WO2025099720A1

WO2025099720A1 - A shrinking and setting bioink

Info

Publication number: WO2025099720A1
Application number: PCT/IL2024/051064
Authority: WO
Inventors: Tal Dvir; Ester-Sapir BARUCH
Original assignee: Ramot at Tel Aviv University Ltd
Current assignee: Ramot at Tel Aviv University Ltd
Priority date: 2023-11-08
Filing date: 2024-11-06
Publication date: 2025-05-15
Anticipated expiration: 2026-05-08

Abstract

Provided is a hybrid hydrogel composition and methods for its preparation and use in 3D printing biomaterials for tissue engineering applications. The hybrid hydrogel composition comprises discrete hydrogel particles of a first crosslinked polymeric network, interpenetrated- through and threaded by a second crosslinked polymeric network, dispersed in a medium. The first polymer is a thermo-responsive polymer, while the second is a protein-based thermo¬ denaturing polymer. The process for preparing this composition involves mixing swelled hydrogel nanoparticles of the first polymer with a non-crosslinked hydrogel of the second polymer in the presence of a crosslinking agent for the first polymer. The 3D printing method utilizes this composition below the LCST of the thermo-responsive polymer, followed by temperature increase wherein the second polymer cures and fixes the printed ink. This approach enables the creation of high-resolution structures, including artificial tissues with capillary blood vessels, featuring dimensions less than 50 μm.

Description

A SHRINKING AND SETTING BIOINK

RELATED APPLICATION

This application claims the benefit of priority of U.S. Provisional Patent Application No. 63/596,997 filed on November 8, 2023, the contents of which are incorporated herein by reference in their entirety.

FIELD AND BACKGROUND OF THE INVENTION

The present invention, in some embodiments thereof, relates to bioengineering and additive manufacturing methods, and more particularly, but not exclusively, to a shrinking and setting bioink.

Tissue engineering (TE) is an evolving field, which integrates life and materials sciences to provide solutions for damaged tissues and organs. This approach involves the combination of cells and 3D biomaterials to develop a living tissue. In particular, this field provides an alternative solution by generating functional substitutes for the injured myocardium. For tissue engineering to succeed, therefore, the intricacies of tissue architecture, from the nanometric structure of the extracellular matrix to the micro- and macroscopic levels of cellular organization, must be precisely recapitulated. In particular, one of the most pressing challenges for tissue engineering is the need to integrate a microvasculature that can efficiently provide cells with necessary nutrients and remove harmful cellular waste products from their environment.

Recent studies have demonstrated that cell organization into tissues with morphological and physiological characteristics that are similar to those in vivo requires a 3D scaffold that closely resembles the extracellular matrix (ECM) of natural tissues in terms of biochemistry, structure, and mechanical properties. These properties are determined by the ECM components, which include mostly collagen in addition to other proteins such as laminin, elastin and fibronectin as well as GAGs and proteoglycans. Therefore, researchers have focused on developing materials and technologies to simulate aspects of this specialized microenvironment. In spite of the fact that synthetic scaffolds can be constructed out of biocompatible materials in a way that closely mimics ECM structure, they still lack the fine, complex architecture and biochemical cues present in native ECM. On the other hand, ECM and ECM-derived materials, such as collagen, can be manipulated, fabricated, and processed in a manner that reliably replicates the natural microenvironment of the cell.

The realm of 3D bioprinting represents a transformative frontier in modern science and medicine, with the potential to revolutionize the way healthcare is approached. This cutting-edge technology has enabled the creation of intricate biological structures, including blood vessels, holding the promise of enhancing organ transplantation, tissue engineering, and disease modeling. However, it becomes apparent that even the most advanced techniques for 3D-printing of biological objects are not without their limitations.

One of the foremost challenges facing 3D bioprinting is the limitation in achieving high- resolution microvascular structures, particularly small capillaries with inner diameters less than 50 microns. These tiny conduits are the lifeblood of circulatory systems, playing a pivotal role in delivering oxygen and nutrients to every cell in the bodies. While existing bioprinting methods have made significant strides in replicating complex vascular networks, achieving the fine detail and precision required for these microvessels remains an ongoing challenge. This limitation is not merely a technical hurdle but a critical one, as the successful bioprinting of microcapillaries is fundamental to creating functional, viable tissues and organs. The ability to recreate these intricate vascular systems is essential for ensuring proper perfusion, which, in turn, is crucial for the longterm survival of 3D-printed tissues.

Three-dimensional (3D) cardiac patches provide a promising regenerative approach for replacing a damaged area of the heart. However, in order for engineered tissue to be functional, several hurdles must still be overcome. Notably, cardiac muscle cells must assemble in the proper locations and with the proper orientation, and an efficient blood vessel network must be present within the tissue to provide oxygen and nutrients to the cells. Many techniques have been used to develop vascularized, engineered cardiac tissues, including co-culturing cardiomyocytes with endothelial and other supporting cells and integrating proangiogenic factor-releasing systems within 3D scaffolds to attract blood vessel-forming cells. Nevertheless, these techniques have two significant shortcomings: they provide no control over blood vessel architecture and location, and they require waiting for upwards of two weeks before the vasculature is sufficiently developed to allow perfusion and support the parenchymal tissue.

To address these challenges, 3D printing has emerged as one of the most promising strategies in tissue engineering. Specifically, extrusion-based technologies are preferred for bioprinting due to the ease with which distinct bio-inks and cell types can be localized within the same tissue. In this way, cardiac patches have been engineered with a fully formed vasculature to provide the necessary transfer of oxygen and nutrients.

Although this technology shows great promise for generating complex tissues, several limitations still jeopardize its success. One of these is its limited printing resolution, which hinders the generation of accurate small-scale tissue units such as the glomerular capsule of the kidney, islets of Langerhans, and small-scale capillaries. When a shear-thinning, cell-containing hydrogel- based bio-ink is extruded from the printer, it immediately expands after leaving the needle. Though, increasing the print resolution by narrowing the needle results in a concomitant increase in shear stress, eventually leading to cell death. There is therefore an urgent need for a technology that can be used to generate, cellularized structures with small-scale tissue units at their correct anatomic scale.

One avenue being explored to overcome this deficit relies on printing “smart” materials in what has been termed “4D printing”. 4D printing is a type of 3D printing that creates objects that can change shape over time in response to stimuli, such as temperature, light, or moisture. 4D printing has the potential to revolutionize the way diseases and injuries are treated, as it can be used to create implants and bioinks that can adapt to the body over time. For example, a 4D printed implant could be used to repair a damaged bone. The implant would be printed with a material that is initially soft and flexible. Once the implant is implanted in the body, it would be exposed to body temperature and fluids. These stimuli would cause the implant to harden and become more rigid, forming a new bone that is compatible with the body.

4D printing is 3D printing that can create objects that can change shape over time. This is useful for making things like implants or tissues that can adapt to the body or materials that can respond to changes in the environment. While some smart materials rely on stimuli that are inherently problematic for biological applications, other materials, such as PNIPAM (poly(N- isopropylacrylamide)), can be readily combined with cells to create biologically relevant smartmaterials. Aqueous solutions of PNIPAM display a sharp sol-gel transition at approximately 32 °C. Below their lower critical solution temperature (LCST), polymeric chains of PNIPAM are extended, minimizing the viscosity of the solution, while above it the PNIPAM molecules collapse and become entangled with one another, forming a hydrogel. These nanometric shifts in the polymers’ conformations lead to several significant changes in the material’s macroscopic properties, including an increase in viscosity, a decrease in hydrophilicity, and a decrease in the volume of the gel. Importantly, these phenomena are observed both in bulk solutions of PNIPAM and when it is incorporated within other polymeric networks to create composite materials with unique, smart properties. While these smart hybrid materials have been demonstrated to be particularly suitable for drug-delivery applications, it has not yet been shown that these technologies can be easily adapted for tissue engineering. For example, Ding et al. [“Thermal and pH dual-responsive hydrogels based on semi-interpenetrating polymer network of poly(N- isopropylacrylamide) and collagen nanofibrils. Polymer International. doi:10.1002/pi.5852] reported the development of thermo- and pH-responsive hybrid hydrogels consisting of semiinterpenetrating polymer networks (semi-IPNs) of poly(N-isopropylacrylamide) (PNIPAM) and collagen nanofibrils. These hydrogels were prepared by first forming a collagen nanofibrillary hydrogel through self-assembly under physiological conditions, followed by in situ polymerization of NIP AM within the collagen network. The resulting semi-IPNs exhibited temperature-sensitive swelling behavior due to the PNIPAM component, as well as pH-sensitivity from the collagen nanofibrils. The swelling ratios and drug release profiles of the hydrogels could be tuned by varying the ratio of PNIPAM to collagen. At physiological temperature (37 °C) and pH (7.2), the semi-IPNs showed controlled release of a model drug (methyl violet) over several hours. The authors proposed that these dual-responsive collagen/PNIPAM semi-IPN hydrogels have potential applications in tissue engineering and drug delivery. Ding’s et al. “smart” drug-delivery system that was fabricated by first allowing collagen to gel at 37°C and then adding NIP AM monomers to be polymerized in situ. The final product could be significantly shrunk by lowering the pH to 3. However, in the context of tissue engineering, this process is impossible to replicate as the NIP AM monomers are toxic (and therefore must be polymerized before cell exposure), and the 4D stimulus-responsive behavior cannot be triggered by such acidic conditions without causing massive cell death.

Ding, C. et al. [“Thermal and pH dual-responsive hydrogels based on semiinterpenetrating polymer network of poly(N -isopropylacrylamide) and collagen nanofibrils” , Polym Int 68, 1468-1477 (2019)] provided hybrid hydrogels with semi-interpenetrating polymer networks (semi-IPNs) for controlled drug release applications. These hydrogels combine PNIPAM as a thermo-sensitive component and self-assembled collagen nanofibrils as a pH- sensitive framework. The semi-IPNs exhibited unique swelling behaviors under various pH values and temperatures, distinct from neat collagen or PNIPAM hydrogels.

SUMMARY OF THE INVENTION

3D tissue printing has significantly evolved over the past several years and can now be used to generate a controlled vascular network within engineered tissues for the proper transfer of oxygen and nutrients. However, a significant remaining challenge when using cell-containing bioink hydrogels is overcoming limited printing resolution, which hinders the generation of controlled, small-scale cellular features such as capillaries. Here, the present inventors disclose a cell-containing nanoparticle-ECM bio-ink that significantly reduces its dimensions, in a controlled manner, after printing and being exposed to temperatures above 32°C. The ability of the bio-ink to interact with cells, shrink, and form accurate cellular structures was demonstrated. The bio-ink was then co-printed with Gelatin beads bio-ink, within pristine ECM bio-ink that were used to create cardiac tissues and large blood vessels. Immediately after heating the printed tissue, selective morphology changes occurred at kinetically controlled sequence, resulting in a functional and perfusable (capable of being perfused) human cardiac patch with varying blood vessel dimensions. Such technology may allow the generation of small-scale tissue units at their correct size.

The present disclosure pertains to a novel hybrid hydrogel. This hybrid hydrogel comprises two polymeric networks: a first crosslinked polymeric network and a plurality of nanoparticles within a second crosslinked polymer network. These nanoparticles are entrapped within the first crosslinked polymeric network. The second crosslinked polymer network contains a thermo- responsive polymer, such as poly(N-isopropylacrylamide) (PNIPAM). The first crosslinked polymeric network can include collagen-based hydrogel and/or ECM-based hydrogel. A unique feature is the formation of an interpenetrating polymeric network where the first network interpenetrates with the nanoparticles.

The present disclosure also covers the process of making this hybrid hydrogel. It involves providing swelled nanoparticles, non-crosslinked collagen-based hydrogel and/or non-crosslinked ECM-based hydrogel, and crosslinking agents. Mixing these components leads to the formation of the hybrid hydrogel. The process operates below the lower critical solution temperature (LCST) of the thermo-responsive polymer.

Furthermore, the present disclosure introduces a method for 3D printing objects using the hybrid hydrogel. The object is formed at a temperature below the LCST of the thermo-responsive polymer and then heated above the LCST. This method enables the creation of high-resolution objects with structural dimensions less than 50 pm, including tubular elements and artificial tissues with capillary blood vessels.

The present disclosure describes a hybrid hydrogel with unique properties and its method of production, along with its application in 3D printing high-resolution objects, especially those with small capillaries.

Thus, according to an aspect of some embodiments of the present invention, there is provided a hybrid hydrogel composition that includes a plurality of discrete hydrogel particles, each of the particles includes a first crosslinked polymeric network formed from a first polymer, and further includes a second crosslinked polymeric network formed from a second polymer, and a dispersing medium, wherein the second crosslinked polymeric network has crosslinked polymeric chains that interpenetrate-through each of the hydrogel particles, forming an interpenetrating polymer network dispersed within the medium, resulting in the hybrid hydrogel composition. In some embodiments, the interpenetrated-through hydrogel particles are maintained in the composition in a non-contacting, spaced-apart arrangement within the second crosslinked polymeric network.

In some embodiments, the interpenetrated-through hydrogel particles and the second crosslinked polymeric network forming the interpenetrating polymer network are not linked by covalent bonds.

In some embodiments, the first polymer includes a thermo-responsive polymer.

In some embodiments, the thermo-responsive polymer is a thermo-contractive polymer.

In some embodiments, the thermo-contractive polymer is characterized by a lower critical solution temperature (LCST) ranging from 30-40 °C.

In some embodiments, the thermo-contractive polymer includes poly(N- isopropylacrylamide), abbreviated “PNIPAM”.

In some embodiments, the second polymer is a thermo-denaturing polymer.

In some embodiments, the thermo-denaturing polymer is a protein-based polymer.

In some embodiments, the second polymer includes collagen and/or ECM-derived polymers, and the second crosslinked polymeric network forms a collagen-based or ECM-based hydrogel.

In some embodiments, the denaturing temperature of the thermo-denaturing polymer ranges from 35-40 °C.

According to an aspect of some embodiments of the present invention, there is provided a process for preparing the hybrid hydrogel composition, the process includes providing the plurality of discrete hydrogel particles in a swelled form, providing a non-crosslinked hydrogel of the second polymer, and mixing the hydrogel particles with the non-crosslinked hydrogel in the presence of a crosslinking agent, wherein the crosslinking agent effects the crosslinking of the second polymer to form the second crosslinked polymeric network and the interpenetrating polymer network, yielding the hybrid hydrogel composition, as provided and described herein.

In some embodiments, the process further includes, after mixing, washing the interpenetrating polymer network.

In some embodiments, the first polymer includes a thermo-responsive polymer, and the mixing occurs at a temperature below the LCST of the first polymer.

In some embodiments, the LCST ranges from 30-40 °C.

In some embodiments, the second polymer is a thermo-denaturing polymer.

In some embodiments, the denaturing temperature of the thermo-denaturing polymer ranges from 35-40 °C. In some embodiments, the progress and completion of the crosslinking of the second polymer are determined by monitoring the viscosity of the hybrid hydrogel composition.

In some embodiments, the completion is defined when the viscosity ranges from 0.03 to 6 10⁵ Pa s.

According to an aspect of some embodiments of the present invention, there is provided a method of 3D printing an object, the method includes printing the object using a first bioink that contains the hybrid hydrogel composition, wherein the first polymer is a thermo-contractive polymer characterized by an LCST, while the second polymer is a thermo-denaturing polymer characterized by a denaturing temperature, with the LCST being lower than the denaturing temperature.

In some embodiments, the first crosslinked polymeric network contains PNIPAM, and the second crosslinked polymeric network forms a collagen-based or ECM-based hydrogel.

In some embodiments, the printing occurs at a temperature lower than the LCST of the first polymer.

In some embodiments, the method includes, after printing, gradually or incrementally increasing the temperature of the object above the LCST of the first polymer.

In some embodiments, the method further includes printing with a second bioink that contains a non-crosslinked hydrogel of a thermo-denaturing polymer. The second bioink is printed at a temperature lower than the denaturing temperature and the LCST of the first polymer.

In some embodiments, the second bioink is printed to envelop at least a portion of the outer surface of the part of the object formed by the first bioink.

In some embodiments, the second bioink includes a non-crosslinked collagen-based or ECM-based hydrogel.

In some embodiments, the method includes, after printing, gradually or incrementally increasing the temperature of the object above the LCST of the first polymer and then above the denaturing temperature of the second polymer to finalize the object.

In some embodiments, the object is characterized by a lumen, in which case the method further includes printing the lumen using a third bioink composed of a low-melting point hydrogel with a melting or gel-sol transition temperature lower than the LCST of the first polymer of the first polymer, wherein the third bioink is printed in the position, shape, and size of the lumen, and essentially, fully or at least partially enveloped by the first bioink.

In some embodiments, the third bioink includes substances like gelatin, a poloxamer, PVA, and/or a carbomer resin. In some embodiments, the method includes, after printing, gradually or incrementally increasing the temperature above the melting temperature of the third bioink, or above the LCST of the first polymer, followed by a gradual or incremental temperature increase above the denaturing temperature of the second polymer to complete the object.

According to an aspect of some embodiments of the present invention, there is provided an object that includes the hybrid hydrogel composition, as provided herein.

In some embodiments, the object is formed by the method of 3D printing described herein.

In some embodiments, the object includes at least one structural lumen characterized by a dimension of less than 30 pm, less than 40 pm, less than 50 pm, less than 60 pm, less than 70 pm, or less than 80 pm.

In some embodiments, the object is an artificial tissue containing capillary blood vessels with an inner diameter of less than 50 pm and a length of at least 0.1 mm, at least 0.2 mm, at least 0.5 mm, or at least 1 mm.

According to an aspect of some embodiments of the present invention, there is provided a method for anisotropic 3D printing of an object, the method includes providing a bioink containing the hybrid hydrogel composition, extruding the bioink to form part of the object, controlling the direction of the printhead to create directional properties, and stimulating the first polymer in the hybrid hydrogel composition (first bioink) to induce anisotropic shrinkage.

In some embodiments, the anisotropic shrinkage is greater perpendicular to the direction of the printhead motion than parallel to it.

In some embodiments, the method includes printing geometric shapes in different patterns to achieve distinct shrinkage behaviors.

In some embodiments, the geometric shapes include circular structures that exhibit radial compaction upon stimulation.

In some embodiments, the radial compaction results in varying rates of change for the inner and outer diameters of the circular structures.

In some embodiments, stimulating the hybrid hydrogel involves applying heat.

According to an aspect of some embodiments of the present invention, there is provided a

3D printed object produced by the method described, exhibiting anisotropic shrinkage in response to stimulation. BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWINGS

Some embodiments of the invention are herein described, by way of example only, with reference to the accompanying drawings. With specific reference now to the drawings in detail, it is stressed that the particulars shown are by way of example and for purposes of illustrative discussion of embodiments of the invention. In this regard, the description taken with the drawings makes apparent to those skilled in the art how embodiments of the invention may be practiced.

In the drawings:

FIGs. 1A-J present a sequence of illustrations demonstrating the formation of the hybrid hydrogel composition, according to some embodiments of the present invention, whereas the steps include the provision of discrete, swelled crosslinked hydrogel particles, the introduction of a crosslinking agent, the mixing of these particles with a non-crosslinked hydrogel of a second polymer, and the subsequent formation of an interpenetrating polymer network;

FIGs. 2A-F depict the step-by-step process of 3D printing a hollow capillary tube using three distinct bioinks, according to some embodiments of the present invention, whereas the illustrations show the sequential deposition of the bioinks, the development of the capillary structure, and the final object after thermal processing, where the lumen is formed, and the bioinks shrink and set;

FIGs. 3A-G illustrate the process and outcomes of the coordinated multi-kinetic 3D printing of vascularized tissue constructs using various bioinks, wherein FIG. 3A provides a schematic of the printing setup and bioink arrangement, FIG. 3B shows images demonstrating the structural integration of the bioinks after thermal activation, FIG. 3C highlights the perfusion capability through the printed vascular structures, FIG. 3D depicts the perfusion and response of the vessels before and after shrinking, FIG. 3E summarizes the extent of shrinkage observed in the dynamic vessels, FIG. 3F presents images comparing shrunk and static vessel-like channels, and FIG. 3G provides measurements of vessel dimensions under different temperature conditions;

FIGs. 4A-F illustrate the anisotropic properties and behaviors of printed structures using the hybrid hydrogel composition, according to some embodiments of the present invention, wherein FIG. 4A presents a schematic diagram showing the anisotropic swelling and deswelling behavior along different axes in printed strands, FIG. 4B shows the distinct shrinkage patterns of pentagonal shapes printed in different orientations, FIG. 4C quantifies the varying shrinkage strains observed in these printed shapes, FIG. 4D depicts a printed ring undergoing radial shrinkage and compaction, FIG. 4E demonstrates the macroscopic shrinkage behavior of circular structures printed with the hybrid hydrogel composition (first bioink), and FIG. 4F provides measurements comparing the inner and outer diameters of the printed rings before and after temperature activation;

FIGs. 5A-C depict various characterizations of the PNIPAM nanogels, according to some embodiments of the present invention, wherein FIG. 5A shows a transmission electron microscopy (TEM) micrograph highlighting the morphology of the synthesized PNIPAM nanogels, FIG. 5B presents photographs of the macroscopic appearance of PNIPAM nanogels at different temperatures, demonstrating their contraction and water excretion behavior when heated, and FIG. 5C depicts rheological measurement results, indicating a significant increase in the viscosity of the PNIPAM hydrogels as a response to temperature elevation; and

FIGs. 6A-N illustrate the basic principles of using the hybrid hydrogel composition as a shrinking and setting bioink, according to some embodiments of the present invention, and present a comprehensive depiction of the evaluation of printed vascularized tissues and their integration with cardiac cells, wherein FIG. 6A illustrates a schematic of the 4D-bioprinting process, emphasizing the sequence of shape and structural changes, FIG. 6B displays an immunostained image of the cardiac patch immediately after printing, FIG. 6C captures the patch following a few hours of incubation, FIG. 6D shows a static blood vessel structure after extended maturation, FIGs. 6E-F illustrate images of dynamic blood vessels post-maturation, FIG. 6G presents a high- resolution image of a dynamic vessel integrated into the cardiac tissue, FIG. 6H quantifies red blood cell retention following perfusion, FIG. 61 demonstrates perfusion results of acellular constructs, FIG. 6J depicts cellularized vessels with minimized thrombin activity, FIG. 6K provides data on thrombin activity, FIG. 6L depicts a cardiac patch implanted on a rat, FIG. 6M shows post-implantation results indicating integration, and FIG. 6N highlights vascular anastomosis between human and rat vessels.

DESCRIPTION OF EMBODIMENTS OF THE INVENTION

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 set forth in the following description or exemplified by the Examples. The disclosure is meant to encompass other embodiments or of being practiced or carried out in various ways.

While conceiving the present invention, the inventors faced the limitations of the methodologies and bioinks for 3D printing scaffolds for live tissues, such as various blood vessels include arteries, veins and capillaries, available before the time of the present invention. The prior art tools, materials and methods, fail to form and provide delicate micrometric tubular 3D printed structures such as needed to form a small blood vessels and capillaries having an inner diameter of less than 50 microns (50 pm).

While further conceiving the present invention, the inventors contemplated the use of postprinting shrinkage to mitigate the printing resolution limitations. In simple words, the inventors contemplated a process in which structures are printed at the resolution limits of the available 3D printing tools and methods, using a combination of bioinks, wherein at least one of component is designed to exhibit thermal contraction post-printing, and at least one component that exhibits post-printing thermal setting/curing.

While reducing the invention to practice, the inventors developed a hybrid hydrogel composition, which is suitable for use as a bioink, and exhibits both thermal contraction, which is mechanically equivalent to shrinking, and thermal denaturing properties, which is mechanically equivalent to thermosetting or curing. The hybrid hydrogel composition provides both properties by comprising an interpenetrating polymer network (IPN) of at least two crosslinked polymeric networks. In some embodiments of the present invention, each crosslinked polymeric network contributes its thermal contraction property or its thermal denaturing property, thus a thermocontracting and thermo-setting bioink is obtained.

Thus, according to an aspect of some embodiments of the present invention, there is provided a hybrid hydrogel composition which includes: a plurality of discrete hydrogel particles, each particle comprising a first crosslinked polymeric network formed from a first polymer; a second crosslinked polymeric network formed from a second polymer; and a dispersing medium, wherein: the second crosslinked polymeric network comprises crosslinked polymeric chains that interpenetrate-through (thread) each of the hydrogel particles, thereby forming an interpenetrating polymer network that comprises interpenetrated-through (threaded) hydrogel particles; the interpenetrated-through hydrogel particles are maintained in a non-contacting, spacedapart arrangement within the second crosslinked polymeric network; and the interpenetrating polymer network of the interpenetrated-through hydrogel particles and the second crosslinked polymeric network is dispersed within the dispersing medium, thereby affording the hybrid hydrogel composition. Interpenetrating Polymeric Network:

In the context of the present invention and in the fields of material science, chemistry and bioengineering, the term "interpenetrating polymer network" or IPN, refers to a polymeric material comprising two or more polymer networks that are at least partially interlaced on a molecular scale, but not covalently bonded to each other. The IPN provided herein is characterized by the presence of distinct polymer networks, with at least one network being synthesized and/or crosslinked in the immediate presence of the other(s). These networks are physically interpenetrated throughout one-another, without forming covalent bonds between the different polymer networks. A key feature of the IPN provided herein is the inability to separate the networks without breaking chemical bonds, despite each individual network being capable of existing as an independent polymer structure in the absence of the other network(s). The IPN provided herein may be formed through various methods, including simultaneous or sequential network formation, and can exhibit unique physical, mechanical, and chemical properties distinct from those of the individual component polymers or simple polymer blends.

According to embodiments of the present invention, the hybrid hydrogel composition provided herein is an IPN that is formed when two or more polymer networks are physically intertwined without chemically reacting with each other. More specifically, the first crosslinked polymeric network and the second crosslinked polymeric network form an IPN without covalent bonding between the first polymer and the second polymer. It is assumed that this lack of covalent bonding allows for unique properties like enhanced mechanical properties, elasticity, and denaturing/contracting behavior, including maintaining stimuli-responsiveness. While covalent bonds may sometimes be present in IPNs for specific applications, as in the case of some specific embodiments of the present invention, the defining characteristic of an IPN is the absence of such bonds between the primary polymeric components thereof.

The second crosslinked polymeric network comprises crosslinked polymeric chains that penetrated through particles (beads) of the first crosslinked polymeric network before being crosslinked, thus preventing the interpenetrated-through hydrogel particles (threaded beads) from slipping off the crosslinked polymeric chains (threads). By being interpenetrated-through, threaded or strung by at least one polymeric chain of the second crosslinked polymeric network, each of the hydrogel particles comprising the first crosslinked polymeric network is being held, secured, fastened, entrapped, and locked within and by the second crosslinked polymeric network.

As used herein, according to some embodiments of the present invention, the term "interpenetrated-through" refers to a structural configuration wherein a string-like element, such as a polymeric chain, a fiber, a filament, or a thread, passes through the core of a bead-like element, such as a hydrogel particle, a bead, a capsule, or a microsphere. This configuration results in the string-like element entering one side of the particle-like element, traversing its internal space, and exiting from another side, thereby creating a continuous path through the particle-like element. The term encompasses arrangements where the string-like element may be movable within the particle-like element or fixed in place, and where single or multiple particle-like elements may be interpenetrated-through by one or more string-like elements in various patterns or sequences.

Structurally, the hybrid hydrogel described herein consists of discrete hydrogel particles, each of which comprises a first crosslinked polymeric network. As oppose to forming clusters or layers, the interpenetrated-through hydrogel particles are distributed along the crosslinked polymeric chains and held in place by a second crosslinked polymeric network that weaves through each interpenetrated-through hydrogel particles. This physical interpenetration-through forms an interpenetrating polymer network, ensuring that the hydrogel particles do not come into direct contact with one another, thereby preserving their individual integrity and functionality. The first crosslinked polymeric network is present essentially as discrete particles that are spatially separated, ensuring that the particles are substantially arranged in a spaced-apart configuration. This non-contacting arrangement prevents the aggregation of particles and maintains substantially uniform distribution of the particles, also referred to herein as nanogels, throughout the second crosslinked polymeric network. The interpenetration of crosslinked polymeric chains of the second crosslinked polymeric network through the hydrogel particles of the first crosslinked polymeric network, enables the particles to remain securely positioned within the second crosslinked polymeric network, which can form a continuous, interwoven scaffold and be used as a bioink.

The second crosslinked polymeric network serves as the primary structural framework, within which the interpenetrated-through hydrogel particles are physically entrapped. During manufacturing of the IPN, the polymeric chains of the second polymeric network pass through the swollen hydrogel particles, forming a three-dimensional matrix that both supports and secures the interpenetrated-through hydrogel particles in place. Despite the lack of covalent bonds between the first and second polymer networks, the physical entanglement and crosslinking of the two networks ensures that the particles remain locked within the overall hydrogel construct, forming a cohesive structure.

The dispersing medium surrounding the hydrogel construct facilitates the overall stability and flexibility of the hybrid hydrogel. This medium allows the networks to remain hydrated, and it enables the mobility of the polymeric chains through the first polymeric network without causing the particles to shift or aggregate. The medium ensures that the overall hydrogel structure maintains its integrity, with the individual particles and the interpenetrating polymer network functioning as a unified, yet highly organized system. In some embodiments of the present invention, the dispersing medium is an aqueous solution.

Thus, the present invention provides a hybrid hydrogel in which discrete hydrogel particles, each comprising a first crosslinked polymeric network, are interpenetrated-through by polymeric chains of a second crosslinked polymeric network. This interpenetration creates an interpenetrating polymer network, where the interpenetrated-through hydrogel particles remain substantially spaced apart and essentially do not contact each other, at least not permanently or by bonding, while being securely entrapped within the second polymeric network, all within a dispersing medium that maintains the structural and functional integrity of the overall system.

When referring to a mass of a hydrogel or a hydrogel component, such as the hydrogel components in the hybrid hydrogel composition provided herein, the reference is made to the dry component, as oppose to the mass of a hydrated, partially swollen or fully swollen hydrogel. The mass ratio of the first crosslinked polymeric network to the second crosslinked polymeric network in the hybrid hydrogel composition provided herein is 20:80, 30:70, 40:60, 50:50, 60:40, 30:70 and 20:80 first crosslinked polymeric network to second crosslinked polymeric network, and any interim mass ratio value. In some embodiments the mass ratio ranges 20:80 to 80:20, 30:70 to 70:30, or 40:60 to 60:40. In a preferred embodiment the mass ratio is 40:60 of the first crosslinked polymeric network to second crosslinked polymeric network.

In some embodiments, the hybrid hydrogel composition provided herein is substantially devoid of hydrogel particles of the first polymer that are not interpenetrated-through by polymeric chains of the second polymer.

The first crosslinked polymeric network:

According to some embodiments of the present invention, the first crosslinked polymeric network comprises the first polymer, which is a stimuli-responsive polymer. Stimuli-responsive polymers are a class of materials that can change their physical or chemical properties in response to external stimuli. These stimuli can include changes in temperature, pH, light, electric or magnetic fields, and the presence of specific chemicals. Stimuli-responsive polymers can be divided into several families based on the type of stimulus that triggers their response. These families include:

Thermo-responsive polymers that respond to changes in temperature. Some examples of thermo-responsive polymers, which are contemplated in the context of the first polymer, include poly(N-isopropylacrylamide) (PNIPAM), poly(N- vinyl caprolactam) (PVCL), and poly(methyl methacrylate) (PMMA); pH-responsive polymers that respond to changes in pH. Some examples of pH-responsive polymers, which are contemplated in the context of the first polymer, include poly(acrylic acid) (PAA), poly (methacrylic acid) (PM A), and chitosan;

Light-responsive polymers that respond to light. Some examples of light-responsive polymers, which are contemplated in the context of the first polymer, include poly(azobenzene), poly(spiropyran), and poly(diacetylene);

Electrically-responsive polymers that respond to electric fields. Some examples of electrically-responsive polymers, which are contemplated in the context of the first polymer, include poly(aniline), poly(pyrrole), and poly(thiophene); and

Magnetically-responsive polymers that respond to magnetic fields. Some examples of magnetically-responsive polymers, which are contemplated in the context of the first polymer, include poly(ferrocenylsilane), poly(cobalt ferrite), and poly(gadolinium-DTPA).

Thermo-responsive polymers are a particularly important class of stimuli-responsive polymers in the context of the present invention, due to their mechanical properties and wide range of potential applications. As used herein, the term "thermo-responsive" refers to a property exhibited by certain materials that respond to changes in temperature by altering their physical and/or chemical characteristics. These materials can undergo reversible changes in response to temperature fluctuations. Typically, thermo-responsive materials have a distinct transition temperature, known as the lower critical solution temperature (LCST) or upper critical solution temperature (UCST), where their properties change significantly. The nature of this change can vary depending on the specific material and its intended application. Some common examples of thermo-responsive materials include certain polymers and hydrogels. Thermo-responsive materials may be defined by a phase transition - thermo-responsive materials often undergo a phase transition in response to temperature changes. For example, a polymer might shift from a hydrophilic (water-attracting) to a hydrophobic (water-repellent) state when heated above its LCST. Thermo-responsive changes are typically reversible; when the temperature is brought back to the appropriate range, the material will revert to its original state.

In the context of some embodiments of the present invention, the first polymer comprises or consists of a thermo-contractive polymer. In the context of “thermo-responsiveness”, the term “thermo-contractive polymer” refers to a subset of the property of thermo-responsiveness of polymers, and specifically refers to a class of polymers that exhibit a decrease in volume or physical dimensions when exposed to an increase in temperature (exposure to heat). This contraction is different than the more common thermal contraction of many substances upon experiencing a decrease in temperature. In the context of the present invention, physical contraction as a response to an increase in temperature occurs above the LCST of a polymeric substance, whereas the polymer chains undergo a conformational change, typically transitioning from an extended, hydrated state to a more compact, dehydrated state, resulting in a macroscopic contraction of the polymer structure, and this property is bestowed to the IPN that comprises the particles of the first crosslinked polymeric network.

The thermo-contracting property is distinct from general thermal expansion/contraction and is characterized by a non-linear, often significant reduction in size or volume in response to heating. In the context of hydrogels, thermo-contraction and deswelling are distinct phenomena. Thermo-contraction is a specific, temperature-driven response characterized by a rapid, significant reduction in hydrogel dimensions, typically occurring at or above the characteristic LCST. It involves an active conformational change of polymer chains from extended to collapsed states, often reversible, and is driven by the inherent properties of the polymer network. In contrast, deswelling refers to the general loss of water or solvent from the hydrogel matrix, which can occur due to various environmental factors such as evaporation, osmotic pressure changes, or mechanical stress. Deswelling is not necessarily temperature-specific, is typically more gradual, and primarily results from water loss rather than active polymer contraction. While both processes may reduce hydrogel volume, their underlying mechanisms, triggers, and characteristics are fundamentally different.

One of the most well-studied thermo-contractive polymers is PNIPAM. PNIPAM is a water-soluble polymer at temperatures below its lower critical solution temperature (LCST) of 32 °C. However, when the temperature is raised above the LCST, PNIPAM becomes insoluble in water and precipitates. This phase transition can be used to create a variety of “smart materials”, such as drug delivery systems and temperature-responsive surfaces.

Another example of a thermo-contractive polymer, which is contemplated in the context of the first polymer, is poly(N-vinyl caprolactam), or PVCL. PVCL has a similar LCST to PNIPAM, but it is more resistant to degradation. PVCL is often used in biomedical applications, such as tissue engineering and drug delivery. PVCL has a number of advantages over PNIPAM, including, higher LCST (32 °C vs. 31 °C for PNIPAM), greater resistance to degradation, and better biocompatibility. Additionally, PNIPAM derivatives, such as Poly(N-n-propylacrylamide) (PNNPAm), which has a LCTS of 10 °C or Poly(N-(2-m-l,3-dioxan-5-yl)methylacrylamide) (PNMM) with LCTS of 22 °C, are also considered as biocompatible, and may be served as substitutes for PNIPAM.

In the context of some embodiments of the present invention, the first polymer is selected such that is exhibits the desired LCST which is suitable for the desired application. The range of LCSTs for known thermo-contractive polymers is approximately 20 °C to 80 °C. In some preferred embodiments, the thermo-contractive polymer is characterized by an LCSTs in the range of 30 °C to 40 °C, which is close to physiological temperatures, and include, without limitation, poly(N-isopropylacrylamide) (PNIPAM) having LCST of about 32 °C, poly(N-vinylcaprolactam) (PNVCL) exhibiting LCST of about 31-38 °C, methylcellulose having LCST of about 50-60 °C, and/or poly(ethylene glycol)-b-poly(propylene glycol)-b-poly(ethylene glycol) (PEG-PPG-PEG) having LCST of about 20-85 °C (depending on composition).

It is noted that a skilled person in the relevant field will appreciate that the LCST of some polymers can be tuned by modifying the polymer structure, incorporating co-monomers, or changing the solution conditions (e.g., pH, salt concentration). In it further noted that some thermo-responsive polymers exhibit upper critical solution temperature (UCST) behavior instead of or in addition to LCST. This range covers most known thermo-contractive polymers, and the skilled artisan would be able to provide alternative polymers or modifications that could extend this range in either direction.

According to some embodiments of the present invention, the first polymer comprises a thermo-responsive polymer. According to some embodiments of the present invention, the thermo-responsive polymer is a thermo-contractive polymer, and in some preferred embodiments the thermo-contractive polymer exhibits an LCST in the range of 30-40 °C. In some embodiments the thermo-contractive polymer is selected from the group consisting of poly(N- isopropylacrylamide) (PNIPAM), poly(N- vinyl caprolactam) (PNVCL or PVCL), poly(methyl methacrylate) (PMMA), and any combination thereof. According to some preferred embodiments, the first polymer comprises or consists of PNIPAM.

One of ordinary skills in the art can slightly change PNIPAM's LCST by varying factors such as the molecular weight of the polymer and the ionic strength of the solution. This allows for precise control over its thermo-responsive behavior, making it suitable for specific applications.

The second crosslinked polymeric network:

According to some embodiments of the present invention, the second crosslinked polymeric network, made from the second polymer, is characterized by the property of stiffening or hardening irreversibly upon heating the polymer above a certain temperature. In some embodiments, the second crosslinked polymeric network hardens upon heating not as a result of covalent bonding or covalent crosslinking, but rather a substance that hardens due to irreversible entanglement of its polymeric chains is selected for the second polymer. This temperaturedependent setting property is bestowed to the IPN that comprises the second crosslinked polymeric network. This hardening process is analogous to the hardening of proteaceous substances as a result of heating, where heat causes the subsequent irreversible entanglement of protein chains without the formation of covalent crosslinks.

In some embodiments, the second polymer is a thermo-denaturing (a.k.a. thermodenaturing) polymer. Thermal denaturation describes the process where polymeric structures lose their native structure due to heat exposure, leading to unfolding and subsequent entanglement or aggregation of the polymeric chains. This process typically results in a firmer, more rigid structure without the formation of new covalent bonds. The hardening effect is primarily due to the rearrangement and interactions of the exposed hydrophobic regions of the unfolded polymer, leading to irreversible entanglement and hardening.

In the context of some embodiments of the present invention, the term “thermal denaturation” or “thermal denaturing”, refers to a property of a substance that undergoes an irreversible physico-mechanical transformation when exposed to heat, resulting in a rigid, intertwined molecular structure. In the context of some embodiments of the present invention, this process does not involve the formation of covalent bonds between polymer chains, but rather a state of irreversible entanglement creating a three-dimensional network that cannot be melted or reshaped once set. The initial state of the substance is typically malleable or liquid, but upon heating, it hardens permanently due to the extensive entanglement. The resulting material exhibits enhanced mechanical strength, thermal stability, and chemical resistance compared to its preheated state. This phenomenon is analogous to the denaturation of proteins, where heat causes the unfolding and subsequent irreversible entanglement of protein chains, altering their structure and function.

While reducing the present invention to practice, the inventors considered that PNIPAM does not have the requisite biological motifs to promote natural cell-matrix interactions. Therefore, in order to exploit the thermo-responsive properties of PNIPAM for tissue engineering, the second polymer was selected as a cell-friendly, natural polymer that can provide cells with the cues and biological motifs necessary for their maturation. Previous work with PNIPAM has shown that the use of nanogels can significantly impact both the kinetics and overall volumetric change of a macroscopic hydrogel. Therefore, the present inventors contemplated integrating the nanogels within a nanofibrous ECM-based hydrogel to create a hybrid hydrogel composition possessing both the dynamic, smart properties of PNIPAM and the native ECM’s biological motifs.

According to some embodiments of the present invention, the second polymer is a proteinbased polymer, namely a proteaceous substance, a polypeptide, a protein-based substance or a protein. In such embodiments, the irreversible hardening of the second crosslinked polymeric network is referred to as “thermal denaturation”. This process is also sometimes referred to as heat-induced denaturation, thermal coagulation, or heat-set gelation.

The term “collagen", as used herein, refers to a family of fibrous proteins that constitute the primary structural component of the extracellular matrix in various connective tissues of animals. More specifically, collagen is characterized as a protein composed primarily of three polypeptide chains, known as a-chains, which are wound together in a triple-helix configuration; each a-chain typically comprises a repeating amino acid sequence of (Gly-X-Y)n, where Gly represents glycine, and X and Y are often proline and hydroxyproline, respectively; a molecule that can self-assemble into supramolecular structures, including fibrils and networks, which contribute to the mechanical properties of tissues; a substance that exists in multiple types (e.g., Type I, II, III, IV, etc.), each with distinct molecular compositions and tissue distributions; a biomaterial capable of undergoing denaturation upon exposure to heat or certain chemical treatments, resulting in the unfolding of its triple-helix structure; a protein that can be extracted from animal tissues and subsequently processed into various forms, including solutions, gels, scaffolds, and powders, for use in medical, cosmetic, and industrial applications. In the context of the present invention, the term "collagen" encompasses all naturally occurring types of collagen, as well as any modified or synthetic forms that retain the essential triple -helical structure and/or the characteristic amino acid composition and/or the mechanical and thermal-denaturing properties of natural collagen.

According to some embodiments of the present invention, the second crosslinked polymeric network comprises a second polymer derived from a decellularized omentum extracellular matrix, or decellularized ECM.

The phrase “extracellular matrix” or “ECM” as used herein, refers to a complex network of materials produced and secreted by the cells of the tissue into the surrounding extracellular space and/or medium and which typically together with the cells of the tissue impart the tissue its mechanical and structural properties. Generally, the ECM includes fibrous elements (particularly collagen, elastin, and/or reticulin), cell adhesion polypeptides (e.g., fibronectin, laminin and/or adhesive glycoproteins), and space-filling molecules (usually glycosaminoglycans (GAG), proteoglycans).

The phrase “extracellular matrix-derived polymer”, or “ECM-derived polymer” as used herein, refer to a member of a family of biomacromolecules or their derivatives that are isolated, extracted, or synthesized based on components naturally found in the extracellular matrix of tissues. These polymeric substances originate from or mimic the structural and functional elements of the native tissue microenvironment and may include materials such as collagen and its various types, elastin, fibronectin, laminin, reticulin, proteoglycans, and glycosaminoglycans (e.g., hyaluronic acid). ECM-derived polymers can be obtained through processes including direct extraction from tissue sources, enzymatic or chemical breakdown of ECM components, decellularization of tissues, recombinant production of ECM proteins, and chemical synthesis based on ECM component structures. These materials can be further modified or processed to enhance their properties through methods such as crosslinking, chemical functionalization, and blending with other natural or synthetic polymers. Substances that retain bioactive properties similar to native ECM, potentially including cell adhesion motifs, growth factor binding sites, and enzymatic degradation sites, are also encompassed within this definition. ECM-derived polymers can be formulated into various forms, including hydrogels, fibrous scaffolds, porous sponges, coatings, and microparticles or nanoparticles. These polymers are often utilized in biomedical applications, tissue engineering, regenerative medicine, and drug delivery systems due to their biocompatibility, biodegradability, and ability to mimic the natural cellular microenvironment.

A person of ordinary skills in the art can find ample guidance and information about such substances. For example, International Patent Application No. W02009/085547 teaches the generation of decellularized omentum scaffolds for tissue engineering; U.S. Patent Publication No. 20050013870 teaches a scaffold comprising decellularized extracellular matrix of a number of body tissues including omentum; U.S. Patent Publication No. 20150202348 teaches decellularized omentum for tissue engineering; Porzionato et al. (Italian Journal of Anatomy and Embryology, 116, 2011 and Eur J Histochem. 2013, 24;57(l):e4. doi: 10.4081/ejh.2013.e4) teaches decellularized omentum; soluble forms of decellularized extracellular matrix are known in the art as described in Acta Biomaterialia, 9(8), 2013, pp. 7865-7873 and Singelyn et al., J Am Coll Cardiol., 21, 2012; 59(8): 751-763; and International Patent Application No. WO2017103930 teaches spherical particles comprising decellularized omentum being between 1 nm to 300 pm in diameter. The contents of all the above references is incorporated herein by reference.

As used herein the phrase “decellularized omentum” refers to the extracellular matrix which supports omentum tissue organization which has undergone a decellularization process (i.e., a removal of all cells from the tissue) and is thus devoid of cellular components. The decellularized omentum comprises extracellular matrix (ECM) components. Omentum may be harvested from mammalian species, such as human, swine, bovine, goat and the like. Following tissue harvesting, the tissue can be either placed in 0.9% saline for immediate processing or stored for later use, preferably at a temperature of about -20° C to about 80° C. Methods of decellularizing omentum may be found in WO2014/207744 and WO2014/037942, the contents of which are incorporated herein by reference.

According to some embodiments of the present invention, the second polymer comprises or consists of a ECM-derived polymer, which form the second crosslinked polymeric network in the form of an ECM-based hydrogel. In some embodiments, the ECM-derived polymer is derived from a decellularized omentum. According to a preferred embodiment, the omentum used to extract the second polymer is derived from a human.

According to some embodiments of the present invention, the second polymer exhibits thermo-denaturation when exposed to heat, namely heated above a certain temperature, referred herein as the denaturing temperature, which is similar in some aspects to the curing temperature of thermosetting substances.

In some embodiments of the present invention, the denaturing temperature of the second polymer, being a thermo-denaturing polymer, is higher than a lower critical solution temperature (LCST) of the first polymer, being a thermo-contracting polymer. According to some embodiments of the present invention, the denaturing temperature of the second polymer ranges 35 °C to 40 °C.

The temperature range for irreversible denaturation of collagen and its derivatives typically falls between 35 °C to 40 °C, though this can vary depending on specific conditions and collagen type. Type I collagen, the most common in mammals, generally denatures around 37 °C to 40 °C under physiological conditions. However, this range can be influenced by several factors. Collagens from different species may denature at varying temperatures, with those from cold- water organisms often having higher denaturation points. The hydration state of collagen, pH conditions, and degree of crosslinking can all affect the denaturation temperature. Collagen derivatives like gelatin usually have lower gelling temperatures, around 30 °C to 35 °C. The concentration of collagen, presence of other solutes, and duration of heat exposure also play roles in determining the exact denaturation point. While the 35 °C to 40 °C range is generally applicable for most native collagen types in physiological conditions, it is noted that specialized applications or modified collagens might exhibit slightly different thermal behaviors.

According to some embodiments of the present invention, the second polymer comprises a naturally occurring, processed and/or synthetic protein-based thermo-denaturing polymer selected from the group consisting of collagen, elastin, fibronectin, laminin, reticulin, and any combination thereof. Process of Preparing of the IPN:

The approach presented herein combines two distinct polymers in an IPN to create a superior biomaterial. The first polymer contributes thermo-responsive properties, and the second polymer adds mechanical strength and structural stability, biocompatibility and cell-supportive features. When these polymers are intricately interwoven in an IPN, they act synergistically to enhance the overall performance of the resulting hybrid hydrogel composition. This unique combination amplifies the individual contributions of each component while mitigating their limitations. Consequently, the hybrid hydrogel composition exhibits not only the capacity to form delicate constructs smaller than the printer’s resolution limits, but also improved mechanical properties, better cell adhesion and proliferation, and enhanced biocompatibility. These characteristics make the hybrid hydrogel composition provided herein particularly well-suited for a wide range of biomedical applications, including but not limited to tissue engineering, drug delivery systems, and regenerative medicine therapies.

According to some embodiments of the present invention, the first step of the process of preparing the IPN described hereinabove includes the provision of a plurality of hydrogel particles, or nanogels (nanometric- sized hydrogel particles).

In the context of some embodiments of the present invention, an emulsion polymerization technique for obtaining polymerized and/or crosslinked polymeric particles comprises dispersing a plurality of monomers, and/or oligomers (short polymers) and/or polymers, or a mixture thereof, in the medium of the internal phase of an emulsion, optionally together with an emulsion stabilizer (e.g., a surfactant), and optionally adding an initiator to the dispersion, thereby forming the solution that constitutes the dispersed (internal) phase of an emulsion. The emulsion is prepared by vigorously mixing the internal phase dispersion with the medium of the continuous (external) phase of the emulsion, optionally adding the initiator at this stage of the reaction, and activating the initiator (by heat, irradiation or otherwise) to initiate polymerization of the dispersion’s contents within the droplets of the internal phase, thereby forming polymeric particles templating in shape the droplets of the dispersed phase. The technique may further include adding a crosslinking agent to any of the emulsion phases to produce crosslinked polymeric particles. The technique allows some control of the size of the particles by adjusting the surfactant concentration or mixing rate, and purifying the resulting particles by filtering, washing, centrifugation, dialysis or other separation methods. The polymerization occurs primarily within the internal phase droplets, which act as nanoreactors, allowing for the production of polymer particles with controlled size and properties suitable for use in applications such as bioinks, drug delivery systems, or other specialized materials. The size of the hydrogel particles that comprises the first polymer, namely the discrete hydrogel particles constituting the first crosslinked polymeric network, is determined by the conditions of the process by which they are produced, namely the controllable polymeric composition and the controllable process parameters. In some embodiments, the hydrogel particles may be synthesized using an emulsion polymerization technique, wherein the first polymer is present in the dispersed phase of the emulsion.

In some embodiments of the present invention, the particle size and size distribution is a function of the process parameters employed during the fabrication process, e.g., in the emulsionbased manufacturing method. In some embodiment, with respect to the hybrid hydrogel composition, the particles in their swollen state exhibit a dimensional range that is at least one order of magnitude (10-times) greater than the diameter of the polymeric chains constituting the second polymeric network. This size differential between the swollen particles and the polymeric chain diameter of the second polymeric network is a feature that ensures the IPN will be formed, which is critical for achieving the desired properties and performance of the hybrid hydrogel composition. In some embodiments, the average size of the hydrogel particles ranges 10-1000 nanometers, or 1-100 nm, or 100-800 nm.

In some embodiments, the hybrid hydrogel composition (the IPN) provided herein includes particles or nanoparticles of a thermo-responsive polymer, such as, e.g., PNIPAM. The description below PNIPAM is used as an exemplary thermo-contractive first polymer, however, the below description should be taken as a general approach to synthesizing plurality of discrete hydrogel particles of the first crosslinked polymeric network using any hydrogel, including any suitable thermo-contractive polymer, not just PNIPAM. In the context of the present invention, PNIPAM is an example of a thermo-responsive polymer that exhibits a reversible phase transition in response to temperature changes. This property has made it a suitable material to construct the presently claimed 3D-printing bioink (hybrid hydrogel composition).

PNIPAM is an exemplary thermo-responsive polymer, which exhibits a distinct LCST at approximately 32 °C. The phase transition of PNIPAM is reversible - if the temperature is adjusted back below the LCST, the polymer will absorb water again, demonstrating its ability to switch between hydrophilic and hydrophobic states, or hydrated (swollen) to dehydrated states. This phase transition makes PNIPAM an excellent thermo-responsive material in the context of some embodiments of the present invention. Additionally, PNIPAM hydrogels are used for cell culture and tissue engineering, where they can act as cell scaffolds with tunable properties.

To manufacture the hybrid hydrogel composition with high contractive ratio, micro-and/or nanoparticles of the first polymer (e.g., PNIPAM) may be synthesized using an emulsion polymerization technique. A mixture of monomers (e.g., NIPAM), a crosslinker (e.g., bisacrylamide), and a initiator (e.g., thermally-activated radical initiator), is encapsulated within nanometric micelles and heated to activate the polymerization process. It is noted that other processes and methods for forming particulate hydrogels are also contemplated within the scope of the present invention.

The hydrogel particles (nanogels), are introduced to a non-crosslinked hydrogel of the second polymer in their hydrated (swollen) state, in order to facilitate the interpenetration of stands from the second polymer into the particles. When preparing the IPN provided herein, attention should be given to the optimal thermal, chemical and mechanical conditions for interpenetration to take place, without compromising the chemical crosslinking of the non-crosslinked hydrogel of the second polymer that follows interpenetration.

When using a thermo-responsive first polymer to produce the nanogels, the introduction is effected at a temperature below the LCST of the polymer, as well as in its hydrated state. This process requirement assures that the polymeric chains of the second polymer of the non- crosslinked hydrogel thereof, will interpenetrate the nanogels.

The chemical crosslinking of the non-crosslinked hydrogel of the second polymer is effected after the polymeric chains interpenetrated through the particles. In addition, the reagents, conditions and mechanism of the crosslinking chemical reaction of the second polymer should not interfere and be different than the contraction stimulus mechanism of the first crosslinked polymeric network, and the denaturing process mechanism of the second crosslinked polymeric network. Hence, the crosslinking reaction that locks the hydrogel particles of the first polymer on the polymeric strands of the second polymer by effecting a crosslinking reaction that is effected below the LCST of the first polymer and below the denaturing temperature of the second polymer, and further use reagents that are not reactive towards the polymers other than towards the functional groups in the second polymer, which participate in the crosslinking reaction.

When the swollen hydrogel particles of the first polymer are combined and thoroughly mixed with the non-crosslinked hydrogel of the second polymer, a specific crosslinking agent is introduced to the mixture. This agent, which may be added to the mixture before the addition of the second polymer thereto, or thereafter, is carefully selected to initiate the crosslinking process in the previously non-crosslinked hydrogel of the second polymer. Subsequently, a crosslinking reaction is initiated and carried out. By effecting the crosslinking reaction of the non-crosslinked hydrogel of the second polymer in the presence of the plurality of discrete hydrogel particles of the first polymer in their swelled (hydrated) form and at an optimal temperature, one can obtain an interpenetrating polymeric network, which is obtained when the non-crosslinked stands of the second polymer are being crosslinked after the hydrogel particles have been interpenetrated- through (threaded) by the non-crosslinked stands of the second polymer, thereby “locking” the threaded particles on the crosslinked polymeric chains.

When using a thermo-denaturing polymer for the second crosslinked polymeric network, the introduction of the hydrated hydrogel particles (first polymer) and the non-crosslinked hydrogel (second polymer) is effected at a temperature below the LCST of the first polymer and the denaturing temperature of the second polymer.

The process may further include a step of washing the IPN to remove any unreacted materials, to remove excess non-crosslinked polymers, and remove excess hybrid hydrogel particles forming the first crosslinked polymeric network that have not been interpenetrated by polymeric chains of the second polymer and locked in the second crosslinked polymeric network.

FIG. 1 A presents illustrations of various steps in the process of forming the hybrid hydrogel composition provided herein, showing left-to-right a plurality of discrete hydrogel particles in a swelled (hydrated) form (also referred to herein as swollen nanogels), followed by introduction of a crosslinking agent into the reaction mixture, which diffused throughout the solution and hydrogels, followed by the introduction of a non-crosslinked hydrogel of the second polymer, followed by interpenetration of chains of the second polymer into the swollen nanogels and the physical entanglement and chemical crosslinking of the second crosslinked polymeric network, leading to the formation of an IPN.

Bioinks:

The term "bioink", as used in the context of the present invention and in field of bioengineering, refers to a printable biological material capable of being seeded with or comprising living cells and/or biomolecules suspended in a carrier medium, wherein said bioink is capable of being extruded through a nozzle or otherwise deposited in a controlled manner to form three- dimensional tissue-like structures. The bioink may be seeded and/or comprise one or more types of living cells, extracellular matrix components, growth factors, and other biological molecules suspended in a hydrogel or other biocompatible material that provides structure and support. The bioink is formulated to maintain cell viability and function during and after the printing process and a seeding step, while also possessing rheological properties suitable for extrusion or deposition using 3D bioprinting techniques. Thus, the term “bioink” implies that the bioink composition exhibits cytocompatibility (viable cell-friendly) as well as printability (mechanical suitability for 3D printing).

In the context of some aspects of some embodiments of the present invention, the hybrid hydrogel composition provided herein, namely the heat-contracting and heat-setting IPN described herewith, is used as a component in a bioink. A bioink based on the hybrid hydrogel composition provided herein can be used to form any 3D objects or parts thereof by any 3D printing techniques at any size - but specifically objects or parts of an object that exhibit structural features characterized by size that is smaller than the highest resolution of the printing apparatus. Namely, the object or a part thereof is printed with the IPN-based bioink at or above the resolution limit of the printing apparatus, and thereafter heated as a step in the printing process, which causes the object or a part thereof to shrink to the final desired size.

It is noted herein that the use of the term “bioink” should not be seen as limiting the provisions of the present invention solely to bioengineering applications, thus the term “bioink” can be replaced with any term that refers to the substance that is being extruded from a printhead during printing, such as “ink”.

Thus, according to embodiments of the present invention, there is provided a bioink, comprising the hybrid hydrogel composition provided herein as a major component thereof. The bioink, which is formulated for 3D bioprinting, comprises the hybrid hydrogel composition as a biocompatible base material, which provides a structural framework. Cells may be incorporated to the bioink prior to the printing process, or incorporated into the printed construct post printing, becoming the functional elements of the printed construct (object). To enhance biocompatibility and cell viability, various additives like growth factors, cytokines, and extracellular matrix components may be added to this bioink. These ingredients work together to create a bioink that can be printed into complex 3D structures, mimicking the natural environment of tissues and organs for applications in regenerative medicine and drug discovery.

A bioink comprising the hybrid hydrogel composition provided herein as a major component thereof is also referred to herein as a “shrinking and setting bioink”, or the “first bioink” in the exemplary printing process described hereinbelow.

Printing 3D-objects:

In a typical 3D-printing process for bioengineering applications, a specialized bioink composed of biocompatible polymers, cells, and supportive growth factors is extruded layer by layer to construct a three-dimensional structure or object. The bioink is deposited through a fine nozzle in precise patterns, controlled by a computer-aided design (CAD) model, to create the desired geometry. During printing, the bioink's composition is carefully regulated to ensure it maintains a suitable viscosity and stability, allowing it to hold its shape while still providing a nurturing environment for the embedded cells. As each layer is printed, it adheres to the previous one, gradually or incrementally building up the object with intricate internal features or vascular networks. Crosslinking agents, UV light, or temperature changes are often employed during or after the printing to solidify the structure, transforming the soft bioink into a stable hydrogel that supports cell growth and tissue formation - this step is sometimes referred to as setting or curing. This approach enables the creation of complex, customized biological constructs, such as tissue scaffolds or organoids, designed to mimic natural tissue environments and support cell viability and differentiation for regenerative medicine or research purposes. As described hereinabove, the present invention pushes the boundaries of presently known 3D-bioprinting by allowing the formation of structural constructs that are smaller than the resolution limit of the printer, using a bioink based on the herein-provided hybrid hydrogel composition.

In general, the bioink based on the hybrid hydrogel composition provided herein, or the shrinking and setting bioink, can be used to print any 3D object - one which will arrive at its final dimensions once the applied bioink has been heated above the LCST and denaturing temperature characterizing the shrinking and setting bioink.

The shrinking and setting bioink is particularly useful in printing small structural features and constructs, characterized by structural features that cannot be achieved due to the resolution limitations imposed by the printing apparatus. The shrinking and setting bioink can also be used to print small hollow constructs, such as tubes and capillaries, mimicking an intricate blood-vessel structures, as demonstrated in the Example section that follows below.

A non-limiting exemplary printing process is described hereinbelow, demonstrating a nonlimiting embodiment of the present invention wherein a hollow capillary tube is printed using three types of bioinks in coordinated steps, using the shrinking and setting bioink provided herein, also referred to in this exemplary embodiment as the first bioink.

FIG. 2A-F presents a series of cross-section illustrations, each depicting a step in a bottom- up additive 3D-printing of a hollow object, e.g., a capillary, using three types of bioinks, smooth light grey denotes the third bioink, smooth dark grey denotes the first bioink, black denotes the second bioink, and grainy grey denotes a non-printed support medium, wherein FIGs. 2A-E present the printing steps of one or more bioinks, and FIG. 2F presents the final product after heating that caused the third bioink to melt and flow outside the lumen at a first temperature (not shown), before the first bioink shrunk at a second temperature (grainy light grey), following setting of the second and second bioinks at a third temperature (grainy black and ), whereas the first temperature is lower than the second temperature, which is lower than the third temperature.

In the initial step of the exemplary 3D-printing process of a hollow object (see, FIG. 2A), the bottom part of the outer wall of the object is printed using a second bioink comprising a noncrosslinked hydrogel composition that comprises a thermo-denaturing polymer, such as, e.g., a non-crosslinked decellularized omentum extracellular matrix hydrogel composition. This composition will stiffen upon heating to a temperature above the denaturing temperature of the second bioink, setting the final structure.

In the following step of the exemplary 3D-printing process of a hollow object, the lower- mid-section of the object is printed (see, FIG. 2B), using the second bioink to further build the outer wall and using a first bioink (the shrinking and setting bioink provided herein) to afford the inner wall of the object. In this step of the 3D-printing process, the pre-shrunk internal wall, or the pre- shrunk endoluminal surface of the capillary, is formed over the printed second bioink.

In the following step of the exemplary 3D-printing process of a hollow object (see, FIG. 2C), the upper-mid-section of the object is printed, using the second bioink to further build the outer wall, using the first bioink to further build the inner wall of the object, and using the third bioink to take the place of the lumen of the capillary.

In the following steps of the exemplary 3D-printing process of a hollow object (see, FIGs. 2D-E), the construction of the capillary is completed, and the construct is ready for the final gradual or incremental heating step.

In the final step of the exemplary 3D-printing process of a hollow object characterized by a lumen (see, FIG. 2F), a gradual or incremental heating step is effected, wherein the third bioink melts at a first temperature and flows away, essentially vacating the lumen of the capillary, or melting to a fluid state that can be pushed away (squeezed-out) by the shrinking walls of the lumen. Thereafter the construct is further heated to a second temperature, equal or above the LCST of the first bioink, higher than the first temperature and lower than the third temperature in which the construct hardens. At this temperature the inner lumen of the capillary shrinks to its final dimensions, pulling with it the outer layer of the capillary which does not shrink yet is still sufficiently pliable, thereby avoiding tears and rips in at or near the interface between the second and the second bioinks. Thereafter the construct is further heated to a third temperature, equal or above the denaturing temperature of the second and second bioinks, which is higher than the first and second temperatures, and effects hardening of the construct, setting it at its final inner and outer dimensions.

The printing process described herein can be modified to include more than three types of bioinks, each exhibiting a different composition, and different mechanical properties before and after the final heating step. For instance, more than one formulation of the first bioink can be controllably designed to exhibit different degrees of shrinking at similar or different LCSTs. This variety of shrinking and setting bioinks can be afforded by providing hybrid hydrogel compositions have a variety of formulations, e.g., varying the chemical composition of the polymers, and/or varying the mass ratio of the first crosslinked polymeric networks to the second crosslinked polymeric network, and/or a variation in the size of the discrete hydrogel particles, and/or the degree of crosslinking of either the first or the second crosslinked polymeric networks. The printing process can be modified to build a construct having portions that shrunk at a gradual degree of shrinkage, according to the variety of first bioink used.

According to some embodiments of the present invention, the method of 3D printing an object, includes printing the object using more than one formulation corresponding to the first bioink, as described herein, wherein each first bioink formulation may exhibit a different LCST and/or a different contraction propensity, and/or any other different mechanical and/or chemical property.

In a non-limiting embodiment of the present invention described below, the present inventors have printed vascularized, patient-specific, cardiac patches. The inventors sought to develop a new approach that would allow to print small-scale capillaries in between the larger blood vessels. To this end, a hybrid hydrogel composition, as disclosed herein, was used to afford a cell-containing bioink that significantly shrinks after printing and heating above 32 °C. It is noted that while the exemplary embodiment refers to specific ingredients for a specific embodiment, other polymers, reagents and conditions can be used under the same principles.

The inventors have demonstrated a hybrid hydrogel composition that is capable of shrinking to more than thirteen times its initial size. Using a 3D bio-printer, the presently disclosed bioink can be easily incorporated within an ECM-based hydrogel that provides support for parenchymal tissue, yielding a process for fabricating a thick functional tissue with an incorporated microvasculature. This technique enables to create the first 3D bio-printed native-sized capillaries, thereby allowing the generation of small-scale tissue units and other fine constructs at their desired final size.

The present inventors demonstrated that it is possible to combine the bioink based on the hybrid hydrogel composition with other hydrogels and bioinks to create multicomponent tissues and organs. In order to demonstrate this aspect of the smart bioink potential, individual polymeric chains of the hybrid hydrogel composition were printed within a bulk structure. The present inventors hypothesized that these hybrid constructs would be thermally triggerable, with the potential to shrink to form capillary-sized blood vessel-like structures within the bulk. For this part of the demonstration, three hydrogel-based bioinks were employed, all of which were printed within a non-thermo-responsive support medium: gelatin that acts as an internal support medium and referred to as the third bioink; the hybrid hydrogel composition comprising PNIPAM particles interpenetrated-through by polymeric chains of a ECM-based second crosslinked polymeric network and referred to as the first bioink; and a non-crosslinked decellularized omentum extracellular matrix hydrogel composition, referred to as the second bioink.

According to some embodiments of the present invention, the method of 3D printing an object, includes printing the object using more than one formulation corresponding to the second bioink, as described herein, wherein each second bioink formulation may exhibit a different denaturing (setting) temperature, and/or any other different mechanical and/or chemical property.

The printing process was completed as follows. First, the bulk of the printed structure was formed by printing the second bioink that acts as an external support medium and the external wall of the finished capillary. Next, the first bioink was printed on certain regions of the external support structure in order to form the inner walls of the blood vessel-like tubes that would be selectively shrunk. Next, features were printed on the certain regions of the printed first bioink using gelatin as the third bioink to form the hollow lumens of the blood vessel-like structures, acting as an internal support medium for the next layers. After printing the gelatin, additional layers of the first bioink and the second bioink were printed on top of the entire structure, enclosing the gelatin with the shrinking and setting bioink, or first bioink, and enclosing the first bioink with the second bioink, thereby printing the entire capillary.

FIG. 3A is a schematic illustration of an exemplary coordinated multi-kinetic process of 3D-printing of a blood vessel construct, wherein the bulk of the printed construct consisted of pristine ECM hydrogel (second bioink), in which the ECM-PNIPAM hybrid hydrogel composition (first bioink) was selectively localized around polymeric chains of gelatin (third bioink), which subsequently liquefy, leaving behind hollow lumens.

The third bioink is used as temporary structural internal support for the lumen of the final construct, filling the void/lumen of the capillary to be formed, and comprising or consisting of a viscoelastic thermo-reversible hydrogel composition. The third bioink comprises a printable substance formulation that transitions from a solid to a flowable liquid form when heated above its melting or gel- sol transition temperature. According to some embodiments, the melting or gelsol transition temperature of the third bioink is lower than or equal to the LCST of the first polymer, and lower than the denaturing temperature of the second polymer. In some embodiments, the melting temperature or the gel- sol transition temperature of the third bioink is lower than 25 °C to 35 °C.

Suitable substances that can serve as a major component of the third bioink include substances that can potentially be used as ink for 3D printing and characterized by a melting temperature or gel- sol transition temperature lower than 25 °C to 35 °C, include certain oils and waxes, such as paraffin, coconut oil, cocoa butter, some vegetable shortening formulations, some polymer blends specifically designed for low-temperature 3D printing, specific polyethylene glycol (PEG) grades, specific formulations of polycaprolactone (PCL), certain bio-based polymers such as gelatin, certain agarose formulations, some alginate -based materials, specific formulations of chitosan-based materials, some silicone formulations, specific formulations of polyvinyl alcohol (PVA), some low-melting point alloys (e.g., Gallium-Indium eutectic), some low-melting point ceramics, and certain metallic pastes or slurries. It is noted that the exact melting point and suitability for 3D printing can vary depending on specific formulations, additives, and processing conditions. Additionally, some of these materials may require special handling or post-processing techniques to achieve desired properties in the final printed object.

According to some embodiments of the present invention, the third bioink may comprise substances that exhibit a gel-to-sol transition in the temperature range of 20-40 °C, such as gelatin (gel-to-sol transition around 35 °C), methylcellulose (gel-to-sol transition upon cooling below its LCST, which can be tuned within the 20-40 °C range), pluronics (some poloxamers formulations transition from gel to sol within 20-40 °C), PNIPAM (LCST about 32 °C), hyaluronan grafted with poly(N-isopropylacrylamide-stat-N-tert-butylacrylamide) (sol-gel transition between 10-35 °C), agarose (low-melting point agarose can transition from gel to sol around 25-30 °C), carrageenan (some types, like kappa-carrageenan, can melt within 20-40 °C), pectin (certain low- methoxyl pectins can undergo gel-to-sol transitions within 20-40 °C), xyloglucan (some modified xyloglucans exhibit thermoreversible gelation within 20-40 °C), and chitosan-based hydrogels (certain formulations can be designed to have gel-to-sol transitions within 20-40 °C).

According to some embodiments of the present invention, the third bioink comprises a substance selected from the group consisting of gelatin, a poloxamer (Pluronic F-127, Poloxamer 407), PVA, and/or a carbomer resin (Carbopol).

In order to ensure proper functionality, the substance of the third bioink is used not as a bulk material but as a slurry of microparticles, rendering the third bioink flowable and quick to respond to changes in the environment, such as a change in temperature, owing to the significantly higher surface-area-to-volume ratio of microparticles.

According to some embodiments of the present invention, the method of 3D printing an object, includes printing the object using more than one formulation corresponding to the third bioink, as described herein, wherein each third bioink formulation may exhibit a different melting temperature, and/or any other different mechanical and/or chemical property.

In some embodiments, the third bioink comprises gelatin. Gelatin and collagen are closely related proteins, with gelatin being derived from collagen through a process of partial hydrolysis. Gelatin is a denatured form of collagen, afforded when collagen is heated in the presence of water - the wet heating unravels its triple-helix structure, and breaking it down into smaller protein fragments. This process, known as hydrolysis, results in the formation of gelatin. Gelatin retains many of the amino acids present in collagen but lacks its organized structure. As such, gelatin exhibits different properties from collagen, most notably its ability to form thermo-reversible gels in water. While collagen is insoluble in cold water and irreversibly denatures when heated, gelatin dissolves in warm water, forms and reforms a gel upon cooling. This property makes gelatin useful in various food, pharmaceutical, and bioengineering applications. Essentially, gelatin can be thought of as a partially broken-down form of collagen, sharing its basic composition but with altered structural and functional characteristics. When an aqueous gelatin solution is cooled, it forms a physical hydrogel through partial recovery of the collagen triple helix structure - this gel is thermo-reversible and will melt when heated above its melting point.

The melting point of gelatin gels typically falls within the range of 25 °C to 35 °C, though this can vary depending on several factors. The concentration of gelatin plays a significant role, with higher concentrations generally resulting in higher melting points. The Bloom strength of the gelatin, which indicates its gel strength, also influences the melting point, with stronger gels melting at higher temperatures. The pH of the solution affects gelling properties, with optimal gelling usually occurring between pH 4-7. The source of the gelatin (such as porcine, bovine, or fish) can lead to slight variations in melting point. Additionally, the presence of additives like salts or sugars, the molecular weight of the gelatin, proper hydration time, and the gel's thermal history all contribute to its final melting behavior. Notably, gelatin gels typically melt below human body temperature (37 °C), contributing to their characteristic "melt-in-the-mouth" property in food applications. The skilled artisan would appreciate that the precise melting point can be tailored for specific uses by adjusting these various factors, making gelatin a versatile ingredient in many industries.

As a proof of concept, the inventors have demonstrated a hybrid hydrogel composition that is capable of shrinking to more than thirteen times its initial size. Using a 3D bio-printer, the presently disclosed bioink containing the hybrid hydrogel composition, was incorporated within an ECM-based hydrogel that provides support for parenchymal tissue, yielding a process for fabricating a thick functional tissue with an incorporated microvasculature, having structural features that are smaller in at least one dimension than the printer’ s resolution limit. This technique enabled the fabrication the first 3D bio-printed native-sized capillaries, and generally allows the fabrication of small-scale tissue constructs at their correct size.

Following the method provided herein, the inventors synthesized nanometric particles of PNIPAM that displayed controlled and triggerable volumetric changes and combined them with a cell-friendly, nanofibrous, extracellular matrix (ECM)-based hydrogel to form a hybrid hydrogel composition as defined herein. Like the PNIPAM particles, the ECM-based hybrid hydrogel composition is thermo-responsive and thermo-denaturing (similar to thermosetting) - above its LCST, the polymeric chains become significantly more intertwined, causing an overall stiffening of the hybrid hydrogel composition. Furthermore, the inventors designed the hybrid hydrogel composition to exploit the kinetic variation between these two thermo-responsive processes. When triggered by heat, the nanogel particles de-swell (shrink), pulling together the ECM-based polymeric elements, thereby compacting the entire volume of the hybrid hydrogel composition. The denaturing process wherein the ECM-based polymer becomes more entangled is slower, however, this stage of the process occurs when the contraction has run its course. When these sequential processes, namely shrinking and setting, occur in the presence of different cell types, a complex tissue with multi-scale tissue units is generated.

As a proof of concept of some embodiments of the present invention, the inventors used this bioprinting strategy to construct fully functional cardiac tissues with a multi-scale vasculature. Tissues were printed from three different bioinks, each with a unique thermo-responsive property, and incorporated two cell types, both of which were differentiated from human induced pluripotent stem cells (iPSCs). First, a cardiomyocyte-laden ECM-based bioink, referred to herein as the second bioink, which slowly stiffens at physiological temperature, was used to form the parenchyma. Next, the microvasculature within the printed tissue was built up by printing two different types of blood vessels: static and shrinking. The static blood vessel was created by directly printing endothelial-laden, sacrificial, gelatin-based bioink that liquefies at physiological temperature and formed the open lumens of the blood vessels - this bioink is referred to herein as the third bioink. The shrinking blood vessels were fabricated by extruding this gelatin-based third bioink inside a shell printed using the hybrid hydrogel composition bioink provided herein and referred to as the first bioink. When incubated at 37 °C, the combined and sequential responses of herein-provided bioinks generated a functional cardiac tissue with large blood vessels and small-scale capillaries formed by the selective shrinking of the hybrid hydrogel composition. The inventors demonstrated that this process is completely safe for viable live cells (cell-friendly) and that the printed tissues function appropriately following their fabrication, generating fully contracting cardiac patches with perfusable blood vessels and a controlled network of capillaries.

Anisotropic Printing:

While reducing the present invention practice and attempting to print delicate structures, the present inventors have noticed that the shear stresses associated with extrusion-based printing techniques can cause polymer alignment within the extruded strips and strands, leading to structures that respond anisotropic to stimuli. This anisotropy results in differential material stiffness and strain when the properties of the shrinking and setting bioink is activated, e.g., heated. To demonstrate the anisotropic activation of the bioink, geometric shapes may be printed according to different patterns, such as longitudinally and transversally. Upon heating, the shapes may exhibit distinct shrinkage patterns based on the print design and direction. It was observed that the direction parallel to the printing pattern may shrink to a lesser extent compared to the perpendicular direction.

Further testing of the smart properties of the hybrid hydrogel composition bioink may involve printing circular structures. As the printer traces a circle, the polymeric elements in the hybrid hydrogel composition may align tangentially to the circle, exposing it to potential radial strain. Upon activation, the circles may constrict, with the inner and outer diameters changing at different rates and/or by different percentages. Interestingly, the thickness of the circular structure exhibited minimal directional variation during the transformation, which can be attributed to the equilibrium between opposing forces within the structure.

These observations suggest that the hybrid hydrogel composition (first bioink) can serve as the basis for printing volumetric structures, while considering the anisotropic shrinkage thereof. The internal strength of the IPN may be sufficient to overcome additional stresses that arise during dimensional changes, allowing for the engineering of precisely detailed structures that can undergo programmed transformations without compromising mechanical integrity. The shrinking and setting bioink provided herein demonstrates potential for various applications, including the creation of complex geometries and structures that can change shape or size in response to specific stimuli. This technology may be particularly useful in fields such as tissue engineering, where precise control over structure and material properties is crucial.

FIG. 4A-F illustrate some of the features of the anisotropic printing aspect of the present invention, and some results obtained from printing 3D objects using a bioink that includes the hybrid hydrogel composition comprising ECM and PNIPAM, whereas FIG. 4A illustrated the process schematics demonstrating the anisotropic swelling/deswelling behavior of printed structures wherein a (longitudinal and transverse) represents the swelling strains within each individual shrinking strand, and r represents the stress applied to the samples during shrinkage; FIG. 4B shows pentagon shapes that were printed either in the (I) longitudinal or (II) transverse direction, and undergone the anisotropic shrinkage that occurred primarily perpendicular to the printed strands (scale bars 2 mm); FIG. 4C presents the normalized shrinking strain for the printed pentagons (data are presented as mean ± s.e.m, n = 6); FIG. 4D illustrates the process of radial compaction of a printed ring; FIG. 4E is an image showing the macroscopic shrinkage of radially printed, anisotropic circles using ECM-PNIPAM hybrid hydrogel composition (scale bars 5 mm); FIG. 4F presents the measurements of the inner and outer diameter of the printed rings at 20 °C and 37 °C (data are presented as mean ± s.e.m, n = 6).

In some embodiments of the present invention, there is provided a method for anisotropic printing of 3D objects, utilizing the hybrid hydrogel composition. For example, a bioink comprising the hybrid hydrogel composition provided herein, can be used to print structures with directional properties, meaning they behave differently along different axes.

The features of an anisotropic printing process include temperature-responsive behavior, where printed structures exhibit swelling and deswelling properties in response to temperature changes. Additionally, the process is characterized by directional shrinkage, with printed structures primarily shrinking perpendicular to the direction of the printhead motion (the direction of the printed strands). Controllable deformation is another process feature, as manipulating the printing direction (longitudinal or transverse) allows for control over the resulting shrinkage patterns. The process also enables radial compaction when printing circular structures, affecting both inner and outer diameters. Finally, the shrinkage and deformation of the printed structures can be measured and quantified, allowing for precise control and prediction of the final shape.

This anisotropic printing process, according to an aspect of some embodiments of the present invention, enables the creation of complex, three-dimensional structures with programmable shape-changing capabilities, which could have applications in tissue engineering, drug delivery, and other biomedical fields.

Definitions:

As used herein the term “about” or “approximately,” refers to ±10 %. For example, the term “about 100 units” encompasses the value 100 units, as well as the values 90 units, 91 units, 92 units, 93 units, 94 units, 95 units, 96 units, 97 units, 98 units, 98 units, 99 units, 100 units, 101 units, 102 units, 103 units, 104 units, 105 units, 106 units, 107 units, 108 units, 109 units, and 110 units.

The terms "comprises", "comprising", "includes", "including", “having” and their conjugates mean "including but not limited to"; namely, as used herein, these terms are intended to be open-ended and not limiting. They indicate that the presence of the listed elements does not preclude the inclusion of additional, unrecited elements or method steps.

The term “consisting of’ means “including and limited to”.

The term "consisting essentially of" means that the composition, method or structure may include additional ingredients, steps and/or parts, but only if the additional ingredients, steps and/or parts do not materially alter the basic and novel characteristics of the claimed composition, method or structure.

The phrase “one or more” as used herein includes one, two, three, or more of the described elements or components and does not exclude any combinations or sub-combinations thereof.

The terms “preferred” or “preferably” indicate an example or embodiment that is more suitable or favorable under certain circumstances, but these terms are not intended to limit the scope of the invention or to suggest that other variations are excluded.

As used herein, the phrase “selected from the group consisting of’ includes all members of the recited group, each member of the recited group, and all possible combinations. For example, selected from the group consisting of A, B, and C, includes A, only, as well as B, only, as well as C, only, as well as A and B, as well as A and C, as well as B and C, and as well as A, B, and C.

The term “substantially,” when used in reference to a characteristic or parameter, means that the characteristic or parameter need not be absolute but is close enough to the specified value or condition so as to achieve the intended purpose or effect.

As used herein, the phrases "substantially devoid of" and/or "essentially devoid of" in the context of a certain substance, refer to a composition that is totally devoid of this substance or includes less than about 5, 1, 0.5 or 0.1 percent of the substance by total weight or volume of the composition. Alternatively, the phrases "substantially devoid of" and/or "essentially devoid of" in the context of a process, a method, a property or a characteristic, refer to a process, a composition, a structure or an article that is totally devoid of a certain process/method step, or a certain property or a certain characteristic, or a process/method wherein the certain process/method step is effected at less than about 5, 1, 0.5 or 0.1 percent compared to a given standard process/method, or property or a characteristic characterized by less than about 5, 1, 0.5 or 0.1 percent of the property or characteristic, compared to a given standard. Further alternatively, the terms "substantially" and/or "essentially " in the context of a characterizing property, means that the characterizing property is expressed to at least 99 %, at least 95 %, at least 90 % of its full or complete expression. For example, the phrase “the particles are maintained substantially in a non-contacting, spaced-apart arrangement” should be read as “at least 99 % of the particles are maintained in a non-contacting, spaced-apart arrangement”.

As used herein, the phrases "substantially devoid of" and/or "essentially devoid of" in the context of a certain substance, refer to a composition that is totally devoid of this substance or includes less than about 5, 1, 0.5 or 0.1 percent of the substance by total weight or volume of the composition. Alternatively, the phrases "substantially devoid of" and/or "essentially devoid of" in the context of a process, a method, a property or a characteristic, refer to a process, a composition, a structure or an article that is totally devoid of a certain process/method step, or a certain property or a certain characteristic, or a process/method wherein the certain process/method step is effected at less than about 5, 1, 0.5 or 0.1 percent compared to a given standard process/method, or property or a characteristic characterized by less than about 5, 1, 0.5 or 0.1 percent of the property or characteristic, compared to a given standard.

When applied to an original property, or a desired property, or an afforded property of an object or a composition, the term “substantially maintaining”, as used herein, means that the property has not change by more than 20 %, 10 % or more than 5 % in the processed object or composition.

The term “exemplary” is used herein to mean “serving as an example, instance or illustration”. Any embodiment described as “exemplary” is not necessarily to be construed as preferred or advantageous over other embodiments and/or to exclude the incorporation of features from other embodiments.

The words “optionally” or “alternatively” are used herein to mean “is provided in some embodiments and not provided in other embodiments”. Any particular embodiment of the invention may include a plurality of “optional” features unless such features conflict.

As used herein, the singular form "a", "an" and "the" include plural references unless the context clearly dictates otherwise. For example, the term "a compound" or "at least one compound" may include a plurality of compounds, including mixtures thereof.

Throughout this application, various embodiments of this invention may be presented in a range format. It should be understood that the description in range format is merely for convenience and brevity and should not be construed as an inflexible limitation on the scope of the invention. Accordingly, the description of a range should be considered to have specifically disclosed all the possible subranges as well as individual numerical values within that range. For example, description of a range such as from 1 to 6 should be considered to have specifically disclosed subranges such as from 1 to 3, from 1 to 4, from 1 to 5, from 2 to 4, from 2 to 6, from 3 to 6 etc., as well as individual numbers within that range, for example, 1, 2, 3, 4, 5, and 6. This applies regardless of the breadth of the range.

Whenever a numerical range is indicated herein, it is meant to include any cited numeral (fractional or integral) within the indicated range. The phrases “ranging/ranges between” a first indicate number and a second indicate number and “ranging/ranges from” a first indicate number “to” a second indicate number are used herein interchangeably and are meant to include the first and second indicated numbers and all the fractional and integral numerals therebetween. As used herein the terms “process” and "method" refer to manners, means, techniques and procedures for accomplishing a given task including, but not limited to, those manners, means, techniques and procedures either known to, or readily developed from known manners, means, techniques and procedures by practitioners of the chemical, material, mechanical, computational and digital arts.

Terms used in the singular form shall also include the plural, and vice versa, unless context clearly indicates otherwise. Furthermore, words of any gender include all genders and are intended to cover all corresponding terms.

Unless otherwise defined, all technical and/or scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which the invention pertains. Although methods and materials similar or equivalent to those described herein can be used in the practice or testing of embodiments of the invention, exemplary methods and/or materials are described below. In case of conflict, the patent specification, including definitions, will control. In addition, the materials, methods, and examples are illustrative only and are not intended to be necessarily limiting.

EXAMPLES

Reference is now made to the following examples, which together with the above descriptions, illustrate some embodiments of the invention in a non-limiting fashion.

The present invention encompasses a hybrid hydrogel composition featuring two crosslinked polymeric networks forming an interpenetrating polymer network: a first crosslinked polymeric network containing thermo-responsive polymers, such as poly(N-isopropylacrylamide) (PNIPAM) and a second crosslinked polymer network containing thermo-denaturing polymers, such as collagen or ECM-derived proteaceous polymers. In the IPN, the first polymeric network is in a form of a plurality of nanoparticles of the crosslinked first polymer that are securely confined within the second polymeric network. The hybrid hydrogel composition can comprise collagen- based or extracellular matrix-based hydrogels, resulting in an interpenetrating polymeric network with the nanoparticles. Below provided methods of crafting a non-limiting example of a hybrid hydrogel composition, which entails combining swelled nanoparticles, non-crosslinked collagen- based or ECM-based hydrogels, and crosslinking agents. The hybrid hydrogel composition has been used as a bioink in a 3D printing process that started at a temperature below the LCST of the thermo-responsive polymer, enabling the creation of high-resolution objects with structural dimensions below 50 pm, including artificial tissues housing capillary blood vessels. In essence, the Example section is a non-limiting proof of concept of the invention provide herein, using an exemplary hybrid hydrogel composition, its fabrication process, and its utility in 3D printing finely detailed structures, particularly those with minute capillaries.

EXAMPLE 1

Synthesis and Characterization of PNIPAM Nanoparticles

Poly(NIPAM) nanoparticles, also referred to herein as PNIPAM nanogels, were synthesized by emulsion free-radical copolymerization in water (see, FIG. 1A) as previously published [Zhang, J. T. et al., “Poly(N -isopropylacrylamide) nanoparticle-incorporated PNIPAAm hydrogels with fast shrinking kinetics, Macromol. Rapid Commun., 2005, 26, 1346— 1350], with several modifications. During the reaction, sodium dodecyl sulfate (SDS, Fisher Scientific, Massachusetts, USA) surfactant was first dissolved in water to form micelles. Thereafter, N-isopropylacrylamide (NIP AM, Sigma-Aldrich, Darmstadt, Germany) monomers, N,N-methylenebisacrylamide (BIS, Sigma-Aldrich) co-monomers and the thermal initiator 2,2- Azobis[2,-(2-Imidazolin-2yl)Propnae] Dihydrochloride (VA-044, Fisher Scientific) were added to the mixture under nitrogen inert environment, while heating to 70 °C for 40 minutes to initiate the radical attack on the methacrylate groups of the monomers and co-monomers. The molar ratios of the reagents were calibrated as 1:0.007:0.1:0.11 NIP AM:SDS:BIS: VA-044. Purification of the product was performed by washing it three times with double distilled water (DDW) at 70 °C immediately after the reaction and dialysis (10,000 MWCO) against 5L of DDW at 4 °C with daily medium changing for 4 days. The polymer was lyophilized, then aliquoted and stored under sealed conditions at -80 °C.

The PNIPAM nanoparticles were characterized using TEM and FTIR.

The size, shape, and uniformity of PNIPAM nanoparticles were studies and measured using transmission electron microscopy (TEM) (JEM- 1400 Plus, JEOL, Pleasanton, CA, USA). Samples were prepared by dropping 5 pL of 0.42 %^IN PNIPAM aqueous solution on a copper grid and then left to dry at RT for 24 hours before imaging. Contrast images were acquired in TEM and captured using SIS Megaview III and iTEM (Olympus, Shinjuku, Tokyo, Japan).

FIGs. 5A-C present the results of the fabrication process of PNIPAM nanogels, wherein FIG. 5A is a TEM micrograph image showing the morphology of the synthesized PNIPAM nanogels (scale bar 1 pm); FIG. 5B are photographs of the he macroscopic appearance of the PNIPAM nanogels at 20 °C (I) and 37 °C (II) (scale bars 5 mm), demonstrating the contraction of the crosslinked hydrogel particles during the heating process accompanied by the excretion of water, and FIG. 5C presenting the results of rheological measurements of the PNIPAM hydrogels showing a significant increase in the complex viscosity after heating (data are presented as mean ± s.e.m, n = 3).

The TEM micrographs of PNIPAM nanoparticles revealed that what appears macroscopically to be a large, uniform PNIPAM hydrogel is, in fact, a collection of PNIPAM nanogels with a mean diameter of 546+18 nm (FIG. 5A). These nanoparticles, or PNIPAM nanogels, constitute an exemplary embodiment of a first crosslinked polymer network.

Having established their successful synthesis, the PNIPAM nanoparticles were hydrated to form a swollen hydrogel, and the stimuli-responsiveness properties of the hydrogel were investigated. First, the hydrogel’s volumetric change was assessed. As expected, the PNIPAM hydrogel shrank as it transitioned above its ECST, excreting water molecules that had been part of the larger, initial hydrogel (FIG. 5B). Alongside the volumetric change, the mechanical properties of the hydrogel were assessed. To that end, the rheological properties of the hydrogel were measured both above and below its ECST. As expected, the collapse of the PNIPAM chains above their LCST led to a significant increase in the viscosity of the hydrogels (FIG. 5C).

The polymerization of NIP AM to PNIPAM was also verified by FTIR, and the results were compared to the FTIR spectrum of the monomer. Chemical analysis of the PNIPAM nanoparticles compared to the NIP AM monomer was performed by using FTIR spectrophotometer (Nicolet iS 10 Mid Infrared FT-IR Spectrometer, Thermo Scientific, Waltham, Massachusetts, USA). The dry powder of both the monomer and the polymer were dissolved with DDW on top of real crystal KBr IR sample cards (Sigma- Aldrich) and dried at RT for 24 hours. The IR absorbance spectrum of the PNIPAM compared to NIP AM monomer was measured between 0-4000 cm ¹. The most significant chemical change that occurs during the polymerization reaction is the saturation of the olefin moieties present in NIP AM. The stretching of the C-H bonds of the monomer’s acrylic hydrogen is known to occur at 3074 cm ¹. The absence of this peak in the obtained spectrum indicated the successful synthesis of PNIPAM.

EXAMPLE 2

Preparation ofECM-Based Hydrogel

In developing the initial proof of concept, the inventors leveraged their extensive knowledge and expertise in extracellular matrix (ECM) and omentum-derived materials. After careful consideration, they chose to utilize omental tissue as the source for the second polymer component. This decision was made, inter alia, due to the tissue's unique properties and potential synergies with the poly(N-isopropylacrylamide) (PNIPAM) nano gels. The combination of these two elements - the PNIPAM nanogels and the omentum-derived polymer - forms the foundation of the hybrid hydrogel composition provided herein. This innovative approach, as outlined in some embodiments of the present invention, aims to harness the benefits of both components, synergistically enhancing the overall performance and applicability of the resulting material in various biomedical applications.

The fabrication of an omentum-ECM hydrogel has been reported and demonstrated its potential to serve as a bioink for 3D printing tissues [Noor, N. et al. 3D Printing of Personalized Thick and Perfusable Cardiac Patches and Hearts. Advanced Science 6, 1900344 (2019); Silberman, E., et al. Post-Maturation Reinforcement of 3D-Printed Vascularized Cardiac Tissues. Advanced Materials (2023); Shevach, M. et al. Omentum ECM-based hydrogel as a platform for cardiac cell delivery, 2015, doi:10.1088/1748-6041/10/3/034106; Soffer-Tsur, N. et al., Optimizing the biofabrication process of omentum-based scaffolds for engineering autologous tissues, 2014, doi:10.1088/1758-5082/6/3/035023; Edri, R. et al. Personalized Hydrogels for Engineering Diverse Fully Autologous Tissue Implants. Advanced Materials 31, 1803895 (2019)]. This ECM hydrogel has biological motifs for blood vessel formation and for the assembly of parenchymal tissues. The present inventors utilized this hydrogel as the bulk material for investigating the impact of incorporating the PNIPAM nanogels therewith.

Porcine omental tissue was obtained and decellularized as previously described elsewhere [Shevach, M., Soffer-Tsur, N., Fleischer, S., Shapira, A. & Dvir, T. Fabrication of omentum-based matrix for engineering vascularized cardiac tissues. Biofabrication 6, (2014); Soffer-Tsur, N., Shevach, M., Shapira, A., Peer, D. & Dvir, T. Optimizing the biofabrication process of omentumbased scaffolds for engineering autologous tissues. Biofabrication 6, (2014)].

Briefly, the omental tissues (Kibbutz Lahav, Israel) were washed with phosphate-buffered saline (PBS), large blood vessels were excised from the tissue, and the tissue was transferred to a hypotonic buffer (10 mm trisaminomethane (Hy- Laboratories), 5 mm ethylenediaminetetraacetic acid (Sigma-Aldrich), 1 pm phenylmethanesulfonylfluoride (Sigma-Aldrich) at pH = 8). After 1 hour at room temperature, the tissue underwent three cycles of freeze-thawing (-80 °C) in the hypotonic buffer. The tissue was then washed for 30 min each with 70 % (v/v) ethanol (Bio-Lab, North Carolina, USA), 96 % (v/v) denatured ethanol (Bio-Lab), and three times with acetone (BioLab). Thereafter, the tissue was soaked for 24 hours in a 40 % (v/v) solution of acetone in n- hexane (Bio-Lab) with three solution changes. The next day, the tissue was washed for 30 min with 96 % ethanol, then incubated overnight at 4 °C in 70 % ethanol. The tissue was then washed four times with PBS and incubated overnight in a 0.25 % solution of trypsin and EDTA (Sartorious Israel LTD, Beit Ha'emek, Israel), followed by four more washings with PBS and incubation for 24 hours in a 1.5 m NaCl (Bio-Lab) solution with three solution changes. The tissue was then washed in a solution of 50 mm trisaminomethane (pH = 8) and 1 % Triton-XlOO (Sigma- Aldrich) for 1 hour. The now decellularized extracellular matrix (dECM) was washed thoroughly with PBS and double distilled water (DDW) and frozen (-20 °C).

The dECM was lyophilized (dried), then ground into flakes with a Wiley Mini-Mill (Thomas Scientific, NJ, USA) and dissolved in a 0.1 M HC1 solution to a level of 1.67 % (w/v). The dECM was then processed enzymatically via the addition of 1 mg porcine pepsin (Sigma- Aldrich) per 10 mg dECM, which was left stirring at room temperature until no large collagen chunks were discernable (3-4 days). Following enzymatic digestion, the solution’s pH was adjusted to 7.4 by the addition of 5 M NaOH. PBS was added at a 10X concentration to reach a final working concentration of IX. This also reduced the final concentration of dECM to 1.5 % (w/v). The entire solution was filtered through a sterile, 70 pm, nylon cell filter (Bar Naor LTD, Israel) before use, and 0.1% Pen/Strep (Sartorious Israel LTD) was added. In order to fabricate 1% (w/v) dECM omentum hydrogel that will be suitable for the incorporation of iPSCs derived cardiac cells within it, the dECM was dissolved in a 0.1 M HC1 solution to a level of 1.11% (w/v), processed enzymatically and titrated until pH was adjusted to 7.4 as previously described. Dried Dulbecco’s modified Eagle medium (DMEM/F12) (Sartorious Israel LTD) was added at a 10X concentration in DDW to reach a final working concentration of IX. This also reduced the final concentration of dECM to 1% (w/v).

EXAMPLE 3

Fabrication of ECM-PNIPAM Hybrid Hydrogel Composition

The innovative approach outlined in this invention harnesses the benefits of two distinct polymers combined in an IPN to create a superior hybrid hydrogel composition. The inventors explored various combinations of the first polymer (e.g., PNIPAM) and the second polymer (e.g., ECM-based materials like collagen) before arriving at the optimal configuration. Initial studies using linear, non-crosslinked PNIPAM or physically entrapping PNIPAM nanogels within noncrosslinked collagen resulted in limited shrinkage and phase separation. To overcome these challenges and maximize volumetric change while retaining the thermo-responsive behavior of each of the polymers, the inventors developed a method to chemically crosslink the nanogels and the ECM-based hydrogel into an IPN. In this configuration, collagen polymers are interwoven with PNIPAM nanoparticles without covalent connections between them. The IPN was created by allowing pre-crosslinked collagen nanofibers to interpenetrate the crosslinked PNIPAM nanoparticles and then crosslinking the collagen, effectively locking the components in place. This approach prevents phase separation and particle aggregation while synergistically enhancing the overall performance and applicability of the resulting material in various biomedical applications.

As the PNIPAM nanoparticles were already crosslinked during their synthesis (see above), the present approach to create the IPN was a two-step process to generate a crosslinked, nanofibrous, hybrid hydrogel composition (see, FIG. 1A). First, PNIPAM nanogels were mixed with a cold, dilute solution of ECM hydrogel. Careful temperature control was necessary to maximally swell the nanogels and to allow the abundant nanofibers of the ECM to move in and through the nanoparticles. After time had elapsed to allow the ECM polymers to interpenetrate the PNIPAM nanogels, EDC/NHS chemistry was used to crosslink the collagen of the ECM hydrogel. This method was selected as EDC/NHS chemistry creates a covalent bond between carboxylic acids and amines, and none of these functional groups are present on PNIPAM, which ensures that the crosslink is specific to the second polymer (collagen). In order to underscore the importance of the temperature control for forming an IPN, heated, hydrophobic, shrunken nanogels were introduced into the ECM solution, which did not allow for the collagen nanofibers to interpenetrate into the particles. As hypothesized, this set of conditions led to the formation of a heterogeneous mixture of PNIPAM and ECM that could be observed. Taken together, these experiments show that the hybrid hydrogel composition requires that both the PNIPAM and the collagen be crosslinked and that the two materials be interwoven as an IPN to achieve optimal functionality.

In contrast to the work of Ding, C. et al., the instantly provided hybrid hydrogel comprises pre-formed nanoparticles of a thermo-responsive polymer, referred to herein as nanogels, whereas Ding, C. et al. uses linear polymeric chains of the thermo-responsive polymer. In addition, instantly provided hybrid hydrogel comprises ECM-derived polymer and not a commercial polymer, which allows to derive it from an autologous source and minimize immune rejection. Thirdly, due to the use of pre-formed nanogel particles, the magnitude of the thermo-responsive phenomenon is much greater, namely the percentages of deswelling are significantly higher than the same exhibited by Ding, C. et al., which uses linear thermo-responsive polymer. Fourthly, deswelling of the Ding, C. et al. hydrogel is effected at non-biocompatible conditions, such as pH of 3, which prevents the incorporation of live cells.

EXAMPLE 4

Characterization of ECM-PNIPAM Hybrid Hydrogel Composition

Using the strategy of combining nanogels and ECM in an IPN, a variety of hybrid hydrogel compositions were synthesized by varying the relative amounts of ECM hydrogel and PNIPAM nanogels. In particular, four concentrations of ECM:PNIPAM were assessed by varying the PNIPAM nanogel content from 20 % by mass to 50 % by mass of the final mass of the composition. As the ECM component is essential for tissue function and maturation, the present inventors chose not to exceed 50 % mass of the first polymer in the hybrid hydrogel composition.

These four samples were examined qualitatively by acquiring scanning electron microscope (SEM) images and quantitatively via the assessment of their rheological properties. Qualitatively, it was observed that the effect of having either too little or too much PNIPAM in the hybrid hydrogel led to a similar outcome. In either case, the PNIPAM nanogels tended to aggregate and form large, amorphous spheres. When using an intermediate concentration of PNIPAM (40 % by mass) did the nanogels refrain from aggregating.

To examine the interaction between the nanofibers and the nanoparticles in the 60:40 hybrid material, the inventors used a Focused Ion Beam - Scanning Electron Microscope (FIB- SEM) technique. In this technique, a beam of ions was used to etch away the surface of a material, exposing interior cross-sections and enabling the full thickness of a sample to be visualized using high resolution SEM. While the similar electronic properties of collagen and PNIPAM make it impossible to distinguish between these two materials within the same sample, the images taken clearly show the formation of an IPN. In particular, it was observed that the collagen nanofibers do not surround PNIPAM nanoparticles, but rather interpenetrate into and through the nanogels|

The substantially even distribution of the nanogels was also reflected in the macroscopic contraction of the hybrid hydrogel, which clearly demonstrated the expulsion of water molecules that occurred above the PNIPAM nanogels’ LCST.

FIGs. 1B-J present the results of the IPN fabrication process, as can be seen in FIGs. 1B- C) presenting high-resolution SEM images showing the nanometric morphology of the 60:40 ECM:PNIPAM hybrid network (scale bar 5 pm and 1 pm respectively), and FIG. ID which is a FIB-SEM image showing the interior of the hybrid hydrogel composition (scale bar 500 nm), where individual collagen nanofibers can be seen interpenetrating into and through the crosslinked PNIPAN nanoparticles (nanogels), while FIG. IE shows a single crosslinked PNIPAM nanoparticle within the IPN (scale bar 200 nm); FIG. IF is a photograph showing the macroscopic appearance of the ECM-PNIPAM hybrid hydrogel composition at 20 °C, and FIG. 1G shown the macroscopic appearance of the IPN at 37 °C after the IPN expels water upon heating (scale bar 5 mm); FIG. 1H presents a comparative plot of the kinetic measurements of the increase in the complex modulus of hybrid hydrogel compositions with varying concentrations of ECM:PNIPAM (first-to-second polymer mass ratios) via heating to 37 °C (data are presented as mean, n= 3); FIG. II are two photographs showing the flexibility and viscoelasticity of the 60:40 ECM:PNIPAM hybrid hydrogel composition at 37 °C (scale bar 10 mm) and FIG. 1 J is a representative microscopy image of live/dead staining of iPSCs-ECs encapsulated in the ECM-PNIPAM hybrid hydrogel composition (live cells appear brighter than the background) (scale bar 50 pm).

Quantitatively, the impact of the aggregates was measured via rheology. While the incorporation of 40 % by mass nanogels led to a small decrease in viscosity compared to the ECM- based hydrogel alone, the inclusion of either a greater or lesser nanogel concentration led to a significant decrease (see, FIG. 1H). Likewise, the hybrid hydrogel compositions with either too much or not enough PNIPAM nanogels were less responsive to changes in temperature. While the storage modulus of the pristine ECM-based hydrogel reached 825.5 ± 43.0 Pa after heating, hybrid hydrogel compositions comprising 20 %, 30 %, or 50 % by mass PNIPAM reached a maximum strength of less than 400 Pa. The hybrid hydrogel composition with 40 % by mass PNIPAM, on the other hand, attained a maximum complex modulus of 758.0 ± 33.3 Pa. Alongside the temperature-dependent changes in the gels’ moduli, all hydrogels demonstrated shear-thinning properties, which indicated that all of them can be used as bioinks in 3D printing.

Alongside the rheological assessment, the inventors performed biological analyses of the hybrid materials to assess their compatibility with tissue engineering applications. In order to assess the biocompatibility of the hybrid hydrogels, induced pluripotent stem cell (iPSC)-derived cardiomyocytes were encapsulated within the varying concentrations of gel. These cells were selected as they are sensitive and non-proliferative. Therefore, if the hybrid hydrogel compositions were to cause their death, the remaining cells would not be able to compensate for their loss and the cell death would remain obvious and measurable. Cardiomyocyte-filled hydrogels were cultured for two weeks during which time all concentrations showed high cell viability and contractile functionality.

Because the biological function of the hybrid hydrogel composition was found unimpacted by the change in mass ratio, a 40:60 composition of PNIPAM nanogels to ECM-based hydrogel was found preferable in terms of mechanical properties and nanoparticle distribution. This formulation was further characterized in the Example section that follows below. It is noted that the optimal mass ratio of first-to-second crosslinked hydrogels should not be seen as limiting the invention to any particular ratio.

Based on these results, the present inventors decided to proceed with the 60:40 composition of ECM-based hydrogel to PNIPAM nanogels, and extensively characterized this hybrid hydrogel composition formulation. As the aim of this study was to develop a shrinking and setting bioink that would shrink at physiological temperature, this aspect of the hybrid hydrogel composition’s behavior was carefully studied. Based on SEM images, it was confirmed that the hybrid hydrogel composition is indeed a combination of two nanometric components: nanogels of PNIPAM and nanofibers of ECM. As expected, the PNIPAM nanogels tended to shrink from a diameter of 867.3 ± 26.8 nm at temperatures below the LCST to 496.5 ± 9.8 nm above the LCST, a change of 42.7 %, as observed by SEM. Meanwhile, the average diameter of the ECM nanofibers remained essentially unchanged.

Beyond the volumetric response of the hybrid hydrogel composition to temperature, several other properties are of particular importance in tissue engineering. As many organs, the heart foremost among these, are in a constant state of dynamic stress, clinically relevant materials must demonstrate high degrees of elasticity and high resistance to shear stress. One of the benefits of working with IPN hydrogels is that they are known for their enhanced viscoelasticity. As shown, the present inventors observed that the hybrid hydrogel composition acted as a viscoelastic material and retained its ductility at physiological temperature, whereas the pristine PNIPAM hydrogel became brittle.

Another property of hydrogels in tissue engineering is how easily cells can manipulate their microenvironment. Cells are constantly re-modelling their ECM environment by enzymatically degrading the existing matrix and producing a new matrix. To ensure that cells embedded within the hybrid hydrogel would be able to perform these necessary functions, samples of the hybrid hydrogel composition were exposed to collagenase. Over the course of two weeks, these enzymes were able to degrade the entire network, which shows that the addition of the synthetic PNIPAM component was sufficiently limited so as not to abolish the beneficial aspects of the ECM-based hydrogel.

The capacity of the hybrid hydrogel composition to support cell growth in 3D was assessed. Common cell sources in tissue engineering are primary cells and differentiated induced pluripotent stem cells (iPSCs), and the inventors have chosen to assess both cell types with the herein-provided hybrid hydrogel. Over the course of ten days, both primary human umbilical vein endothelial cells (HUVECs) and iPSC-derived endothelial cells (iPSC-ECs), placed in the exemplary hybrid hydrogel composition of PNIPAM/ECM at 40:60 mass ratio, showed high viability and characteristic elongation (see, FIG. 1H). Additionally, a significant increase in cellular metabolism was recorded over the course of the study. The ability of iPSCs to proliferate and to differentiate into cardiomyocytes in an embryonic-like pathway, within a 3D hybrid hydrogel composition was also investigated. The high cellular coverage of the iPSCs within the 3D hybrid hydrogel composition consequently allowed successful cardiac differentiation, resulting in mature contracting tissues (results not shown). EXAMPLE 5

4D-Printing Shrinking and Setting Bioink

Having successfully created a cell-friendly, shrinking and setting bioink based on the hybrid hydrogel composition provided herein, the present inventors used the bioink for 3D printing an organ-mimicking construct having viable cells incorporated therein - whereas this type of 3D- printing is referred to herein as 4D printing due to the shrinking and setting bioink’s capacity to undergo a notable and controllable volume change post-printing. Specifically, the present inventors incorporated the cellular building blocks necessary to form functional tissue with the hybrid hydrogel.

To demonstrate the utility of the shrinking and setting bioink, the present inventors chose to engineer cardiac tissue with an ab initio vasculature. In order to print functional tissue with its inherent vasculature, two mature cell types were required: the cardiomyocytes (CMs) that formed the parenchyma and endothelial cells (ECs) to line the blood vessels. Both cell types were differentiated from human induced pluripotent stem cells (iPSCs). iPSCs were carefully maintained according to standard protocols and were regularly assessed via flow cytometry and immunohistology. To be used for these studies, cells showed high levels of expression of Oct4, a nuclear pluripotency marker, and Ki67, a marker of proliferation. Once differentiated, CMs were assessed for expression of a-cardiac sarcomeric actinin and cardiac troponin. Endothelial cells populations were enriched by MACS and the level of CD31and CD144 expression was assessed.

The present inventors next sought to optimize cell concentration for achieving efficient EC monolayer formation after shrinkage. Therefore, both HUVECs and iPSC-ECs were seeded onto the hybrid hydrogel composition prior to shrinking, and the cell confluence post-shrinking was experimentally determined. Initially, four concentrations of primary endothelial cells (pre-stained with cytostain) were seeded on the shrinking and setting bioink, and the samples were imaged before and after shrinkage was triggered. Additionally, images were taken after a week of cultivation to assess the extent of endothelial cell growth and their morphology (raw data not shown). As was observed in the cell viability experiments, the shrinkage of the hybrid hydrogel led to a significant increase in cell density, which in turn resulted in contact inhibition. As a result, while the overall extent of the endothelial-cell coverage increased for all cell concentrations, the largest increase was observed for cells seeded at only 2xl0⁴ cells/mm². The experiment was then repeated with iPSC-ECs. Since these cells are smaller, cells were seeded at an initial concentration of 3xl0⁴ cells/mm², which amounted to 4.0 ± 0.2 % of the bioink sample’s area. However, after one-hour incubation at 37 °C, the contraction of the bioink sample’s area resulted in an increase in cell density such that 64.6 ± 3.2 % of the bioink sample’s area was now covered by cells. After seven days of incubation, the cells continued their spreading, reaching 99.2 ± 0.2 % confluency with a classic elongated morphology and high levels of CD31 expression (data not shown).

Finally, the differentiated cardiomyocytes (CMs) and endothelial cells (ECs) were encapsulated in their respective bioinks, and the tissue-mimicking constructs were printed using essentially the method described above. A simplified cardiac patch was designed for this experiment, incorporating a single blood vessel with two branches: one dynamic and one static (see, FIGs. 3A-G and FIG. 6 A-N).

FIGs. 3B-G present the results 4D-prining of a simplified cardiac patch, wherein FIG. 3B are hrSEM images showing the continuous interface between the pristine ECM and the ECM- PNIPAM hybrid after heating to 37 °C (scale bars = 1 pm), FIG. 3C shows a demonstration of the perfusion through the printed construct, wherein dynamic blood vessels are marked with arrows (scale bar = 5 mm), and FIG. 3D shows the perfusing of fluorescent dye through the dynamic vessels before and after triggering their shrinkage (scale bar = 200 pm); FIG. 3E presents quantification of dynamic blood vessel shrinkage (data are presented as mean ± s.e.m, n = 10); FIG. 3F is a hrSEM image showing a selectively shrunk dynamic blood vessel-like channel (white arrow) alongside a static blood vessel-like channel (blue arrow)(scale bar = 100 pm), and FOG. 3G presents quantification of the diameter of both dynamic and static blood vessel-like printed structures at 20 °C and 37 °C (data are presented as mean ± s.e.m, n = 10).

To assess the dynamics of the coordinated sequential multi-kinetic 4D printing, the printed tissues were investigated at three different time points. First, immediately after printing, samples were fixed and stained for appropriate cellular markers, CD31 for the endothelial cells and a- cardiac sarcomeric actinin for the CMs.

059097832sFIGs. 6A-N present various aspects of a coordinated 4D bioprinting of vascularized tissues comprising selective fast- shrinking capillaries, according to some embodiments of the present invention, wherein FIG. 6A presents an illustration of some process steps demonstrating the sequential changes to the printed vascularized cardiac patch, FIG. 6B presents an image of immunostained cardiac patch immediately after printing (scale bar = 200 pm); FIG. 6C presents an image of the immunostained cardiac patch after three hours of incubation (scale bar = 50 pm); FIG. 6D presents an image of a static blood vessel after two weeks of maturation (scale bar = 10 pm); FIGs. 6E-F present images of a dynamic blood vessel after two weeks of maturation (scale bar = 10 pm); FIG. 6G presents an hrSEM image showing a dynamic blood vessel within the vascularized cardiac patch after two weeks of cultivation (scale bar = 10 pm); FIG. 6H is a quantification of intact red blood cells following perfusion shows no significant reduction in cell count even after 20 minutes of perfusion through fully cellularized capillaries (data are presented as mean ± s.e.m, n = 5) , FIG. 61 presents a fluorescent image taken after 20 minutes of perfusion of whole blood through acellular constructs shows significant thrombin activation (scale bar = 10 pm); FIG. 6J presents an image of a fully cellularized blood vessels showing minimal thrombin activation even after 20 minutes of perfusion (scale bar = 10 pm); FIG. 6K is a quantification of thrombin activity as evidenced by the fluorescent signal of a Anorogenic substrate (data are presented as mean ± s.e.m, n = 10) , FIG. 6L presents an image of a cellularized patch, immediately after being implanted on a rat’s omentum, FIG. 6M presents an image of the tissue taken one week after implantation, where the high level of anastomosis can be visualized even macroscopically; and FIG. 6N presents immunostaining of an extracted patch showing clear anastomosis of the human endothelialized blood vessels to the rat’s vasculature (scale bar = 100 pm).

The immediate printing process did not allow time for the bioinks' thermo-responsive behaviors to manifest, resulting in endothelial cells filling the blood vessels. A thick layer of the acellular hybrid hydrogel composition was observed surrounding the dynamic branch of the blood vessel. After three hours of incubation, additional samples were fixed and stained to visualize the hydrogels' thermo-responsive behavior. At this stage, lumens were visible as the gelatin had evacuated the blood vessels' interior, although cell proliferation and elongation had not yet begun. The activated hybrid hydrogel was no longer visible between the endothelial cells and cardiomyocytes. While the static branch of the blood vessel, printed with pristine bio-ink, maintained its diameter and lumen structure, the branch printed with hybrid bio-ink had shrunk (see FIG. 6C). Samples were cultured for an additional two weeks, allowing cells to grow and self-organize into tissue before fixation and staining. At this point, the lumens displayed high shape-fidelity, and endothelial cells showed high levels of CD31 expression in both the blood vessel printed with pristine hydrogel and the small-scale blood vessel printed with shrinkable hydrogel (see FIGs. 6D-F). Moreover, high-resolution scanning electron microscopy (hrSEM) images of sliced epoxy-based resin fixed tissues revealed a diameter of less than 5 pm (see FIG. 6G).

The creation of a fully endothelialized layer on the luminal surface of blood vessels, particularly capillaries, is crucial for various vascular properties. Therefore, hemocompatibility studies were conducted to assess this aspect. The endothelial layer's smoothness is vital for red blood cell transport, as these cells tend to rupture when exposed to rough, non-endothelialized surfaces. Moreover, blood exposure to any surface other than endothelial cells quickly leads to coagulation. Both of these aspects were evaluated in the engineered capillaries. Initially, blood was perfused through the entire blood vessel network, with red blood cell counts measured before and after (see FIG. 6H). In an acellular control, the red blood cell concentration decreased by 25.6 ± 6.5 %, attributed to cell rupture upon encountering the rough, non-endothelialized surface. Conversely, blood perfused through fully endothelialized tissue showed no significant change in red blood cell concentration.

The blood vessels' thrombogenicity was then assessed by monitoring thrombin activity evolution. Blood was supplemented with a Anorogenic thrombin substrate (N-T-Boc-Val-Pro- Arg-AMC), which emits a Auorescent signal when enzymatically cleaved by thrombin. Thrombin was chosen as the thrombogenesis indicator because it is activated by both extrinsic and intrinsic coagulation pathways. As with the hemolysis assessment, an acellular control was used for comparison. After 20 minutes of perfusion, the Auorescent signal in blood contacting acellular channels increased almost threefold, while no increase was observed in cellularized blood vessels (see FIGs. 6I-K).

Beyond in vitro experiments, in vivo assessments were performed to determine the engineered blood vessels' functionality. Engineered blood-vessel networks were implanted in a murine model (rat) (see, FIG. 6L), and their integration was evaluated after one week. Macroscopically, a high level of integration and anastomosis was evident (see, FIG. 6M). Postextraction, tissue samples were fixed, sliced, and stained. A generic antibody was used to mark all endothelial cells (murine and human) via CD31 expression, along with an antibody to stain only human nuclei. This allowed differentiation between host and implanted tissue blood vessels. As shown, the host's vasculature, stained only for CD31, had anastomosed with the printed human blood vessels, which were co-stained for CD31 and human nuclei (see FIG. 6N).

Lastly, the cardiac tissues' morphology and functionality were evaluated through immunostaining and calcium transient visualization via video (not shown). Results revealed that the printed iPSC-derived cardiomyocytes were elongated and aligned with striated sarcomeres (see, FIG. 6H). The integration of cardiomyocytes with the printed blood vessels was evident from the smooth passage of calcium transients throughout the vascularized tissue (see, FIGs. 6LK). High-magnification videos of the cardiac tissue surrounding a dynamic blood vessel demonstrated the calcium transient movement across it (see, FIGs. 6L-N). This observation indicates that the cardiomyocytes established three-dimensional intercellular communication, allowing action potentials to transmit around the interference presented by a hollow blood vessel.

Following printing, the structures were incubated at 37 °C to stimulate (trigger) each bioink's volumetric and/or viscoelastic transformation. To successfully integrate all three bioinks, the kinetics of the volumetric and/or viscoelastic transformation were manipulated to ensure each response occurred in the appropriate order. The sequence of events, namely the gradual and/or sequential increase of temperature, according to some embodiments of the present invention, is as follows:

First, upon heating to less than the LCST of the first bioink and at least to its melting temperature, the third bioink (gelatin) undergoes a significant reduction in stiffness and elastic modulus. Next, upon heating to at least its LCST, the first bioink (the shrinking and setting bioink comprising the hybrid hydrogel composition), no longer experiencing resistance from the third bioink, constricts and contracts, thereby squeezing the gelatin out of the lumen. Finally, the second bioink (ECM-based hydrogel), initially viscous enough to be dragged and pulled along by the first bioink's shrinkage, undergoes physical entanglement, creating the final, mechanically robust printed tissue by setting/curing.

If this order of heat-steps is not properly followed, shrinkage of the first bioink would either be ineffective due to resistance provided by the gelatin, or lead to tearing in the construct (tissue) as the first bioink separates from the bulk of the second bioink (ECM-based hydrogel).

In order to ensure proper functionality, gelatin was printed not as a bulk material but as a slurry of microparticles, such that the gelatin quickly equilibrated to the changing temperature. Additionally, the nanometric size of the PNIPAM nanogels ensured that the shrinking of the hybrid hydrogel composition occurred before the collagen nanofibers of the second polymer became more entangled with one another. Because the collagen nanofibers were present in both the bulk, pristine, ECM-based second bioink and the first bioink, a seamless interface was ultimately created of interlocking nanofibers from both bioinks (see, FIG. 3B).

The final product of this printing process was a thick, fully perfusable structure (see, FIG. 3C). In order to demonstrate the selective shrinking of the first bioink, constructs were printed with both dynamic blood vessel-like structures that incorporated the first bioink and static blood vessel-like structures in which gelatin was printed directly into the second bioink without the intermediary layer of the first bioink. Following heating, the dynamic blood vessel-like structures that had been printed with the first bioink experienced significant shrinkage. The average diameter of these vessels decreased from 252 ± 7 pm to 20.0 ± 0.8 pm (see, FIG. 3D and FIG. 3E). Meanwhile, the vessels printed without the layer of the first bioink showed no dimensional changes during heating. The final products, therefore, incorporated both large-scale arteriole-like structures, for which the diameter was >200 pm, and micrometric capillary-like structures, for which the diameter was < 20 pm (see, FIG. 3F).

Anisotropic printing aspects were also observed and monitored during this 4D-printing experiment. As previously discussed, shear stresses associated with extrusion-based printing methods cause polymer alignment within printed strands, leading to anisotropic structures. This anisotropy results in differential material stiffness in perpendicular and parallel directions to the printed strand, causing differential strain upon activation of the shrinking and setting bioink properties (see, FIG. 4A). To demonstrate anisotropic activation, elongated pentagons were printed in two patterns: longitudinal and transversal. Initially nearly identical, each pentagon shrank distinctly based on its print design (see, FIG. 4B). The direction parallel to the printing pattern shrank less than 10 %, while the perpendicular direction shrank over 75 % (see, FIG. 4C).

To further assess the hybrid hydrogel composition's properties in the shrinking and setting bioink, a circle was printed. As the printer traced the circle, nanofibers of the hybrid hydrogel aligned tangentially, exposing the circle to potential radial strain (see, FIG. 4D). Circles were printed with an inner diameter of 9.65 ± 0.05 mm and an outer diameter of 13.12 ± 0.13 mm (see, FIG. 4E). After heating, the circles constricted to final dimensions of 1.86 ± 0.16 mm (inner diameter) and 5.15 ± 0.31 mm (outer diameter), representing changes of 81 % and 61 %, respectively (see, FIG. 4F).

Interestingly, the circle's thickness showed minimal variation during transformation, shrinking only 5 % from 3.47 ± 0.12 mm to 3.29 ± 0.26 mm. This observation is attributed to the printed structure's bulky nature. As the inner radius constricts, polymers oppose the motion, dragging the receding area toward the center of mass due to changes in the ring's bending moment. Equilibrium between these forces results in a more compact overall structure while maintaining total thickness.

These results suggest the shrinking and setting bioink can serve as a basis for printing volumetric structures. A potential concern when working with such smart bioinks is that preprogrammed behavior might cause undue stress, leading to mechanical failure. However, this experiment demonstrated that the IPN's internal strength is sufficient to overcome additional stresses arising during heat-induced shrinkage. Consequently, it is shown herein that it is possible to engineer precisely detailed organ units at native size by first extruding volumetric structures and then shrinking them without compromising mechanical integrity.

In conclusion, the present inventors have demonstrated a novel biocompatible, stimulus- responsive hybrid hydrogel composition for use as a bioink in tissue engineering. The hybrid hydrogel composition comprises both synthetic nanoparticles and natural ECM nanofibers, which synergistically create a hybrid with both the supportive microenvironment cells needed to develop properly and the augmented, 4D behavior of smart materials. The present inventors have demonstrated that this hybrid hydrogel composition can serve as a bioink for the 3D bioprinting of functional, vascularized cardiac tissue. As a result of the smart bioink’s unique 4D behavior, the present inventors have demonstrated the first successful printing of a cell-lined capillary-sized blood vessel.

EXAMPLE 6

Materials and Methods

Crosslinked PNIPAM nanoparticles were either physically entrapped to the collagen fibers of the ECM-based hydrogel or chemically crosslinked to them by carbodiimide activation using EDC/NHS chemistry, to create a hybrid thermo-responsive hydrogel composition. For the physical encapsulation of PNIPAM nanoparticles within the ECM-based hydrogel, the crosslinked PNIPAM nanoparticles were mixed at 40 % of the hybrid hydrogel composition's final mass, at the pH values of 5.7, 6.0, 6.3 and 6.7. Hybrid hydrogel composition samples at RT were subjected to physiological temperature of 37 °C, and were macroscopically imaged after 30 minutes, 24, and 48 hours in order to follow the deswelling process. Furthermore, the strains in XY plane of each sample were calculated. For the chemical conjugation of the PNIPAM nanoparticles within the ECM-based hydrogel, carbodiimide chemistry was adopted, with modifications.

Briefly, lyophilized PNIPAM nanoparticles (1 equivalent) were dissolved to 1 % (w/v) in 0.1 M MES buffer (Sigma- Aldrich), pH 6.0. Then, the activator ethyl-3-[3- dimethylaminopropyl]carbodiimide hydrochloride (EDAC) (Thermo Fisher Scientific) (4 equivalents) was added, and N-hydroxysulfosuccinimide (Sulfo-NHS) (Sigma-Aldrich) (6 equivalents) was added as well to stabilize the reactive EDAC intermediate against competitive hydrolysis, thereby achieving a high efficiency of chemical crosslinking. The solution was mixed at 4 °C for 20 minutes to create activators laden PNIPAM nanogels. Then, 1.5 % ECM-based hydrogel was added at different omentum-dECM/PNIPAM percentage mass ratios (80:20, 70:30, 60:40 and 50:50, respectively). The collagen fibers within the ECM-based hydrogel were diffused through the PNIPAM nanogels, while carboxylic acid and amine terminals of it were conjugated via amide bond during the EDC/NHS carbodiimide activation for 2 hours of reaction, so that PNIPAM nanoparticles were chemically entrapped within the crosslinked collagen fibers. Purification of the hybrid hydrogel composition was performed by dialysis (10,000 MWCO) against 5L of XI PBS at 4 °C with daily medium changing for 4 days. The hybrid was dried in a lyophilizer, then suspended in DDW at 4 °C to achieve a final concentration of 1.5 % (w/v) of ECM.

Directional relative shrinking strains measurements were conducted as follows. In order to evaluate the shrinking strains of the physically encapsulated ECM-PNIPAM hybrid hydrogel composition in each direction at the XY plain, three samples of the exemplary hybrid hydrogel composition, 500 pL each, were measured for each condition. The samples were imaged using a binocular microscope (SMZ18, Nikon), at 20 °C, and during the incubation period of 24 and 48 hours at 37 °C. The length in each direction were measured using ImageJ (FUI), and the directional relative shrinking strains were calculated using the following formulae:

Wherein siength, I and are the calculated strain value and the sample length at 20 °C and 37 °C in the y axis, respectively. Swidth, and w are the calculated strain value and the sample length at 20 °C and 37 °C in the x axis, respectively.

Enzymatic degradation test with collagenase was also conducted. The exemplary ECM- PNIPAM hybrid hydrogel composition samples were prepared by casting 500 pL into disposable base molds (Leica Biosystems, IL, USA), and incubating for 45 minutes at 37 °C. Thereafter, the samples were immersed in a solution of 1 U -mL”¹ collagenase Type II (Worthington, Lakewood, NJ, USA). The samples mass was measured using analytical scale every 48 hours between day 0- 12, and the normalized mass compared to the initial mass on day zero was calculated. The collagenase solution was refreshed every 48 hours.

For the high-resolution scanning electron microscopy (hrSEM) imaging, samples of exemplary hybrid hydrogel composition were fixed with 2.5 % glutaraldehyde 2 hours at room temperature followed by a graded incubation series in ethanol-water solutions (50-100 % v/v). All samples were dried using critical point drying (Balzers), sputter-coated with gold (Polaron E 5100, Quorum Technologies, Lewis, UK) and observed using a Gemini 300 hrSEM (Zeiss, Germany). For the small-scale capillaries imaging via hrSEM, structure preservation was conducted by epoxy-based resin infiltration technique, and sliced with ultra-microtome (UC7 ultra-microtome, Leica Biosystems) prior the imaging. Briefly, samples were infiltrated by a graded soaking incubation series of increasing concentration of Epon resin (Sigma Aldrich) in absolute ethanol, at RT in a sealed container, using these ratios: 1:3 (2 x 3 hours), 1:2 (2 x 3 hours), 1:1 (overnight), 1:2 (2 x 3 hours), 2:1 (2 x 3 hours), 3:1 (2 x 3 hours). Resin-ethanol mixtures were replaced with freshly 100% resin and samples were infiltrated at RT overnight. Excess of resin removal was conducted by rinsing the samples in absolute ethanol prior to polymerization in oven at 70 °C overnight. For the focused ion beam- scanning electron microscopy (FIB-SEM), samples of the exemplary hybrid hydrogel composition were fixed with 2.5 % glutaraldehyde 2 hours at RT followed by a graded incubation series in ethanol-water solutions (50-100 % v/v). All samples were dried using critical point drying (Balzers), sputter-coated with gold (Polaron E 5100, Quorum Technologies, Lewis, UK) and observed using a Dual Beam HELIOS 5 UC (Thermo Fisher Scientific).

Rheological measurements (n = 3) were performed using a Discovery HR-3 hybrid Rheometer (TA Instruments, DE, USA) with 8 mm diameter parallel plate geometry and a Peltier plate to maintain the sample temperature. PNIPAM samples were loaded at a temperature of 20 °C, and their complex modulus was measured by performing a frequency sweep between 0.1 and 100 rad/s at a constant 1 % strain at both 20 °C and 37 °C. ECM-PNIPAM hybrid hydrogel composition samples were loaded at a temperature of 20 °C and their storage, loss and complex modulus was measured by performing a frequency sweep between 0.1 and 1 rad/s at a constant 1 % strain at both 20 °C and 37 °C. In order to evaluate the kinetics gelation of the different concentrations of the ECM-PNIPAM hybrid hydrogel composition samples, the samples were loaded at a temperature of 20 °C, and their complex modulus was measured by performing a time sweep at a constant angular frequency of 1 Hz rad/s and 1 % strain for 2500 seconds, at a temperature range of 20-37 °C. iPSC cultures were generated from omental stromal cells and were a kind gift from Dr. Rivka Ofir from Ben Gurion University. The undifferentiated cells were cultivated on 10-cm culture plates pre-coated with Matrigel (BD, Franklin Lakes, NJ, USA) diluted to 250 pg mL"¹ in DMEM/F12 (Sartorious Israel LTD). Cells were maintained in NutriStem (Sartorious Israel LTD) medium containing 0.1 % penicillin/streptomycin (Sigma- Aldrich) and cultured under a humidified atmosphere at 37 °C with 5 % CO2. Medium was refreshed daily, and cells were passaged at 80 % confluence by treatment with ReLeSR (Stemcell Technologies, Vancouver, Canada).

CM Differentiation from iPSCs - prior to differentiation, cells were dissociated with Accutase (StemCell Technologies) and passaged to 6-well plates coated with Matrigel as before. NutriStem (Sartorious Israel LTD) was refreshed daily until iPSCs reached 100 % confluence. At that point (Day 0), medium was changed to RPMI (3 mL) (Sartorious Israel LTD), supplemented with 0.5 % 1-glutamine (Sartorious Israel LTD), B27-Insulin (Gibco, NY, USA), and 5.0 pm CHIR-99021 (Tocris, Bristol, UK). On Day 2, the medium was changed to RPMI (3 mL) supplemented with 0.5 % 1-glutamine, B27-Insulin, and 5 pm IWP-2 (Tocris). On Day 4, the medium was changed to RPMI (3 mL) supplemented with 0.5 % L- glutamine and B27-Insulin, and this medium was refreshed on Day 6. On Day 8 the medium was changed to RPMI (2.5 mL) supplemented with 0.5 % 1-glutamine and B27 (Gibco). On Day 10 the medium was changed to RPMI-glucose (Sartorious Israel LTD) (2.5 mL) supplemented with 0.5 % 1-glutamine and B27 (starvation medium) and this medium was refreshed on Day 12. On Day 14, medium was changed to RPMI (2.5 mL) supplemented with 0.5 % 1-glutamine and B27. On day 16 medium was changed to M-199 (Gibco), supplemented with 0.1 % penicillin/streptomycin, 5 % fetal bovine serum (FBS, Sartorious Israel LTD), 0.6 mm CuSO4, 0.5 mmZnS04, and 1.5 mm vitamin B 12 (Sigma- Aldrich). This medium was refreshed every other day.

Primary human umbilical vein endothelial cells (HUVECs) were purchased commercially (Angio-Proteomie, MA, USA), and maintained in Endothelial Growth Medium (EGM-2) (Lonza, Basel, Switzerland) supplemented with an additional 1.5 % (v/v) FBS. The medium was refreshed every other day.

For the study of EC differentiation from iPSCs, cells were differentiated as previously described. Briefly, human iPSCs were dissociated on day 0 with Accutase (STEMCELL Technologies) and replated on MatrigelTM (BD), diluted to 50 pg mL^-1 in DMEM/F12 (Sartorious Israel LTD), coated plates. Cells were seeded at a density of 47,000 cells/cm² and maintained in NutriStemTM (Sartorious Israel LTD) medium containing 1 %

Penicillin/Streptomycin (Sartorious Israel LTD) and 10 pM Y-27632 ROCK inhibitor (Tocris). On day 1, the medium was replaced with mesoderm induction medium containing a 1:1 (v/v) mix of Neurobasal (Gibco) and DMEM/F12 supplemented with L-Glu (Sartorious Israel LTD), N2 (Gibco) and B27-retinoic acid (Gibco) with 25 ng mL-1 BMP4 (PeproTech, NJ, USA) and 8 pM CHIR99021 (Tocris). The media was not changed for 3 days to induce a mesoderm state. On day 4, the medium was changed to EC induction medium consisting of StemPro-34 SFM medium (Gibco) supplemented with 200 ng mL-1 VEGF165 (PeproTech) and 2 pM forskolin (PeproTech). The EC induction medium was changed daily. On day 7, the cells were dissociated with Accutase and magnetic-activated cell sorting (MACS) was used to separate for CD31+ CD144+ cells. The sorting was performed using a manual MACS magnetic separator and magnetic beads conjugated antibodies (Miltenyi Biotech, MD, USA). The CD31+/CD144+ cells were seeded onto cell culture treated flasks and cultured in EGM-2 supplemented with 20 pM SB431542 (PeproTech). Media was replaced every other day. When the cells reached about 90 % confluency, they were either passaged using 0.25 % Trypsin-EDTA solution or cryopreserved.

Support medium was prepared according to previously reported protocols. Briefly, a solution of sodium alginate, xanthan gum, and sodium chloride was prepared with uniformly distributed calcium carbonate. To this solution, gluconic acid <5-lactone was added, the solution was mixed thoroughly, and the entire mixture was allowed to sit overnight. After 24 hours, the mixture was dissolved in DDW and the entire contents were homogenized. This homogenized stock solution was then set aside at 4 °C. Prior to use, the stock solution was centrifuged at 15 800 g for 20 minutes and the supernatant was removed. The pellet was washed three times by resuspension in DMEM (Sartorious Israel LTD) with an addition of 20 mM HEPES (Gibco). The final pellet formed the working support medium.

Gelatin microparticles were prepared heating a solution of 0.9 % Gelatin Type B 225 Bloom (Sigma- Aldrich) with 0.1 % carboxymethylcellulose sodium salt (Sigma- Aldrich) in DDW to 60 °C while stirring. After 2 hours, the temperature was lowered to 45 °C. A solution of 1 % acetic acid was added dropwise with stirring until reaching the clouding point, and the solution was then allowed to continue stirring for 15 minutes. The solution was then placed in an ice bath and stirring continued for another 15 minutes. An excess of acetone was then added, and the solution was stirred for 15 more minutes. The solution was then centrifuged at 4 °C and 3000 g for 15 minutes. The pellet was washed via resuspension in PBS and another centrifugation as before. The wash was repeated, and the final pellet was kept at 4 °C until its use.

For the perfusion experiments, samples were printed directly into a custom plastic chamber containing support medium, which was designed using opensource computer-aided design software and printed with a Max X DLP 3D printer (Asiga; Sydney, Australia). The walls of the chamber contained a small hole matching the diameter of a 27G needle that lined up precisely with the printed lumen, so that a needle could be easily inserted directly into the blood vessel. To image the perfusion, fluorescent yellow-green carboxylate-modified polystyrene latex beads (Sigma- Aldrich) were dissolved in 4.5 % Bovine Serum Albumin solution (BSA, MP Biomedicals, Santa- Ana, California, USA) to achieve a final concentration of 0.5 % v/v. A peristaltic pump was connected to the chamber and the FITC latex beads solution was pumped through the lumens. In order to image the selective thermo-responsive behavior of the vessel composed of the hybrid hydrogel composition, the printed samples were exposed to FITC latex beads solution perfusion via temperature gradient of 20-37 °C, and imaged using a binocular microscope (SMZ18, Nikon). Video clips were acquired using an ORCA-Flash 4.0 digital complementary metal-oxide- semiconductor (CMOS) camera (Hamamatsu Photonics, Hamamatsu city, Japan) at a rate of 100 frames s’¹.

The viability of cells in the hybrid hydrogel composition implants, containing either HUVECs or iPSCs derived ECs was determined using a live/dead fluorescent staining assay with fluorescein diacetate (7 pg-mL^-1, Sigma- Aldrich) and propidium iodide (5 pg-mL^-1, Sigma- Aldrich) for 30 min at 37 °C. Live and dead cells within the different implants were visualized by an inverted fluorescence microscope (Nikon ECLIPSE TLE, Melville, NY, USA), on day 0, and 3-, 6-, and 10-days post encapsulation.

For cellular metabolism and proliferation assay, ECM hydrogel and ECM-PNIPAM hydrogel were mixed with HUVECs at a concentration of 20M (cells- mL"¹). 3D hydrogel- HUVECs laden droplets were prepared at a volume of 5 pL, inside 24-well plates, 7 droplets per each well. For each time point, 12 measurements were recorded per group. PrestoBlue™ reagent (Fisher Scientific) was added to each well in a 1:10 (v/v) ratio with cell medium and incubated for 2 hours, which appeared to be the color changing time on day 0. The fluorescence was measured at 560 nm (590 nm serving as the reference wavelength) using an InfiniteM200Pro plate reader (Tecan, Mannedorf, Switzerland ). All values were normalized to Day 0.

For biocompatibility assessment of iPSC-CMs within ECM-PNIPAM bioinks at different ratios, cardiomyocytes were differentiated as previously described. On Day 16 of the differentiation, cells were dissociated from the Matrigel and encapsulated in ECM-PNIPAM hybrid hydrogel compositions of varying concentrations (80:20, 70:30, 60:40, 50:50 ECM:PNIPAM). Within each experimental group, the gel was extruded through a syringe in order to simulate the shear stresses present during the printing process. All droplets were matured for 14 days. On Days 3 and 14 of the maturation, live/dead assays were performed. Bright field video of spontaneous contractions were recorded on day 14 by using inverted fluorescence microscope (Nikon ECLIPSE TLE).

For calcium imaging, samples were incubated with 10 pm Fluo-4 AM (Invitrogen, Massachusetts, USA) and 0.1 % Pluronic F-127 (Sigma-Aldrich) in Tyrode’s solution (1.8 mm CaCh, 5 mm glucose, 10 mm HEPES, 1 mm MgCh, 5.4 mm KC1, 135 mm NaCl, 0.33 mm NasPO4) for 45 minutes at 37 °C. Samples were then imaged using a binocular microscope (SMZ18, Nikon, Melville, NY, USA). Video clips were acquired using an ORCA-Flash 4.0 digital complementary metal-oxide- semiconductor (CMOS) camera (Hamamatsu Photonics, Hamamatsu city, Japan) at a rate of 100 frames s-1. For the spatiotemporal heat maps, movies of the calcium signals were filtered using ImageJ (FIJI) and analyzed using MATLAB software (MathWorks, MA, USA), where a custom script was employed for detecting the time-point of maximum change in intensity for every pixel.

For immunofluorescence staining, 2D and 3D Samples were fixed in 3.5 % formaldehyde (Bio-Lab) for 30 minutes and 1 hour at RT, respectively. Then, samples were washed 3 times with PBS, permeabilized with 0.1% (v/v) Triton X-100 (Sigma-Aldrich) for 10 min and blocked with 2% BSA (MP Biomedicals) in PBS for 1 hour at RT. Samples were then stained with primary antibodies as listed in the antibodies list, diluted in 2% BSA blocking solution at 4 °C overnight. After three washes with PBS, samples were incubated with secondary antibodies as listed in the antibodies list, diluted in 2% BSA blocking solution for 1.5 hours at RT. For nuclei detection, D API ready-made solution (1:50; Sigma- Aldrich) was added along with the secondary antibodies. Samples were imaged using a confocal microscope (Nikon Eclipse NI-E)). Images were processed and analyzed using NIS elements software BR 3.2 (Nikon Instruments, Melville, NY, USA).

Antibodies for stem cells included: Mouse aOct3/4 (IgG2b) (SC 5279) 1 :250 (Santa-Cruz). Rabbit to Ki67 (abl6667), 1:250 (Abeam, Boston, USA). Antibodies for cardiac cells: Rabbit to Sarcomeric Alpha Sarcomeric Actinin (ab68167), 1 :200 (Abeam). Antibodies for endothelial cells: Mouse to CD31/PECAM-1 (P8590), 1:250 (Sigma- Aldrich), Rabbit to CD31 (ab28364), 1:100 (abeam). Antibodies for human cells: Mouse to human nuclei (abl91181), 1:300 (abeam), Rabbit to CD31 (ab28364), 1:100 (abeam). Secondary antibodies: Goat Anti-Rabbit (Alexa Fluor 488) (ab2338046), 1:250 (Jackson ImmunoResearch, PA, USA). Goat Anti-Mouse (Alexa Fluor 555) (abl50118), 1:500 (Abeam). Goat Anti-Mouse (Alexa Fluor 647) (ab2338902), 1:250 (Jackson ImmunoResearch). For detection of nuclei, cells were incubated with DAPI readymade solution, 1:50 (Sigma- Aldrich). Cell Proliferation Staining: CytoPainter Cell Proliferation Staining Reagent- Green Fluorescence (abl76735) (Abeam).

For flow cytometry assay of iPSCs, cells were dissociated with Accutase™ (StemCell Technologies), centrifuged at 300 g, and resuspended in Flow Cytometry Staining Buffer (R&D Systems). Cells were aliquoted and compared against a control aliquot and an aliquot stained with a control isotype. The antibodies used were (Miltenyi Biotech): TRA-1-60 (REA157), SSEA-1 (REA321), SSEA-4 (REA101), Control antibody (REA293). Data was collected on a CytoFEEX S Flow Cytometer (Beckman Coulter) and analyzed using their CytExpert software.

For flow cytometry assay of iPSC-CMs, cells were dissociated with TrypEE™ Express (Gibco, Waltham, Massachusetts), centrifuged at 300 g, and resuspended in eBioscience™ Permeabilization Buffer and Fixation/Perm Diluent Buffer (Invitrogen). Cells were subsequently blocked with Bio-Pure Human Serum Albumin 10 % solution (Biological Industries) diluted to 0.1 % in PBS. 2 The cells were aliquoted for controls and isotype and stained for (Miltenyi Biotech): Cardiac Troponin (REA400) and REA Control (REA293). Data was collected on a CytoFLEX S Flow Cytometer (Beckman Coulter) and analyzed using their CytExpert software.

For flow cytometry assay of iPSC-ECs, cells were dissociated with Accutase (StemCell Technologies), centrifuged at 300g, and resuspended in eBioscience™ Permeabilization Buffer and Fixation/Perm Diluent Buffer (Invitrogen). Cells were subsequently blocked with Bio-Pure Human Serum Albumin 10% solution (Sartorious) diluted to 0.1% in PBS. The cells were aliquoted for controls and isotype and stained for CD31 (REA730) (Miltenyi Biotech) and REA control antibody (REA293) (Miltenyi Biotech). Data was collected on a CytoFLEX S Flow Cytometer (Beckman Coulter) and analyzed using their CytExpert software.

To optimize the cell confluency within the ECM-PNIPAM hybrid hydrogel composition at physiological conditions, HUVECs and iPSCs derived ECs were used. HUVECs were cultured as previously described. First, dissociated cells were resuspended in HBSS medium (Sartorious) and with additional 2 pL CytoPainter Cell Proliferation Staining Reagent- Green Fluorescence (Abeam) per IM cells, for 30 min. of incubation at 37°C. The marked cells were centrifuged at 300g for 5 min, and the marked cells were seeded at a concentration range of 10-70 cells- mm^-2, onto 50 pL of hybrid hydrogel sheets, inside 24 well plate, to create cells monolayers. The sheets were immediately imaged at 20°C by an inverted fluorescence microscope (Nikon ECLIPSE TI- E), and then incubated at 37°C in a humidified, 5% CO2 incubator for Ih for the shrinking and crosslinking of the hybrid hydrogel, as well as cells attachment. Next, the cells different monolayers were imaged at 37°C by an inverted fluorescence microscope (Nikon ECLIPSE TI- E), to measure the change in cells coverage within the shrinking process of the ECM-PNIPAM hydrogel. EGM-2 medium was added to the samples, which were then cultured in a humidified, 5% CO2 incubator for 7 days, until fixed for immunostaining for nuclei and CD31 endothelial marker, to evaluate the cells confluency 7 days post incubation. According to the results, the experiment was repeated with iPSCs derived endothelial cells only with the optimal cells concentration of 30xl0³ cells- mm^-2.

Measurements of the cellular coverage to evaluate the optimal cells confluency postshrinking within the hybrid ECM-PNIPAM hydrogel were performed on both the HUVECs and the iPSCs derived ECs loaded hybrid monolayers. The cellular images of the cyto-painted cells that were taken at 20°C and at 37°C, Ih post incubation were analyzed using ImageJ (FIJI), and the cellular coverage was calculated according to the number of cells, normalized to the size of the measured field. The immuno staining images of the nuclei and CD31 cellular expression 7 days post incubation were analyzed using ImageJ (FIJI), and the cellular coverage was calculated according to the number of nuclei, normalized to the size of the measured field. The results on day 7 were normalized to the cellular coverage at 20°C, to present the fold coverage on day 7.

Vascularized patches (both cellular and acellular) were printed using a 3Ddiscovery Evolution® printer (RegenHU, Villaz-St-Pierre, Switzerland). For the acellular printing process, three different thermo-responsive bioinks were utilized to fabricate the 3D-printed vascularized cardiac tissue: 1. Pristine ECM bioink; 2. gelatin microparticles bioink; 3. ECM-PNIPAM bioink (smart hybrid-ink). Both the gelatin microparticles and the smart hybrid hydrogel bioinks were printed using a high-precision printheads, while the pristine-ECM bioink was printed using pneumatic pressure controlled printhead. The Pristine-ECM and the smart-hybrid bioinks were printed through a 25G needle. The gelatin microparticles bioink was printed through a 30G needle. The patches were printed onto 12-wells filled with a support medium bath. During the printing process the pristine-ECM bioink was first being extruded in a crisscross geometry, to fabricate the six lower layers of the patch bulk. Embedded within the sixth layer, the smart hybrid bioink was localized to certain regions of the structure to generate the walls of the blood vessel-like tubes that would be selectively shrunk. Then, strands of the gelatin microparticles bioink were printed to form the hollow lumens of the blood vessel-like structures (both the static and the dynamic vessels). Finally, another six layers of the pristine-ECM bioink were printed on top of the patch.

For the fabrication of the vascularized cardiac patches via cellular printing process, tissues were printed similar to the acellular constructs, whereas iPSCs derived cardiomyocytes were encapsulated within the Pristine ECM bioink (CM-ink) and iPSCs derived endothelial cells were encapsulated within gelatin microparticles bioink (EC-ink). For the preparation of the CM-ink, iPSC-derived CMs grown on Matrigel coated plates were incubated for 10 min with TrypLE Express (Gibco, Waltham, MA, USA). Colonies were then mechanically triturated, and cells were centrifuged at 300 g for 5 min. The supernatant was removed, and ECM-based hydrogel was added at a ratio of 1 mL per 200 million cells. For the preparation of the EC-ink, iPSC-derived ECs were incubated with a 0.25% solution of trypsin and EDTA (Biological Industries) for 5 min to dissociate cells. The cells were collected in an excess of DMEM and centrifuged at 300 g for 5 min. The supernatant was removed and the cells were resuspended in the gelatin microparticles at a ratio of 40M cells per 1 mL of gelatin microparticles.

Following printing, both acellular and cellular patches were placed in a humidified incubator (37 °C, 5% CO2) for 45 min, during which time each bioink’s thermo-responsive behavior was triggered, at a coordinated kinetically controlled sequence. The acellular patches were then utilized for further analysis of the selective shrinkage of the smart printed vessels and their ability to be perfused. The vascularized cardiac tissues were suspended with M-199 medium (gibco) with the addition of the EGM-2 bullet factors kit (Lonza), and alginate lyase 1 U mL-1 (Sigma- Aldrich) for support medium dissolving. This medium was refreshed without the alginate lyase every other day, for 14 days.

Evaluation of hemolysis, an indicator for red blood cell destruction, caused by the printed blood vessels, was performed by quantifying the concentration of red blood cells using an inverted fluorescence microscope (Nikon ECLIPSE TLE, Melville, NY, USA). Briefly, branched a-cellular and HUVECs laden vascularized patches (n=5 in each group) were printed and cultivated for two weeks. The patches were then perfused with defibrinated sheep's blood diluted in a 1:1 ratio with HEPES buffer for 20 minutes, 50 pL- min’¹ flow rate, at physiological temperature (37 °C). The blood was collected, and the fractions from each patch were imaged using inverted fluorescence microscope. Red blood cells concentration was analyzed using ImageJ (FIJI). The red blood cells concentration from all groups was normalized to the red blood cells concentration post perfusion of the blood through the pump’s tubes only.

The thrombogenicity of the printed blood vessels was assessed by monitoring the evolution of thrombin activity. Branched a-cellular and HUVECs laden vascularized patches (n=10 in each group) were printed and cultivated for two weeks. Defibrinated sheep's whole blood was diluted in a 1:1 ratio with HEPES buffer and mixed with Boc-Val-Pro-Arg-AMC to create 125 pg- mL ¹ solution. This solution was then perfused through the branched a-cellular and cellular vascularized patches, for 20 minutes, at physiological temperature (37 °C) with a flow rate of 50 pL-min"¹. The perfusion was performed using a peristaltic pump, connected to the outlets of the perfusion chamber, that is fixed to the stage of a binocular microscope (SMZ18, Nikon). Movies of the perfusion were acquired using an ORCA-Flash 4.0 digital complementary metal-oxide- semiconductor (CMOS) camera (Hamamatsu Photonics, Hamamatsu city, Japan) at a rate of 100 frames s’¹. The level of fluorescence post perfusion, indicating on thrombogenicity, was analyzed using ImageJ (FUI). The fluorescence from all groups was normalized to the initial fluorescence value.

Testing angiogenesis processes in engineered tissues implanted on an omentum tissue of an animal model, was conducted as follows. Recipient SD male rats (260-280 g) were anesthetized with a combination of ketamine (100 mg/kg) and xylazine(10 mg/kg). After a midline abdominal incision was made, the printed vascularized patches were placed on the omentum and were secured in place with a 6-0 prolene suture. Post-surgery, the rats received heparin subcutaneous injections (50 U-Kg’¹) every 48 hours to prevent blood coagulation. Seven days after implantation, the animals were killed and the patches were gently cut and cleaned with scalpel. The patches were then fixed with formaldehyde 3.5 % solution for 40 minutes, and immunostained for endothelial CD31 and human nuclei markers. Samples were imaged using a confocal microscope (Nikon Eclipse NI-E)). Images were processed and analyzed using NIS elements software BR 3.2 (Nikon Instruments, Melville, NY, USA).

All rats were treated according to ethical regulations of Tel Aviv University. Permission was granted by the Institutional Animal Care and Use Committee (IACUC) of Tel Aviv University, protocol number 2403 - 113 - 3 “Testing angiogenesis processes in engineered tissues implanted on an omentum tissue of an animal model”. Statistical analyses were presented as mean ± standard error of mean on the basis of at least three replicates. Differences between samples were assessed by the relevant tests (details for each experiment were provided alongside the data), and p< 0.05 was considered significant. Analyses were performed using GraphPad Prism 8 (Version 8.4.2) for Windows (GraphPad Software).

It is expected that during the life of a patent maturing from this application many relevant bioinks for 4D printing will be developed and the scope of the phrase "a bioink for 4D printing" is intended to include all such new technologies a priori.

It is appreciated that certain features of the invention, which are, for clarity, described in the context of separate embodiments, may also be provided in combination in a single embodiment. Conversely, various features of the invention, which are, for brevity, described in the context of a single embodiment, may also be provided separately or in any suitable subcombination or as suitable in any other described embodiment of the invention. Certain features described in the context of various embodiments are not to be considered essential features of those embodiments, unless the embodiment is inoperative without those elements.

Various embodiments and aspects of the present invention as delineated hereinabove and as claimed in the claims section below find experimental support in the following examples.

Although the invention has been described in conjunction with specific embodiments thereof, it is evident that many alternatives, modifications and variations will be apparent to those skilled in the art. Accordingly, it is intended to embrace all such alternatives, modifications and variations that fall within the spirit and broad scope of the appended claims.

All publications, patents and patent applications mentioned in this specification are herein incorporated in their entirety by reference into the specification, to the same extent as if each individual publication, patent or patent application was specifically and individually indicated to be incorporated herein by reference. In addition, citation or identification of any reference in this application shall not be construed as an admission that such reference is available as prior art to the present invention. To the extent that section headings are used, they should not be construed as necessarily limiting.

The claims should be interpreted to cover all modifications, equivalents, and alternatives falling within the scope and spirit of the invention as defined by the claims. The scope of the claimed subject matter is intended to be as broad as legally permissible. In addition, any priority document(s) of this application is/are hereby incorporated herein by reference in its/their entirety.

Claims

WHAT IS CLAIMED IS:

1. A hybrid hydrogel composition comprising: a plurality of discrete hydrogel particles, each of said particles comprises a first crosslinked polymeric network that comprises a first polymer; a second crosslinked polymeric network that comprises a second polymer; and a dispersing medium, wherein: said second crosslinked polymeric network comprises crosslinked polymeric chains that interpenetrate through each of said hydrogel particles, thereby forming an interpenetrating polymer network that comprises a plurality of interpenetrated-through hydrogel particles and said second crosslinked polymeric network; and said interpenetrating polymer network is dispersed within said dispersing medium, thereby forming the hybrid hydrogel composition.

2. The composition of claim 1, wherein said interpenetrated-through hydrogel particles are maintained substantially in a non-contacting, spaced-apart arrangement within said second crosslinked polymeric network.

3. The composition of any one of claims 1-2, wherein said interpenetrated-through hydrogel particles and said second crosslinked polymeric network forming said interpenetrating polymer network are not linked to each other by covalent bonds.

4. The composition of any one of claims 1-3, wherein said first polymer comprises a thermo-responsive polymer.

5. The composition of claim 4, wherein said thermo-responsive polymer is a thermo- contractive polymer.

6. The composition of claim 5, wherein a lower critical solution temperature (LCST) of said thermo-contractive polymer ranges 30-40 °C.

7. The composition of claim 6, wherein said thermo-contractive polymer comprises poly(N -isopropylacrylamide) (PNIPAM) .

8. The composition of any one of claims 1-7, wherein said second polymer is a thermo-denaturing polymer.

9. The composition of claim 8, wherein said thermo-denaturing polymer is a proteinbased polymer.

10. The composition of claim 9, wherein said second polymer comprises collagen and/or ECM-derived polymers, and said second crosslinked polymeric network comprises collagen-based hydrogel and/or an ECM-based hydrogel.

11. The composition of claim 10, wherein a denaturing temperature of said thermodenaturing polymer ranges 35-40 °C.

12. A process of preparing the composition of any one of claims 1-11, comprising: providing said plurality of said discrete hydrogel particles in a swelled form; providing a non-crosslinked hydrogel of said second polymer; mixing said hydrogel particles with said non-crosslinked hydrogel of said second polymer in the presence of a crosslinking agent, said crosslinking agent is selected to effect crosslinking of said second polymer to thereby obtain said second crosslinked polymeric network and said interpenetrating polymer network; thereby providing the hybrid hydrogel composition.

13. The process of claim 12, further comprising, subsequent to said mixing, washing said interpenetrating polymer network.

14. The process of any one of claims 12-13, wherein said first polymer comprises a thermo-responsive polymer, and said mixing is effected at a temperature below an LCST of said first polymer.

15. The process of claim 14, wherein said LCST ranges 30-40 °C.

16. The process of any one of claims 12-15, wherein said second polymer is a thermodenaturing polymer.

17. The process of claim 16, wherein a denaturing temperature of said thermodenaturing polymer ranges 35-40 °C.

18. The process of any one of claims 12-17, wherein a progress and completion of said crosslinking of said second polymer is determined by a monitoring an increase in viscosity of the hybrid hydrogel composition.

19. The process of claim 18, wherein said completion is determined when said viscosity ranges 0.03 - 6- 10⁵ Pa- s.

20. A method of 3D printing an object, comprising printing the object using a first bioink that comprises the hybrid hydrogel composition according to any one of claims 1-11, wherein said first polymer comprises a thermo-contractive polymer characterized by an LCST, and a thermo-denaturing polymer characterized by a denaturing temperature, said LCST is lower than said denaturing temperature.

21. The method of claim 20, wherein said first crosslinked polymeric network comprises PNIPAM and said second crosslinked polymeric network comprises a collagen-based crosslinked hydrogel and/or an ECM-based crosslinked hydrogel.

22. The method of any one of claims 20-21, wherein said printing is effected at a temperature lower than said LCST.

23. The method of any one of claims 20-22, further comprising subsequent to said printing, increasing the temperature of the object above said LCST.

24. The method of any one of claims 20-22, the method further comprising printing the object using a second bioink that comprises non-crosslinked hydrogel composition that comprises a thermo-denaturing polymer characterized by a denaturing temperature, said second bioink is printed at a temperature lower than said denaturing temperature and lower than said LCST.

25. The method of claim 24, wherein said second bioink is printed such that it envelops at least a portion of an outer surface of the part of the object comprising said first bioink.

26. The method of any one of claims 24-25, wherein said second bioink comprises a non-crosslinked collagen-based hydrogel and/or a non-crosslinked ECM-based hydrogel.

27. The method of any one of claims 24-26, further comprising subsequent to said printing, increasing the temperature of the object above said LCST, and gradually or incrementally increasing the temperature of the object above said denaturing temperature to thereby obtaining the object.

28. The method of any one of claims 20-26, wherein the object is characterized by a lumen, the method further comprising printing the object using a third bioink that comprises a low- melting point hydrogel composition characterized by a melting temperature or a gel-sol transition temperature, said melting temperature or said gel-sol transition temperature is lower than said LCST, said third bioink is printed at a temperature lower than said melting temperature or said gel-sol transition temperature in the position, shape and size of said lumen, and substantially enveloped by said first bioink.

29. The method of claim 28, wherein said third bioink comprises a substance selected from the group consisting of gelatin, , a poloxamer, polyvinyl alcohol, and/or a carbomer resin (Carbopol).

30. The method of any one of claims 28-29, further comprising subsequent to said printing, increasing the temperature of the object above said melting temperature, or increasing the temperature of the object above said LCST, and increasing the temperature of the object above said denaturing temperature to thereby obtaining the object.

31. An object comprising the hybrid hydrogel composition of any one of claims 1-11.

32. An object formed by the method of any one of claims 20-31.

33. The object of claim 32, comprising at least one structural lumen, said lumen is characterized by at least one dimension of less than 50 pm.

34. The object of any one of claims 32-33, being an artificial tissue that comprises at least one capillary blood vessels characterized by an inner diameter of less than 50 pm and a length of at least 0.5 mm.

35. A method for anisotropic 3D printing of an object, comprising: a) providing a bioink comprising the hybrid hydrogel composition of any one of claims 1- 11; b) extruding said bioink through a printhead to form at least a part of the object; c) controlling the direction of printhead motion during extrusion to create directional properties in the object; and d) stimulating said hybrid hydrogel composition forming said at least a part of the object to induce anisotropic shrinkage.

36. The method of claim 35, wherein said anisotropic shrinkage is characterized by greater shrinkage in a direction perpendicular to said direction of said printhead motion compared to shrinkage parallel to said direction of said printhead motion.

37. The method of claim 35, further comprising printing geometric shapes according to different patterns to achieve distinct shrinkage patterns upon activation.

38. The method of claim 37, wherein said geometric shapes include circular structures that exhibit radial compaction upon stimulating said hybrid hydrogel composition.

39. The method of claim 38, wherein the radial compaction results in different rates of change for inner and outer diameters of said circular structures.

40. The method of claim 35, wherein stimulating said hybrid hydrogel composition comprises applying heat.

41. A 3D printed object produced by the method of any one of claims 35-40, wherein the object exhibits anisotropic shrinkage in response to stimulating said hybrid hydrogel composition.