EP4669369A1

EP4669369A1 - PERSONALIZED SUPPORT DEVICE FOR MECHANICALLY INFLUENCING THE DYNAMIC BEHAVIOR OF A PULSE INGREDIENT ORGAN

Info

Publication number: EP4669369A1
Application number: EP24709530.0A
Authority: EP
Inventors: Daniel Cohn; Amir OFFER; Dean NACHMAN; Nicola MANSOUR; Ziv PESELEV
Original assignee: Hadasit Medical Research Services and Development Co; Yissum Research Development Co of Hebrew University of Jerusalem
Current assignee: Hadasit Medical Research Services and Development Co; Yissum Research Development Co of Hebrew University of Jerusalem
Priority date: 2023-02-23
Filing date: 2024-02-22
Publication date: 2025-12-31
Also published as: WO2024176233A1

Abstract

A support device for a pulsating organ is comprised of a structure formed at least partially of one or more polymers. The structure is configured to transition, in response to a stimulus, between a placement shape, in which the support device is dimensioned larger than the dimensions of the organ, and a treatment shape, in which the support device provides therapeutic support to the organ; wherein the treatment shape is smaller than the placement shape. Preferably, the polymers are shape memory polymers. The support device may be manufactured in an initial shape in which the one or more shape memory polymers are in their permanent shapes, and the placement shape is larger than or equivalent to the initial shape. Optionally, the stimulus is application of an aqueous medium. Embodiments of support devices may be used for treatment of heart failure with reduced ejection fraction and with preserved ejection fraction.

Description

PERSONALIZED SUPPORT DEVICE FOR MECHANICALLY AFFECTING A PULSATING ORGAN’S DYNAMIC BEHAVIOR

RELATED APPLICATIONS

This Application claims priority to U.S. Provisional Patent Application No. 63/486,539, filed February 23, 2023, entitled “Personalized Support Device for Mechanically Affecting a Pulsating Organ Dynamic Behavior,” the contents of which are incorporated by reference as if fully set forth herein.

TECHNOLOGICAL FIELD

The present disclosure relates to the field of medical devices for treatment of pulsating organs, more specifically heart failure, and even more specifically, but not exclusively, to personalized shape memory devices that are configured to transition between different shapes, and that may be applied onto a pulsating organ, and particularly the left ventricle of a heart, in order to mechanically affect or modulate the heart’s dynamic behavior.

BACKGROUND OF THE INVENTION

Heart failure (HF) is a pathologic condition that leads to serious clinical symptoms, such as shortness of breath and peripheral edema, substantially lowers the quality of life, and engenders a high mortality rate. Heart failure is generally categorized into two different types, each with a different presentation, etiology, and pathophysiology. These types are heart failure with reduced ejection fraction (HFrEF) and heart failure with preserved ejection fraction (HFpEF). In heart failure with reduced ejection fraction, the left ventricle loses its ability to contract normally, and as a result the heart is unable to pump enough blood into circulation. In heart failure with preserved ejection fraction, the left ventricle pumps normally, but is too stiff to fill properly.

Ventricular remodeling is a particularly deleterious physical response to heart failure. Ventricular remodeling refers to changes in ventricular geometry, mass, and volume in response to myocardial injury or alterations in load. Significant ventricular dilatation or remodeling in patients with heart failure is a strong predictor of both morbidity and mortality. Treatments for heart failure may include medications, coronary bypass surgery, heart valve repair or replacement, insertion of a ventricular assist device, or cardiac resynchronization therapy.

Ventricular Restraint Therapy is an emerging investigational approach for treatment of heart failure, and in particular heart failure with reduced ejection fraction. In this type of therapy, a net-like device (known as a “ventricular support device” or “ventricular wrap”) is placed around one or both ventricles, exerting a small amount of pressure on the epicardium, to assist the ventricle in its pumping function. Ventricular restraint therapy aims to prevent, and reverse, ventricular remodeling resulting from heart failure.

Over approximately the past 25 years, various implementations of ventricular support devices have been investigated for heart failure with reduced ejection fraction. These devices may be categorized as “passive” devices, which exert suitable pressure due to intrinsic properties of the devices, and “active” devices, which operate through active control of the shape of the device (e.g., through pumping of a fluid or air during the heart cycle). Passive devices exhibit several advantages compared to active devices. These include lack of a need for a power supply, and the lack of contact with circulating blood, which thereby avoids complications of blood clot formation and bleeding.

Two “passive” ventricular support devices have been studied extensively. The CorCap device (Acorn Cardiovascular, USA) is a polyester mesh that encloses the entire heart surface from the apex to the atrioventricular groove. The CorCap device is placed surgically, and is adjusted snugly by the surgeon at end diastole, to thereby provide circumferential support to the heart. The HeartNet device (Paracor Medical, Sunnyvale, CA) is comprised of a metallic alloy (e.g., nitinol) coated in silicone and is placed against the epicardium of the right and left ventricles, to provide pressure to the epicardium. The flexibility of the nitinol enables the device to exert continuous elastic force on the heart throughout the cardiac cycle. Both HeartNet and the CorCap devices were shown in clinical trials to be somewhat successful in reversing ventricular remodeling, but each device failed to generate long term improvements in survival.

Recently, U.S. Patent Publication 2020/0085579 proposed adapting the shapes of devices to the particular heart that is being treated. That disclosure discusses preparing a ventricular support device that is sized and dimensioned both according to the shape of the patient’s heart and particularly to the strain that is estimated at each 3D location in the heart. The device provides increased support to those areas of the heart which have been determined to experience increased strain, relative to those portions of the same diseased heart which undergo less strain.

The equivalent ventricular support device for heart failure with preserved ejection fraction is called a “left ventricle expander.” To date, left ventricle expanders have achieved even less success than ventricular wraps. CorAssist Inc. has developed two devices for expansion of the left ventricle. The ImCardia device is an elastic, selfexpanding device made of a biocompatible metallic alloy that is attached to the exterior surface of the myocardium. ImCardia was discontinued after a single clinical trial, in part due to the aggressive nature of the metallic mesh, and the unacceptable supra- physiological stiffness. The second device, the CORolla, is a spring-like metallic device that is implanted within the left ventricle in order to improve its diastolic function. The spring captures energy during systole, and uses this captured energy to augment expansion of the heart during diastole. CORolla has been studied in only a small trial consisting of ten patients. Since the device is implanted intraventricularly, it raises severe concerns regarding compatibility with the blood, as well as concerns regarding the consequences of retaining the device in the heart should it fail mechanically.

Structures consisting of a shape memory (SM) displaying polymer have a thermodynamically stable “permanent shape” that is able to adopt a metastable “temporary shape” following the application of a suitable stimulus, such as temperature or hydration. By applying an appropriate stimulus to the frozen-in “temporary shape”, the polymer can be actuated in due time, to revert to its “permanent shape.” The relevant external stimulus may be for example, temperature, hydration, ion concentration, or an electric or magnetic field. The two fundamental elements responsible for the shape memory mechanism of polymers are the net-points and the switching segments, with the former storing the latent “permanent shape” and the latter being responsible for actuating the macroscopic transformation. Typically, the net-points are covalent or physical crosslink junctions and the switching segments are based on traversing either the glass transition or melting temperature of the switching segments.

Shape memory polymers have been applied to a variety of biomedical applications. These include: stents, sutures, hemostatic devices, contraceptive devices, devices for embolization, cardiovascular grafts, wound dressings, and 3D scaffolds for tissue engineering. 3D printing, or additive manufacturing, refers to the construction of a three- dimensional object from a computer aided design or digital 3D model. Several 3D printing technologies exist. Common 3D printing technologies for polymers include Fusion Deposition Modelling (FDM), stereolithography (SLA) and digital light processing (DLP).

SUMMARY OF THE INVENTION

The present disclosure describes ventricular support devices that are made of polymers, and, in preferred embodiments, at least partially of a shape memory polymer. The disclosed devices are capable of implementation for treatment for both heart failure with reduced ejection fraction and treatment of heart failure with preserved ejection fraction. In both instances, the devices are placed over the heart in a large size that enables non-injurious implementation. The device is then shrunk over the heart, either relying on intrinsic properties of the shape memory polymer (for HFrEF) or active stretching (for HFpEF), to thereby adapt the shape of the device to the shape of the heart and enable the device to apply inward or outward force to the heart muscle.

The ventricular support devices are manufactured and implemented with various technical innovations that, in combination, enable the devices to achieve a more effective therapeutic treatment than previously known passive ventricular support devices. These technical innovations relate to the materials that are used, the personalization of the devices, the pattern in which the polymer is printed, the dynamic interplay between the device and the pulsating heart, and the manner that the device is deployed onto the heart.

As used in the present disclosure, the term “ink” refers to the materials that are 3D printed, and may include polymers, more specifically shape memory polymers, as well as other components. The shape memory inks allow for the deploying of the device in a placement shape that is sufficiently large not to damage the organ’s surface during deployment, and conversion of the device to a treatment shape that closely fits the external surface of the organ. The transition between the placement shape of the device to the treatment shape may be induced by an external stimulus such as hydration, dehydration, increase or decrease of temperature, application of a magnetic field, irradiation, pH changes, etc., or may be induced by the mere physiological conditions present at the implementation site, namely humidity, pH, temperature, biological cues, etc. In preferred embodiments, the transition is enabled through application of an aqueous medium such as water or saline. Optionally, following application of the stimuli discussed above, the conversion of the device to the treatment shape may occur through stretching, e.g. by the surgeon or a surgical robot. Once the device is implanted or fitted around the organ forming an intimate fit, the forces encoded into the structure mechanically affect the organ’s dynamic behavior, preventing the organ from deviating from its clinically desired parameters, and restoring the organ’s normal function.

The shape memory polymers that may be included in the inks cover a broad range of compositions and mechanical properties, and can be actuated by a number of different environmental stimuli, such as temperature or hydration. In some embodiments, the shape memory polymers are sufficiently hydrophilic to enable a transition in state when introduced into an appropriate aqueous environment. The polymers are sufficiently flexible to allow the ventricle to expand and contract normally when the device is applied around the ventricle. Identification of inks that meet all these criteria required significant experimentation and itself represents a significant advance. The inks may comprise also moieties aimed at bestowing to the printed structure additional functionalities. For example, the inks may also include biodegradable components such as aliphatic esters, thereby potentially rendering the device more flexible over time as the biodegradable inks selectively degrade and are leached out. The biodegradable components may be included within the polymeric chain, to thereby render the device selectively, partially, or totally biodegradable. Alternatively, part of the inks may comprise one or more biodegradable components which are co-printed and covalently bonded to the other components of inks in the structure. The inks may also be PH sensitive or enzymatically degradable, or have various other combinations of features which render them suitable for biomedical application. The inks may also include additional printable mono -functional molecules within the polymeric chain, not capable of forming cross -linkable bonds, biodegradable and/or water-soluble sacrificial components that are not covalently bonded to the polymeric chains, and biologically active molecules for drug delivery.

With respect to the pattern in which the polymer is printed, in preferred embodiments, the inks are printed in a mesh pattern, as opposed to a continuous surface. The pattern may be a polygonal pattern. The mesh pattern allows the device to exhibit sufficient flexibility to maintain pressure when the heart expands and contracts. In some embodiments, the pattern is an auxetic pattern, meaning it has a negative Poisson’s ratio. That signifies that when stretched in one direction, the devices expand, not shrink, also orthogonally to the applied stress. In some embodiments, the pattern is substantially diamond- shaped. This shape is particularly advantageous for devices for treatment of HFpEF, in which the device is stretched longitudinally onto the heart following placement, while shrinking radially. In other embodiments, the device is configured in any pattern that allows the device to shrink radially onto the heart when stretched longitudinally following placement.

With respect to the manner of application, the ventricular support device is printed in an initial shape, applied to the heart in a placement shape, and then, once the device is dressed on the heart, reshaped to a treatment shape. The placement shape and treatment shape may be larger or smaller than the initial shape of the device. For treatment of heart failure with reduced ejection fraction, the device is printed optimally undersized, for example a size that is about 80% of the actual size of the subject’s organ it intends to cover. Prior to placement in the patient’ s body, a stimulus, typically thermal, may be used to render the shape memory polymers in the device flexible. If thermal, the temperature used is above the switching temperature of the shape memory ink. The stimulus is removed while keeping the device in its stretched shape. In the case of a thermal stimulus, this may be achieved by lowering the temperature below the switching temperature of the expanded device, while keeping it stretched, whereby the temporary placement shape is fixed. This allows for expansion of the device from the optimally undersized permanent shape to a temporary placement shape that is larger than the size of the subject’s organ, e.g., 120% relative to the size of the organ. While in its expanded temporary form, the device is deployed around the heart in a non-injurious manner due to its oversized shape, allowing also optimal positioning of the device around the heart. The device may then be fixated to the heart by triggering the reversion of the device from its expanded temporary placement shape to its treatment shape by one or more relevant stimuli, such as temperature, hydration, irradiation, or a biological cue, and optionally through suturing to the heart muscle. The device may further include microneedles that embed into the heart muscle during this shrinking to the treatment shape. The tight fit about the organ and the relative mechanical properties of the heart and the device prevent the organ from deviating from its clinically desired parameters. The inherent tendency of the device to continue to shrink even further, to the undersized permanent shape, imparts a compressive force to the heart. For treatment of heart failure with preserved ejection fraction, the placement shape may be the initial shape of the device, which is also the permanent shape, and which is initially printed larger than the treatment size. Following an appropriate environmental stimulus, in one embodiment hydration of the device, which renders the device flexible, the device is adapted to the shape of the heart. In one embodiment, this is achieved through application of longitudinal tension by a surgeon or surgical robot, which causes the device to shrink radially and attach to the heart. The device is then sutured to the heart in this adapted state, and/or attached to the heart with microneedles, when the heart is at peak systole. The microneedles may include tips that are hydrophilic, and which swell upon entry into the heart muscle, to thereby lock the device on the heart. The inherent tendency of the device to revert to the larger permanent shape imbues within the device an expansive force that acts on the ventricle to increase ventricular volume during diastole, and that also helps reduce and/or optimize cardiac filling pressures.

Although, in the example discussed above, the organ onto which the device is attached is the heart, the device may be attached onto any suitable pulsating organ. These may include, for example, lungs, blood vessels, bladder, etc. The device may provide support to a partial or full, internal or external, surface of the organ.

In some embodiments, the ventricular support devices are personalized devices that are manufactured from an initial digital 3D model of the device generated based on the subject’s organ’s anatomy. For the purpose of achieving a device that can intimately fit about the exterior of the organ, or envelop the organ to the degree desired, data relevant to a person’s organ is collected by, e.g., scanning the organ using one or more imaging techniques known in the art. The scanning provides a detailed mapping of the organ’s exterior, including the shape or contour of the organ’s exterior and the overall surface area of the organ’s exterior, taking into consideration the organ’s anatomy and any other physical attributes that may be relevant for the construction of the device. Additionally, physiological aspects of the organ as well as pathologies affecting it can be taken into consideration when printing the personalized device. One of the techniques that may be utilized for scanning of the organ is echocardiography which harnesses high frequency sound waves (ultrasound) to generate a model of the patient’s organ, e.g., heart. Additional or alternative techniques may include MRI, Computed Tomography (CT), Computed Tomography Angiography (CTA), echocardiography, Positron Emission Tomography (PET) combined with CTA, and others. Typically, the data collected is then processed using a Computer-Aided Design (CAD) software, generating a Stereolithography (STL) file. The digital object is sliced to obtain the geometric data for 3D computer graphics and the personalized device is thereafter printed.

Device manufacturing may comprise any manufacturing technique, including additive manufacturing techniques. Examples of additive manufacturing techniques relevant to shape memory polymers include VAT polymerization such as stereolithography (SLA) and Digital Light Processing (DLP), material jetting, binder jetting, extrusion, powder bed fusion, directed energy deposition, and fused deposition modeling (FDM). The deposition technique may comprise layer-by-layer deposition of an ink formulation, such that each successive layer bonds to a preceding layer of deposited, melted or partially melted materials. For some of the techniques enumerated above the ink may comprise a thermoplastic polymer.

The device may be digitally defined by computer-aided-design (CAD) software used to create .stl files that "slice" the object into ultra-thin layers, providing guidance to a nozzle or print head path so it precisely deposits a material upon a preceding layer. In the case of other additive manufacturing techniques, the printer may follow the relevant digital information fed into the printer, and print the device according to the specific additive technique being used and the corresponding printer.

In some deposition methods, deposition may involve a laser or an electron beam to selectively melt or partially melt in a bed of powdered material. As materials cool or are cured, they fuse together to form the 3D device of the invention.

Devices of the invention may be tailored, structured and manufactured to fit around or inside any pulsating organ as disclosed herein. In some configurations, a device may be formed as a cardiac device configured to fit about a subject’s heart.

In some embodiments, the ventricular support device comprises regions of different composition and properties such as thickness, morphology, pattern, and rigidity. In some embodiments the different regions may be configured to be solid or porous, or continuous or discontinuous, or be configured as a mesh. The device may have different thicknesses, different degrees of hydrophilicity, different degrees of biodegradability, and different degrees of water- solubility in different parts thereof. In some embodiments the different regions may be configured to be thermoplastic in some areas and thermoset in others. These differences in properties enable differential treatment for differently diseased areas of the heart. The properties of the device and its different components, such as its base region and its surface elements, may change over time due to chemical, physical or mechanical, or biological changes, encoded in the device and/or actuated by the environment

Advantageously, the device may be optimally sized for achievement of therapeutic benefit. Previously known versions of ventricular support devices for HFrEF were applied onto the heart, and subsequently provided therapeutic treatment, in a single shape. In order for the device to be fitted around the ventricle, it was necessary to make this shape the same as the shape of the ventricle at end-diastole (i.e., at its largest possible size). However, it is not always possible to size the device exactly to the size of end diastole, while still safely installing the device on the heart. As a result, in practice, the compressive force that the devices could apply was limited. By contrast, the present device is able to achieve its treatment shape after the device is deployed on the heart. As a result, this shape is optimally tailored to the size of the heart to render effective treatment. Additionally, to date, no effective passive restraint device has been successfully implemented for heart failure with preserved ejection fraction. The devices of the present disclosure address this unmet need.

Additional advantages may also be achieved. The device may be used to harness the spatial resolution inherent to 3D printing technologies, to enable engineering of a personalized cardiac restraint device tailored to each patient’s anatomy and pathology. The inks may be selected so as to minimize post- implantation tissue adhesions. Such adhesions, which may firmly glue together tissue planes and organs not meant to be connected, are highly disfavored. In addition, the 3D printing process may be used to implement customizable biodegradability in the devices, if so dictated by the patient’s pathology.

According to a first aspect, a support device for a pulsating organ is comprised of a structure formed at least partially of one or more polymers. The structure is configured to transition, in response to a stimulus, between a placement shape, in which the support device is dimensioned larger than the dimensions of the organ, and a treatment shape, in which the support device provides therapeutic support to the organ; wherein the treatment shape is smaller than the placement shape.

Optionally, the one or more polymers are shape memory polymers.

The structure may comprise a mesh pattern. Optionally, the structure is an auxetic pattem. The stimulus may be a mechanical stimulus. The use of a mechanical stimulus may be particularly advantageous when the at least one polymer does not include shape memory polymers.

Optionally, the one or more polymers are shape memory polymers. The support device is manufactured in an initial shape in which the one or more shape memory polymers are in their permanent shapes, and the placement shape is larger than or equivalent to the initial shape. Advantageously, this size of the placement shape enables the device to be applied over the heart in a safe and non-injurious manner.

Optionally, the stimulus for transitioning between the placement shape and the treatment shape is application of an aqueous medium. Advantageously, application of an aqueous medium is safe for treatment within a live patient.

The device may be 3D printed from at least one ink displaying shape memory behavior. The device may be 3D printed through curing with ultraviolet energy. Examples of such 3D printing methods include SLA and DLP. In such embodiments, the ink comprises one or more central polymeric segments and at least one endcap molecule, wherein each endcap molecule includes at least one double carbon bond suitable for reacting with a corresponding double bond of a different endcap molecule by curing with the ultraviolet energy. The endcap molecule may include acrylate or methacrylate.

The ink may include a shape memory polymer formed of a segment of repeating units of a first monomer. A cumulative molecular weight of the segment may be between 500 and 30,000. These molecular weights may be selected to impart the device with suitable strength and flexibility characteristics.

The first monomer may be hydrophilic. Optionally, the hydrophilic first monomer is polyethylene glycol. Optionally, the shape memory polymer further comprises one or more segments of a second monomer which is hydrophobic. A ratio of the length of the first segment and the second segment may be controlled, thereby enabling imparting of a desired hydrophilicity to the ink.

The shape memory polymer may further include one or more biodegradable components covalently bonded to a polymeric chain of the shape memory polymer. The shape memory polymer may include one or more ethoxysilane moieties covalently bonded to a polymeric chain of the shape memory polymer, wherein said moieties are able to undergo in vivo hydrolysis. The shape memory polymer may further include one or more biodegradable components included within the polymeric chain of the shape memory polymer. In these embodiments, the device changes its rigidity characteristics over time, while applying treatment to the heart muscle.

Optionally, the ink is comprised of one or more additional molecules that are not bonded to the shape memory polymer prior to the printing process. In such embodiments, the first monomer is hydrophilic and the one or more additional molecules may be hydrophobic, or vice versa, thereby enabling imparting of a desired hydrophilicity to the ink. Optionally, the one or more additional molecules comprise one or more of a water- soluble component, a biodegradable sacrificial component, or a bioactive agent.

A plurality of microneedles may be arranged at an organ-facing surface of the support device, the plurality of microneedles configured to latch onto a surface of the organ. Optionally, the microneedles comprise substantially rigid bodies and substantially hydrophilic tips, wherein the hydrophilic tips are configured to swell when placed within a surface of the organ. The microneedles may be used to affix the device to the organ, thereby helping the device perform its therapeutic function on the organ.

According to a second aspect, a method of applying a support device onto a pulsating organ, wherein the support device comprises a structure formed at least partially of at least one polymer, is disclosed. The method includes: deploying the support device on the organ in a placement shape, in which the support device is dimensioned larger than dimensions of the organ; and transitioning the polymer from the placement shape to a treatment shape, in which the support device provides therapeutic support to the organ; wherein the treatment shape is smaller than the placement shape.

Optionally, the polymer is a shape memory polymer.

The method may further include preparing the support device. The preparing step may include: imaging the organ in three dimensions; determining therapeutic goals for different regions of the organ; and printing the support device in a 3D printing process with attributes configured to implement a specific therapeutic benefit onto different regions of the organ consistent with the therapeutic goals. The attributes may include one or more of: materials used in the support device at different locations, thickness of the support device at different locations; pattern of the patterned structure at different locations, and strain and stress constants of the device at different locations. The attributes may include hydrophilicity of the support device at different locations. In such cases, the printing step includes utilizing inks having different degrees of hydrophilicity in different locations of the support device. The attributes may include biodegradability of the support device at different locations. In such cases, the printing step includes utilizing inks having different percentages of biodegradable components in different locations of the support device.

The organ may be a mammalian heart diseased with heart failure with reduced ejection fraction, and the at least one polymer is a shape memory polymer. In such case, the preparing step may include: printing the support device in the 3D printing process in an initial undersized shape corresponding to a permanent shape of the at least one shape memory polymer; applying a stimulus to the support device to thereby raise a temperature of the support device to above a glass or a melting temperature of the switching segment of the at least one shape memory polymer; stretching the support device to the placement shape; and removing the stimulus from the stretched support device while the support device is being kept in its stretched configuration.

Optionally, the applying step further comprises placing the device on an expandable mold; heating the mold to thereby transfer heat to the support device; expanding the mold to thereby stretch the support device; and cooling the mold. Further optionally, the expanding step may include expanding the mold unevenly, to thereby stretch portions of the support device to different degrees.

Optionally, the placing step comprises layering the support device on the mold unevenly, to thereby stretch portions of the support device to different degrees.

Optionally, the transitioning step comprises applying an aqueous medium to the support device to thereby cause the support device to contract from the placement shape to the treatment shape.

The organ may be a mammalian heart diseased with heart failure with preserved ejection fraction. In such embodiments, the preparing step includes printing the support device in the 3D printing process in an initial oversized shape which is the same as the placement shape. Optionally, in such embodiments, the transitioning step comprises applying a mechanical stress. Optionally, the at least one polymer comprises a shape memory polymer, the transitioning step comprises applying an aqueous medium to the support device to thereby render the support device amorphous, and the method further comprises: stretching the support device when in an amorphous state to thereby shape the device around the heart; and fixing the stretched support device to the heart while the device is in said flexible state. The fixing step may include applying one or more of sutures, staples, or adhesive. The device may include a plurality of microneedles arranged on an organ-facing surface of the patterned structure, the plurality of microneedles configured to latch onto a surface of the exterior of the heart. In such embodiments, the fixing step may include embedding the microneedles within an exterior surface of the heart. Optionally, the microneedles comprise substantially rigid bodies and substantially hydrophilic tips, and the fixing step further includes causing the microneedles to swell after the microneedles are placed within a surface of the heart.

The fixing step may be performed when a left ventricle of the heart is at peak systole.

According to a third aspect, a method of preparing a support device for a pulsating organ is disclosed. The method includes: imaging the organ in three dimensions; determining therapeutic goals for different regions of the organ; and printing the support device in a 3D printing process with attributes configured to implement a specific therapeutic benefit onto different regions of the organ consistent with the therapeutic goals.

According to a fourth aspect, a method of embedding a medical device into an organ is disclosed. The medical device includes microneedles having rigid bodies and hydrophilic tips. The method includes: deploying the medical device onto the organ in a dry state; embedding the tips into the surface of the organ; and allowing the tips to swell within the organ, thereby locking the tips within the organ.

Brief Description of the Drawings

FIG. 1A illustrates a bottom view and a perspective view of a ventricular support device for treatment of HFrEF, in an initial shape, according to embodiments of the present disclosure;

FIG. IB schematically illustrates stretching of the ventricular support device of FIG. 1A following application of heat or any suitable stimulus, according to embodiments of the present disclosure;

FIGS. 1C and ID illustrate comparisons of the ventricular support device in the initial and stretched shapes;

FIG. IE illustrates dressing of the stretched ventricular support device onto the heart;

FIG. IF illustrates the stretched ventricular support device on the heart; FIG. 1G illustrates application of a stimulus such as an aqueous liquid to the ventricular support device while the device is dressed on the heart, to thereby enable transitioning of the device to the treatment shape;

FIGS. 2A-2G illustrate steps of application of a second embodiment of a ventricular support device, for treatment of HFpEF, according to embodiments of the present disclosure;

FIG. 2H illustrates a mesh pattern for the device of FIGS. 2A-2G that enables a scissoring effect during stretching, according to embodiments of the present disclosure;

FIG. 21 illustrates a mesh pattern for the device of FIGS. 2A-2G that enables an enhanced scissoring effect during stretching, according to embodiments of the present disclosure;

FIGS. 3A-3E illustrate components of the 3D printed inks for the ventricular support devices, according to embodiments of the present disclosure;

FIG. 4A illustrates steps in a method of preparing a ventricular support device;

FIG. 4B illustrates an echocardiogram of a rat ventricle;

FIG. 4C and FIG. 4D illustrate CAD slicing of the recorded ventricle;

FIGS. 5A and 5B illustrate structures of the ventricular support devices as mesh patterns of different sizes;

FIGS. 5C and 5D illustrate auxetic patterns;

FIG. 5E illustrates a process of stretching an auxetic pattern;

FIG. 6A illustrates steps in a method of applying a ventricular support device in a case of heart failure with reduced ejection fraction; and

FIG. 6B illustrates steps in a method of applying a ventricular support device in a case of heart failure with preserved ejection fraction.

Detailed Description of Embodiments of the Disclosure

The present disclosure relates to the field of medical devices, and more specifically, but not exclusively, to personalized shape memory devices that are configured to transition between different shapes, and that may be applied onto a pulsating organ, and particularly the left ventricle of a heart, in order to mechanically affect or modulate the heart’s dynamic behavior.

Before explaining at least one embodiment of the invention in detail, it is to be understood that the invention is not necessarily limited in its application to the details of construction and the arrangement of the components and/or methods set forth in the following description and/or illustrated in the drawings and/or the Examples. The invention is capable of other embodiments or of being practiced or carried out in various ways.

As used in the present disclosure, the term “shape memory polymer” refers to a polymer that, after a suitable stimulus such as heat or hydration is applied to it, and after being subjected to a deformation and being fixated in the deformed state, resumes its original shape upon being subjected to an external stimulus. The stimulus may include heating above the glass-transition or melting temperature of the switching segment of the ink, application of an aqueous medium such as water or saline, pH, light, ion concentration, or application of an electric or magnetic field. Shape memory polymers contain both “switching segments,” which are the parts that are deformed, fixated, and restored to their initial shape following application of the stimulus, and “net points” which are parts which retain their shapes and sizes even when the stimulus is applied. Unlike in standard uses of the term “shape memory polymer,” the use of this term does not necessarily imply that the polymer in fact undergoes all of the transitions listed above. Indeed, as will be discussed herein, the method of treatment may specifically include a shape memory polymer that remains in its deformed shape during treatment, pulsating with the organ that is being treated, without ever resuming its permanent shape. In this regard, the “shape memory polymer” does not reach a single “permanent” shape, and may be understood to fluctuate continuously between two shapes throughout treatment.

As used in the present disclosure, the term “3D printing” refers to the construction of a 3-dimensional object from a computer aided design model or digital 3D model. Several specific technologies have been developed for 3d printing of shape memory polymers. One common 3D printing technology is stereolithography (SLA). In SLA, light energy is used to cause chemical monomers and oligomers to cross-link together to thereby form polymers. Digital light processing (DLP) is similar to stereolithography, except that DLP uses a digital light projector screen to flash a mask of each resin layer at once, as opposed to portions of a layer in SLA. In both SLA and DLP, low molecular weight liquid precursors containing reactive C=C bonds are rapidly photo -polymerized by UV radiation into solid layers. DLP is a preferred method of printing and polymerization of the devices of the present disclosure.

Theoretically, other methods of 3D printing are also possible. One common method is fused deposition modeling (FDM). Fused deposition modeling refers to a process of building parts layer by layer by selectively depositing melted material in a predetermined path. Other pertinent 3D printing technologies include selective laser sintering (SLS), selective laser melting (SLM), material jetting, binder jetting, extrusion, powder bed fusion, and directed energy deposition. The deposition technique may comprise layer-by-layer deposition of an ink formulation, such that each successive layer bonds to a preceding layer of deposited, melted or partially melted materials.

As used in the present disclosure, the term “ink” refers to the materials that are 3D printed, and may include shape memory polymers as well as other components. The term “shape memory ink” refers to an ink that includes shape memory polymers, so that the item printed from it exhibits shape memory characteristics.

As used in the present disclosure, the term “pulsating organ” refers to any organ which undergoes cyclic expansion and contraction as part of its routine function. The most common such organ is a mammalian heart, and this is the example used throughout the present disclosure. However, the principles described herein could be applied to treatments for other pulsating organs, such as the lungs, trachea, blood vessels or bladder.

As used in the present disclosure, the term “permanent shape” refers to the shape in which a shape memory polymer is printed and to which it is thermodynamically biased to return. The term “temporary shape” refers to a shape to which the shape memory polymer may be temporarily deformed following application of a stimulus and fixing (such as by lowering the temperature to below the switching temperature). The terms “permanent shape” and “temporary shape” are used in their conventional meanings as applied to shape memory polymers. The term “initial shape” refers to the shape in which the ventricular support device is initially printed; the term “placement shape” refers to the shape of the device at the time of placement over the heart, and the term “treatment shape” refers to the shape to which the device is transitioned while on the heart. As will be discussed herein, the device changes shape while on the heart as a matter of course through pulsation of the heart, and may also gradually change shape over time when in place over the heart. The terms “initial shape,” “placement shape,” and “treatment shape” are used specifically in reference to different shapes of the ventricular support devices.

The devices described below may operate principally on the left ventricle, but may also operate on the right ventricle. The devices may improve inter-ventricular dependency by optimizing right ventricular pressure and volume. The devices may also support the cardiac annulus to prevent valvular regurgitation. The devices described below are described as fitting an external region of the heart. In particular, the devices are illustrated as wrapping around (at least) the apex of the left ventricle. In principle, however, the devices may be implemented onto any portion of the heart. The devices may also be implemented intraventricularly, with suitable modifications.

The devices of the present disclosure may be specifically useful for:

1. mechanically affecting a pulsatile behavior of a pulsating organ to restore one or more clinically desired parameters associated with the behavior of the organ;

2. imposing dimensional restrictions to the heart, thereby modifying left ventricular chamber volume and function, in both HFpEF and HFrEF diseased patients, and HF with supra normal ejection fraction;

3. improving cardiac function or preventing heart failure with reduced ejection fraction (HFrEF);

4. improving cardiac function or preventing heart failure with preserved ejection fraction (HFpEF);

5. preventing ischemia due to a coronary artery disease;

6. preventing pathologic remodeling and dilatation of the heart in patients with heart failure;

7. treating and preventing pathologic remodeling including dilatation and/or geometric and/or functional changes of the myocardium in patients with heart failure;

8. improving valvular heart dysfunction;

9. relief of elevated cardiac systolic and/or diastolic pressure; and

10. minimizing post-surgical adhesion formation.

I. Devices for Treatment of Heart Failure With Reduced E jection Fraction

FIGS. 1A-1G illustrate the application of a ventricular support device onto a heart suffering from heart failure with reduced ejection fraction. It is understood herein, that reduced ejection fraction includes mildly reduced ejection fraction, meaning that it refers to a heart with an ejection fraction smaller than 50% of the normal ejection fraction. The ventricular support device is configured to apply compressive force on the heart, and more specifically on the left ventricle, to assist in the pumping of the blood. FIG. 6A lists steps in a method 600 for application of the ventricular support device of FIGS. 1A-1G. Referring now to FIG. 1A, device 100a is a structure printed from a polymer, such as a shape memory ink. The left side of FIG. 1A depicts a bottom view of the device 100a, and the right side of FIG. 1A depicts a perspective view of device 100a. The patterned structure is sized and shaped to fit an interior or exterior of a pulsating organ. In the illustrated embodiment, device 100a is shaped to fit the exterior of a mammalian heart. Although, in the illustrated embodiment, the device is fitted substantially around the entire heart, the device may be suitably fitted over any portion of the heart. In the illustrated embodiment, the pattern is a mesh structure with a honeycomb pattern. The use of a mesh structure is advantageous in that it enables the device to have both strength and flexibility. However, other embodiments are possible, including a continuous structure with no internal gaps.

Like all shape memory polymers, the shape memory polymers of device 100a (when present) are configurable between a “permanent” shape and a “temporary” shape. The permanent shape is the position in which the device 100a is manufactured, and to which it would theoretically return after being stretched, following application of a stimulus such as heat or a liquid, as discussed above. The temporary shape is an expanded position which is achieved by stretching the device 100a following application of a stimulus such as heat. The composition and specific properties of the shape memory polymers will be discussed in detail below, in connection with FIGS. 3A-3E.

In the context of the clinical treatments described herein, device 100a may be configurable between three different sizes. The transition between these sizes is illustrated in FIGS. 1A-1G, in an illustration of application of device 100a for treatment of heart failure with reduced ejection fraction. Reference numeral 100a is used to refer to the device in its initial shape, reference numeral 100b refers to the device in a placement shape, and reference numeral 100c refers to the device in a treatment shape.

Referring to FIG. 1A and step 601 in FIG. 6A, device 100a is initially printed in an undersized “initial shape.” The “initial shape” is equivalent to the “permanent shape” of the shape memory polymer, in the sense that the shape memory polymer is biased to return to that shape.

Referring to FIG. IB and step 602 in FIG. 6A, a stimulus is applied to device 100a. In one example, sufficient heat is applied to raise the temperature of the polymers of the device 100a to above the relevant glass transition temperature (T_g) or melting temperature (T_m) of the switching segment of the polymer. Following heating, the device 100a is stretched through application of any suitable means, and cooled in order to preserve the device in the enlarged shape. Stimuli other than heat may also be used to render the device stretchable; however, as this process is performed prior to implantation in the body, heat is an effective and safe mechanism.

In one example, the heating and stretching are performed by placing the device 100a on an expandable mold. The mold may be conductive of heat and may be connected to a power source. The mold may consist of multiple wings arranged substantially radially around a central core. The mold may be heated, then expanded, and then cooled while the mold is kept in its expanded form, to thereby heat, stretch, and fix the support device. The wings may be individually expandable. As a result, the device 100a may be stretched to different degrees in different directions. In addition or in the alternative, the device may be layered over the mold unevenly, so the mold will expand different regions of the device to different extents. As a result, the restoring restricting forces encoded in different regions of the device will be different, as dictated by the spatially selective required therapeutic function of the device. For example, the device may be placed with a portion of the device folded over itself, or with some slack in the device, or with a “bubble” preformed in the device during the printing process. The portion with folding or slack is stretched less than the portion which is tightly fit over the mold. The variation in stretching enables the final device to be configured with different dimensions, and thus to confer a different therapeutic benefit onto different regions of the heart, according to the individual needs of the patient.

After stretching, the device 100b is cooled in order to preserve the device in a “placement shape.” The term “placement shape” refers to the size of the device 100b when it is deployed on the pulsating organ. The “placement shape” is equivalent to what is typically called the “temporary shape” of the shape memory polymer. FIGS. 1C and ID illustrate the differences in size between the initial shape and the placement shape.

In FIG. IE and at step 603 of FIG. 6A, the device 100b, still in its placement shape, is deployed over the heart 110. The device is typically deployed while dry. FIG. IF depicts the device 100b loosely fitted over heart 110.

In FIG. 1G and at step 604 of FIG. 6A, an aqueous medium 120 (such as water, a saline solution, or a buffer solution) is applied to the device 100b. The use of an aqueous medium as a stimulus is available when the device is printed using inks of suitable hydrophilicity; otherwise, a different stimulus (such as heat or magnetism) is employed. One advantage of the use of an aqueous medium as a stimulus is that it is safe for application onto a mammalian heart. The water molecules act as plasticizers of the suitably hydrophilic device. Additionally, in the case of semi-crystalline switching segments, water molecules destroy the order of hydrophilic crystallites, rendering them amorphous, and increase chain mobility at physiological temperature. As a result, the device is now able to return to a smaller shape. The same effect may also be generated simply through the moistness of the cardiac environment, but the aqueous medium may accelerate the process.

The aqueous medium thus serves as a stimulus for transitioning the device from the placement shape of device 100b at the top of FIG. 1G, to a treatment shape of device 100c at the bottom of FIG. 1G. The treatment shape is the shape assumed by the device when the device applies therapeutic treatment to the heart.

In the embodiment of FIGS. 1A-1G, the initial shape of the device is smaller than that of the heart on which it is to be deployed, and the placement shape of the device is larger than that of the heart on which it is to be deployed. For example, the initial shape may be approximately 80-90% of the shape of the heart, and the placement shape may be 110% to 120% or even 130% of the shape of the heart. When the aqueous medium is applied to the stretched device 110b, the device is biased to shrink to the initial shape, which is smaller than the heart. However, due to the presence of the heart itself, device 110c achieves a treatment shape which is only the same size as the heart. The bias of the device 110c to return to the initial shape causes device 110c to exert a constant inward pressure on the ventricle, not only during diastole (when the ventricle is at its biggest) but also during systole (when the ventricle is smallest). Preferably, the inward pressure is at least 3 mm Hg (400 Pa), during both systole and diastole.

Optionally, the device 110 includes microneedles. The microneedles are embedded into the wall of the heart muscle when the device shrinks over the heart. A full description of the structure and function of the microneedles is discussed in connection with the device of FIGS. 2A-2F.

The device of FIGS. 1A-1G is designed to achieve several clinical benefits. This inward pressure assists the ventricle in its pumping function and also serves to prevent, and even reverse, cardiac remodeling resulting from the reduced ejection fraction. Other benefits include: application of focused mechanical pressure to thereby improve secondary valvular insufficiency; restoring spherical characteristics of the heart alongside its spring and twist capabilities; and applying support to aneurismal cardiac parts.

II. Devices for Treatment of Heart Failure With Preserved Ejection Fraction

FIGS. 2A-2I illustrate a second embodiment of a device 200 for treatment of heart failure with preserved ejection fraction, and FIG. 6B depicts steps in a method 610 of treatment of HFpEF. In order to provide treatment for this condition, it is necessary to provide an outward (pull) force on the myocardium, to thereby enable the ventricular walls to expand during diastole, as physiologically required. In particular, it is desirable to provide a minimum of 6 mm Hg (800 Pa) outward pressure on the left ventricle. Device 200 applies this treatment through the mechanisms described herein. In FIGS. 2A-2G, the reference numeral 200a refers to the device in the initial shape, and the reference numeral 200b refers to the device in the treatment shape.

In FIG. 2A and at step 611, device 200a is prepared in the initial shape. The left side of FIG. 2A represents a top perspective view, and the right side of FIG. 2A represents a side view of device 200a. When the device includes shape memory polymers, the initial shape is the “permanent shape” of the shape memory polymers and is also the placement shape. In the illustrated embodiment, device 200a is printed in a mesh structure 202. In one embodiment, the mesh structure is formed of a pattern of diamond- shaped holes 232, as illustrated in FIG. 2H. Optionally, as shown in FIG. 21, the mesh structure consists of a more rigid polymer in the body 242 of the mesh, and a more flexible polymer at junction points or vertices 244 of the mesh. This configuration enables the mesh to flex more easily during stretching. The use of diamond- shaped holes particularly enables stretching of the device in a single direction, as will be explained further herein. Other configurations are possible. Although the diamond- shaped devices have a positive Poisson’s ratio (one dimension expands as the other contracts), it is also possible to use auxetic structures, in which one dimension contracts as the other also contracts, and expands when the other expands.

In addition to including the mesh structure, device 200a may also include multiple microneedles or “hooks” 214 arranged on the surface of the patterned structure facing the pulsatile organ, as shown schematically in FIG. 2E. The microneedles 214 may be configured in a regular pattern, or in a specific pattern that matches the anatomy, physiology, and pathology of the individual patient. The number and distribution of the microneedles, and their size, geometry and mechanical properties may likewise be tailored to the needs of the individual patient. In one embodiment, part of the microneedles may be biodegradable. The microneedles 214 are configured to latch onto a surface of the organ (here, the exterior of the myocardium). Optionally, the microneedles 214 are at least partially formed of a hydrophilic material. Specifically, the microneedles 214 may be formed of a first section 215 closer to the patterned structure (a “body” section) that is substantially hydrophobic and rigid, and a second section 216 that is further from the patterned structure (a “tip” section), and closer to the wall of the organ, which is made of a hydrophilic material. Additional compositions and structures of the microneedles are possible.

In FIG. 2B and at step 612 of FIG. 6B, the device 200a is deployed over the heart 210. Again, the left side of FIG. 2B illustrates a top perspective view, and the right side illustrates a side view. The device 200a is printed larger than the heart 210 onto which it is placed, for example, 110% or 120% or 130% of the size of the heart, thereby enabling easy and safe placement and optimal positioning of the device over the heart. Due to the oversized nature of the device, the device may be deployed without embedding the microneedles into the heart muscle.

At step 613 of FIG. 6B, a suitable stimulus, preferably an aqueous medium (e.g., water, saline solution, or a buffer solution) is applied to the device 200 when the device is deployed over heart 210. As described in connection with the device of FIGS. 1A-1G, the aqueous medium plasticizes the suitably hydrophilic device and, in the case of crystalline inks, it also renders the switching segments of the shape memory polymers amorphous, thereby permitting stretching of the device 200. In embodiments in which the polymer is not a shape memory polymer, the stimulus may be purely a mechanical stimulus, and this step is combined with the subsequent step.

In FIG. 2C and step 613 of FIG. 6B, the device 200b is gently stretched longitudinally when over the heart 210 (e.g., by the surgeon, with suitable surgical tools, or by a surgical robot) to bring the device to the treatment shape. For patterns such as the diamond shape of FIGS. 2H and 21, the stretching in one direction may contract in a perpendicular direction. This pattern of stretching may be imagined as a “scissoring” deformation. This stretching causes the device 200b to reduce its size from up to around 130% of the heart to around 100% of the heart, so that it may be snugly secured over the heart muscle. In one embodiment, aiming at further enhancing the “scissoring” capability of the device, the scissoring elements are not connected directly but via an additional ink (see dots 244 in FIG. 21) characterized by being especially flexible and displaying superior fatigue properties.

In FIGS. 2D and 2E and step 614 of FIG. 6B, once device 200b is stretched in place, the device 200b is fixed to the heart. The fixing may be performed using any suitable device, such as staples, adhesive, or sutures 212, as shown in FIG. 2D. In one example, the fixing may be performed using sutures that are routinely used during heart surgery, such as 3-0 to 5-0 Prolene sutures. Suturing the myocardium is standard practice for many types of treatments, and is utilized in various medical devices that are attached directly to the myocardium. The fixing may also be performed using the microneedles 214, as discussed above, and as shown in FIG. 2E. During or following the stretching process, the dry and stiff microneedles 214 are embedded into the myocardium. The hydrophilic tips 216 of the microneedles 214 (when present), now in the aqueous environment of the heart muscle, swell and expand, to thereby lock the needles in place. This locking of the needles in place confers a key therapeutic benefit, as it assures that, when the device 200 expands, the heart muscle is expanded with it. The use of microneedles also has distinct advantages, including: better stress distribution throughout the connection between the device and the heart; [b] achieving a more intimate contact between the heart and the device; [c] avoiding or minimizing stress concentration due to the number of needles and the superior mechanical isocompliance between the microneedles and the heart muscle, as opposed to the exceedingly stiff sutures or staples. The microneedles may also be used for sustained or slow release of bioactive molecules into the heart muscle.

Parenthetically, the method of embedding a device onto an organ with microneedles described herein is not limited to a ventricular support device, and may be performed for any type of device onto which the microneedles are attached. Specifically, a method of embedding a medical device into an organ, wherein the medical device includes microneedles having rigid bodies and hydrophilic tips, may include: deploying the medical device onto the organ in a dry state; embedding the tips into the surface of the organ; and allowing the tips to swell within the organ, thereby locking the tips within the organ. The device achieves therapeutic benefit so long as it is fixed to the heart during the systolic phase. The suturing is preferably performed while the left ventricle is at peak systole, i.e., is in the smallest size throughout the heart cycle.

When the device is printed from shape memory polymers, the fixing thus secures the device to the heart while the device is in its “temporary” state, as relates to the shape memory polymers. As a result, device 200, which remains in an aqueous environment so long as it is attached to the heart, is biased to return to the thermodynamically preferred initial size, which is the “permanent shape.” Device 200 cannot return all the way to its initial size, however, because it is prevented from doing so by the sutures, microneedles, or other fixing components. As a result, the shape memory polymer causes the device to exert a constant outward pull. Because the device is fixed to the heart muscle, when the ventricle contracts in systole, the contraction is against the outward pull of the device 200, although not so much to detrimentally affect pumping of the blood. Most significantly from a therapeutic perspective, when the ventricle expands in diastole, it is assisted in its expansion by the pull of the device 200.

Thus, device 200 utilizes the principles of shape memory polymers in an unconventional manner. Typically, a shape memory polymer in its permanent shape is designed to morph to a temporary state, typically by expansion, and then revert back to its permanent state. By contrast, device 200 starts in the permanent state, is converted to a temporary state, and remains in the temporary state, while constantly striving to return to the permanent state. The shape memory quality thus interplays with the pulsating heart muscle, with the device constantly driving to reverting to its oversized “permanent shape”, and the heart restoring the device’s “temporary shape”, cyclically, once and again, as the heart pulsates. This repetitive transition is illustrated in FIGS. 2F and 2G, with arrows 222 of FIG. 2F illustrating the outward expansion of the device 200b as the heart expands from peak systole to diastole, and arrows 224 of FIG. 2G illustrating the contraction of the device as the heart contracts from peak diastole to systole. In classic shape memory displaying structures, the transition between the “permanent” and “temporary” shapes take place only once in each direction. By contrast, device 200 undergoes this transition around 100,000 times a day, due to its interaction with the beating heart. The devices of FIGS. 2A-2I may exert outward pressure of approximately 6 mm Hg (800 Pa) on the affected cardiac zones. This pressure represents the minimal outward pressure required to improve diastolic function.

The foregoing description of devices for treatment of HFpEF was explicated with respect to devices formed at least partially of shape memory polymers. The use of shape memory polymers is indeed advantageous, in that the shape memory quality enables a strong expansive force, and also provides a controllable expansion force, as will be discussed further herein. That said, similar benefits are achievable with any device that is biased to return to a larger shape following contraction over the heart, such as through a mechanical contraction, so long as it is securely attached to the heart and the restorative force is sufficient to provide the required therapeutic treatment.

The devices described above enjoy multiple advantages over known ventricular support devices.

First, with respect to both types of devices, the engineering of the device as polymeric, as opposed to metallic, has substantial advantages concerning clinical performance. These include not being overly stiff and preventing injury to the myocardial muscle. In addition, because the devices are 3D-printed, the devices may be custom- manufactured to the clinical needs of each patient. Additional advantages of polymers over metals include (a) the ability to tailor/modify their surface to enhance their biocompatibility, blood compatibility, and prevention post-implantation tissue adhesions that could curtail the performance of the device (b) imparting functionalities for release drugs and the like; and (c) selective biodegradability of the polymers. The specific manner in which the device may be customized to the needs of given patients is discussed further herein.

With respect specifically to the device of FIGS. 2A-2I, the device actively applies the outward forces required to improve the diastolic behavior of the heart, as it cyclically and reversibly reverts to its oversized placement shape. The device thus meets the long- felt need for a device capable of treatment of heart failure with preserved ejection fraction.

In addition, it should be understood based on the above description that the shape memory polymers used in the described devices serve a different, and more sophisticated, function than shape memory metals (such as nitinol) of prior art ventricular support devices. In known devices, the nitinol is not chosen because of its shape memory behavior but for qualities such as its ability to remain elastic over a range of strain of up to 4%, its corrosion resistance, fatigue resistance, and creep resistance. By contrast, in the implementation of the disclosed devices, the shape memory capabilities serve different functions. These include: [a] easy deployment of the device when in its oversized temporary shape, aiming at not injuring the heart when being deployed and optimizing its positioning around the heart, and [b] secure attachment of the device to the heart. For treatment of HFrEF, the shape memory functionality plays a role only during deployment, then it is the elasticity and fine-tuned flexibility /rigidity of the device that play a more dominant therapeutic role. In addition, in the case of HFpEF, the shape memory material is used to cause the device to exert an outward force at unison with the pulsating heart muscle.

It is also apparent from the preceding discussion that the microneedles play a different, more sophisticated role than merely attaching the device to the heart muscle. In prior art devices, the function of needle-like attachments is merely to prevent the device from migrating. By contrast, the microneedles of the presently disclosed devices perform a dynamic function for treatment of HFpEF, actively expanding the left ventricle during the diastolic phase of the cardiac cycle. This active expansion is achieved through the outward “pull” exerted on the muscle of the left ventricle by the device, which is delivered through the rigid microneedles. The rigid microneedles, in turn, are held in place in the myocardium by the swollen hydrophilic tips, when present. The advanced bi-component type of microneedles thus ensures effective stress transfer between the heart and the device both during the diastolic and systolic phases of the cardiac cycle.

III. Shape Memory Polymers and Other Components of the Ink

It is apparent from the previous discussion that the inks from which the devices are printed must exhibit certain characteristics. Many inks that are currently used for 3D printing are stiff materials, which are not appropriate for placement over soft tissues such as the heart. The inks that are used in the devices of the present disclosure must be suitably elastic and flexible to meet the needs of the cardiac treatment. At the same time, the inks must exhibit suitable strength and durability to ensure that they are able to provide the therapeutic treatment for as long as needed.

As discussed above, structures consisting of a shape memory polymer have a thermodynamically stable “permanent shape” that is able to adopt a metastable “temporary shape” following the application of a stimulus, such as temperature or hydration. By applying an appropriate stimulus to the “temporary shape”, the polymer can be actuated to revert to its “permanent shape.” The two fundamental elements responsible for the shape memory mechanism of the polymers are the net-points and the switching segments, with the former storing the latent “permanent shape” and the latter being responsible for actuating the macroscopic transformation. Typically, the net-points are covalent or physical crosslink junctions and the switching segments are based on traversing either the glass transition or melting temperature of the switching segment of the polymer. The transition of the switching segments may be actuated through various mechanisms, such as heat, water, UV energy, magnetism, acidity, biological cues, etc.

As stated, shape memory polymers contain both “switching segments,” which are the parts that are deformed, fixated, and restored to their initial shape following application of the stimulus, and “net points” which are parts which retain their shapes and sizes even when the stimulus is applied. Any moieties able to perform as “switching segments,” and any moieties able to act as “net points” are covered by this disclosure, regardless of their composition. In some embodiments, the “net points” are covalent crosslink junctions as in the case of C=C capped inks while in other embodiments the “net points” consist of physical crosslink domains, such as, for example, crystalline components having a melting temperature above the temperature of performance of the device, typically above 37 °C, for example at 50 or 60 °C or above. In other embodiments, the “net points” may consist of amorphous components having a glass transition temperature above the temperature of performance of the device, typically above 37 °C, for example around 50 or 60 °C or above. In some embodiments, the “switching segments” may be crystalline or amorphous, transitioning between the “permanent” and “temporary” shapes by crossing their melting temperature or glass transition temperature, respectively, provided said melting or glass transition temperatures are above the temperature of performance of the device, typically above 37 °C, for example at 50 or 60 °C or above, where said polymer may be a block polymer or any moiety of any composition displaying this feature. In all cases, where the “net points” are not covalent crosslink junctions, the melting or glass transition temperature of the ’’net points” is higher than the melting or glass transition temperature of the ’’switching segments.” The above may be exemplified by a shape memory displaying ink where the “net points” consist of a P(L)LA block with a melting point above 100 °C, and the “switching segments” consist of a PCL segment with a melting point around 50 °C. The example given relates to biodegradable inks but it will be clear to a person with skill in the art that numerous examples exist for non-biodegradable polymers. The above may be further exemplified by a shape memory displaying ink where the “net points” consist of an amorphous component having a high glass transition temperature, such as polymethyl methacrylate (PMMA) having a glass transition temperature around 110 °C, and the “switching segments” consist of polyethyl methacrylate (PEMA) with a glass transition temperature around 60 °C. While the first example given relates to a biodegradable ink, the second one describes a non-biodegradable ink. Additionally, the first example describes a crystalline ink while the second one relates to an amorphous ink. Inks combining the two are also covered by this disclosure, where, for example, the “switching segment” is amorphous while the “net points” are crystalline. Additionally, inks for which different phases of the same component differ in their degree of crystallinity, with one phase having a high melting point and a second phase having a melting point at a lower temperature, still above the temperature of performance of the material, will display shape memory behavior. In this case, the phase having a high melting point generates the “net points” and the phase having a lower melting point creates the “switching segments,” as in the case of some PCL-containing polymers.

In a preferred embodiment, the inks that meet the requirements of the disclosed devices may be synthesized by end-capping shape memory polymers with molecules having one or more C=C double bonds. The C=C double bonds may be polymerized (in the case of one C=C bond) or cross-linked (when more than one C=C double bond is present in the ink molecule) during the process of printing the ventricular support device. In addition, the inks may include other features, such as biodegradable components, whether within the polymeric chain or as separate molecules, so as to render the device with biodegradability capabilities or enable adjustment of the rigidity of the device after it is put into place on the heart. The inks may also include other components, as will be discussed further herein. The inks cover a broad range of compositions, morphological features, hydrophilicity levels and mechanical properties.

The inks may comprise one or more monomers, oligomers, or prepolymers that are reacted under printing conditions to provide the polymers described herein; or the ink may include the polymers described herein, already preformed prior to the printing process. FIG. 3A illustrates a basic architecture of the inks that is suitable for 3D printing with digital light processing (DLP) or stereolithography (SLA), according to embodiments of the present disclosure. The ink includes multiple polymeric units 300. Each polymeric unit 300 includes a block 302 of a polymer, which serves as a switching segment. The polymeric unit 300 further includes endcaps 304. The endcaps are formed from molecules having at least one C=C bond following formation on the polymeric unit 300. The endcaps render the units polymerizable or cross-linkable under UV radiation during the 3D printing process. Specifically, in the case of cross -linking, each double bonded carbon at an end of a first endcap molecule is crosslinked to a double bonded carbon of a different endcap molecule by curing with the ultraviolet energy, in the presence of a suitable photo-initiator. The cross-linked endcaps serve as netpoints.

In alternative embodiments, instead of all of the endcaps having two double bonds, some endcaps may have one double bond. Partial use of endcaps one double bond may be useful for separating the crosslink junctions. In addition, theoretically, it is possible to have inks with all the endcaps having only one double bond. In such instances, the rest of the polymer is composed of different segments that provide both the net points and the switching segments.

The inks may alternatively be comprised of thermoplastic polymers, and may be printed with fused deposition modeling (FDM).

In addition to the polymeric units 300, the inks may include other materials, which are not cross-linked with, or otherwise bonded to, the polymeric units 300 during the 3D printing process. For example, the printing system may include, in addition to the ink or inks, also water-soluble sacrificial components 312. The water-soluble sacrificial components 312 may be used to impart the device with a certain degree of rigidity during the processes of printing and deployment, while allowing the device to soften over a course of treatment as the sacrificial components dissolve. The printing system may include, in addition to the ink or inks, also water-soluble or initially non-water soluble biodegradable sacrificial components. In some embodiments, said water-soluble or initially non-water soluble biodegradable sacrificial components may be able to become bonded covalently or otherwise to the ink or inks, as the printing process proceeds.

In addition, the ink may include bioactive agents 314. The bioactive agents may be, for example, antibiotics or anti-inflammatory medication, among numerous others. These agents may be delivered to the heart muscle or to the environment surrounding it to help prevent adverse reactions to the device from the heart muscle or surrounding tissues, or for any other clinical need. The bioactive agents 314 may be included as separate components, as illustrated here, or, alternatively, may be bonded to the inks via a biodegradable spacer. Additionally the drug may be covalently or otherwise bonded to the sacrificial water soluble component or to the biodegradable component, and/or blended with them.

FIG. 3B illustrates various segments of the polymeric block 302. The polymeric segments include one or more hydrophilic segments 306 and one or more hydrophobic segments 308.

Generally, shape memory polymers may be hydrophilic, hydrophobic, or of a tunable hydrophilicity. In FIG. 3C, polyethylene glycol (also known as polyethylene oxide, PEG, or PEG) 306a is provided as an example of a hydrophilic segment, and polypropylene glycol (also known as polypropylene oxide, PPG, or PPO) 308a is provided as an example of a hydrophobic segment. The terms “oxide” and “glycol”, as for example in polypropylene oxide (PPO) and polypropylene glycol (PPG), respectively, will be used throughout this document interchangeably.

The use of hydrophilic “inks” has distinct advantages especially when aiming at fully complying with the multifaceted set of requirements posed by medical devices combining shape memory behavior and high flexibility, such as the device being disclosed. As discussed, the hydrophilic nature of the shape memory “inks” bestows the structures printed with the ability to revert from their temporary expanded shape to their treatment shape, when actuated by water or an aqueous medium. Additionally, engineering devices consisting of “inks” comprising crystalline hydrophilic components allows: [i] To have a suitably somewhat rigid dry device during deployment that makes handling easier, and that then, once hydrated after being optimally positioned around the heart, generates a highly flexible, often isocompliant device; [ii] To safely actuate the device’s shape memory response using an aqueous medium, not heat, that converts the crystalline domains of the “ink”, that act as switching segments, into an amorphous matrix, increasing drastically the flexibility of the device; and [iii] By converting the crystalline, often opaque or translucent structure into an amorphous one, the device becomes fully transparent, yet another feature of clinical significance. That said, hydrophobic inks may also be used, with the switching being triggered through bio-safe mechanisms, such as slightly supra-physiological temperature or magnetism. In preferred embodiments, polyethylene glycol (PEG)-based switching segments are end-capped with C=C bonds, in some instances forming urethane-containing crosslinkable “inks.” This may be achieved by reacting the OH terminal groups of the PEG switching segments with isocyanate ethyl methacrylate (IEMA) or isocyanate ethyl acrylate (IEA), which is known to photopolymerize faster. The PEG segment may be end capped also with other functional groups, such as amine and carboxylic acid moieties, which may result, following the end-capping reaction, in the generation of amide or urea groups, for example. The degree of cross-linking may be controlled by fine tuning the composition and molecular weight of the PEG-containing chains and their degree of methacrylation (or acrylation), namely the number of double bonds per molecule. The use of PEG as a basis for the switching segments has multiple advantages. First, PEG is hydrophilic, as discussed. Second, the use of PEG minimizes formation of detrimental tissue adhesions following surgery.

In FIG. 3B, block 302 consists of multiple sub-blocks. In particular, the polymers are rendered of tunable hydrophilicity by including within the block 302 different polymeric blocks, some of which are hydrophilic 306, and some of which are hydrophobic 308. In the example of FIG. 3B, the block consists of a central hydrophilic block 306 of repeating units of a first molecule. The block has a molecular weight of between 300 and 30,000 daltons. The block may have an even larger molecular weight, of even up to 40,000 or 60,000 daltons. Peripheral to central block 306 are additional blocks 308. The outer blocks are made of a different material which is hydrophobic. The length of the hydrophobic chains, and the ratio between the lengths of the hydrophobic and hydrophilic chains, may be selected in order to control the hydrophilicity of the ink. For hydrophobic inks, the central block may be hydrophobic, and the outer blocks may be hydrophilic. In addition to including segments of different hydrophilicity within a single polymeric chain, it is also possible to prepare separate components within the ink having different hydrophilicity, and to mix or bond them together during the printing process, to thereby generate the 3D printed device with the desired degree of hydrophilicity. In principle, it is possible to achieve a device of specific hydrophilicity by performing a polymerization reaction on two different monomers in different ratios.

Additional blocks may be added to the polymeric block 302. For example, a biodegradable polymeric segment 307, such as an aliphatic ester, may be added to a periphery of the block 302 or otherwise located within the ink molecule. The aliphatic ester may be polycaprolactone, polyglycolic acid, or polylactic acid and combinations thereof. In such instances, the blocks 307 may include a relatively limited number of repeating molecules in a polymeric chain (for example, n = 4 to 40). Biodegradable elements may be introduced into the polymeric chain, as shown here, or may be covalently bonded to other segments of the polymeric block 302. In another example, the blocks may have sacrificial components designed to change the water solubility of the polymer over time. These may be exemplified by a block copolymer comprising polyethylene oxide (PEO) and polylactic acid (PLA) blocks. Depending on the length of each of the blocks and the weight ratio between them, the biodegradable polymer may be initially non-water soluble or water soluble. This type of material may also be a hydroxylterminated PLA-PEO-PLA triblock, that may be rendered polymerizable or crosslinkable by end-capping with one or two C=C double bonds, by reacting it, for example with isocyanate ethylmethacrylate (IEMA). As degradation takes place, the initially non-water soluble molecule becomes water soluble.

Examples of blocks that are suitable for inclusion in the shape memory polymers are provided below. For some of these examples, methods of manufacture and properties of the resulting polymers are illustrated in FIGS. 3D-3E. As may be recognized by those of skill in the art, polymeric blocks having similar characteristics may be manufactured in similar ways. Providing a specific example with respect to one type of polymer block should not be understood as excluding synthesis of a similar polymer.

FIG. 3D illustrates a reaction of a polymerized, typically hydroxyl-terminated block 302a of PEG with 2-Isocyanoethyl methacrylate (IEMA) to create an end-capped polymer. The end-capping reaction is performed by reacting the terminal hydroxyl groups of the chain with the IEMA in a 1:2 ratio of the chain with the end-capping molecule. This end-capped polymer is referred to as PEG-dMA (i.e., polyethylene glycol with methyl acrylate on both ends). Although not shown, the same reaction may be performed with 2-Isocyanoethyl acrylate (IEA) and a similar reaction with methacryloyl chloride, for example. Because PEG is hydrophilic, the resulting end-capped polymer is substantially hydrophilic. Chains of PEG were evaluated having molecular weight of between 600 to 24,000 daltons. The PEG chains were optionally coupled together via bifunctional molecules such as diisocyanates and diacyl chlorides, among others, generating urethane or ester groups along the backbone, respectively. Referring now to FIG. 3E, aiming at gaining further freedom of design, in order to better control their molecular weight between crosslink junctions (Me) and improve the shape memory capabilities and also render the inks biodegradable, “inks” consisting of PEG segments and polycaprolactone (PCL) blocks, were synthesized.

Novel photo-curable (meth)acrylate-capped “inks”, consisting of a flexible PEG chain as their middle segment, and semi-crystalline PCL blocks of controllable length on both sides were synthesized. The PCL-PEG-PCL triblocks were then end-capped with photo-curable carbon double bonds, to render them printable. The composition of the triblocks was optimized by fine tuning the molecular weight of each of the two components as well as the ratio between them. In some instances the crystalline PCL or other crystalline molecules may perform as the net-points, with the crystallites performing as physical crosslink domains. In other embodiments, di(meth)acrylate “inks” were designed to comprise suitably semi-crystalline PCL segments that performed as switching segments, while the crosslink junctions functioned as the net-points, both components largely defining the shape memory behavior of the device to be printed. By judiciously fine tuning the length of both the PEG segment and the PCL blocks incorporated into the “inks”, the mechanical properties of the device and its shape memory capability were optimized. This design strategy is exemplified here for PEG and PCL segments, but can be implemented with a variety of correspondingly suitable segments.

FIG. 3E illustrates formation of a triblock copolymer. In the first step, the central triblock is formed by the ring opening polymerization of 8-capro lactone (CL), initiated by the terminal hydroxyl groups of the PEG molecule. This results in a triblock (referred to herein as “PECA”) with a central unit of PEG, which is flexible, and external blocks of PCL, which is semi-crystalline and biodegradable. In the second step of the synthesis, the PCL-b-PEG-b-PCL triblock is end-capped with reactive carbon double bonds by reacting it with IEA or IEMA or other compounds comprising a group able to react with the terminal group of the triblock, typically hydroxyl, and a C=C bond. This is the same reaction as that shown in FIG. 3D. The composition and molecular weight of the PECA triblocks generated during the first step of the synthesis determine the morphology, mechanical properties, and shape memory behavior of the ink.

Although not shown, the exact same reactions may be performed using a hydrophobic segment, such as PPG or poly tetramethylene oxide (PTMO). Monofunctional monomers that may be added to the other components of the inks include hydrophilic hydroxyethyl methacrylate (HEMA) and hydrophobic n-butyl methacrylate (BuMA). Furthermore, the same reactions may be performed following separate preparation of a hydrophilic backbone and attachment thereto of hydrophobic blocks, for purposes of tuning both the hydrophilicity of the ink as well as the molecular weight between crosslink junctions.

This modular approach allows for accurate tailoring of the composition of each of the two components of the “inks” and the ratio between them and, therefore, their respective morphologies. As a result, inks displaying different viscosity values and printability levels prior and during printing, and constructs exhibiting a wide range of shape memory behavior and mechanical properties once printed, may be developed. The inks may also be rendered pH sensitive by adding, for example, acrylic acid to the ink molecule or to the printing bath.

Other components may be added to the polymeric blocks to achieve other effects. For example, in some embodiments, the addition of a mono-functional co-monomer such as N-vinyl pyrrolidone (NVP), acrylamide (AAm), PEG methacrylates (PEGMA), among others, allows better control of the hydrophilicity and crosslink density of the ink and, as a result, of its various properties. Experimental results show that increasing weight percentages of these co-monomers within the blocks 302 correlates with an increase in water uptake. Similarly, the addition of hydrophobic mono-functional monomers will result in a decrease in water uptake.

In still another example, the “inks” may also comprise two types of reactive moieties that will enable the “ink” to crosslink following two consecutive, orthogonal mechanisms. This concept is exemplified hereby as follows: (1) the C=C double bonds necessary for the photo-crosslinking of the “ink” during the printing process, as discussed above, and (2) ethoxysilane moieties are able to undergo in vivo hydrolysis, whereby silanol groups that are formed will then undergo a condensation reaction, further crosslinking the structure and increasing its rigidity, as clinically desired. Similarly, click chemistry mechanisms, as well as others, can be used to engineer into the device this staged crosslinkability. The initial suitable UV-triggered low crosslink density attained during the printing of the components and their concomitant enhanced chain mobility, enables the efficient and straightforward deployment of the printed device. It also allows minimally invasive implantation of the device. Then, after being deployed on the heart as described above, and when hydrated by physiological fluids, the second crosslinking reaction sets in, gradually increasing the overall crosslink density of the device, rendering it with the desired clinical performance.

Following their synthesis, compositional analysis of the inks may be performed by different techniques, including Infrared Spectroscopy (IR) and Nuclear Magnetic Resonance (NMR). For example, the degree of methacrylation of the PEO-based inks may be determined by NMR spectroscopy, by ratioing the large singlet at 3.6 ppm due to the protons of the EO repeating units and the vinyl protons appearing at 5.6 and 6.2 ppm.

Other embodiments of polymeric inks may be used to print the ventricular restraint device, without departing from the scope of the present disclosure. The netpoints and switching segments may be made of different monomers than those listed above, or combinations of the materials listed above and the examples listed below. In some of these embodiments, the inks are printed using mechanisms that do not involve UV crosslinking (e.g., fused deposition modeling). In some embodiments, suitable net points may be made of poly-L-lactic acid (P(L)LA), P-Lam (the block obtained by the ringopening polymerization of lactams such as caprolactam), hard blocks of polyether and polyester urethanes, polyamides, polyurea polymers, aromatic rings, polycaprolactone (PCL). Suitable switching segments may be made of polycaprolactone (PCL), poly-DL- lactic acid (P(DL)LA), poly-lactic-co-glycolic acid (PLGA), P(L)LA, soft segments of polyether and polyester urethanes, polycarbonates, polyolefins, polyamides, or polysiloxanes (e.g., polydimethyl siloxane). It should further be noted that certain polymers may function either as a net point or as a switching segment, depending on the manner of transition (e.g., heat or aqueous medium). In addition, the shape memory polymers may include one or more pendant crystalline chains. The polymers may also have a higher functionality and, for example, be tetra-armed molecules, with one or more of the four arms being end-capped with a molecule containing C=C double bond. In addition, in the foregoing examples, each of the hydrophilic or hydrophobic blocks was described as a single block. These blocks may alternatively consist of multiple segments, typically in the Ik - 4k molecular weight range, which are coupled via diisocyanates such as HDI, MDI, or H-MDI, or other suitable coupling agents.

Without limitation, the shape memory polymers used to manufacture devices according to the present disclosure may include, in addition or in the alternative to those already mentioned: olefinic copolymers such as copolymers of 1,4 butadiene and styrene; block PET-PEO copolymers, sometimes crosslinked with crosslinkers such as maleic anhydride, dimethyl 5-isophthaltes, trimethylol propane or glycerin; ABA triblocks of poly(2-methyl-2-oxazoline and poly tetrahydrofuran; organic-inorganic thermoplastic polymers of linear polynorbornene and polyhedral oligosilsesquioxane; copolymers of polycyclooctene and poly(5-norbomene-exo-exo-2,3-dicarboxylic anhydride); polyurethane crosslinked with trimethylol propane or glycerin; AA/MAA copolymers crosslinked with N,N’-methyelene-bis-acrylamide; MAA/N-vinyl pyrrolidone crosslinked with ethylene glycol dimethacrylate; PMMA/N-vinyl pyrrolidone crosslinked with ethylene glycol dimethacrylate; block copolymers of trans-polyisoprene which performs as switching segments and urethane blocks that function as net points; polycaprolactone/polybutylene terephthalate copolyesters; poly(ketone-co-alcohol); polyethylene and its copolymers crosslinked via ionizing radiation such as g radiation or neutrons; crosslinked polycaprolactone/polymethylvinylsiloxane blends, typically crosslinked via radiation; chemical crosslinking of poly[ethylene-co-(vinyl acetate)] with peroxides such as dicumyl peroxide; crosslinked crystalline polycyclooctene; soybean oil copolymerized with styrene and divinyl benzene, crosslinked typically using norbornadiene or dicyclopentadiene as the crosslinker; copolymers of stearyl acrylate, methacrylate crosslinked using N,N’-methyelene-bis-acrylamide as the crosslinker, with the crystalline stearyl side chains functioning as the switching segments; and/or copolymers of poly(octadecyl vinyl ether) diacrylates or dimethacrylates with butyl acrylate, with the crystalline octadecyl pendant chains performing as the switching segments.

Additionally, one or more of the inks or the sacrificial components may be environmentally responsive, for example, reverse thermo-responsive. This may be achieved by incorporating N-isopropyl acrylamide (NIPAAm) or poly-N-isopropyl acrylamide (PNIPAAm), with the former being able to co -print with the inks, and the latter acting as a sacrificial component. Another example of a reverse thermo-responsive polymer is based on macromolecules comprising PEO and PPO segments.

IV. Personalizing the Devices through 3D Printing

Due to the complex, multi-factorial nature of the pathologic process of heart failure, every patient’s heart exerts a unique injury pattern, both spatially and temporally, to which the device has to be responsive. For example, myocardial hypertrophy tends to spread unevenly throughout the heart with different anatomical locations being more afflicted than others in different patients, depending on the specific etiology and comorbidities. It is apparent, therefore, that there is a therapeutic benefit in personalizing a ventricular support device to the needs of each individual patient. Since the heart of each patient exerts a distinct behavioral pattern both in space as well as over time, developing a custom-made cardiac device having these capabilities is of the utmost importance.

The devices of the present disclosure, being made of shape memory polymers, are especially well-suited to personalization. This is due to a number of factors, including the different types of polymers that may be included in the device, as discussed above, as well as the different shapes into which the polymers may be printed.

In a method of manufacture according to the present disclosure, physiological cardiac imaging is used for the precise spatial and hemodynamic characterization of the failing heart. This digital information is then fed into a 3D printer, which will generate the personalized device using the specially designed “inks”.

Referring now to FIG. 4A, the methodology of the 3D printing may proceed in steps of a method 400. At step 401, the patient’s heart is imaged through a suitable imaging process, such as an echocardiogram or a chest CT. The results of the imaging are used to create a volumetric reconstruction of the heart at maximum diastole. From this reconstruction, a personalized computer aided design (CAD) of the ventricular support device may be generated. At step 402, strain data is measured for the patient’s heart. The strain data is used to evaluate the patient’s therapeutic needs at different three- dimensional locations of the heart and to determine therapeutic goals at different regions. The strain estimates may be mapped onto different 3D locations of the heart. Based on this information, at step 403, a computerized 3D model of the ventricular support device is generated. The computerized 3D model may take into account the strain estimates for each location in the heart. The personalized 3D model may include information such as desired flexibility and strength of the device at different locations. At step 404, the device is then 3D printed, with attributes configured to implement a specific therapeutic benefit onto different regions of the organ, and optionally at different time points, consistent with the therapeutic goals.

Optionally, in an interim stage, a 3D model of the patient’s heart may also be generated, in order to enable testing of the ventricular support device and design optimization. For example, an in vitro test may be implemented, consisting of the cyclic pulsation of the device on the 3D printed heart model inflated by a pump.

Referring now in more detail to step 401, imaging data may be obtained by one or more imaging techniques, such as echocardiography, MRI, Computed Tomography (CT), Computed Tomography Angiography (CTA), or positron emission tomography (PET) combined with CTA. FIG. 4B illustrates an image of a mammalian heart obtained through echocardiography. Echocardiography harnesses high frequency sound waves (ultrasound) to generate a model of the patient’s heart.

In preferred embodiments, the echocardiography is performed in 4D, meaning that the heart is measured in 3D over time, throughout the systole and diastole cycle. In one clinical test performed on a rat heart, a 4D cardiac echocardiography capability was added to the Vevo 3100 (Fujifilm, VisualSonic) Echocardiography instrument. This technology allowed obtaining of 3D volumetric files of the rat's heart during systole and diastole. The images were converted into a DICOM file, and using a commercial 3D reconstruction software (Meshmixer, Autodesk), a 3D printable volumetric model was assembled. The volumetric parameters were converted at a 1mm resolution and enabled to conduct volumetric assessment.

In order to better evaluate cardiac physiologic and hemodynamic properties of the heart, Flow and Tissue Doppler and Strain Imaging using the echocardiographic assessment was performed. This data is used to fine tune the properties of the “ink” and optimize the ventricular support device’s size and dynamic response.

The echocardiography data is processed using a Computer-Aided Design (CAD) software. As a result of this process, a stereolithography (STL) file may be generated, in which the digital object is sliced, preparing the geometric data for 3D computer graphics and printing of the custom-made ventricular support device. FIGS. 4C and 4D show internal and external slicing of the device model.

As discussed previously, for treatment of heart failure with reduced ejection fraction, the devices are printed in an initial undersized shape, which is the permanent shape of the shape memory polymer. For treatment of heart failure with preserved ejection fraction, the devices are printed in an initial oversized shape, which is the permanent shape of the shape memory polymer.

As discussed previously, the tailoring of the device to achieve specific therapeutic benefits at specific regions, consistent with the identified therapeutic goals, may be achieved in various ways. These include: varying the materials used in the support device. As discussed previously, the devices may be printed with different inks. The composition of the inks may be selected in order to impart a desired strength and flexibility at different locations and at different time points. For example, experimental results performed on hydrophobic inks demonstrated that synthesizing inks comprising longer PPO segments and/or shorter PCL blocks drastically enhanced their flexibility. The strength and flexibility may additionally be tailored at the individual locations through inclusion of sacrificial water-soluble components, through inclusion of biodegradable elements in the inks, through 3D printing different dimension differentials between the permanent shape of the device at different points in space and the heart at those respective points, and through application of different degrees of stretching at different part of the device, as discussed in connection with FIGS. 1A-1G. In addition, as will be discussed in the following section, the attributes of the device may also be tailored based on the shapes of the patterns in which the device is printed.

Not only may the devices be personalized at the stage of initial printing, but they may also be personalized to provide differential therapy over time. As discussed above, this may be implemented through having inks printed with biodegradable components that may or may not be bonded to the other inks. This results in an initially stiffer device that becomes more flexible as the biodegradable inks selectively degrade. In addition, this may be implemented by incorporating a water-soluble sacrificial component during the printing process that will gradually dissolve in situ. Conversely, or in parallel, the hydrophilic inks of the device may be selected to absorb water to different degrees. As a result, the device may selectively or in its entirety gradually become more flexible.

As discussed above, the specific technology used for the 3D printing (e.g., SLA, DLP, or FDM) may be selected based on the characteristics of the ink. In preferred embodiments, as discussed, the printing mechanism involves curing with ultraviolet energy.

The devices may be printed in their initial three-dimensional shape. That is, the devices may be printed in the shape illustrated in FIG. 1A and FIG. 2A, already sized for the heart. Alternatively, the device may be printed as a flat sheet, and may be subsequently bonded or otherwise closed up into their three-dimensional shapes. Yet in another embodiment, different sections of the device having different properties may be printed separately, and bonded together following various techniques, prior to, during or following deployment.

When the device includes microneedles, the microneedles are printed together with the device, in the same printing process.

V. Personalizing the Devices Through Selecting the Pattern of Printing

In addition to the methods described above, one way to tailor the device to the needs of a particular patient is to select a particular pattern in which the device is to be printed. The device may be printed in a mesh pattern, with alternating regions of printed material and openings. The mesh pattern may be a repeating geometric pattern. The various parameters of the patterned structure may change in space, such as, for example, consisting of different inks at different locations, and/or the size and/or orientation of the openings may change at different regions of the heart, as dictated by clinical requirements.

The morphology of the material printed may be studied by Differential Scanning Calorimetry (DSC) and X-rays Diffraction (XRD) analyses. Scanning Electron Microscopy (SEM) may be used to study the topography of the surface of the device, while X-ray photoelectron spectroscopy (XPS) is an analytical tool that may be utilized to determine the surface chemistry of device. The mechanical behavior of the devices printed may be determined by an Instron Universal Testing Machine.

FIG. 5A and FIG. 5B illustrate devices printed in honeycomb patterns. FIG. 5A represents computer models of the devices, and FIG. 5B illustrates images of devices that were actually printed. The numbers next to each device represent the diameter of each of the holes in the devices, measured in pm (from 800 pm to 4,000 pm).

Experimental results were conducted on solid devices as well as devices printed in the honeycomb pattern with cells of different diameters. Generally, the Young’s modulus of the device decreased as the size of the cells became larger. For example, a dog-bone shaped device printed from a polymer having a PPO-based backbone with molecular weight 12,000 daltons was prepared in honeycombs with different diameters. The relevant strength and flexibility data is summarized in the table below:

Table 1. Strength Data for Honeycomb Shapes Having Different Diameters

Thus, a solid device had a 470 kPa Young’s modulus, versus a 7 kPa Young’s modulus for a device printed in a honeycomb shape with a 1600 pm cell diameter. It is apparent from the data presented that a remarkable two orders of magnitude increase in the flexibility of the device, approaching isocompliance with the beating heart, was achieved. These results proved that the mechanical performance of the device may be fine-tuned over a vast range, by optimally designing its architecture in addition to selecting the ink being printed, enabling the engineering of personalized devices.

It should be noted that although the patterns of FIGS. 5A and 5B are regular, with the same diameter of each cell throughout the structure, this need not be the case. The device may be formed of repeating units having different sizes or different shapes. For example, the honeycombs of FIGS. 1A-1G have smaller diameters near the apexes and larger diameters corresponding to the body of the ventricle.

Still another manner of personalizing the devices is to vary the ratio of the size of the device as printed relative to the placement shape and the treatment shape. For example, for treatment of HFrEF, instead of printing the entire device at 80% of the final treatment shape, some parts of the device may be printed at 80% of the final shape and some at 85%. The part that is printed at 80% size will exert a stronger compressive force on the heart during peak diastole than a part that is printed at 85%. This variation may be implemented even if the mold that is used for stretching the device is not capable of differential stretching. A similar principle applies for treatment of HFpEF. In that instance, as well, printing the initial, placement shape at different degrees of oversize causes the corresponding restorative outward force to differ after the device is affixed to the heart.

Referring now to FIGS. 5C-5E, an additional capability may be imparted to the pattern of the device by rendering the device with auxetic behavior. An auxetic structure refers to a structure that has a negative Poisson’s ratio. This means that when stretched or compressed in a given direction, the auxetic structure is also correspondingly stretched or compressed in a perpendicular direction. This occurs due to a particular internal configuration of the auxetic structure and the way this internal configuration deforms when the material is uniaxially stretched. The same unique response is obtained also when under compression, with the structure contracting transversely when compressed longitudinally. Of course, in order to take advantage of the auxetic structures, it is necessary for the inks to be able to undergo the required deformations under the physiological stresses in place.

Auxeticity is of value in enabling the ventricular support device to optimally respond to the expanding and contracting stresses applied by the heart during the systolic and diastolic phases of the cardiac cyclic, since lateral shrinking may significantly curtail its clinical efficacy. Auxetic constructs typically display superior mechanical properties.

Thus, in some embodiments, the ventricular support device is characterized by having an auxetic structure. The auxetic structure is constructed of a plurality of unit cells that are arranged in such a way that the overall structure expands orthogonally to the stretching direction when stretched and contracts orthogonally to the compression direction when compressed. The unit cell architecture may be any of the known architectures, including a re-entrant bow- shaped architecture, honeycomb architecture, and complex forms thereof. The unit cells and their properties can be modified by varying the unit cell architecture, such that upon stretching or compressing of the device, e.g., in response the cyclic pulsating activity of a pulsating organ, enhanced mechanical properties are achieved. It is important to note that such enhanced properties derived from the auxetic behavior may generally be independent of any particular material, but be mainly due only to geometrical considerations. That said, the mechanical properties of the ink/s used are of the utmost importance, since they will enable the auxetic structure to undergo its unique dimensional behavior, when under the physiological stresses applied by the heart.

FIGS. 5C and 5D illustrate auxetic structures 3D printed using flexible shape memory ink, as described above. FIG. 5E illustrates the auxetic properties of one such structure. As the user stretches the device in the horizontal connection, the device also expands in the vertical direction.

It should be noted that some of the benefits of auxetic structures may be independently obtained through use of the microneedles as described in the embodiment of FIGS. 2A-2I. Due to the secure and widespread fixation to the heart achieved by the microneedles, the microneedles could thereby prevent foreshortening orthogonal to the radial direction. This advantage is thus potentially relevant both for treatment of heart failure with reduced ejection fraction and for treatment of heart failure with preserved ejection fraction. In addition, the microneedles may be applied onto an auxetic structure, in order to magnify this effect.

VI. Implementation and Clinical Study

Following printing of the device, the device may be deployed over the heart, in the manner described in connection with FIGS. 1A-G and FIGS. 2A-I. The device is implemented during a sternotomy procedure. This type of surgery, while invasive to some degree, is far more common and less risky than implementation of a device within the left ventricle, as has been done to date for devices for treating HFpEF.

Clinical tests were performed on rats for devices similar to those described in connection with FIGS. 1A-1G, for treatment of heart failure with reduced ejection fraction. The clinical tests included deploying the device, and then, after 7, 14, and 28 days, performing echocardiography of the treated ventricles. In at least a portion of the devices, the devices remained attached in their desired clinical locations, and were functioning synchronously with the heart. Furthermore, the devices prevented remodeling of the left ventricle beyond what was already present prior to implementation. Testing further demonstrated that the devices did not cause any deleterious effect on diastolic function, and did not give rise to any unwanted adhesions.

In a preferred embodiment the device will be deployed following minimally invasive procedures, such as, among others, minithoracotomy procedures, and/or being navigated to the site endoluminally.

Claims

CLAIMS:

1. A support device for a pulsating organ, comprising a structure formed at least partially of one or more polymers, wherein the structure is configured to transition, in response to a stimulus, between a placement shape, in which the support device is dimensioned larger than the dimensions of the organ, and a treatment shape, in which the support device provides therapeutic support to the organ; wherein the treatment shape is smaller than the placement shape.

2. The support device of claim 1, wherein the one or more polymers are shape memory polymers.

3. The support device of any of the preceding claims, wherein the structure comprises a mesh pattern.

4. The support device of claim 3, wherein the structure is an auxetic pattern.

5. The support device of any of the preceding claims, wherein the stimulus is a mechanical stimulus.

6. The support device of any of the preceding claims, wherein the one or more polymers are shape memory polymers, wherein the support device is manufactured in an initial shape in which the one or more shape memory polymers are in their permanent shapes, and wherein the placement shape is larger than or equivalent to the initial shape.

7. The support device of claim 6, wherein the stimulus for transitioning between the placement shape and the treatment shape is application of an aqueous medium.

8. The support device of claim 6 or 7, wherein the device is 3D printed from at least one ink displaying shape memory behavior.

9. The support device of any of claims 6-8, wherein the device is 3D printed through curing with ultraviolet energy.

10. The support device of claim 9, wherein the ink comprises one or more central polymeric segments and at least one endcap molecule, wherein each endcap molecule includes at least one double carbon bond suitable for reacting with a corresponding double bond of a different endcap molecule by curing with the ultraviolet energy.

11. The support device of claim 10, wherein the endcap molecule comprises acrylate or methacrylate.

12. The support device of any of claims 6-11, wherein the ink comprises a shape memory polymer formed of a segment of repeating units of a first monomer, wherein a cumulative molecular weight of the segment is between 500 and 30,000.

13. The support device of claim 12, wherein the first monomer is hydrophilic.

14. The support device of claim 13, wherein the first monomer comprises polyethylene glycol.

15. The support device of claim 13 or 14, wherein the shape memory polymer further comprises one or more segments of a second monomer which is hydrophobic, wherein a ratio of the length of the first segment and the second segment is controlled, thereby enabling imparting of a desired hydrophilicity to the ink.

16. The support device of any of claims 12-15, wherein the shape memory polymer further comprises one or more biodegradable components covalently bonded to a polymeric chain of the shape memory polymer.

17. The support device of any of claims 12-16, wherein the shape memory polymer further comprises one or more ethoxysilane moieties covalently bonded to a polymeric chain of the shape memory polymer, wherein said moieties are able to undergo in vivo hydrolysis.

18. The support device of any of claims 12-17, wherein the shape memory polymer further comprises one or more biodegradable components included within the polymeric chain of the shape memory polymer.

19. The support device of claim 12, wherein the ink is further comprised of one or more additional molecules that are not bonded to the shape memory polymer prior to the printing process.

20. The support device of claim 19, wherein the first monomer is hydrophilic and the one or more additional molecules are hydrophobic, or vice versa, thereby enabling imparting of a desired hydrophilicity to the ink.

21. The support device of any of claims 19-20, wherein the one or more additional molecules comprise one or more of a water-soluble component, a biodegradable sacrificial component, or a bioactive agent..

22. The support device of any of the preceding claims, further comprising a plurality of microneedles arranged at an organ-facing surface of the support device, the plurality of microneedles configured to latch onto a surface of the organ.

23. The support device of claim 22, wherein the microneedles comprise substantially rigid bodies and substantially hydrophilic tips, wherein the hydrophilic tips are configured to swell when placed within a surface of the organ.

24. A method of applying a support device onto a pulsating organ, wherein the support device comprises a structure formed at least partially of at least one polymer, the method comprising: deploying the support device on the organ in a placement shape, in which the support device is dimensioned larger than dimensions of the organ; and transitioning the polymer from the placement shape to a treatment shape, in which the support device provides therapeutic support to the organ; wherein the treatment shape is smaller than the placement shape.

25. The method of claim 24, wherein the polymer is a shape memory polymer.

26. The method of claim 24 or 25, further comprising preparing the support device, the preparing step comprising: imaging the organ in three dimensions; determining therapeutic goals for different regions of the organ; and printing the support device in a 3D printing process with attributes configured to implement a specific therapeutic benefit onto different regions of the organ consistent with the therapeutic goals.

27. The method of claim 26, wherein the attributes comprise one or more of: materials used in the support device at different locations, thickness of the support device at different locations; pattern of the patterned structure at different locations, and strain and stress constants of the device at different locations.

28. The method of claim 26 or 27, wherein the attributes comprise hydrophilicity of the support device at different locations, and the printing step comprises utilizing inks having different degrees of hydrophilicity in different locations of the support device.

29. The method of any of claims 26-28, wherein the attributes comprise biodegradability of the support device at different locations, and the printing step comprises utilizing inks having different percentages of biodegradable components in different locations of the support device.

30. The method of any of claims 24-29, wherein the organ is a mammalian heart diseased with heart failure with reduced ejection fraction, the at least one polymer is a shape memory polymer, and the preparing step further comprises: printing the support device in the 3d printing process in an initial undersized shape corresponding to a permanent shape of the at least one shape memory polymer; applying a stimulus to the support device to thereby raise a temperature of the support device to above a glass or a melting temperature of a switching segment of the at least one shape memory polymer; stretching the support device to the placement shape; and removing the stimulus from the stretched support device while the support device is being kept in its stretched configuration.

31. The method of claim 30, wherein the applying step further comprises placing the device on an expandable mold; heating the mold to thereby transfer heat to the support device; expanding the mold to thereby stretch the support device; and cooling the mold.

32. The method of claim 31, wherein the expanding step comprises expanding the mold unevenly, to thereby stretch portions of the support device to different degrees.

33. The method of claim 31, wherein the placing step comprises layering the support device on the mold unevenly, to thereby stretch portions of the support device to different degrees.

34. The method of claim 30, wherein the transitioning step comprises applying an aqueous medium to the support device to thereby cause the support device to contract from the placement shape to the treatment shape.

35. The method of claim 24, wherein the organ is a mammalian heart diseased with heart failure with preserved ejection fraction, and the preparing step comprises printing the support device in the 3D printing process in an initial oversized shape which is the same as the placement shape.

36. The method of claim 35, wherein the transitioning step comprises applying a mechanical stress.

37. The method of claim 35, wherein the at least one polymer comprises a shape memory polymer, the transitioning step comprises applying an aqueous medium to the support device to thereby render the support device amorphous, and the method further comprises: stretching the support device when in an amorphous state to thereby shape the device around the heart; and fixing the stretched support device to the heart while the device is in said flexible state.

38. The method of claim 37, wherein the fixing step comprises applying one or more of sutures, staples, or adhesive.

39. The method of claim 37, wherein the device comprises a plurality of microneedles arranged on an organ-facing surface of the patterned structure, the plurality of microneedles configured to latch onto a surface of the exterior of the heart, and the fixing step comprises embedding the microneedles within an exterior surface of the heart.

40. The method of claim 39, wherein the microneedles comprise substantially rigid bodies and substantially hydrophilic tips, and wherein the fixing step further comprises causing the microneedles to swell after the microneedles are placed within a surface of the heart.

41. The method of any of claims 35-40, further comprising performing the fixing step when a left ventricle of the heart is at peak systole.

42. A method of preparing a support device for a pulsating organ, comprising: imaging the organ in three dimensions; determining therapeutic goals for different regions of the organ; and printing the support device in a 3D printing process with attributes configured to implement a specific therapeutic benefit onto different regions of the organ consistent with the therapeutic goals.

43. A method of embedding a medical device into an organ, wherein the medical device comprises microneedles having rigid bodies and hydrophilic tips, the method comprising: deploying the medical device onto the organ in a dry state; embedding the tips into the surface of the organ; and allowing the tips to swell within the organ, thereby locking the tips within the organ.