC coatings with an ice-like hydration layer can provide high mobility of liquid molecules near the solid-liquid interface. On a macroscopic scale, the ice-like hydration layer manifests as water or blood droplets sliding across the surface at low sliding angles, exhibiting slipperiness. On a microscopic scale, the ice-like hydration layer may manifest as a finite sliding length. On a molecular scale, the ice-like hydration layer can manifest as high lateral mobility of water molecules, providing substantial lateral force for the removal of adsorbed blood proteins.

[0101] Fabrication and characterization of SLIC coatings.SLIC coatings can be fabricated on acrylic and polyurethane substrates through consideration of surface activation parameters, molecular construction, and reaction coordinates. Acrylic and polyurethane have been used in conventional LVADs and can be adapted to future additive manufacturing technologies. However, the application of SLIC coatings is not limited to the above substrates. Further optimization of exemplary SLIC coatings for maximum antithrombotic response (i.e., achieving lower physical and chemical heterogeneity) can be carried out by systematically tailoring hydroxylation parameters, molecular construction, and synthetic reaction coordinates. The optimized SLIC coating protocol can then be seamlessly adapted to efficiently coat blood-contact surfaces of exemplary LVAD designs.

[0102] The surfaces of LVAD polymers (i.e., acrylics and polyurethanes) can be modified via liquid-phase silanization [24'], [25'], because it is a simple, versatile, and scalable technique and because a wide variety of organosilanes with hydrophilic chemicals are commercially available.

[0103] Figure 4D shows an organosilane consisting of two chemical moieties: a functional organic group and a reactive group. The functional organic group may include a non-reactive moiety (e.g., a hydrophilic chemical such as PEG, an ionic or zwitterionic oligomer) that can impart a desired property (e.g., hydrophilicity) to the substrate. The reactive group may include a hydrolyzable moiety (e.g., chloro, amino, methoxy, ethoxy, etc.) that reacts with the underlying hydroxylated substrate (i.e., the surface with the -OH group) to form a durable covalent siloxane bond. The degree of silanization on the surface (or graft density of the functional group) depends on the hydroxylation of the substrate, the molecular construction of the silane [26']-[28'], and the silanization reaction conditions (e.g., type of solvent, silane concentration, type of catalyst, catalyst concentration, silanization time, water concentration, temperature, pH, etc.) [24'], [25'], [29']-[31'].

[0104] In some embodiments, hydroxylation times of 60 to 3600 seconds may be used with oxygen flow rate (1 cc / min to 50 cc / min) and plasma output (1 watt to 50 watts) and oxygen pressure (1 psi to 30 psi), followed by silanization to graft brushes of different hydrophilic chemicals.

[0105] Effects of hydroxylation: Substrates can have hydroxyl (-OH) groups on their surface for silanes to react with them [24'], [25']. Higher substrate hydroxylation can lead to higher silanization or higher graft density, which can ensure high chemical homogeneity, then a design requirement for SLIC coatings [28'], [32'], [33']. To achieve higher hydroxylation, LVAD polymer surfaces (e.g., made of acrylic and polyurethane) can be modified using oxygen plasma. Oxygen plasma for hydroxylation is simple, portable, and scalable. Too low oxygen plasma exposure results in insufficient hydroxylation, while too high oxygen plasma exposure can lead to substrate degradation or roughening (i.e., physical heterogeneity) [34']~[38'], which is undesirable for designing SLIC coatings. Oxygen plasma parameters (e.g., hydroxylation time, oxygen pressure, oxygen flow rate, plasma power, etc.) can be used to obtain the desired degree of substrate hydroxylation. The hydroxylated substrate can be modified with silane [42']~[52'].

[0106] Influence of Molecular Construction: The molecular construction of silanes (i.e., the type and length of functional groups, the type and number of reactive groups, etc.) can affect the graft density of silanes on a substrate [25']~[28']. To obtain the high graft density required to design SLIC coatings, the molecular construction can be adjusted for graft density, which can then be used to determine the effectiveness of SLIC coatings and, subsequently, their antithrombotic properties. Smaller molecular size and chain length (i.e., the number of repeating units in the oligomer or the length of the molecule) of functional groups (e.g., PEG, ionic or zwitterionic), as well as a higher number and reactivity of reactive groups, can allow for higher graft density (i.e., lower chemical heterogeneity), while larger molecular size and chain length of functional groups, as well as a higher number and reactivity of reactive groups, lead to lower physical heterogeneity [44'], [48'], [53']~[56']. By systematically adjusting the molecular construction, the optimal combination of low physical and chemical heterogeneity can be determined, resulting in an effective SLIC coating.

[0107] Influence of synthetic reaction coordinates: Although silanation is a simple technique, it is also a versatile technique that allows for a great deal of adjustment to obtain effective SLIC coatings by changing the synthetic reaction coordinates (e.g., type of solvent, silane concentration, type of catalyst, catalyst concentration, silanation time, water concentration, temperature, pH, etc.) [24'], [25'], [28'], [57'].

[0108] Wireless energy transmission and communication Figure 5 shows an exemplary charger 136 (indicated as wireless energy transmission and communication system 136a) that can be implemented in an exemplary centrifugal LVAD. The wireless energy transmission and communication system 136a can reduce the risk of infection, enable free movement of patients, and improve quality of life.

[0109] The wireless energy transmission and communication system 136a includes an external power supply for the SLIC LVAD pump. A rechargeable energy storage device in the embedded controller is designed to provide uninterruptible power when the external wearable power supply (e.g., provided in a belt or vest) is briefly removed. The communication link can provide feedback to the external power supply system for improved control of the electrophysiological function and charging circuit, thereby reducing overheating and improving both mechanical and electrical cardiac function.

[0110] Wireless energy transmission. The LVAD100 (e.g., 100d) may utilize an RF wireless energy transmission subsystem that operates with an externally rechargeable device coupled to a belt or vest to reliably supply and charge its energy storage. The externally rechargeable device can inductively charge the LVAD through an embedded internal antenna. The embedded charging components include circuitry that supplies power to the LVAD and charges the internal energy storage device, including a combination of a rechargeable battery, supercapacitor, and hybrid supercapacitor, when the external belt or vest is removed and / or unavailable.

[0111] The charging subsystem may include communication between the embedded LVAD100 and other internal sensors and an external vest or belt controller for system monitoring and / or feedback control. In some embodiments, the communication may be based on low-energy Bluetooth or other communication described herein to collect sensor readings, e.g., ECG data, and transmit them to the charging subsystem. The antenna may include backplane material to assist in focusing energy transmission. The antenna may additionally include temperature and position sensors to monitor coupling efficiency. The antenna may be constructed of a flexible material to conform to the body shape to improve comfort.

[0112] A control algorithm for improving dynamic LVAD performance.Wireless energy and communication can function as controllers for the LVAD. Calibration values ​​can be modified during operation, for example, via ML control, to respond to different dynamic physical load events. Combined with the modified LVAD calibration, the control algorithm can pulse the transmitted RF energy to provide sufficient power to the SLIC LVAD to monitor and reduce tissue heating.

[0113] In Figure 5, the charger 112 (shown as 112a) includes a rechargeable external power supply 136, a transmitter coil 134, and a receiver coil 132, and powers an internal rechargeable energy source 502 (shown as "embedded receiver" 502) which includes a rechargeable battery and a hybrid supercapacitor 504 (shown as "energy storage"), and powers the embedded LVAD 100 (shown as 100d). In some embodiments, the internal energy source may power the LVAD motor and magnetic levitation bearing system for at least 20 minutes (e.g., 5W).

[0114] In the embodiment shown in Figure 5, the external transmitter system 136 includes a microcontroller 506 (e.g., an ARM or Xilinx Zynq FPGA). An oscillator 508 may be used to generate an alternating current (AC) signal, which is supplied to a digital upconverter circuit 510 to generate a radio frequency (RF) signal. The RF signal may be routed to a power amplifier (PA) (also shown in 510). The output of the PA 510 may then be connected to the transmitting coil antenna 134 through an electrical matching network to ensure proper signal matching. The power level may be adjusted to provide sufficient power to the embedded controller 502, LVAD(100d), and rechargeable energy storage device 504. The controller 506 may adjust power delivery based on feedback from the embedded controller. The external transmitter system 136 includes a rechargeable battery 512, which may be packaged to be coupled to a waist belt or vest, for example, to be placed in close proximity to the embedded receiver coil 132.

[0115] The LVAD100d may include an embedded internal electric drive lead wire 514 that connects the receiver coil 132 to the control unit and associated circuitry. The drive lead wire 514 may continue along a path beneath the skin. The embedded control unit 502 may be similar in size and shape to a pacemaker control unit and may be positioned beneath the skin together with the receiver antenna 132. The embedded control unit 502 includes an antenna coil 132 for capturing energy through inductive coupling.

[0116] The embedded control unit 502 may include a rectifier circuit 516 for converting the received RF AC current signal into a DC signal to power an internal low-power control processor 518 (e.g., a Texas Instruments MSP430 or an application-specific integrated circuit (ASIC) chip). The processor 518 may be connected to a controller 520 configured to perform charging operations for an internal rechargeable energy storage device 504.

[0117] One or more rechargeable batteries and energy storage components, such as supercapacitors or hybrid supercapacitors, may be used. A supercapacitor is a rechargeable energy storage device that enables significantly more charge / recharge cycles than a rechargeable battery. The processor 518 may interface with motor control and magnetic levitation control electronics (e.g., also located in the controller 520). The processor 518 may connect to a communication module 522 (indicated as "LP Comm" 522) to transmit pump performance and power delivery data to an external power transmitter. The communication module 522, as an internally embedded transmitter such as a low-energy Bluetooth link, can transmit LVAD status information along with ECG data to the external controller to improve control algorithms and enable long-term electrophysiological monitoring. The communication module 522 may use other communication protocols, such as Bluetooth, ZigBee, and Z-Wave communication, or power line carrier (PLC) communication protocols.

[0118] Both the internal control unit 502 and the external power supply 136 may each have integrated position and temperature sensors and corresponding circuitry to monitor heat resulting from antenna misalignment and inefficiencies in wireless energy transmission. The controller 520 may include control software to enable module pairing and transmission of data from various internal sensors. A Bluetooth module 524 (indicated as "LP Comm" 524) on the external power supply may be additionally paired with a smartphone for monitoring the LVAD100d (e.g., by both the patient and a cardiologist).

[0119] Electrophysiological analysis to control LVAD function.Electroanatomical mapping may be performed before and after LVAD implantation to establish a baseline for comparison, assess acute changes, and evaluate long-term outcomes in changes to the innate conduction pattern. Additionally, 12-lead surface ECG may be used for overall conduction assessment. Extensive electroanatomical conduction data across various time points and significant cardiovascular events can provide much-needed insights into longer-term patient management. This level of understanding of conduction before and after LVAD implantation can provide clues as to whether additional rhythm management, such as ablation, would be beneficial to the patient in the long term.

[0120] Electronic control and LVAD pump motors can generate undesirable electrical noise that may affect sensors used to monitor ECG and pacing. In addition to shielding, active DSP noise cancellation algorithms may be used to improve signal quality. Machine learning (ML) algorithms (e.g., running on charger 136 or a cloud infrastructure coupled via charger 136) may be developed to model LVAD time-series data and outlier events.

[0121] LVAD control. The LVAD control algorithm may include a trained machine learning model configured to adapt LVAD motor operation to the dynamic load caused by patient activity and stress. The algorithm may also monitor cardiac electrical conduction and modify LVAD operation to allow left ventricular rest and enable healing and improved cardiac function.

[0122] The control algorithm can adjust LVAD operation to reduce energy consumption during periods of reduced need, such as when the system is idle. The communication data rate from the internal controller and sensors to the external controller and vest can also be reduced during periods of reduced activity. Reducing the load on the LVAD system also reduces the need for wireless energy transmission to maintain the charge of the internal rechargeable battery and hybrid supercapacitor.

[0123] LVAD drive system design with magnetic levitation Figure 6 shows an exemplary LVAD 100 (indicated as 100e) comprising a rotor 120 (indicated as “impeller” 120b) having mounted rotor permanent magnets 602, a stator 604 embedded in a pump housing 114 adjacent to the inlet member 102, and a drive system (also referred to as the “magnetic levitation system”) having magnetic levitation, comprising an active magnetic bearing 606 (indicated as “magnetic levitation bearing” 606) located at the bottom of the pump housing. The pump housing 114 has a position sensor 608 (indicated as “eddy current sensor” 608) configured to monitor the gap between the impeller tip and the rotor 120b. The drive system includes an axial magnetic bearing and a radially passive magnetic bearing with independent BLDC motors for preventing radial touchdown of the rotor and for axial adjustment of the rotor's position in response to force changes under normal and extreme conditions. The drive system can perform active control in the axial direction, as well as passive control of radial position and tilt. Fewer control shafts can reduce the outer diameter and the number of sensors and electronics required. In Figure 6, the impeller 120b is fixedly mounted to the rotor's permanent magnet 602. Inside the LVAD housing, the stator 604 is embedded adjacent to the inlet cannula 116.

[0124] An active magnetic levitation bearing 606 located at the bottom of the LVAD, equipped with an eddy current sensor 608, can monitor the gap between the impeller tip and the magnetic levitation device. The eddy current sensor 608 may be analyzed to improve the control of the LVAD pump speed. In the radial direction, the rotor permanent magnets 602 align with the stator 604, and in the axial direction, the magnetic levitation bearing 606 resists the electromagnetic force between the motor stator 604 and the rotor 120b, as well as the force from the rotor due to gravity. In some embodiments, the eddy current sensor 608 is coupled to electronic equipment 117 to operate the control of the LVAD pump speed based on the sensor readings.

[0125] When implanted in large animal models, SLIC LVADs can contribute up to 5 L / min of total cardiac output in calf models and maintain physiological arterial pressure and low hemolysis rates.

[0126] Additional embodiments and descriptions of LVAD drive systems having magnetic levitation can be found in the published PCT application WO2023 / 189970A1, which is incorporated herein by reference in its entirety.

[0127] Experimental results and additional examples Several studies have been conducted on various features of the LVAD system described herein. Studies have been conducted to evaluate the inlet member design, flexible blades and casings, and SLIC coatings.

[0128] Research on entrance components The design was computationally modeled and optimized via finite element simulation, for example, in Abaqus / CAE (Dassault Systèmes). Mesh 204 was fabricated by laser cutting nitinol tubing. After suturing the Dacron fabric to the stent, the inlet member 102 was deployed onto a patient-specific silicone left ventricle model and attached to the flow loop of a left heart simulator. Three-dimensional particle tracking kinetics was used to thoroughly evaluate the flow and potential blood damage in the stented inlet and LVAD model. The results can be compared to the same model with a straight inlet cannula to confirm that the stented inlet eliminates stagnant areas around the inlet and thus completely reduces the risk of thrombosis. Test conditions ranged from a healthy beating heart to a pathologically failing heart.

[0129] Heart rate synchronized velocity modulation for pediatric LVAD Increasing the pulsality of the combined LVAD and cardiac output can be achieved by modulating the pump driver speed in sync with the cardiac rhythm. Increased pulsality can approach physiological standards, reduce blood stagnation, and improve feedback about pressure receptors for pressure regulation. Figure 7A shows a previously validated computational model of a simulated circulatory loop that can be used as a platform for observing the hemodynamic effects of synchronized pump speed modulation to the heart. The study established a model of a brushless DC motor driving a pediatric LVAD using the speed, torque, and inertia constants from the manufacturer's specifications to represent the motor's ability to vary speed. The LVAD impeller pressure and flow rate HQ curves are used to calculate the pump output flow rates at different pressures and impeller speeds. The set-speed LVAD flow rate is 3 L / min, and the speed-modulated LVAD flow rate varies between 1 L / min and 5 L / min during diastole and systole, respectively. The study used a PID controller to control the pump speed to reach the desired flow output.

[0130] Figure 7B shows left ventricular and aortic pressures controlled by a set-velocity LVAD (subpanel A) and a velocity-modulated LVAD (subpanel B). The modulated LVAD flow resulted in greater changes in aortic pressure. The velocity-modulated LVAD had an aortic pressure difference of 20 mmHg between minimum diastolic and maximum systolic, while the set-velocity LVAD had a difference of 7 mmHg. The set-velocity LVAD and velocity-modulated LVAD had similar mean aortic pressure values ​​of 93 mmHg and 91 mmHg, respectively. The changes in aortic pressure during systole and diastolic periods resulting from velocity-modulated flow indicate that synchronized motor velocity modulation results in physiological conditions closer to the standard.

[0131] Evaluation of flexible blades and casings The study evaluates the integration of additionally flexible blades and casings into LVAD designs. In the study, the blades were fabricated from cast polyurethane rubber for their durability and chemical stability. The acrylic casings were coated with layers of the same material. The study developed an exemplary LVAD using acrylic for its low cost, blood compatibility, and suitability for potential additive manufacturing techniques. Furthermore, polyurethane rubber can be readily coated with SLIC coatings. Performance and blood loop tests may be performed to evaluate the reduction of blood damage due to different material hardnesses. Best performance was evaluated using particle image kinetics. Blade and casing deformation was measured under different flow conditions. Material durability was thoroughly evaluated.

[0132] Flexible rotor design using machine learning.The LVAD design was optimized using machine learning. Based on industry-proven design concepts[23'], the pump geometry was characterized by nine parameters. As shown in Figure 3G, the parameters controlling the geometry and its range are the inlet diameter (D1=10-15mm), rotor diameter (D2=20-45mm), leading edge angle (β1=5-90°), trailing edge angle (β2=5-90°), number of blades (z=3-6), blade thickness (d=1-4mm), blade tip gap size (distance between blade tip and drain) (h=1-4mm), blade wrap angle (φ=10°-600°), and outlet diameter (D3=10-20mm). Based on these parameters, a program was developed in MATLAB to generate an LVAD model. Subsequently, the 3D geometric shape was constructed in Solidworks. The hemolysis index (HI), a measure of blood damage, was optimized for design conditions of 5 L / min and 90 mmHg. HI was determined as a function of all parameters in the problem, HI = f(RPM, D1, D2, β1, β2, z, d, h, φ, D3), and was evaluated using computational fluid dynamics (CFD). The study began by simulating approximately 500 cases as a training dataset, covering all extremes of the parameters. A separate 200 cases were randomly generated and used to validate the model. The entire process was automated using Matlab and Ansys Workbench.

[0133] The study utilized a deep neural network (DNN) architecture to predict HI based on 10 input parameters. The DNN was trained using a training dataset and evaluated using a validation set. After the model was trained, another machine learning algorithm (stochastic gradient descent) was used to predict the parameter combination with the lowest possible HI. Efficiency was also optimized through a similar process.

[0134] Numerical simulation.Using the Ansys Fluent 3D pump model, the study simulated auxiliary flow conditions of 90 mmHg and 5 L / min. The rotor speed was therefore adjusted to match these conditions. Performance data from in vitro experiments were used to validate the CFD simulations, and high-resolution PIV experiments were used to calibrate the turbulence model. A hemolysis model based on stress and exposure time was applied to assess blood damage. The parameters of the hemolysis model were calibrated using preliminary blood loop experiments.

[0135] Designed for off-point operation of LVADs using flexible rotor blades. The operating speed of LVADs is often adjusted to adapt to patient variability, which can lead to the LVAD operating outside of its design point, potentially causing loss of efficiency and an increased risk of thrombosis / hemolysis. Flexible rotor blades can deform in response to complex inflow conditions, and with appropriate geometry and flexibility, they can produce a flatter performance curve.

[0136] The study experimented with the hydraulic performance of various impeller geometric shapes of LVADs equipped with flexible rotor blades under pediatric and adult pressure / flow conditions. Figure 8A shows a schematic diagram of the hydraulic test setup used to evaluate the hydraulic performance of the rotor and impeller. The study fabricated the flexible rotor blades by casting polyurethane resin (Easton, PA) with a Shore hardness of 60A into a four-piece mold. The study 3D printed the mold parts using a J850 Pro (Stratasys Ltd, MN, USA). The blade inlet angle was set to 20°, and the outlet angle was varied from 10° to 70°. The blade thickness was 1.2 mm, and the blade height was 7 mm.

[0137] Flexible rotor blades were evaluated in a hydrodynamic test setup using a 40 wt% glycerin solution as the working fluid. The setup included a motor to drive the rotor, a flow probe, upstream and downstream pressure transducers of the rotor, and a pinch valve to vary the downstream resistance. The mounted rotor was driven at 3000 RPM, and the flow generated by the rotor was measured as the pinch valve gradually increased the downstream resistance. The study then scaled the hydrodynamic results to achieve pediatric and adult pressure / flow rate conditions of 70 mmHg-2 L / min and 90 mmHg-5 L / min, respectively. To analyze out-of-design-point operation, the study calculated the slopes of the pressure / flow rate curves under the two operating conditions using equations 1 and 2, respectively.

number

number

[0138] For flexible rotor blades with a 30° outlet angle, the operating speeds for pediatric and adult pressure / flow conditions were 2550 RPM and 3230 RPM, respectively. The gradient under pediatric and adult operating conditions was -6.14 mmHg / (L / min). -1 and -11.12 mmHg / (L / min) -1 Figure 8B shows the results for a flexible rotor blade with an exit angle of 30°.

[0139] Analysis of the differences between rigid rotor blades and flexible rotor blades.The study evaluated the performance between flexible and rigid rotor blades. Three designs were selected for the production of rigid and flexible rotor blade LVAD prototypes, respectively. The blade thickness and height were fixed at 1.2 mm and 7 mm, respectively. As shown in Figure 8C, the blade inlet angle was set to 20°, while the outlet angle was varied in 20° increments from 30° to 70°. The rigid rotor blades were 3D printed using J850 Pro (Stratasys Ltd, MN, USA), while the flexible rotor blades were cast using polyurethane resin with a Shore hardness of 60A. The results are presented and discussed in relation to Figure 3F.

[0140] In vitro blood compatibility assessment. Figure 8D shows an evaluation of blood damage caused by both rigid and flexible blade LVAD prototypes using pig blood in a blood circulation loop. As shown in Figure 8D, the test system includes a motor to drive the rotor, a flow probe, upstream and downstream pressure transducers of the rotor, and a pinch valve to vary the downstream resistance.

[0141] The blood circulating loop (BCL) was adjusted to adult physiological conditions, and blood samples were taken every 60 minutes over a 6-hour period. Blood samples were centrifuged to prepare plasma. Hemolysis and platelet activity were quantified using a commercially available enzyme-linked immunosorbent assay (ELISA).

[0142] Fluid dynamics performance evaluation. The study also evaluated the hydrodynamic performance of the LVAD prototype in a benchtop setup. The study used a 40 wt% glycerin solution mimicking blood as the working fluid. The mounted rotor was driven at 3000 RPM, and the flow generated by the rotor was measured as the pinch valve gradually increased the downstream resistance. To characterize the hydrodynamic performance, the study measured the gradient and efficiency of the pressure-flow curve using Equation 2.

[0143] For a rigid rotor blade with a 30° exit angle, the gradient under adult operating conditions is -11.52 mmHg / (L / min). -1 In contrast, the flexible equivalent was -11.12 mmHg / (L / min). -1 The gradient was as follows. Figure 8E shows that the pressure / flow gradient for the rigid rotor blade prototype decreased as the outlet angle increased.

[0144] SLIC coating evaluation The study characterized the performance and flow of LVADs constructed with SLIC coatings. Optimized flexible LVADs, with and without SLIC coatings, were tested in benchtop flow loop experiments. In addition to performance characterization, the study utilized an advanced 3D time-resolved particle tracking velocimetry system to directly quantify the hemolysis index of LVADs with and without coatings, calibrating results obtained from CFD. Four cameras were used to record particle images from four different field angles, and the state-of-the-art particle tracking algorithm, Shake-the-box, was used to track particle trajectories as particles entered and exited the pump. HI was directly calculated and compared by integrating the shear stresses experienced by individual particles along their trajectories.

[0145] Characterization of physical and chemical heterogeneity of SLIC coatings. The study characterized SLIC coatings fabricated to evaluate physical and chemical heterogeneity and the resulting slipperiness.

[0146] Atomic force microscopy (AFM) and electrostatic force microscopy (EFM). The study utilized a Bruker MultiMode 8-HR AFM with a silicon nitride probe in ScanAsyst mode to characterize surface roughness, and then determined physical heterogeneity. The study scanned a 4 μm × 4 μm area at a speed of 1 Hz to acquire at least 30 images. The images were used to measure the root mean square roughness R rmsTo obtain the data, the samples were analyzed using Nanoscope Analysis software. Additionally, using EFM accessories, the study characterized the surface charge to assess the uniformity of the surface charge distribution.

[0147] X-ray photoelectron spectroscopy (XPS). The study characterized the near-surface chemical composition at multiple locations using XPS to assess chemical heterogeneity. The study then characterized the surface chemistry using a Physical Electronics PHI-5800 XPS. Analysis can be performed using a monochromatic Al Kα X-ray source operating at 15 kV, with photoelectrons collected at a 45° extraction angle relative to the sample surface. XPS data can be acquired from at least 15 spatially distinct locations on the surface, and spectral analysis is performed using PHI Multipak software.

[0148] Ellipsometry. The study assessed chemical heterogeneity by estimating graft density through thickness measurements using polarization analysis. The thickness of the SLIC coating was measured using a variable-angle spectrophotometric polarization analyzer (VASE-VB-250). Surface spectral scans could be collected at 5° intervals between 500 nm and 900 nm for incidence angles between 55° and 75°. The thickness of the PEG layer (refractive index = 1.45) could be determined using a three-layer planar model (air / PEG / silica) of the solid surface from the collected spectra. For each sample, the study performed at least 15 measurements at different locations.

[0149] Raman spectroscopy. The study characterized the presence or absence of an ice-like hydration layer on SLIC coatings using angle-resolved TIR Raman spectroscopy with a UV laser (Horiba XploRA PLUS). Spectral scans were performed at laser wavelengths of 266 nm and 366 nm. At an penetration depth of approximately 2 nm, Raman spectra can provide a detailed chemical fingerprint of the sample. The study used 3000–3400 cm⁻¹ to detect a clear peak difference between the more ordered ice-like hydration layer and the more disordered water-like hydration layer. -1The wavelength range used was as follows:

[0150] Contact Angle Measurement and Force-Tensiometer Measurement: The study evaluated quasi-static slipperiness by measuring the contact and sliding angles of whole blood and its components using optical angle measurement, and by measuring shear force using a force-tensiometer. The study measured the contact and sliding angles of whole blood and its components using 20 μL of adhesive droplets with a contact angle goniometer / tensiometer (Rame-Hart 260) inside a chamber with controlled humidity. The study evaluated dynamic slipperiness by measuring lateral adhesion using a force-tensiometer (Kruss Force Tensiometer K 100C). Through the measurement of lateral adhesion, substrates with a SLIC coating having an ice-like hydration layer may exhibit lower lateral adhesion compared to substrates without a SLIC coating. The study may perform at least 10 measurements for each substrate.

[0151] SLIC coating material interaction. The study evaluated the blood-material interactions of SLIC coatings fabricated under both static and dynamic conditions.

[0152] Preliminary results. Based on hydrophilic and slipperiness design principles, the study fabricated SLIC coatings on a wide variety of polymers, including rigid acrylic and flexible polyurethane (materials selected for LVAD). The study used polyethylene glycol (PEG; molecular weight approximately 500 daltons) brushed onto a polyurethane substrate via TEOS modification followed by silanization.

[0153] Figure 9A shows that PEGylation was evident from the high-resolution C1s XPS spectrum. PEGylation made the surface hydrophilic at a contact angle of θ = 35°. Furthermore, the surface was nm 2The ultra-high graft density of approximately 1.3 PEG brushes per unit area resulted in low chemical heterogeneity (based on polarization analysis measurements). Subpanel C of Figure 9A shows that blood-like droplets cannot slide off conventional hydrophilic surfaces, while subpanel E of Figure 9A shows that they can easily slide off the SLIC coating due to low contact angle hysteresis of Δθ < 10°, which results from low physical and chemical heterogeneity. Subpanels D and F of Figure 9A show that the LVAD surface with the SLIC coating had minimal adhesion of blood-like droplets. This was a macroscopic result of an ice-like hydration layer at the molecular scale, resulting in the removal of adsorbed proteins due to slipperiness and low adhesion.

[0154] Preliminary studies observed negligible fibrinogen adsorption with SLIC-coated surfaces incubated for one week, as shown in subpanels G and H of Figure 9A. Furthermore, the preliminary studies also showed virtually negligible platelet adhesion in subpanels I and J of Figure 9A, and leukocyte adhesion in subpanels K and L of Figure 9A, for the SLIC coating. This contrasted sharply with the rapid (approximately 1 minute) and at least an order of magnitude larger amounts of fibrinogen adsorption, platelet adhesion, and leukocyte adhesion observed on polyurethane surfaces. These in vitro preliminary studies constitute progress, as even state-of-the-art coatings cannot prevent protein adsorption for more than a few minutes.

[0155] In vitro analysis of thrombosis under quasi-static conditions.Some interactions occur after protein absorption and after the substance comes into contact with blood. The study characterized interactions under quasi-static conditions, with and without SLIC coating. The study obtained platelet-poor plasma (PPP) or platelet-rich plasma (PRP) by centrifugation from human citrated whole blood (ZenBio, NC). Whole blood could be centrifuged at 300 g for 15 minutes to obtain platelet-rich plasma (PRP, plasma containing platelets and leukocytes, and not containing erythrocytes). PRP was further centrifuged twice at 500 g at 35°C for 10 minutes to obtain platelet-poor plasma (PPP, plasma without platelets, leukocytes, and erythrocytes). The study utilized the obtained PPP in protein binding that occurs when blood comes into contact with a surface. The study used PRP to evaluate the later stages of blood-substance interactions. All biological experiments were performed on at least five surfaces with at least three different cell populations (nmin=15) from different donors.

[0156] Fibrinogen and albumin binding from PPP on different surfaces.Thrombin, a key protein responsible for platelet activation, initiates the coagulation cascade. The enzyme thrombin converts fibrinogen to fibrin monomers, which, under normal conditions, can form high-molecular-weight fibrin fibers and thus blood clots (thrombi). Therefore, fibrinogen is associated with blood coagulation, and preventing its adsorption can prevent thrombus formation. To evaluate procoagulation activity, the study used a human fibrinogen ELISA assay to assess fibrinogen binding from PPPs on different surfaces. Albumin also acts as a carrier for (1) antithrombin, which prevents the coagulation enzyme thrombin from acting unless needed, and (2) heparin cofactor, which was necessary for the anticoagulant effect of heparin. Quartz crystal microbalance (QCM) was used to quantitatively assess total protein adsorption from PPPs on the surface. QCM was a highly sensitive technique that measured nanograms of adsorbed protein on the surface based on changes in the vibration frequency of the quartz crystal. Additionally, protein adsorption from PPPs on different surfaces was measured by quantitative XPS, determining the contribution of the NC=O (amide) peak, which is characteristic of proteins adsorbed on the surface.

[0157] The adhesive force of proteins as measured by atomic force microscopy. The adhesion between the protein and the surface was a measure of how strongly the protein was adsorbed or bound to the surface. To estimate how strongly the protein was adsorbed or bound to the surface, the study estimated the adhesion using atomic force microscopy (AFM). Silicon nitride AFM tips were adsorbed with known concentrations (e.g., 10 mg / ml) of fibrinogen and albumin solutions. Each modified AFM tip with fibrinogen or albumin was then brought into contact with (i.e., engaged with) and then separated from (i.e., disengaged from) different surfaces, and the force required to separate from the surface could be measured by immersing the surface in PBS. The study then imaged the protein morphology in tapping mode using PBS and air for comparison. The study performed measurements at 30 random locations on each surface.

[0158] Platelet adhesion, activation, and platelet-leukocyte complex formation. The study stained cells with calcein-AM biostaining to evaluate platelet adhesion on surfaces different from those of PRP. The study utilized an LDH assay to characterize platelet viability. Platelet activation was assessed via SEM. Platelet and leukocyte activation was investigated for specific marker proteins, P-selectin (for platelets) and CD-45 (for leukocytes), by immunofluorescence staining and Western blotting. P-selectin expression by platelets can play a role in the initial recruitment of leukocytes to inflammatory sites, leading to platelet aggregation. CD-45 is a transmembrane protein present in human leukocytes that plays a role in signaling.

[0159] Contact and complement activation. Contact activation can interfere with the blood compatibility of blood-contacting substances. Proteins involved in the contact activation system (e.g., factor VII, factor IX, prekallikrein) are fibrinolytic, anti-adhesion, procoagulant, and pro-inflammatory. In this study, all kallikrein was released from the surface of the material due to exposure to PRP, so the activity of the kallikrein-α2-macroglobulin complex was evaluated using an antigen assay to determine the degree of contact activation.

[0160] Thrombin generation from PRP. Thrombin is an enzyme in the coagulation cascade whose measurement provides direct information about the thrombogenicity (i.e., its ability to form a thrombus) of a material. In normal plasma, thrombin can be trapped in fibrin meshwork and is rapidly inactivated by antithrombin III or other antiproteases. Thrombin's short half-life can interfere with its accurate enzymatic measurement. The study assayed surface-exposed PRP to determine thrombin activity using a thrombin generation assay.

[0161] Whole blood coagulation and hemolytic activity.To evaluate whole blood coagulation, the study dropped 5 μl of blood onto different surfaces. The blood allowed to coagulate for up to 60 minutes, and the concentration of free hemoglobin in the uncoagulated blood was measured at 15-minute intervals. The absorbance values ​​were positively proportional to the concentration of free hemoglobin in deionized water and were an indirect measure of the surface's ability to promote coagulation. The study investigated the hemolytic activity of different surfaces using a material hemolysis assay.

[0162] In vitro analysis of thrombosis under dynamic flow conditions. The study evaluated the ability of SLIC coatings to inhibit thrombus formation in real time under physiologically relevant flow conditions using a custom fluid engineering-based microscopy approach. Glass slides and PDMS devices with antithrombotic coatings were assembled on the slides. The devices were then mounted on a wide-field microscope with an automated stage for imaging, during which whole blood, plasma, or isolated coagulation components (i.e., fibrinogen and thrombin) could be introduced into the devices. The study included fluorescently labeled fibrinogen to enable monitoring of thrombus formation / growth. The study involved 1, 10, 100, and 1000 s. -1 Images were acquired at 5-minute intervals for 12 hours to monitor thrombus formation, growth, and flow velocity at physiologically relevant shear rates.

[0163] Ex vivo analysis of thrombosis in rabbit models.Through an iterative approach between the fabrication and characterization of SLIC coatings and in vitro analysis of thrombosis, the study identified and evaluated two promising SLIC coatings in a rabbit model of extracorporeal circulation. The study tested three experimental groups using rigid acrylic and flexible polyurethane tubing (depending on the results of the in vitro study) with and without the two SLIC coatings. The SLIC coating can coat the entire blood-contact area of ​​the tubing. The study utilized New Zealand white rabbits (2.5–3.5 kg) with groups evenly divided between male and female rabbits. After induction of anesthesia, the left carotid artery and right jugular vein were isolated, and the arteriovenous / venous connections could be made by cannula insertion into the left carotid artery and right jugular vein, as shown in Figure 9B of the extracorporeal circuit. Blood flow through the shunt was initiated by releasing the clamps on the arterial and venous sides of the extracorporeal circuit, and blood flow was then monitored for 2 hours, at which point the tubing was removed. The study photographed cross-sections of the tubing to determine the occlusion rate of the circuit. The study involved collecting, weighing, and preparing thrombi in tubes for SEM analysis.

[0164] Characterization of blood damage performance. The study used in vitro experiments to evaluate the performance and potential for blood damage of flexible rotors. Rotor blades with hardnesses ranging from Shore 10A to 100A were studied. Flexible, colorless, transparent polyurethane (BJB Enterprise, Tustin, CA) was used for the rotors to observe the rotor passages. Acrylic casings were machined and coated with a layer of the same polyurethane. Acrylic is a low-cost material with blood compatibility and may be fabricated by additive manufacturing in the future to further reduce costs. A shaft-mounted torque meter was used to monitor input power, and efficiency was calculated using pressure and flow rate data. All flexible rotors were tested at 2000–6000 RPM and 2–8 L / min to determine their performance, and the results were compared to those obtained from rigid rotors. The study focused on the performance and efficiency of flexible rotors outside of design conditions and compared them to their rigid counterparts.

[0165] Durability testing of flexible rotors. The blades are flexible, raising concerns about their durability. The study used custom-made accelerated fluid dynamics testing equipment. Briefly, the rotor was continuously operated at 1.5 times its normal speed (>4500 RPM) and under twice the normal pressure rise (>180 mmHg). Performance was continuously monitored by a data acquisition system, and the time to failure was recorded.

[0166] Performance under pulsating inflow conditions. The study tested a flexible LVAD with a stent-placed inlet and SLIC coating under pulsating inflow conditions. The study deployed and fixed the model downstream of a left heart simulator. In the simulator, the bladder pump was driven by compressed air and controlled by a solenoid valve to generate pulsating flow. Physiological and pathological waveforms were obtained by adjusting the resistor and compliance chamber downstream of the pump. The study used flow and pressure waveforms that mimicked those from end-stage heart failure patients. The LVAD inlet was subjected to these pulsating inflow conditions, but the pump pressure rise and efficiency were recorded at 3000 RPM. The average efficiency over at least 100 cardiac cycles can be used to determine the effectiveness of the flexible rotor.

[0167] Flow field measurement and direct blood damage quantification.The study used a selection of 10 flexible rotors with the highest average efficiency. The study used high-resolution 2DPIV to visualize the flow in the rotor passages and measure the shear rate in the region of interest. Briefly, the flow was seeded with microparticles coated with a fluorescent dye. To facilitate visualization, a blood-mimicking fluid (water-glycerin-NaI) with the same refractive index (n=1.49) as the acrylic casing and the same viscosity as blood could be used. The laser beam emitted from a high-speed Nd:YLF laser was transformed by a series of optics through a thin sheet (less than 1 mm) illuminating the flow. A high-speed camera positioned perpendicular to the laser sheet recorded images of the fluorescent particles through a bandpass filter that blocked all other undesirable reflections. The high-resolution flow field was calculated using the cross-correlation-based coding package, DaVis 10 (LaVision GmbH). The study obtained measurements in a series of axial planes covering the inlet-outlet of the LVAD model. For each plane and rotor blade orientation (phase), 1000 instantaneous visualizations were captured to ensure convergence of statistics such as turbulent kinetic energy and Reynolds stress. These were used to calibrate the turbulence model in CFD. The mean shear rate was calculated directly by applying the correlation sum algorithm described by Westerweel et al. The resolution of the measurements can reach up to the size of a single camera pixel (approximately 10 μm), which is more than sufficient for accurate shear rate quantification, even at the rotor tip gap.

[0168] Characterization of the antithrombotic response. The study characterized the thrombotic response of all flexible LVADs with SLIC coatings fabricated under quasi-static and dynamic conditions. The study used a minimum sample size of n=10 in each test and compared results from rigid rotors with and without SLIC coating.

[0169] Quasi-steady condition.Under quasi-steady-state conditions, human whole blood was centrifuged to separate plasma from red blood cells. A flexible rotor was incubated in plasma. Adsorption of fibrinogen, albumin, and immunoglobulin-G onto the LVAD material was evaluated using ELISA to understand how serum proteins interact with the surface, leading to an assessment of the coagulation properties of the sample and their interaction with whole blood. Human whole blood was placed on the sample to allow coagulation, and the free hemoglobin concentration could be measured. The study imaged the surface via SEM to visualize fibrin thrombus formation.

[0170] Under dynamic conditions, approximately 150 mL of human whole blood was anticoagulated with 3.2% sodium citrate (blood-to-citrate volume ratio = 10:1), introduced into a flow loop, and remineralized (citrate-to-Ca2+ molar ratio = 8:1). Pressure transducers upstream and downstream of the test LVAD measured the pressure. The pump was operated at a constant RPM, varying the resistance conditions of the loop to simulate the complex inflow in a real patient. The loop resistance was regulated by a pneumatic pinch valve placed around the tubing upstream of the pump, and a sinusoidal waveform may be used to control the valve. The study used D-dimer (fibrin degradation product) and thrombin-antithrombin complex (TAT, a protein complex of thrombin and antithrombin) to quantify thrombogenicity. The study measured concentrations using commercially available ELISA kits. Plasma free hemoglobin (pfHgb) and plasma lactate dehydrogenase (LDH) levels were used to assess hemolysis.

[0171] Dynamic conditions.The study also characterized the above thrombotic reactions under dynamic conditions. Briefly, approximately 150 mL of human whole blood was anticoagulated with 3.2% sodium citrate (blood-to-citrate volume ratio = 10:1), introduced into a flow loop, and remineralized (citrate-to-Ca2+ molar ratio = 8:1). The study used a linear actuator to physiologically drive the blood flow. The piston was housed in a titanium tube connected downstream to a suitable test chamber containing an LVAD. A bypass tube facilitated the return of blood to the titanium tube during diastole. Pressure transducers upstream and downstream of the test valve measured the pressure. A latex bag housed in an acrylic chamber (downstream of the leaflet) served as a compliance chamber. The study demonstrated that a pulsating blood loop could capture the initial biochemical changes associated with blood coagulation.

[0172] Design and testing of wireless energy transmission and communication systems The study evaluated the design, placement, and control of antennas for energy transmission systems. Figure 10A shows a COMSOL6.0 coil model with a backplane layer for improved coupling coefficient. Improved design of internal and external antennas, as well as the backplane, can increase energy efficiency and reduce heating losses.

[0173] Finite element analysis was used to design a flexible antenna that conforms to the body's contours to reduce losses due to gap variations from the outside to the inside. Preliminary studies included finite element analysis of the electromagnetic field using COMSOL 6.0, SPICE circuit analysis of the rectifier design (Figure 10C), and verification by physical testing. Each simulation was performed with different gaps between the two antennas to represent varying skin and tissue thicknesses across the population. The antenna geometry was measured and modeled using COMSOL. The antennas were set parallel to each other with gap distances of 1 cm to 5 cm, at 1 cm intervals. To reduce computation time by utilizing axial symmetry, an axisymmetric geometry was used, as shown in Figure 10A.

[0174] A field-programmable gate array (FPGA) programmable hardware board equipped with a variable-gain power amplifier (PA) was used to generate the input signal to the transmitting antenna (Figure 10D).

[0175] The study used two identical antennas with the same geometric shape and inductance values. Three features were extracted from the results for each gap size: the inductance of the antenna (L), the mutual inductance between the two antennas (M), and the coupling parameter (k). The k parameter is the ratio between the mutual inductance between the antennas and the inductance of the antennas. The closer the k parameter is to 1, the more efficient the power transmission between the antennas becomes.

[0176] Wireless transmission utilizes non-flexible materials in its design. Flexible antennas can allow for better conformity to the human body, increasing comfort. The effect of curvature on the coupling between antennas can be modeled using COMSOL 2D geometric shapes. The curvature of the antenna can be done along a circular arc while keeping the antenna length the same. To change the curvature, the radius of the circle can be adjusted. The smaller the radius, the greater the degree of curvature.

[0177] Figure 10B shows results from finite element analysis simulations, indicating that the gap distance between the two antennas increases as the coupling coefficient decreases. This means that the efficiency of power transmission from transmitter to receiver may also decrease. The study developed a custom flexible antenna with appropriate backplane material to improve efficiency and conform to the patient's body. The study determined the size, shape, and number of coil windings for improved energy transmission.

[0178] The study collected data from several animal studies, including (i) conduction velocity (the speed at which depolarization waves travel through the myocardium; this may aid in identifying areas of slow conduction requiring correction), (ii) myocardial voltage (the voltage of the myocardium may aid in identifying the degree of injury or scarring in a quantifiable manner), (iii) inductive threshold (the lowest level of either pacing or medication that produces a persistent arrhythmia), and (iv) depolarization and repolarization heterogeneity: differences in conduction between different cardiac regions, helping to specify areas of injury and scarring.

[0179] The study performed data acquisition using standard electrophysiological studies utilizing a high-resolution mapping catheter (HDGrid, Abbott, IL) and its associated electrophysiological mapping study system (Ensite, Abbott, IL). Electroanatomical mapping provides high-resolution data for evaluating the electrical conduction of the myocardium around the LVAD and other chambers of the heart, with the LV and LA being of greatest interest.

[0180] The study performed electroanatomical mapping before and after LVAD implantation, and at the end of longer chronic implantation time points, to establish a baseline for comparison, assess acute changes, and evaluate long-term outcomes in changes to the native conduction pattern. Additionally, 12-lead surface ECG may be utilized for overall conduction assessment over the duration of all animal studies. Finally, 12-lead data may be used as a surrogate to map data in future LVAD iterations, and LVAD may use the sensed 12-lead data to optimize pump settings and improve cardiac health.

[0181] Evaluation of heart rate-synchronized velocity modulation in continuous flow LVADs using a simulated circulatory loop. The study evaluated cardiac-gated velocity modulation. Current continuous-flow LVADs (CF-LVADs) generate flow that is contrary to the natural pulsating flow of the heart and vascular system, leading to a decrease in aortic blood pressure, which is associated with increased internal bleeding. The study used sensor inputs to modulate the velocity of the CF-LVAD in synchronization with the cardiac rhythm. The study also used a Frank-Stirling controlled simulated circulatory loop (MCL) to evaluate the hemodynamic response of velocity modulation.

[0182] The study used a 3D-printed centrifugal CF-LVAD driven by a brushless DC motor. The study developed an algorithm to control the CF-LVAD velocity using a left ventricular pressure sensor input to time velocity changes. The CF-LVAD controller and MCL controller operated independently of each other. The CF-LVAD inlet and outlet were attached to the left ventricle and aorta of the MCL, respectively. The MCL heart rate was set to a severe heart failure condition with a heart rate of 60 BPM.

[0183] Figure 11 shows three timing test cases in addition to the baseline without the LVAD. In the single-velocity flow, the pump operated at a set velocity of 2200 RPM throughout the entire cardiac cycle. The velocity was set to reach a cardiac output of 6 L / min. In the co-pulse case, the LVAD reached a peak velocity of 2600 RPM during systole and a minimum velocity of 1900 RPM during diastole. In the counter-pulse case, the LVAD reached a minimum velocity of 2000 RPM during systole and a peak velocity of 2500 RPM during diastole. In the velocity modulation case, the minimum velocity was set to prevent backflow through the CF-LVAD, and the peak velocity was set to reach an average cardiac output of 6 L / min.

[0184] To evaluate performance, the study performed MCL for 2 minutes each in a baseline case and in three CF-LVAD velocity timing cases. The study used an MCL sensor to measure and record aortic pressure, total cardiac output, and left ventricular volume and pressure.

[0185] As shown in Figure 11, all three CF-LVAD cases were able to increase total cardiac output without placing a load on the heart. By using copulse velocity modulation on the CF-LVAD, the aortic pulse pressure more than doubled compared to the other two CF-LVAD cases and reached 87% of the baseline aortic pulse pressure. Copulse velocity modulation achieved increased cardiac output without placing a load on the left ventricle, and in addition, maintained pulsating flow and pulse pressure closer to physiological standards.

[0186] The study evaluated control algorithms for improved dynamic LVAD performance, including electrophysiological investigations to provide data on the effect of LVADs on electrical conduction patterns before and after implantation. Data analysis can be used to improve LVAD pump speed control under various dynamic conditions, leading to better long-term outcomes.

[0187] Numerical modeling of BLDC motors. The study modeled a BLDC motor using the FEM software package (Comsol 5.0). The outer diameter of the BLDC stator was constrained to 35 mm. The study evaluated different gap sizes between the rotor and stator to analyze the motor torque, as well as the forces generated by radial and axial displacements. The diameters of the stator and rotor were optimized to ensure that flux saturation does not occur at a motor operating speed of 3,000 rpm. The study optimized the length of the BLDC motor to ensure that an appropriate motor torque could be generated to meet the hydraulic torque requirements (a hydraulic torque of 10 mNm was generated based on hydraulic tests using test data). The FEM analysis provides an induced voltage that can be used to calculate the back electromotive force (EMF) constant kE and estimate the motor torque during operation. The study then calculated the torque constant. The air gap magnetic field Bg is also recorded for the analysis calculation of the magnetic force generated due to rotor eccentricity. The study then analyzed the magnetic forces present in incremental axial and radial rotor displacements to calculate the radial and axial stiffness of the BLDC motor.

[0188] Design and fabrication of magnetic bearing systems (MBS). The rotation controlled by the motor is defined as θz, the translation along the axis of rotation is z, the radial translation is x and y, and the rotation (tilt) of each axis is θ x , θ y The BLDC motor was chosen to control rotation, and magnetic levitation technology was selected to control the other five degrees of freedom. A radially passive magnet was fixed to the rotating impeller, and three copper wire permanent magnets were mounted near the inlet cannula on the opposite side of the stator. The study magnetized all permanent magnets radially for the passive magnetic bearing so that the repulsive force between the rotor and stator permanent magnets was radial (x, y) and tilt (θ). x , θ y ) is passively stabilized. According to Earnshaw's theorem, a magnetic system that is passively stable in the radial direction can be unstable in the axial direction. If the axial position is stabilized by an active magnetic bearing characterized by magnetoresistance, the passive stiffness should be calculated for the rotor dynamic model. Furthermore, these stiffness values ​​depend on the coil current, and the negative radial stiffness of the motor should be included in the design. The design goal of the passive magnetic bearing was that the rotor and stator would not come into contact even when the rotor was subjected to accelerations exceeding 5G. With a rotating impeller weight of 5g (0.005kg) and a rotor magnet weight of 50g (0.05kg), the force required to generate the axial magnetic bearing under an acceleration of 5G was 0.055kg × 9.8m / s² × 5 = 27N. In summary, all six degrees of freedom are stabilized, θz is rotated by BLDC, and x, y, θx, and θ y z was passively stabilized, while z was actively stabilized.

[0189] Figure 12A shows an overview of the test stand for motor force measurement. To obtain precise force measurements with high spatial fidelity, a force transducer (ATI Industrial Automation, Apex, NC, USA) with a resolution of 1 / 512N was mounted on a micrometer-driven xyz stage (MiSUMi Group, Inc., Tokyo, Japan).

[0190] The stator section was fixed to a force transducer. The rotor was set up on a highly precise xyz stage moved by a micrometer head. Rotor displacement was measured with a laser displacement sensor with accuracy higher than micrometers. In this experiment, it was possible to accurately measure radial stiffness and negative axial stiffness. Several different permanent magnet designs were tested with various rotor thicknesses and diameters, all fabricated with rare earth magnets (NeFeB). The permanent magnets were measured for static performance. Accurate measurement of the radial and axial stiffness of the permanent magnets was crucial for the design and control of axial magnetic bearings.

[0191] The maximum axial displacement of the rotor must be ±0.2 mm, and the axial repulsive force was the sum of the electromagnetic force between the motor stator and rotor and the rotor force due to gravity. An axial active magnetic bearing should balance this force to achieve stable non-contact rotor suspension. To maintain the patient's quality of life, the rotor should remain stably suspended during daily human activities, and the magnetic levitation system should operate at 5W or less to improve battery life. FEM magnetic field analysis allows for precise testing, including stator geometry, stator material, number of coil windings, diameter of winding wires, and size of permanent magnets. Based on the simulation results, the study designed the entire magnetic levitation motor, including the displacement sensing system, using a 3D CAD system.

[0192] Stabilization control and performance evaluation of magnetic levitation motor systems. The research implemented a magnetic levitation and rotation controller. Figure 12B shows a diagram of the magnetic levitation motor control system. The control system consists of a magnetic levitation motor, a displacement sensor, an analog-to-digital converter (ADC), a controller, a digital-to-analog converter (DAC), and a power amplifier.

[0193] Position feedback is provided by sensors to actively control the axial direction. Displacement signals were fed into the controller via an ADC. A high-speed digital signal processor calculates the magnetic levitation current to reposition the rotor and feed its value back to a current amplifier via a DAC. The amplifier then returns a specified control current to the magnetic levitation coil to dynamically generate the regulated suspension force. The study developed the controller in a MATLAB / Simulink environment and implemented it on a dSPACE controller board. The study measured the static and dynamic characteristics of the magnetic levitation motor system in response to a series of perturbations, including both rigid (patient motion) and fluid (hemodynamic) stimuli.

[0194] Experimental testing of a motor drive system. Based on the numerical results, the study developed and manufactured the motor rotor and stator separately. The motor was driven by a pulse-width modulation (PWM) amplifier (Koford Inc., S24V10A-H3) and a power supply (TKD-Lambda Inc., CME350A-24). The PWM amplifier was used when the motor speed was 0 min -1 ~Approximately 26,000 min -1 Because the range is 0V to 5V, the motor's target speed is specified by a reference input voltage. A PWM amplifier supplied control current to the BLDC motor via a terminal box. The terminal box connected a power meter (Yokogawa Electric Corporation, WT-1800), a dynamic motor torque test tool, the BLDC motor, and the PWM amplifier. A motor mounting jig fixed the motor to a base with a motor coupling connected to the motor and a torque meter (Sugawara Laboratory, TB-200NM). A controller (Sugawara Laboratory, DMC-2) controlled the torque meter, specifying an automatic measurement sequence. The study allows for evaluation of motor torque and efficiency.

[0195] Characterization and verification of the design.To test the levitation characteristics, the study measured the impulse response in 40 wt% glycerol at a rotational speed of 3,000 rpm at 24°C. An external electromagnet generated an impulse disturbance in the levitation rotor via the pump casing. The controller's response was evaluated by measuring the time it took for the vibration amplitude to converge to less than 5% of the maximum fluctuation after the disturbance was applied. Since the axial suspension was actively controlled, the axial responsiveness could be adjusted until it was sufficiently fast. In the radial direction, the settling time could be longer because the passive permanent magnet bearings providing radial assistance do not generate damping force. However, the damping effect of the glycerol solution should improve radial stability.

[0196] To test the levitation-rotation characteristics, the study measured the vibration amplitude in a glycerol solution environment while the rotational speed was varied. Eddy current sensors measured axial displacement, and additional laser displacement sensors measured radial position and tilt motion. The study gradually increased the rotational speed up to a maximum of 5,000 rpm. After the rotor reached a steady speed, the study recorded the vibration amplitude of the rotor's motion. Axial displacement, radial displacement, and tilt amplitude increased at each resonant frequency. However, fluid viscosity can dampen rotor vibration. From the viewpoint of the stability of magnetic levitation control, a viscous liquid was desirable. The study adjusted the magnetic levitation control gain in a static state. However, when the rotor levitates and rotates, vibrations of a frequency synchronized with the rotational speed occur, which can destabilize the control. Therefore, it may be necessary to apply a tracking filter or adaptive control to reduce vibrations synchronized with the rotational speed.

[0197] System evaluation of SLIC LVAD on a simulated circular loop. Simulated circulatory loops (MCLs) enable in vitro hemodynamic assessments[58'][59'] and are used to evaluate both total artificial hearts[60'] and LVADs[61'][62'] implanted in various locations.

[0198] Figure 12C shows that a pulse test of the MCL provides dynamic fluid perturbations that emulate what the pump may experience due to hemodynamic effects. The MCL has characteristic resistance, arterial compliance, peripheral resistance, inertial components, and venous compliance. The heart has four chambers, whose performance is characterized by their volume and contractility. The volume of each ventricle can be calculated in real time by using a magnetostrictive level sensor. Combining the volume with the measured pressure allows for the determination of instantaneous pressure-volume relationships throughout the cardiac cycle. The left and right ventricles are contracted by injecting compressed air into the chambers. Contractility was controlled by an electro-pneumatic regulator (ITV2010-21N2S4, SMC Corporation, Tokyo, Japan). For clinically relevant results, the MCL should be able to replicate the Frank-Stirling mechanism. Passive filling enhances the functionality of the Frank-Stirling mechanism. Because the chambers are passively filled, the end-diastolic volume (EDV) depends on the state of the rest of the system. Contractility was calculated by a proportional controller that depended on EDV measurements and was updated with each cardiac cycle.

[0199] The pumps were mounted on a voice coil shaker system (MB Dynamics, OH, USA), and each was shaken radially and axially during operation to simulate patient movement. Each perturbation was tested at multiple frequencies to identify potential resonant frequencies or flutter instabilities. Impeller rotation speed and power consumption were monitored as metrics of success; deviations from the objective indicated that the levitation system failed to maintain system stability. This experiment provided a comprehensive evaluation of the MBS performance in the device.

[0200] In vitro blood compatibility evaluation of a combined SLIC LVAD device. Figure 12D shows the in vitro blood circulation loop for blood compatibility testing. Blood compatibility testing was performed in the circulation loop at a steady flow rate according to ASTM standards. The study utilized a total blood volume of 450 ± 50 mL in each test and used a temperature-controlled water bath to ensure that the blood temperature was maintained at 37 ± 1 °C. The blood flow and pressure differences between pumps were maintained at 5.0 ± 0.5 L / min and 100 ± 5 mmHg, respectively. The SLIC LVAD pumped citrated blood for 6 hours, and blood samples were collected every 60 minutes at intervals. Blood pressure, temperature, flow rate, rotor speed, and power consumption were recorded throughout the experiment. The study performed blood compatibility testing on the SLIC LVAD. The study also performed blood compatibility testing using a Rotaflow blood pump as a control pump for comparison with the SLIC LVAD device. Additionally, the remaining blood was stored in a blood bag and completely submerged in a water bath as a static control.

[0201] The study collected loop blood samples, centrifuged them, and transferred the plasma supernatant from each aliquot to three cuvettes, which were then diluted in 0.1% Na2CO3 solution. Absorbance could be measured at three wavelengths (380, 415, and 450 nm), from which plasma-free hemoglobin and normalized hemolysis index (NIH) could be calculated. The study utilized a flow hematology calculator for the quantitative measurement of activated platelets and leukocytes. The study also stained whole blood samples with CAPP2A, CD42b, and CD62P antibodies to quantify platelet activation, and with CD45, CD11, CD14 antibodies, and 7AAD dye solution to quantify leukocyte activation. In addition, platelet function was assessed by agglutination assays using standard agonists of collagen, ristocetin, ADP, and TRAP-6. For vWF multimer analysis, plasma protein sizes can be separated using an agarose gel for electrophoresis and then transferred to a polyvinylidene difluoride membrane using capillary blotting techniques. The membrane can be incubated with anti-human vWF antibodies and finally visualized and quantified using concentration assays. For data analysis, the study calculated the mean and standard deviation for each hemolysis, platelet activation, and function, leukocyte activation, and vWF multimer degradation. Analysis of variance can be applied to determine significant differences between the tested geometric shapes compared to the differences in bovine blood used for each test.

[0202] SLIC LVAD devices can introduce nominal levels of blood trauma due to surface roughness or retention points in the system. Therefore, blood was initially driven by a Rotaflow blood pump through a SLIC LVAD device with the rotor not in place, so that a blood trauma baseline could be established. The Rotaflow blood pump has a well-documented blood compatibility profile, allowing for the separation of blood damage from the SLIC LVAD device and blood damage generated by the Rotaflow blood pump. A bovine model was chosen due to the required blood volume for each experiment (450 mL) and concerns regarding human bloodborne pathogens.

[0203] In vivo studies of SLIC LVAD in large animal models. To evaluate the hematopoietic compatibility of partial vascular assist devices (SLIC LVADs) in vivo, studies can utilize large animal models. Fifteen calves (10 for acute and 5 for long-term 30-day studies) can undergo open implantation of SLIC LVADs (one at a time to allow for prototype replication between studies) to evaluate hemodynamic performance. Calves were chosen as the animal model because (i) they are more similar in size to humans, (ii) they share similarities in hemodynamic physiological function and microcirculation function, and (iii) they are widely used in biomedical studies to model other human conditions. Clinical (human) investigations do not allow for a systematic investigation of organs without confounding effects such as prior illness, varying degrees of organ failure, concomitant medications and diseases, and age.

[0204] The study can utilize 4-8 month old calves weighing 75-100 kg. The calves can undergo intravenous anesthesia, intubation, and mechanical ventilation following standard procedures. The study can place venous and arterial lines for blood sampling and pressure monitoring in the jugular vein and mammary artery, respectively. After probe and catheter placement, the study can position the calf on the cardiopulmonary bypass for SLIC LVAD implantation. The SLIC LVAD can be implanted via thoracotomy and positioned transapically within the left ventricle. The outflow graft can be anastomosed to the descending aorta. The SLIC LVAD can be connected as desired, the lungs can be slowly reperfused, and ventilation can be resumed. Once the animal is successfully withdrawn from the bypass, the natural heart can assist circulation through the reduced workload achieved by the SLIC LVAD. Flow probes can be placed on the pulmonary veins, ascending aorta, and outflow graft. Pressure lines can be placed in the left atrium and ascending aorta for continuous monitoring. The chest can then be closed to maintain normal physiological function and chest pressure. Calves may remain anesthetized during a series of tests to thoroughly investigate the pump-host interaction for up to 24 hours. The study may collect serial blood samples for arterial blood gas estimation and perform blood compatibility tests to assess hemolysis, platelet activation, leukocyte activation, and vWF degradation and its activity. The hemodynamic performance of the SLIC LVAD can be measured at different pump speeds. Rotational speed, power consumption from the BLDC motor, power consumption, pump temperature, and pump graft flow can be recorded continuously. In addition, innate cardiac performance may be manipulated with vasoactivators and anisotropic agents to evaluate the performance of the SLIC LVAD under different physiological conditions. A phenylephrine (2 ug / kg) bolus may be given to evaluate the pump's performance in high systemic vascular resistance. This study may create an overactive state using dobutamine infusion (10-40 ug / kg / min). The study will evaluate each hemodynamic change for up to two hours, and may allow time between each physiological conditional manipulation to allow steady state to be restored.

[0205] In a 30-day non-GLP study, animals will recover after implantation. Throughout the postoperative course, calves may receive medication to prevent pain and antibiotic therapy to maintain and balance the intestinal microbiota and prevent gastric ulcer formation. The study may collect data postoperatively and may include measurements of pump variables such as velocity, output, and flow rate through the outflow graft. The study collected biometric data to monitor calf conditions, including heart rate, body temperature, arterial / venous blood gas, hematocrit and total protein levels, activated clotting time, and AoP. Blood samples, including standard blood tests, plasma free hemoglobin (PFHb) levels, and coagulation and chemical tests, may be taken as appropriate to monitor blood trauma or organ injury throughout the 30 days.

[0206] Machine Learning Optimization LVAD Determining patient suitability and the appropriate timing of intervention can be crucial for LVAD implantation and patient recovery. Deep learning tools that can help assess the severity of cardiac conditions in patients may assist professionals in decision-making. The study utilized electrocardiogram (ECG) signals, a fundamental tool in cardiology, as input data for the system.

[0207] Deep learning for ECG classification has been used in recent years, with convolutional architectures being widely employed for this purpose, demonstrating the potential and necessity of interpretability and uncertainty recognition for real-world applications. This study focused on the candidate evaluation problem and proposed ECG criteria for training a model. The study combined three models to obtain a report highlighting useful information about physicians. Furthermore, interpretability and uncertainty estimation were implemented to enhance the reliability and applicability of the system.

[0208] ECG diagnostic criteria.To identify ECG diagnoses, the study looked for right heart problems (up to 53% of LVAD patients have right heart failure after implantation) to select LVAD candidates for ECG diagnosis identification, while simultaneously determining that the problem is related to the left ventricle. The study divided ECG diagnostic criteria into three groups. The primary criteria were QRS duration, which had a risk / association with candidates starting at a duration of 110 milliseconds and increasing linearly, as well as left bundle branch block (LBBB), ventricular premature contractions (PVCs), left ventricular hypertrophy (LVH), premyocardial infarction (AMI), and congestive heart failure (CHF). Minor criteria were inferior myocardial infarction (IMI) and atrioventricular block (AVB), as well as potential contraindications related to right heart problems, including right ventricular hypertrophy (RVH), right bundle branch block (RBBB), and right atrial hypertrophy / overload (RAE). The output was the estimated probability of having each of these diagnoses. The study did not use a single numerical score as the final measure of candidates, but rather used more informative reports.

[0209] A system for evaluating LVAD candidates from ECG. The ML system included i) a single-direction classifier, ii) a 12-direction classifier, and iii) several parts of a semantic segmentation classifier. All datasets used were publicly available datasets on Physionet. The single-direction and 12-direction classifiers output predicted probabilities, which are reported as probability bins: 0-30%: not detected, 30-45%: not ruled out, 45-60%: consider, 60-75%: possible, 75-100%: consistent. The study used bins to provide a more flexible approach to decision-making by having the system's output in a more nuanced way, similar to real-world annotations that are often not binary. Interpretability results from Grad-CAM are available for the 12-direction classifier, and uncertainty recognition results from Monte Carlo Dropout are shown in the output report.

[0210] Single-induction model.Several diagnostics of interest can be correctly classified from single-lead ECGs, and high-quality datasets are available for them, along with pulsation level annotations. The model used for the single-lead classifier is based on a 1D convolutional neural network with residual blocks. Figure 13A shows the single-lead ECG classification mode. The model includes five layers with 1D convolutional layers, max pooling, dropout, and skip connections, followed by fully connected layers and S-shaped activation.

[0211] After using Keras Tuner to find the best combination of hyperparameters, a model with 64 filters, a kernel size of 6, a MaxPool size of 4, and 48 / 80 units in the last two high-density layers was used. Training was performed using the Adam optimizer for 30 epochs with a learning rate scheduled at exponential decay (initial learning rate of 0.0001, rate of 0.75) and a callback-monitored validation loss. The loss function used was binary cross-entropy.

[0212] 12-guided model. The model uses a single-inductance classifier, adding a sixth level of depth with a 1D convolutional layer, implementing a group of layers for each input channel, i.e., each induction. After concatenating the extracted features of 12 channels, a high-density layer was added along with a final S-shaped activation for multi-label classification. Training was performed using the Adam optimizer for 15 epochs with a learning rate scheduled at exponential decay (initial learning rate of 0.0001, rate of 0.75) and a callback-monitored validation loss. The loss function used was binary cross-entropy.

[0213] Semantic segmentation model.For semantic segmentation of ECG signals, the study fitted a U-net model to process the 1D signal of each ECG lead and determine the mean QRS duration in milliseconds. It consists of five levels: a grouped max pooling layer, a 1D convolutional layer, and a dropout layer, with upsampling performed in the 1D transform convolutional layer. It receives single-lead waveforms of 1000 samples as input and outputs a segmentation mask of background, P-segment, QRS-segment, and T-segment for each of the 1000 samples from the last Softmax activation layer. Training was performed using the Adam optimizer for 85 epochs with a batch size of 64, at a learning rate scheduled with exponential decay (initial learning rate of 0.0001, rate of 0.75), and a callback-monitored validation loss. Classification cross-entropy was used as the loss function.

[0214] Dataset. The study selected public ECG datasets from Physionet to evaluate diagnoses of interest. To enhance the robustness of the model, each diagnosis of interest had observations from at least two datasets included in the model training. The datasets used included: (i) MIT-BIH arrhythmia database: normal, other heartbeats, LBBB, RBBB, PVC; (ii) MIT-BIH supraventricular arrhythmia database: normal, other heartbeats, PVC; (iii) PTB diagnostic ECG database: normal, AMI, IMI, CHF; (iv) PTB-XL: normal, LBBB, RBBB, LVH, RVH, AVB, RAE / RAO, AMI, IMI; (v) Lobachevsky University electrocardiogram database (LUDB): normal, LBBB, RBBB, LVH, RVH, AVB, RAE, PVC, AMI, QRS duration (semantic segmentation of P, QRS, T segments); and (vi) BIDMC congestive heart failure database: CHF.

[0215] The study ran training, validation, and test split sets with approximately 0.7 / 0.2 / 0.1 splits for single-lead and semantic segmentation models, and 0.8 / 0.1 / 0.1 splits for 12-lead models. In all cases, the study stratified the splits, maintaining the same ratio across all classes in each subset of data. In all cases, the study used a patient-to-patient paradigm, i.e., patient ECG data (heart rate) could be used in one of the training, validation, or test sets to avoid data leakage. For two minority classes (RVH and RAE), augmentation was performed by making a slight stretch to extend the signal in the time domain (with a random factor of 1.05 to 1.3) and by performing scaling (with a random factor of -0.875 to 1.125). Augmentation was applied only to 12-lead classifiers.

[0216] As a preprocessing step, for single-lead classifiers, the study detected R peaks and segmented the pulses. For both single-lead and 12-lead classifiers, the study removed noise from both using a fourth-order Butterworth high-pass filter and wavelet filtering. For single-lead and semantic segmentation models, the study applied minimum-maximum normalization.

[0217] Model evaluation results.The study evaluated three models following the recommended metrics for each case. Precision per class, recall, F-1 score, and AUC were assessed. Table 1 shows the results obtained for the three models. Additionally, the crossed overunion (IoU) score was calculated for the semantic segmentation model, taking the length ratio instead of the region to measure the overlap between the predicted region and the ground truth region. The IoU scores obtained from the test set using the semantic segmentation model were 0.902, 0.741, 0.867, and 0.786 for the Other (background), P, QRS, and T segments, respectively. The weighted IoU score was 0.871. [Table 1]

[0218] output. The study implemented interpretability using Grad-CAM, applying it to the last 1D convolutional layer of each of the 12 heads of the model, one per input lead. Figure 13B shows an exemplary result of interpretability plotted on the 1D signals of the leads being tested in a heatmap.

[0219] The Monte Carlo (MC) dropout technique can provide a good approximation of the posterior probability distribution of the model. This was implemented in the system and plotted in a notched box plot. Observations were forward-passed 100 times through the model. During inference, the system output a report highlighting major contraindications, minor contraindications, and potential contraindications, along with the resulting spleness map. Physicians can also examine the box plots of reported criteria of interest, which also show the estimated uncertainty of the outcomes. Figure 13C shows examples of one high candidate outcome (a), a second one with a heart predicted as normal (b), a third one with a high potential contraindication (c), and the last one (d), which was a sample of high uncertainty observations. Specifically, subpanel A of Figure 13C read: Major - consistent with LBBB, possible AMI, QRS duration 167 ms. Minor - AVB cannot be ruled out. Subpanel B of Figure 13C read: Major - Consistent with NORMAL, QRS duration 112.5 ms (the next predicted probability is AVB with a lower value, shown for comparison). Subpanel C in Figure 13C read: Major - AMI cannot be ruled out, QRS duration 128.6 ms. Minor - Possible IMI, AVB considered. Potential contraindication - Consistent with RBBB. Subpanel D in Figure 13C read: Major - NORMAL (high uncertainty) considered, LVH considered.

[0220] The study outlined a step-by-step approach to evaluating LVAD embedding candidates directly from physiological signals, in this case, ECG. A multi-model deep learning system was constructed, achieving state-of-the-art results in each model and combining predictions into a report. The study confirmed the importance of interpretability and implemented methods to demonstrate the importance of each lead and segment. To enhance confidence in the model, the study implemented uncertainty computation and graphically reports the predicted probabilities and uncertainties.

[0221] Machine learning.Various analytical systems may be implemented using one or more artificial intelligence and machine learning operations. The term “artificial intelligence” may include any techniques that enable one or more computing devices or computing systems (i.e., machines) to mimic human intelligence. Artificial intelligence (AI) includes, but is not limited to, knowledge bases, machine learning, representation learning, and deep learning. The term “machine learning” is defined herein as a subset of AI that enables machines to acquire knowledge by extracting patterns from raw data. Techniques of machine learning include, but are not limited to, logistic regression, support vector machines (SVMs), decision trees, naive Bayes classifiers, and artificial neural networks. The term “representation learning” is defined herein as a subset of machine learning that enables machines to automatically discover representations needed for feature detection, prediction, or classification from raw data. Techniques of representation learning include, but are not limited to, autoencoders and embedding. The term “deep learning” is defined herein as a subset of machine learning that uses layers of processing to enable machines to automatically discover representations needed for feature detection, prediction, classification, etc. Deep learning techniques include, but are not limited to, artificial neural networks or multilayer perceptrons (MLPs).

[0222] Machine learning models include supervised, semi-supervised, and unsupervised learning models. In supervised learning models, the model learns a function that maps inputs (also known as features) to outputs (also known as objectives) during training with labeled data sets. In unsupervised learning models, the algorithm discovers patterns between data. In semi-supervised models, the model learns a function that maps inputs (also known as features) to outputs (also known as objectives) during training with both labeled and unlabeled data.

[0223] Neural network.An artificial neural network (ANN) is a computing system comprising multiple interconnected neurons (also referred to as, for example, “nodes”). In this disclosure, it is assumed that the nodes may be implemented using computing devices (e.g., the processing units and memory devices described herein). These nodes may be arranged in multiple layers, such as an input layer, an output layer, and one or more hidden layers having arbitrarily different activation functions. An ANN with hidden layers may be referred to as a deep neural network or a multilayer perceptron (MLP). Each node is connected to one or more other nodes in the ANN. For example, each layer may consist of multiple nodes, and each node is connected to all nodes in the previous layer. Nodes in a given layer are not interconnected with each other; that is, nodes in a given layer function independently of each other. As used herein, nodes in the input layer receive data from outside the ANN, nodes in the hidden layer(s) modify the data between the input and output layers, and nodes in the output layer provide the results. Each node is configured to receive an input, execute an activation function (e.g., a binary step function, a linear function, an S-shaped function, a hyperbolic tangent function, or a rectified linear unit (ReLU)), and provide an output according to this activation function. Additionally, each node is associated with its respective weights. The ANN is trained to maximize or minimize an objective function using a dataset. In some implementations, this objective function is a cost function, which is a measure of the ANN's performance during training (e.g., an error such as an L1 or L2 loss), and its training algorithm adjusts the node weights and / or biases to minimize the cost function. This disclosure is intended to show that any algorithm that finds the maximum or minimum value of the objective function may be used to train the ANN. Training algorithms for the ANN include, but are not limited to, backpropagation. It should be understood that the ANN is provided only as an illustrative machine learning model. This disclosure is intended to show that the machine learning model may be any supervised, semi-supervised, or unsupervised learning model.The machine learning model is, optionally, a deep learning model. Machine learning models are known in the art and are therefore not described in further detail herein.

[0224] A convolutional neural network (CNN) is a type of deep neural network that has been applied, for example, to image analysis applications. Unlike traditional neural networks, each layer of a CNN has a plurality of nodes arranged in three dimensions (width, height, and depth). A CNN can include different types of layers, such as convolutional layers, pooling layers, and fully connected (also referred to herein as "densely connected") layers. A convolutional layer includes a set of filters and performs most of the calculations. A pooling layer is optionally inserted between convolutional layers to reduce the computational load and / or control overfitting (e.g., by downsampling). A fully connected layer includes neurons, and each neuron is connected to all the neurons in the previous layer. These layers are stacked in the same way as traditional neural networks. A GCNN is a CNN adapted to operate on structured data sets such as graphs.

[0225] Other supervised learning models. A logistic regression (LR) classifier is a supervised classification model that uses a logistic function to predict the probability of an object, and the model can be used for classification. During training, the LR classifier is trained using a data set (also referred to herein as a "dataset") to maximize or minimize an objective function, such as an evaluation of the performance of the LR classifier (e.g., an error such as L1 or L2 loss). The present disclosure contemplates that any algorithm for finding the minimum value of the cost function can be used. LR classifiers are known in the art and are therefore not described in further detail herein.

[0226] A naive Bayes (NB) classifier is a supervised classification model based on Bayes' theorem, which assumes independence between features (i.e., the presence of one feature in a given class is independent of the presence of any other feature). An NB classifier is trained on a dataset by computing the conditional probability distribution of each feature given a label, and then applying Bayes' theorem to compute the conditional probability distribution of labels given observations. NB classifiers are already known in the art and are therefore not described in further detail herein.

[0227] A k-NN classifier is an unsupervised classification model that classifies new data points based on a similarity measure (e.g., a distance function). A k-NN classifier is trained on a dataset (also referred to herein as the “dataset”) to maximize or minimize a measure of its performance during training. This disclosure assumes any algorithm for finding the maximum or minimum value. k-NN classifiers are already known in the art and are therefore not described in further detail herein.

[0228] In addition, various analysis systems can be implemented using encoders, transformers, and other convolutional deep neural network architectures.

[0229] Consideration Cardiovascular disease (CVD) is the leading cause of death in the United States.[1'] Common cardiovascular conditions include heart attacks, strokes, arrhythmias, and valvular stenosis, all of which can ultimately lead to heart failure (HF). The total direct medical costs of HF are estimated to reach $53 billion by 2030.[2'] When HF reaches its end stage, there are no known drugs that are effective in treating it. Heart transplant surgery is the most standard treatment for end-stage heart failure, but it is severely limited by the availability of donor hearts. Furthermore, there are no readily available ways to increase the pool of donor hearts. As a promising alternative, left ventricular assist device (LVAD) implants have steadily increased from 2010 to 2021, reaching a total of 25,551 in the United States and are projected to reach 7,000 per year by 2030.[3'][4'] While small, implantable pumps help failing hearts maintain blood circulation throughout the body, current generations of LVADs are invasive, primarily due to percutaneous drive lines and frequent emergency pump replacement surgeries to treat thrombosis-related pump failures[5'][6']. Common complications, including infections, thrombosis, stroke, and bleeding, significantly reduce the quality of life for LVAD patients. Significant research efforts are needed to make LVAD therapy less invasive, more effective, and more accessible to the general public and military. The comprehensive challenges facing the development of less invasive LVADs are as follows: 1) Blood damage caused by mechanical forces and associated side effects, including bleeding and thrombosis requiring hospitalization and emergency surgery. Blood damage is associated with hyperphysiological shear stress in the device[3'][7'], where high mechanical forces destroy blood cells, causing hemolysis and activating platelets, leading to thrombosis. 2) Pump thrombosis, which often leads to emergency pump replacement and surgery. One of the underlying causes of LVAD failure is the presence of blood stasis points, which cause blood protein adsorption, followed by a chain reaction including thrombosis, ultimately leading to sepsis and death. 3) Percutaneous drive lines, which cause infection and usually lead to readmission and surgery. Furthermore, drive lines significantly reduce the patient's mobility, thereby lowering their quality of life.

[0230] To address the excess problems facing the current generation of LVADs, more than one innovation is needed to significantly improve clinical outcomes. The exemplary device incorporates hemodynamic, surface science, electronics, and blood compatibility innovations to achieve the shared goal of developing a less invasive, more durable, and effective LVAD to improve the quality of life for HF patients. This includes (i) improving blood compatibility and eliminating thrombosis issues at the inlet cannula by employing innovative design changes, including flexible blade / casing and a stent-placed inlet; (ii) using a slippery hydrophilic (SLIC) coating with unprecedented antithrombotic properties under both quasi-static and dynamic conditions of the LVAD surface to address thrombosis issues within the device; and (iii) using an external power supply and transmitter structure to provide wireless power to the communication link between the implanted pump and the centrifugal pump and to provide feedback to an external power supply system.

[0231] The elimination of percutaneous drive lines and the extremely low risk of thrombosis-related pump replacement surgery make long-term support with LVADs far less invasive. It also provides a significantly improved quality of life for patients supported by LVADs. Furthermore, the flexible rotor and casing design opens the door to the future development of minimally invasive LVADs that can be implanted via catheters. This innovation will have a significant impact on the treatment of heart failure patients and will benefit the American public.

[0232] Review of the literature. An exemplary device may use fabric to promote endothelialization in order to reduce the risk of thrombosis. In contrast, U.S. Patent No. 9,114,034, B2 appears to disclose coating a stent with an antithrombotic agent to reduce the likelihood of thrombosis. U.S. Patent Application Publication 2016 / 0303287 appears to disclose coating a stent with an antithrombotic agent to reduce the risk of thrombosis in stented heart valves.

[0233] Exemplary devices can be used to reduce flow stagnation and thrombosis near the LVAD inlet. In contrast, WO2020 / 127616A1 discloses an artificial heart valve having a stent structure with a conical convex inlet region and a linear cylindrical outlet region.

[0234] An exemplary device may utilize a flexible rotor to reduce blood damage and outcomes in LVAD patients. In contrast, U.S. Patent No. 8,449,443, B2 discloses a pump having a flexible rotor for assisting Fontan circulation. The propeller appears to be flexible to facilitate intravascular implantation by crimping.

[0235] conclusion It should also be noted that, as used herein and in the appended claims, the singular forms “a,” “an,” and “the” include multiple references unless the context expressly indicates otherwise. In this specification, ranges may be expressed as from a certain value, such as “about” or “approximately 5,” to another specific value, such as “about” or “approximately.” Where such ranges are expressed, other exemplary embodiments include one specific value and / or the other specific value.

[0236] "Comprising," "containing," or "including" means that at least the compound, element, particle, or method step of the name is present in the composition, article, or method, but does not exclude the presence of other such compounds, materials, particles, or method steps, even if those other compounds, materials, particles, or method steps have the same function as the one named.

[0237] In describing exemplary embodiments, technical terms will be used for clarity. Each term is intended to be understood by those skilled in the art in its broadest sense and to include all technical equivalents that operate in a similar manner to achieve similar purposes. It should also be understood that reference to one or more steps of a method does not preclude the presence of additional or intervening method steps between those explicitly identified steps. The steps of a method may be performed in an order different from that described herein without departing from the scope of this disclosure. Similarly, reference to one or more components within a device or system should also be understood that it does not preclude the presence of additional or intervening components between those explicitly identified components.

[0238] The following patents, applications, and publications, listed below and extending throughout this Spec, are incorporated herein by reference in their entirety. List of References Part 1 [1]US2016 / 0303287 [2]US10442166 B2 [3]US4906237 [4]US5662960 [5]US11285312B2 [6]US9814611 B2 [7]US2011 / 0276123 [8]US10434235B2 [9]US11065462B2

[10] US7520850B2

[11] Tsao CWet al.Circulation,vol.145,no.8,pp.e153-e639,2022.

[12] Sahni A.et al.Annals of Biomedical Engineering,2023

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Claims

1. An implantable left ventricular assist device (LVAD), A pump housing that forms a volume inside, the pump housing having an inlet and an outlet, A rotor is disposed within the pump housing so as to rotate within the volume for pumping blood, A motor coupled to the rotor to drive the rotation of the rotor, An implantable left ventricular assist device (LVAD) comprising: an inlet member coupled to the inlet of the housing, the inlet member having (i) an external structure configured to conform to and contact the outer wall at the base of the left ventricle; and (ii) an internal expandable body configured to move between a housing configuration and a deployable configuration positioned and extending within the left ventricle, the internal expandable body being defined by (i) a first section having a first circumference sized to conform to the inlet of the pump; and (ii) a second section having a second circumference sized to substantially contact the inner wall of the left ventricle in the deployable configuration to eliminate flow stagnation within the left ventricle.

2. The rotor is The rotor body and The device according to claim 1, comprising a plurality of curved flexible blades extending therefrom, wherein the plurality of curved flexible blades are flexibly molded and formed from a flexible material to reduce shear stress on the components of the blood.

3. The device according to claim 1 or 2, wherein the rotor includes an antithrombotic coating containing a hydrophilic and slippery polymer.

4. A drive system configured to drive the rotor so as to rotate within the volume of the pump housing, The drive system and the embedded control unit electrically coupled to the drive system for controlling the rotation of the rotor are further comprising, One or more energy storage devices, Charging circuit and The device according to any one of claims 1 to 3, comprising a wireless power transmission circuit having an antenna coil and electronic equipment for controlling inductive charging operation using an external RF source.

5. The aforementioned motor is One or more permanent magnets located within the rotor body, A stator is disposed in the pump housing at a position close to one or more of the aforementioned permanent magnets, The device according to any one of claims 2 to 4, comprising a magnetic levitation drive subassembly comprising one or more magnetic bearings disposed at one or more positions in the pump housing adjacent to the plurality of curved flexible blades of the rotor.

6. The device according to any one of claims 1 to 5, wherein the internally expandable body of the entrance member comprises a flexible mesh surrounded by or embedded in a fabric, and the flexible mesh is configured to be self-expanding or balloon-expanding.

7. The device according to claim 6, wherein the internally expandable body of the inlet member is antithrombotic and anti-restenotic drug eluting.

8. The device according to claim 6, wherein, after deployment, the internally expandable body is defined by (i) a first section having a first circumference sized to fit to the inlet of the pump, and (ii) a second section having a second circumference sized to fit to the inner wall of the left ventricle in continuous contact with it and to conform to the shape of the left ventricle.

9. The device according to claim 6, wherein the internally expandable body is configured to deform with ventricular motion over a predetermined set of deformation cycles.

10. The device according to any one of claims 1 to 9, wherein the fabric includes an absorbent material for promoting endothelialization.

11. The device according to any one of claims 1 to 9, wherein the internally expandable body includes an absorbent material for promoting endothelialization.

12. The device according to claim 10, wherein the fabric comprises polyester, polytetrafluoroethylene, or a combination thereof.

13. The device according to any one of claims 1 to 12, wherein the internal expandable body, including a flexible mesh and fabric, is patient-specific, so as to be sized and molded to match a scan of the patient's left ventricle.

14. The pump housing and / or the inlet member are The device according to any one of claims 1 to 13, comprising a quick-connect connector for removably and reattachably connecting the pump housing to the inlet member (for example, the device can be installed and deployed with the pump housing cut off from the inlet member, and the inlet member can be capped as necessary).

15. The device according to any one of claims 2 to 14, wherein the pump housing comprises (i) an external structure and (ii) an internally deformable structure disposed within a portion of the surface defining the volume of the pump housing, the internally deformable member comprising a flexible material.

16. The device according to claim 15, wherein the flexible material of the internally deformable member is the same as the flexible material of the rotor.

17. The device according to claim 15, wherein the flexible material of the internally deformable member is different from the flexible material of the rotor.

18. The device according to claim 15, wherein the portion of the surface defining the volume of the pump housing having the flexible material corresponds to a region of contact with the plurality of curved flexible blades during rotation of the plurality of curved flexible blades.

19. The device according to claim 15, wherein the portion of the surface defining the volume of the pump housing having the flexible material corresponds to a region of contact with the plurality of curved flexible blades between (i) the rotation of the plurality of curved flexible blades and (ii) the off-axis movement of the plurality of curved flexible blades configured to move off-axis via magnetic bearings and electromagnetic drive subsystems.

20. The device according to claim 15, wherein the external structure is made of a flexible material.

21. The device according to any one of claims 15 to 20, wherein the inner surface of the pump housing, including the internally deformable structure, includes an antithrombotic coating comprising a hydrophilic and slippery polymer.

22. The device according to any one of claims 3 to 21, wherein the antithrombotic coating on the interface surface of the pump housing and / or the antithrombotic coating on the rotor are optimized for maximum antithrombotic response through a systematic tailoring of hydroxylation parameters, molecular structure, and synthetic reaction coordinates.

23. The device according to claim 22, wherein the antithrombotic coating on the interface surface of the pump housing and / or the antithrombotic coating on the rotor are formed by liquid-phase silanization of the flexible material on the interface surface of the pump housing and / or the antithrombotic coating on the rotor to form a silane.

24. The device according to claim 23, wherein the silane is adjusted to (i) obtain a high graft density to the flexible material on the interface surface of the pump housing and / or to the antithrombotic coating of the rotor, and (ii) ensure antithrombotic properties.

25. The device according to any one of claims 22 to 24, wherein the coating is prepared by a synthesis reaction.

26. It also features an implantable subcutaneous coil, The device according to any one of claims 4 to 24, wherein the wireless power transmission circuit of the embedded control unit is operably coupled to the implanted subcutaneous coil via a drive line conductor, and the implanted subcutaneous coil is configured to operate as a pair of coupling coils with an external coil configured to be mounted (i) in the patient's skin area and (ii) in close proximity to the implanted subcutaneous coil.

27. The device according to claim 26, wherein the one or more energy storage devices include (i) one or more rechargeable batteries and (ii) at least one of a supercapacitor and a hybrid supercapacitor, and the one or more rechargeable batteries, and at least one of the supercapacitor and the hybrid supercapacitor, have combined energy storage for at least 20 minutes.

28. The embedded control unit comprises (i) a processor and (ii) a memory having instructions stored thereon, The device according to claim 26, wherein the execution of the instruction by the processor causes the processor to execute a dynamic charging algorithm.

29. The device according to any one of claims 4 to 28, wherein the embedded control unit comprises a wireless communication interface configured to operably connect to a remote controller.

30. The remote controller comprises (i) a processor and (ii) a memory having instructions stored thereon, The execution of the instruction by the aforementioned processor means that the processor Run a dynamic charging algorithm that includes a trained machine learning model or a model derived therefrom, and The device according to claim 29, wherein, in order to adjust the control operation, the output of the dynamic charging algorithm is transmitted to the embedded control unit.

31. The device according to claim 29, wherein the embedded control unit is configured to execute a dynamic charging algorithm.

32. The device according to claim 30 or 31, wherein the dynamic charging algorithm includes a dynamic motor control loop that reduces the LVAD speed during a period of less patient activity, and the dynamic motor control loop has an output for driving the operation of the motor.

33. The device according to claim 32, further comprising a sensor configured to acquire the electrical signals of the heart, wherein the dynamic motor control loop includes one or more inputs, including a first input for receiving the acquired electrical signals.

34. The pump housing is equipped with a position sensor for the rotor, The device according to claim 32, wherein one or more inputs of the dynamic motor control loop include a second input for receiving an electrical signal acquired from the position sensor.

35. The device according to any one of claims 30 to 25, wherein the trained machine learning algorithm is used to evaluate long-term LVAD performance data.

36. The device according to any one of claims 1 to 26, further comprising a magnetic levitation drive system including a magnetic levitation ventricular assist device (VAD), wherein the magnetic levitation ventricular assist device (VAD) comprises an impeller attached to a rotor permanent magnet, a stator embedded in the VAD housing adjacent to the inlet cannula, and an active magnetic levitation bearing located at the bottom of the VAD housing with an eddy current sensor configured to monitor the gap between the impeller tip and the magnetic levitation device.

37. The implantable left ventricular assist device according to any one of claims 1 to 36, wherein the implantable left ventricular assist device is selected to be implanted in a patient identified via a trained ML algorithm used to evaluate candidate evaluations.

38. The device according to claim 37, wherein the trained ML algorithm is configured to estimate the likelihood of the existence of candidate criteria for an LVAD implant.

39. The device according to any one of claims 5 to 38, wherein the stator together with the rotor forms a brushless DC motor, the optimization of the diameters of the stator and the rotor is performed via numerical simulation, and the length of the stator is optimized to generate motor torque that sufficiently satisfies hydraulic torque requirements.

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