Tunable degradation in hydrogel microparticles

Abstract

Non-alginate hydrogel microparticles for local delivery and sustained release of therapeutic cells and / or tissues (including homogeneous or heterogeneous cell clusters) at the implantation site. The microparticles include a 3-dimensional matrix of a covalently cross-linked non-alginate polymer compound and a therapeutically effective amount of cells and / or tissues embedded therein, wherein the cells have a viability of at least about 50%, and wherein the microparticles have a size greater than about 30 μm. Compositions comprising such microparticles and methods of using such microparticles for treatment are also described.

Description

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS

[0002] This application claims the benefit of priority to U.S. Provisional Patent Application Serial No. 62 / 858,578, filed on June 7, 2019, entitled “TUNABLE DEGRADATION IN HYDROGEL MICROPARTICLES,” which is incorporated herein by reference in its entirety. Background Art

[0003] Field of the Invention

[0004] The present invention relates to hydrogel-based microparticles for encapsulating biologics, and methods for modulating the physical and chemical characteristics of the hydrogel matrix, for example, to increase or decrease the degradation profile of the microparticles.

[0005] Description of Related Technology

[0006] Cell microencapsulation is a rapidly expanding field with broad potential applications in treating diseases such as diabetes, cancer, heart disease, and immune disorders, as well as in tissue engineering research and regenerative medicine. Traditional alginate microspheres still suffer from poor biocompatibility, and microencapsulation using more advanced hydrogels is challenging due to their slow gelation rate. What is needed are non-cytotoxic, non-emulsion-based hydrogel microparticles that can be manufactured from a variety of hydrogel materials. Summary of the Invention

[0007] Described herein are non-alginate hydrogel microparticles for local delivery and sustained release of therapeutic cells and / or tissues (including homogeneous or heterogeneous cell clusters) at the implant site. The microparticles include a 3-dimensional matrix of a covalently cross-linked non-alginate polymer compound and a therapeutically effective amount of cells and / or tissues embedded therein, wherein the cells have a viability of at least 50%, and wherein the microparticles have a size greater than about 30 μm. Exemplary polymer compounds are selected from slow-gelling polymer precursors selected from the group consisting of: branched or unbranched hyaluronic acid, branched or unbranched functionalized hyaluronic acid, branched or unbranched functionalized polyethylene glycol, hyaluronan, fibrin, chitosan, collagen, polylactic acid, poly(L-lactic acid), polylactic-co-glycolic acid, polycaprolactone, polyvinyl alcohol, and combinations thereof. Described matrix can further include the cross-linking agent cross-linked with described non-alginate polymer compound, and described cross-linking agent is selected from the group of free following composition: dithiothreitol, side chain or non-side chain functionalized polyethylene glycol, dithiol, ethylene glycol bis-thioglycolate and combination thereof.Exemplary functionalized polyethylene glycol includes but is not limited to polyethylene glycol dithiol, polyethylene glycol diacrylate, polyethylene glycol divinyl sulfone, polyethylene glycol dimaleimide and combination thereof.Described matrix can be homogeneous, includes the polymer backbone (with or without the cross-linking agent of extra different polymer types) of a type of covalent cross-linking.Described matrix can also be heterogeneous, includes the mixture of two or more polymer precursors, such as high ( > 500kDA, preferred > 100kDA) and low (<100kDA, preferably < 50 kDa) molecular weight polymer precursors, and / or mixtures of two or more cross-linking agents.

[0008] Also described herein are compositions comprising a plurality of such microparticles. Such compositions may also include a plurality of "empty" microparticles dispersed together with the microparticles of the loaded cells or tissues. The compositions may include a mixture of different microparticles, such as a mixture of degradable and durable hydrogel microparticles, and / or a mixture of microparticles of different sizes (e.g., a plurality of larger microparticles and a plurality of smaller microparticles), and / or a mixture of microparticles each comprising different types of cells and / or tissues (e.g., insulin-producing cells in one group of microparticles and supportive stem cells in another group of microparticles). Two different compositions may also be combined or co-administered, rather than two different types of particles being mixed into a single composition.

[0009] Also contemplated herein are various methods for local delivery and sustained release of therapeutic cells and / or tissues at the site of implantation by administering the microparticles and / or compositions comprising such microparticles (alone or, if desired, in combination with other types of therapeutic compounds. Typically, the microparticles or compositions are injected or implanted directly at or near the treatment site (adjacent to the treatment site).

[0010] Also described herein are various methods and techniques for modulating the degradation rate of microparticles by varying various parameters of the microparticles, including formulation and / or processing parameters, such as polymer precursor mass fraction or molecular weight, crosslinker molecular weight, ratio of crosslinker to polymer precursor, crosslinker hydrolysis, crosslinking kinetics, crosslinking time (e.g., UV exposure time), and combinations thereof.

[0011] Further, embodiments described herein contemplate the use of empty microparticles for joint lubrication and / or viscosupplementation by implanting a plurality of empty microparticles in a joint. BRIEF DESCRIPTION OF THE DRAWINGS

[0012] Figure (Fig.) 1 is a schematic representation of an example of covalent cross-linking of hydrogel precursor compounds.

[0013] Figure 2 is a schematic diagram of the process for forming microparticles, including (A) droplet formation, (B) incubation in an alginate bath, (C) formation and thickening of the alginate shell, and (D) cross-linking of the hydrogel core followed by removal of the alginate shell to produce hydrogel microparticles.

[0014] Figure 3 Shown are images of hydrogel microparticles for (A) the alginate shell surrounding the microparticles; and (B) after removal of the alginate shell.

[0015] Figure 4 Images of (A) PEG-based spherical durable microparticles loaded with islets, (B) PEG-based elliptical microparticles loaded with islets, (C) MeHA-based flat-edged microparticles loaded with stem cells, and (D) MeHA-based elliptical microparticles loaded with stem cells are shown.

[0016] Figure 5 Shown are results from cytotoxicity studies and the effects of (A) calcium concentration and exposure time on the viability of canine islets, and (B) the effects of photoinitiator concentration and UV light exposure time on cell viability.

[0017] Figure 6 is a graphical depiction of the diameters of 100 representative microparticles composed of ThHA, MeHA, PEGDA or AHA and their good monodispersity profile.

[0018] Figure 7 Images showing the diffusion of fluorescently labeled dextran into PEGDA, ThHA, and MeHA hydrogel microparticles.

[0019] Figure 8A is a graph of the degradation of different particles based on the remaining micropolymer mass over time.

[0020] Figure 8B It is a graph showing the change of Q value (swelling ratio) of different particles over time.

[0021] Figure 8C is a degradation graph of different particles based on the mass loss of the particles over time.

[0022] Figure 9 is a graph of the mean daily blood glucose levels for each group during the initial study cohort period following transplantation in mice.

[0023] Figure 10 Shown are images of microparticles recovered from the IP space of mice, (A) PEDGA, (B) MeHA, (C) PEDGA, where viability is indicated by deep red disulfide hydrazone staining, and (D) PEDGA, where viability is indicated by co-staining with deep red disulfide hydrazone and green calcein.

[0024] Figure 11 Images of healthy fat cells are shown, demonstrating the absence of fibrotic tissue following implantation with PEGDA microparticles.

[0025] Figure 12 is a graph of mean non-fasting blood glucose from measurements obtained in dogs each morning and afternoon.

[0026] Figure 13 The following images are shown: (A) HEK293 cells aggregated in durable PEGDA microparticles, (B) little change in morphology or number of cells from (A) in PEGDA microparticles after 34 days, (C) HEK293 cells dispersed in degradable MeHA microparticles, and (D) proliferation of cells from (C) in MeHA microparticles after 34 days.

[0027] Figure 14 Shown are images from microplate wells where spent media was collected showing (A) very few HEK293 cells diffusing from PEGDA microparticles, and (B) a large number of viable cells diffusing from MeHA microparticles.

[0028] Figure 15 is a graph summarizing average data from studies on cumulative cellular release.

[0029] Figure 16Images of (A) MSCs encapsulated in PEGDA microparticles, (B) MSCs encapsulated in MeHA microparticles, (C) islets encapsulated in PEGDA microparticles, (D) islets encapsulated in MeHA microparticles, and (E) MSCs encapsulated in AHA microparticles are shown.

[0030] Figure 17 Figure 2 is a graph showing cell viability of rat insulinoma cell lines encapsulated in PEGDA and AHA microparticles.

[0031] Figure 18 Shown are images of pancreatic islets encapsulated in PEGDA or MeHA and stained with disulfide hydrazone to demonstrate functional insulin-producing islets.

[0032] Figure 19 Shown are images of (A) PEDGA microparticles infused into and around the abdominal omentum, (B) microparticles attached to tissue in the first dog, and (C) microparticles attached to tissue in the second dog, demonstrating the "stickiness" of the particles.

[0033] Figure 20 The following images are shown: (A) blue-stained PEGDA microparticles in a rat knee joint, visible through the muscle; (B) a dissected knee joint showing blue-stained PEGDA microparticles in the knee capsule area; (C) a further dissected knee joint showing blue-stained PEGDA microparticles remaining where they were originally implanted.

[0034] Figure 21 Hematoxylin and eosin-stained images of explanted tissues of the three microparticle groups are shown.

[0035] Figure 22 is a graph showing the response of (without) foreign matter to particles in the retina.

[0036] Figure 23 is a graph showing the effect of the mass fraction of the core (backbone) polymer species used in the hydrogel precursor solution on the Q value.

[0037] Figure 24A is a graph showing formulations where increasing cross-linker concentration results in lower Q values.

[0038] Figure 24B is a graph showing formulations where increasing cross-linker concentration results in higher Q values.

[0039] Figure 25A is a graph showing formulations where increasing pH results in higher Q values.

[0040] Figure 25B is a graph showing formulations where increasing pH results in lower Q values.

[0041] Figure 26A is a graph showing formulations where increasing the molecular weight of the cross-linker results in higher Q values.

[0042] Figure 26B is a graph showing a preparation in which increasing the molecular weight of the cross-linking agent does not cause a change in the Q value.

[0043] Figure 27A is a graph showing the effect of varying the molecular weight of a core polymer species with constant reactive sites / gram on the Q value.

[0044] Figure 27B is a graph showing the effect of varying the molecular weight of core polymer species with different reactive sites / gram on the Q value.

[0045] Figure 28A is a graph showing the effect of UV cross-linking time on the Q value (swelling ratio) and diameter of the particles for different microparticle formulations.

[0046] Figure 28B is a graph showing the effect of UV cross-linking time on the Q value of MeHA microparticles. DETAILED DESCRIPTION

[0047] The present invention relates broadly to hydrogel microparticles for local delivery and sustained release of therapeutic cells and tissues. The hydrogel microparticles protect the implanted cells from initial degradation and flushing (biodegradation) under physiological conditions and control the kinetics of the release of cells (and their therapeutic proteins and molecules) to local or systemic regions over a period of several hours to several weeks or months. The hydrogel microparticles designed for degradation ultimately degrade in vivo over a period of several days or weeks and typically within less than 3 months through normal physiological processes. This means that the entire hydrogel has been degraded and washed away by normal physiological conditions within this timeframe, leaving substantially no hydrogel in the implantation region (it should be understood that trace amounts of gel or polymer may still remain in the region, but are generally no longer sufficient to serve as an encapsulating agent for the implanted cells). It should also be understood that in some indications, particularly in the case of implantation at joints, the microparticles are also subjected to mechanical degradation by movement and impact at the joint site.

[0048] The hydrogel microparticles include a hydrogel matrix in which a therapeutically effective amount of cells is encapsulated. The cells may be eukaryotic or prokaryotic. Examples of prokaryotic cells include bacteria and archaea. Eukaryotic cells include animal cells, plant cells, and fungi. The cells may be non-proliferative or proliferative. Before implantation and / or during the implantation period, or after the proliferative cells are released from the degrading hydrogel, the proliferative cells will further proliferate (increase) from the initial number of cells encapsulated in the hydrogel. Non-proliferative cells have limited in vitro expansion capacity and generally do not increase from the initial number of cells. However, due to the favorable conditions in the hydrogel microparticles, there is very little cell death in the encapsulated cells, and the initial number of non-proliferative cells encapsulated by the hydrogel will not be significantly reduced. In one or more embodiments, the initial viability of the cells after encapsulation in the microparticles is maintained at least 50%, preferably at least about 70% viability, more preferably at least about 80% viability, even more preferably at least about 90% viability, and even more preferably at least about 95% viability after encapsulation, compared to the viability of the starting cell population included in the precursor solution. In other words, less than 50% of the cell population is lost during encapsulation, preferably less than 30%, more preferably less than 20%, even more preferably less than 90%, and even more preferably less than 5% of the cell population is lost during the encapsulation process. This is due to the gentle encapsulation process and the ability of cell culture medium and nutrients to diffuse into the microparticles during storage. In other words, in one or more embodiments, the microparticles have a low diffusion barrier.

[0049] In one or more embodiments, the cell population included in the hydrogel precursor solution will include about 1 million to about 500 million cells / mL of hydrogel precursor solution, preferably about 4 million cells to about 300 million cells / mL of hydrogel precursor solution, and in some embodiments, about 5 million to about 150 million cells / mL of precursor solution. In one or more embodiments, the hydrogel precursor solution will include about 15% by volume of the cells, cell clusters, or tissue to be encapsulated / mL of hydrogel precursor solution.

[0050] Examples of proliferative cells include stem cells and activated T cells, while examples of non-proliferative cells include pancreatic islets and differentiated neuronal cells. In either case, cells (and their related cell products) are slowly released from the hydrogel to the local area of implantation. Exemplary cell products that can be secreted together with the cells include signaling molecules, therapeutic proteins, vesicles, antibodies, viruses, extracellular bodies, etc. These protein-based components can be in their natural form or can be genetically modified. Similarly, cell product components can be natural or genetically modified. Compared with other carriers or unencapsulated cells, hydrogel microparticles allow cells to maintain a longer period of time in the local area of implantation and provide a temporary but sustainable release of cells and / or their signaling molecules, proteins, etc. to the area to improve the therapeutic effect of implanted cells. It should be understood that the cell loading will depend on the size of the cells used and the size of each microparticle. In one or more embodiments, up to 50% of the volume of the microparticles is composed of encapsulated cells, cell clusters and / or tissues, preferably, the cell, cluster and / or tissue loading range is about 10% to about 20% of the volume of the microparticles. Typically, about 50 to about 1x10 5 Cells can be encapsulated in each 1-mm microparticle, preferably about 100 to about 1×10 5 will be encapsulated in each 1-mm microparticle.

[0051] In one or more embodiments, the microparticles are formed from weaker hydrogels that ultimately degrade in vivo within about 3 months (90 days) or less from the date of implantation and are referred to herein as "degradable" hydrogels (it should be understood that most hydrogels eventually degrade). As used herein, "degradable" hydrogel microparticles are defined as "sustained release" but "fast degrading" or "short-term" hydrogels that will maintain their self-sustaining microparticle size for at least 24 hours in PBS and up to 3 months when implanted in vivo. Such degradable hydrogel microparticles will also decompose and degrade within about 6 months under in vitro storage conditions (PBS and 37°C). In other words, the hydrogel matrix has an "in vitro storage stability" of less than 6 months, which means that it begins to hydrolyze and break apart from its self-sustaining size within 6 months or less when stored in PBS at 37°C. In one or more embodiments, the degradable hydrogel microparticles are characterized by having a Q value greater than 150, preferably greater than 200. However, as explained below, additional modifications to the cross-linking chemistry (e.g., using hydrolytically unstable cross-linkers) can produce microparticles with low Q values (strong initial matrix) but which decompose fairly rapidly and are still characterized as degradable hydrogel microparticles.

[0052] These degradable microparticles are in contrast to alternative embodiments involving durable hydrogel microparticles. Such durable hydrogel microparticles are also sustained release, but for longer term applications, such that they do not break down or degrade under storage conditions in PBS at 37°C for at least 6 months. In other words, the durable microparticles have an "in vitro storage stability" of greater than 6 months. In one or more embodiments, the durable hydrogel particles are storage stable in PBS at room temperature (27°C) for a year or more. Preferably, when such durable hydrogel microparticles are implanted, they will not break down under normal physiological conditions (also known as normal foreign body clearance by phagocytosis, degradation, adsorption, etc.) for at least 3 months and more preferably at least 6 months.

[0053] In either case, the hydrogel microparticles can also be characterized as microspheres, microbeads, or hydrogel microparticles. They are in the form of individual, self-supporting bodies comprising a 3-dimensional hydrogel matrix and cells suspended, embedded, or encapsulated in the hydrogel matrix. Also described herein are sustained-release compositions for implantation into a subject. The sustained-release compositions comprise a plurality of 3-dimensional hydrogel microparticles suspended in a pharmaceutically acceptable delivery vehicle.

[0054] Suitable hydrogel precursor compounds for use in forming microparticles include polymers, oligomers and / or monomers that form hydrogels, and therefore can form cross-linked or network structures or matrix (i.e., " hydrogels ") by cross-linking, wherein liquid and cells can retain, suspend, embed and / or be encapsulated in the interstitial space or pores of the structure or matrix body of the resulting gelation. Hydrogel precursor compounds for the present invention are preferably non-alginate hydrogel precursor compounds. That is, hydrogel precursor solutions are preferably substantially free of alginate compounds (i.e., compounds based on alginate, alginic acid, or its salt or derivative). As used herein, term "substantially free of" means that composition is not intentionally added to the composition, although incidental impurities may occur, or residues / trace amounts may be left in the manufacturing process. In such embodiments, the hydrogel precursor solution composition includes less than about 0.05% by weight, preferably less than about 0.01%, and more preferably about 0% by weight of such composition, based on the gross weight of the solution being 100% by weight.

[0055] Hydrogel microparticles can be advantageously made from a variety of slow-gelling polymer precursors, such as hyaluronic acid (HA) and polyethylene glycol diacrylate (PEGDA), polyethylene glycol maleimide (PEGMAL), multi-arm PEG, hyaluronic acid, fibrin, chitosan, collagen, heparin, polylactic acid (PLA), poly (L-lactic acid) (PLLA), polylactic acid-co-glycolic acid (PLGA), polycaprolactone (PCL), polyvinyl alcohol (PVA), polyvinyl pyrrolidone (PVP), methacrylic acid (MAA), 2-hydroxyethyl methacrylate (HEMA), polyacrylamide (PAM) extracellular matrix and functionalized species thereof (e.g., acrylated, methacrylated, thiolated, etc.) or their multi-arm species (e.g., 4-arm PEG, 8-arm PEG). However, any cross-linkable hydrogel precursor compound will be suitable for use with the present invention, wherein preferred compounds are biocompatible non-alginate homopolymers or copolymers, and particularly non-alginate block copolymers, as well as other types of cross-linkable monomers and / or oligomers. Typically, the amount of polymer precursor included in the precursor solution will be less than 50% w / w. > 500kDa, preferred > 100 kDa), polysubstituted polymers such as hyaluronic acid, the amount will range from about 1% to about 50% w / w, and for lower molecular weight polymers (<100 kDa, preferably < 50 kDa), such as PEGDA or PEGMAL, at about 5% to about 50% w / w.

[0056] The cross-linking profile of the microparticles can be adjusted by adjusting the molecular weight of the precursor compound and the selected cross-linking agent, cross-linking conditions and cross-linking process. For example, in some embodiments, "tighter" or faster cross-linking may be required to obtain a smoother bead surface and / or a stronger gel by limiting the ability of molecules to seep out of the droplet / core into the surrounding environment during cross-linking. This can be achieved by increasing the molecular weight of the precursor substance (e.g., ~>40kDa) and / or by reducing the cross-linking time (by adjusting the cross-linking chemical or initiation method). However, increasing the molecular weight of the bifunctional precursor substance (i.e., PEGDA) generally results in a more diffuse gel due to the reduced cross-linking density. Further, hydrolytically unstable cross-linkers can be used to produce initial strong microparticles with a low initial Q value to facilitate cell implantation into the site. However, such cross-linkers will degrade rapidly under normal physiological conditions, thereby rapidly releasing their payload to the implantation site. Depending on the specific hydrogel precursor compound, cross-linking can be performed by various mechanisms. In one or more embodiments, the core / shell microparticles are combined with a hydrogel matrix cross-linking agent, preferably in solution. Cross-linking agent is leached into the core / shell microparticle by alginate shell, causes hydrogel precursor compound gelation (cross-linking), to form 3-dimensional hydrogel matrix.Cross-linking agent will correspond to hydrogel precursor compound, but can change to control the cross-linking speed and the level realized in gained cross-linked matrix.Usually, according to selected cross-linking agent and polymer system, for forming the cross-linking amount range of microparticle will be about 0.5mM to about 30mM, preferably about 1mM to about 20mM, more preferably about 2mM to about 15mM, even more preferably about 2.5mM to about 10mM, and even more preferably about 2.5mM to about 5mM.

[0057] Suitable crosslinking agents include light or heat initiated crosslinkers, chemical crosslinkers such as acrylates, methacrylates, acrylamides, vinyl sulfones, dithiols, etc., which are included as part of the alginate bath. Self-crosslinking hydrogel precursors can also be used. In these embodiments, a photoinitiator such as 2959 (2-hydroxy-4′-(2-hydroxyethoxy)-2-methylpropiophenone) or lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP) can be included in the hydrogel precursor solution as a catalyst. A photoinitiator can also be included in the alginate bath. For chemical crosslinking systems, the crosslinker is typically provided in the alginate bath. For UV-initiated crosslinking systems, the crosslinker can be included in the alginate bath or hydrogel precursor solution along with the primary polymer (backbone) species.

[0058] Table 1 provides examples of reactive groups from different backbone chemistries that can be used to form hydrogels with appropriate crosslinkers. Some of the reactions are initiated by UV light, while others are chemical reactions.

[0059] Table 1. Reactive groups and crosslinking chemistries

[0060]

[0061] Additional exemplary precursor compounds include but are not limited to non-alginate polysaccharides, collagen / gelatin, chitosan, agarose, etc. These precursor compounds can be branched polymers with multiple arms or unbranched / linear polymer chains of a single main chain. They can be further functionalized by attaching proteins, drugs or enzymes to the hydrogel to achieve control of cell attachment, function, proliferation or viability. Particularly preferred hydrogel precursor compounds are hyaluronic acid, or hyaluronic acid / PEG mixtures. One or more precursor compounds can also be functionalized with various chemical entities added to the precursor compound when the hydrogel is manufactured. These chemicals will be combined with the hydrogel matrix, rather than being encapsulated in the interstitial voids, so as to simply diffuse out the microparticles. Examples will include adding cell support agents that can be functionalized so that they react with cross-linking agents or the hydrogel main chain. In rapidly degradable microparticles, these agents will be locally released together with the cells and cell products encapsulated therein.

[0062] Similarly, preferred hydrogels are slow-gelling hydrogels, which as used herein mean a hydrogel having a gelation rate that is insufficient to form a generally smooth, generally spherical construct when the gel precursor droplets are in contact with a solution comprising a cross-linking agent. Slow-gelling hydrogels are generally considered to be those that do not form gel droplets immediately (or nearly immediately, for example, in about 5-10 seconds) when exposed to a cross-linking agent. Biocompatible hydrogels are also particularly preferred, depending on the specified end use of the hydrogel. As used herein, "biocompatibility" means that it is harmless to the living tissue or cells encapsulated in the matrix, and more particularly, it is biologically or otherwise not undesirable in that it can be administered to a subject without excessive toxicity, irritation, or allergic or immunogenic reactions, and does not cause any undesirable biological effects or interact in a harmful manner with any other component of the composition comprising it. As will be well known to those skilled in the art, biocompatible hydrogels will be selected to minimize any adverse side effects in the subject. Additional optional ingredients that may be included with the hydrogel precursor include fibronectin, laminin, collagen, other components of the extracellular matrix, and the like, including synthetic forms thereof.

[0063] Generally, techniques for preparing microparticles are described in detail in U.S. Patent No. 9,642,814, filed June 3, 2015, which is incorporated herein by reference. The method includes preparing a hydrogel precursor solution comprising a hydrogel precursor compound, cells to be encapsulated, and divalent cations (e.g., calcium, barium, strontium, and combinations thereof) dispersed or dissolved in a solvent system. The divalent cations are dispersed or dissolved in the solvent system along with the hydrogel precursor compound. The divalent cations should be included in the solution at a level of about 0.025 mol / L to about 0.25 mol / L, based on the total volume of the solution being 100%.

[0064] The hydrogel precursor solution may also include optional hydrogel cross-linking agents, catalysts, additives, culture media, nutrients, pH buffers, density modifiers, viscosity modifiers, etc. The hydrogel precursor solution is then combined with alginate to initiate gelation of the alginate around the hydrogel precursor solution (via an "inside out" gelation process) to produce core / shell microparticles. Each core / shell microparticle includes an alginate shell surrounding a liquid core, which includes the hydrogel precursor solution. Figure 2 As shown, this generally involves (A) adding a precursor solution dropwise to (B) an alginate bath, such as by producing / extruding a precursor solution droplet that is dripped or sprayed into an alginate bath. The amount of alginate in the solution can vary, but can range from about 0.1% to about 2.0% weight / volume, based on taking the total volume of the solution as 100%. Typically, the viscosity of the alginate solution should be less than the viscosity of the hydrogel precursor solution. The viscosity of the alginate solution depends on the alginate concentration and the average molecular weight of the alginate polymer (i.e., the length of the alginate molecule or the number of monomer units in the chain). At similar concentrations, the longer the chain, the higher the viscosity. In one or more embodiments, the viscosity of the alginate solution will range from about 1 to about 20 cP at room temperature (~20 to 25 ° C), and preferably from about 1 to about 4 cP. More particularly, at room temperature, the ratio of the viscosity of the hydrogel precursor solution to the viscosity of the alginate solution should be greater than 1. In another embodiment, the ratio of the viscosity of the hydrogel precursor solution to the viscosity of the alginate solution is from about 1:1 to about 1000:1. In one or more embodiments, the viscosity ratio of the hydrogel precursor is about 20:1. In one or more embodiments, the viscosity of the hydrogel precursor solution is about 1 to about 500 cP, with about 40 to about 100 cP at room temperature being particularly preferred. The pH range of the alginate bath should be about 6.2 to about 7.8, and preferably about 6.6 to about 7.4.

[0065] As shown in (B) and (C), the alginate shell is formed from the inside out and thickens around the droplet as cations leach out from the precursor solution droplet. In other words, the presence of cations in the droplet causes the alginate in the bath to coalesce to the surface and crosslink around the droplet. Next, the gelation of the hydrogel precursor compound in the liquid core is initiated, such as by crosslinking and / or polymerization to produce core / shell crosslinked microparticles. Each core / shell crosslinked microparticle includes an alginate shell and a gel core, which includes a crosslinked 3D hydrogel matrix and embedded cells. Crosslinking can be chemically induced, thermally induced or photoinduced, depending on the specific precursor solution prepared. For example, a hydrogel crosslinker can be included in an alginate bath and diffused through the alginate shell to crosslink the hydrogel microparticles in the core. Alternatively, the core / shell microparticles can be subjected to a UV radiation source to initiate crosslinking in the hydrogel microparticle core. The UV exposure time range is from about 1 min. to about 10 min., preferably from about 2 min. to about 8 min., and even more preferably from about 2.5 min. to about 5 min. The wavelength of the UV exposure will depend on the photoinitiator and / or polymer system selected, but will generally range from about 100 nm to about 400 nm, preferably from about 315 nm to about 400 nm, and more preferably about 365 nm.

[0066] Unlike the ionic bonds of alginate-type microparticles, the microparticles of the present invention involve covalent cross-linking of hydrogel precursor compounds, e.g. Figure 1 The PEGDA example shows the diacrylate chemistry at the end of the PEG backbone (black line), with the reactive sites represented by the double red lines. The PEGDA molecules crosslink to each other to produce the final hydrogel formulation. In contrast, the MeHA backbone (blue line) with reactive groups shown as dispersed on the chain (red double line) is covalently crosslinked with PEGDA ( Figure 1 ). Therefore, it should be understood that the precursor compounds in the final product are chemically altered by the cross-linking reaction to produce a cross-linked hydrogel matrix. For example, in one or more embodiments, the hydrogel matrix is essentially composed of (or primarily composed of) a cross-linked network of PEG and HA polymers containing thioether and ester cross-links (e.g., for thiolated HA). However, it should be understood that the matrix may include trace amounts of unreacted precursor compounds, functional groups, etc. that remain in the network. Table 1 provides other examples of possible hydrogel matrices and cross-links. This list provides examples, but is not exhaustive. Examples of backbone molecules that can be used for the chemicals listed in Table 1 include chitosan, agarose, chondroitin sulfate, or a combination of molecules. The bonds formed between the backbones include chemical bonds to thioether plus ester, thioether plus ester and methyl on the β carbon, thioether sulfone, or thioether succinimide. However, other chemical bonds can be used. Photo (free radical) bonds can include ester plus ether or amide plus ether. It should be understood that there are increasingly complex combinations that can be utilized, making the possible combinations virtually infinite.

[0067] Once the core has been gelled, the alginate shell is then removed (e.g., with a chelating agent and / or mechanical agitation, such as ultrasonic treatment) to produce self-sustaining hydrogel microparticles or microbeads and cells embedded therein, as illustrated in (D). In other words, the alginate shell is not part of the final product and is always removed before using the microparticles. The resulting hydrogel microparticles can be collected from the solution using a mesh screen or other device and can be rinsed or suspended in a culture medium or appropriate nutrients as needed. Similarly, the resulting degradable hydrogel microparticles are preferably substantially free of alginate.

[0068] This technology can be used for embedding / encapsulating cells, cell clusters, tissues, combinations thereof and fragments thereof in hydrogel microparticles. The cell can be a primary cell of natural origin or fresh separation, a cell cultured or amplified, a cell line set up, differentiated or undifferentiated, engineered or genetically modified etc. The non-limiting examples of cells and tissues include islets of Langerhans, islet clusters, hepatocytes, stem cells and related cells and tissues, and endocrine cells, stem cell clusters, thyroid clusters, adrenal clusters, pituitary clusters and other 3-dimensional cell clusters for tissue engineering or cell-based treatment. The combination of cell types and / or tissues (for example, the primary cell types combined with stem cells) can also be used in microparticles. Depending on the cell type, microparticles can also be included in the hydrogel matrix component of the cell culture medium that supports cell maintenance and growth. Exemplary culture medium components include one or more of the following: serum, nicotinamide, antibiotics (penicillin, amphotericin B, streptomycin, gentamicin sulfate), amino acids (alanine, arginine, aspartic acid, glutamine, etc.), pH buffer (sodium bicarbonate), inorganic salts (sodium, potassium, magnesium and calcium ion sources), carbon sources (glucose, galactose, fructose, maltose, sodium pyruvate), proteins and peptides (albumin, transferrin, fibronectin, activin, insulin), fatty acids and lipids, vitamins (A, D, E, K, B, nicotinamide), minerals and trace elements (zinc, copper, selenium), hormones, growth factors, etc. Complete culture media, basal culture media, and culture medium components are commercially available and include, but are not limited to, Eagle's Minimum Essential Medium (EMEM), Dulbecco's Modified Eagle's Medium (DMEM), RPMI-1640, Ham's Nutrient Mixtures, Iscove's Modified Dulbecco's Medium (IMDM), CMRL-1066, and the like.

[0069] The characteristic of the obtained hydrogel matrix is a semi-rigid network that is permeable to liquids and gases, but it shows as not flowing and maintains its integrity under steady state. Hydrogel can be considered as a kind of solid and can also be described as having a viscosity close to infinity. Hydrogel matrix is a 3-dimensional self-sustaining body. The term "self-sustaining body" means that once the hydrogel matrix is formed, it does not need an external support structure to maintain its shape, and is not easily deformed only due to its own internal force or weight. The self-sustaining body is not as flexible, permanently deformable or flowable as jelly, putty or paste, but is elastic so that the matrix body can temporarily yield or deform under force. In other words, the self-sustaining body will retract or rebound to its original shape after slight compression and / or bending-it should be understood that the hydrogel matrix can split, break or shear (and will not recover afterwards) when enough external pressure or force is applied.

[0070] Hydrogel microparticles are a type of matrix capsule that holds the filler material throughout the bead, rather than having different shells like core-shell capsules. As described above, hydrogel microparticles are also self-sustaining. In one or more embodiments, the shape of the resulting microparticles is substantially spherical. Advantageously, the particle size is highly customizable, depending on the capabilities of the selected droplet generator. In one or more embodiments, the resulting hydrogel microspheres or microparticles have an average (average) maximum cross-sectional surface specific surface size (i.e., in the case of spherical or ellipsoidal microspheres, their diameter) greater than 30 μm, and in some cases greater than 300 μm. In one or more embodiments, the resulting hydrogel microspheres or microparticles have an average (average) maximum surface specific surface size of less than about 5 mm. Preferably, the resulting hydrogel microspheres or microparticles have an average (average) maximum surface specific surface size of less than about 2 mm, more preferably from about 30 μm to about 2 mm, even more preferably in the range of from about 50 μm to about 1.5 mm, more preferably from about 150 μm to about 1.5 mm, even more preferably from about 300 μm to about 1.4 mm. In some cases, smaller microparticles ranging in size from about 30 μm to about 750 μm or 500 μm can be formed. For ease of reference, this cross-sectional dimension is referred to herein simply as the "size" of the microparticle. In any case, the hydrogel microparticles of the present invention are not nanoscale and are not considered to be nanoparticles or any other type of nanocrystalline shape.

[0071] Hydrogel microparticles are particularly suitable for local therapeutic treatment of acute or chronic conditions, wherein short-term treatment regimens can be used to alleviate such conditions.Examples include local delivery (e.g., direct injection) of cells, cell clusters and / or tissues for bone or cartilage repair (e.g., in joints) and the general protection of the therapeutic cells, cell clusters and / or tissues of transplantation from the host immune system.Microparticles protect encapsulated cells, cell clusters and / or tissues from the immune system in a short period of time (e.g., a few days, a few weeks, up to three months), releasing their therapeutic benefits until the hydrogel is decomposed and washed away by normal physiological processes. Systemic treatment is also envisioned. In one or more embodiments, microparticles are not used to deliver small molecule therapeutic agents and are preferably free of such medicines, unless such medicines are included in microparticles as cell support active agents (e.g., to ensure that cells and implants are successful, such as immunosuppressants), but not as the extent of active agents or therapeutic agents themselves.

[0072] For therapeutic methods, the microparticles are suspended or dispersed in a suitable delivery vehicle for administration to the subject. Exemplary delivery vehicles will include biocompatible liquid suspensions, viscous solutions, putties, pastes or gels in which the microparticles are distributed, and preferably include nutrients and components that support the maintenance and growth of cells, cell clusters and / or tissues to be administered. Cell culture media as described above that can be further supplemented according to cell or tissue type are preferred carriers. Saline solutions or other buffered solutions can also be used as delivery vehicles. The method generally includes (or is composed of) the following: administering a therapeutically effective amount of the obtained microparticle composition to the patient locally, such as administering a therapeutically effective amount of the obtained microparticle composition to the position of the patient's inflammation, injury, arthritis, degeneration, etc. Administration generally includes direct injection of the microparticle composition at or near the site of inflammation, injury, arthritis, degeneration, etc. Advantageously, the sustained-release microparticle composition limits the administered cells, cell clusters and / or tissues to a local area for a sustained therapeutically effective period of time, so that the slowly released cells, cell clusters and / or tissues and / or their cellular components or products are released in the local area to alleviate or mitigate the severity of the patient's condition. As used herein, the term "therapeutically effective" refers to an amount and / or time period that will elicit the biological or medical response of a tissue, system, animal, or human being that is being sought by the researcher or clinician, and in particular, an amount and / or time period that elicits some desired therapeutic effect. For example, in one or more embodiments, a therapeutically effective amount and time period are those that reduce inflammation and induce or promote healing of sites of inflammation, injury, arthritis, degeneration, and the like. Those skilled in the art recognize that an amount or time period may be considered therapeutically effective even if the condition is not completely eradicated but only partially improved.

[0073] If necessary, the treatment can be repeated by additional injections or infusions. It will be understood by those skilled in the art that treatment regimens can vary according to the specific inflammation, injury, arthritis, degeneration, illness or healing state and the preferences of the medical or veterinary practitioner or researcher. For example, neuronal repair, such as the repair of spinal cord injury or degenerative disease, can be potentially treated by injecting microencapsulated stem cells into the epidural space or epidural lesion area. Degradable hydrogel microparticles are intended to release stem cells into the area immediately, allowing stem cells to interact directly with the environment, or degradable microparticles designed to last longer (e.g., ~1-3 months) can microencapsulate stem cells to release growth-promoting factors into the area while slowly degrading away, releasing cells for destruction and degradation until nothing is left. Similarly, cardiac repair after a heart attack can be improved by locally releasing cells or growth factors from cells, cell clusters and / or tissues that are microencapsulated and placed in the damaged area. Another example includes treating a non-functional gland by hormone therapy, such as restoring thyroid function (if scar tissue is present after the thyroid is removed or reduced) by placing microencapsulated thyroid cells, cell clusters and / or tissues in the neck or nearby / adjacent areas. Additional examples of durable hydrogel microparticles include local delivery of neurotransmitters associated with neuroendocrine disorders, release of bone-promoting factors near non-healing fractures, systemic release of molecules such as atrial natriuretic peptide to treat congestive heart failure, release of coagulation factors in the treatment of hemophilia, or encapsulation of engineered or transgenic cells that produce various bioactive substances.

[0074] In one or more embodiments, "empty" hydrogel microparticles can also be administered as part of a joint lubrication treatment. In other words, the presence of the hydrogel itself provides therapeutic relief to the patient without the need for cells or any other active agents within the particle itself.

[0075] A particular advantage of the compositions and microparticles of the present invention when implanted is their improved "stickiness" compared to alginate or other types of microparticles. Specifically, the microparticles have a tendency to adhere or attach to tissues in the body at the site of implantation, as demonstrated in the photographs in the following examples. This further enhances the effectiveness of the treatment by maintaining localized release of cells, cell clusters, and / or tissues and their cellular products at the implantation site. Furthermore, the microparticles demonstrated no inflammatory response and / or foreign body reaction, including lymphocyte or collagen ring formation, at the implantation site.

[0076] After reviewing the disclosure herein and the working examples below, additional advantages of the various embodiments of the present invention will be apparent to those skilled in the art. It should be understood that, unless otherwise indicated herein, the various embodiments described herein are not necessarily mutually exclusive. For example, features described or depicted in one embodiment may also be included in other embodiments, but not necessarily. Therefore, the present invention encompasses various combinations and / or combinations of the specific embodiments described herein.

[0077] As used herein, the phrase "and / or," when used in the context of a list of two or more items, means that any one of the listed items may be employed alone or in any combination of two or more of the listed items may be employed. For example, if a composition is described as comprising or excluding components A, B, and / or C, the composition may comprise or exclude: A alone; B alone; C alone; A and B combined; A and C combined; B and C combined; or A, B, and C combined.

[0078] This specification also uses numerical ranges to quantify certain parameters associated with various embodiments of the present invention. It should be understood that when a numerical range is provided, such a range will be interpreted as providing literal support for the claimed limitation of only the lower limit of the range and the claimed limitation of only the upper limit of the range. For example, a disclosed numerical range of about 10 to about 100 provides literal support for the claim of listing "greater than about 10" (no upper limit) and the claim of listing "less than about 100" (no lower limit).

[0079] Example

[0080] The following examples illustrate the process according to the present invention. However, it should be understood that these examples are provided by way of illustration and nothing in them should be considered as limiting the overall scope of the present invention.

[0081] We show that one or more HA formulations used in the examples degraded in less than 3 months. In contrast, PEGDA formulations were more durable. We compared the properties of products made in this manner by comparing degradable microparticles to durable microparticles. These comparisons included physical and chemical differences, in vitro and in vivo stability, cell morphology, function, and migration within the final product microparticles, and the ability of the final product microparticles to attach to tissues in vivo, thereby localizing their application in vivo.

[0082] Example 1

[0083] Durability comparison of physical and chemical differences in the final product of degradable hydrogel microparticles

[0084] introduce

[0085] The concept of cell encapsulation was first popularized by Lim and Sun in 1980, who showed that pancreatic islets embedded in alginate hydrogel microparticles could reverse diabetes in rats without the need for immunosuppression, albeit only for a few weeks. This initial success was followed by research aimed at better understanding and improving the process. Over time, the field has made great progress, however, it is well known that alginate microspheres continue to suffer from long-term biocompatibility issues. Despite its limitations, alginate continues to be the obvious material choice for cell encapsulation because its unique, near-instantaneous cross-linking kinetics enable the direct manufacture of convenient injectable microparticles.

[0086] In contrast, other hydrogels can usually only be prepared into bulk macrostructures due to their slow cross-linking rates. Some approaches have been developed to produce microspheres using alternative hydrogels such as polyethylene glycol, agarose, chitosan, or hyaluronic acid. However, these methods are based on oil emulsion technology and are generally not suitable for cell encapsulation and transplantation due to their reliance on non-aqueous solvents and poor process control.

[0087] As the number and complexity of cell encapsulation applications continue to grow, the ability to manufacture convenient, injectable, biocompatible microparticles using advanced biomaterials will be decisively beneficial. Here, a new method for producing hydrogel microparticles, termed core-shell spherification (CSS), described in U.S. Patent No. 9,642,814, filed on June 3, 2015 and incorporated herein by reference, was tested. This method is strategically designed to utilize hydrogels that gel much more slowly than alginate, enabling the production of hydrogel microparticles with a variety of chemical, physical, and biologically active properties. Furthermore, rather than using oil emulsion technology, the method was developed for standard GMP-ready equipment and materials to better facilitate accessibility, scalability, and regulatory compliance. We demonstrate microparticle production via CSS using two popular hydrogels: hyaluronic acid (HA) and polyethylene glycol diacrylate (PEGDA).

[0088] method

[0089] Cytotoxicity of the CSS cross-linking procedure

[0090] The cytotoxicity of calcium (a component in the CSS method) was tested by exposing canine islets to a solution comprising 100mM or 200mM calcium chloride, 10mM HEPES buffer and 20% PEGDA, MW 3,400Da (Laysan Bio, Inc.) for 5, 10 or 15 minutes. The group was also exposed to 100mM calcium chloride without PEGDA to assess the effect of calcium alone. Islets were suspended in the test solution at approximately 5,000 IEQ / mL to simulate the high cell loading density associated with the encapsulation process. After calcium exposure, the islets were washed twice with supplemented CMRL 1066 islet culture medium and then incubated at 37°C and 5% CO2 for 3 hours before evaluation.

[0091] Dead cells were identified using propidium iodide staining and fluorescence microscopy. Fluorescence micrographs were captured using Cytation 5 Imaging Multi-Mode Reader (Biotek Instruments, Inc, 531 / 647nm ex / em). Cell death was quantified by calculating the ratio of red (dead) pixels to total islet pixels using Adobe Photoshop. Twenty-five individual islets were analyzed per group and distinguished from non-islet tissues by disulfide hydrazone staining. The results were reported as average viable cell fractions and normalized to the islet viability of the untreated control (i.e., islets from the same batch in culture medium) of coupling.

[0092] For PEGDA and methacrylated hyaluronic acid (MeHA) microparticles, photocrosslinking was used. Therefore, studies were first conducted to examine the toxicity of ultraviolet (UV) light and photoinitiators. Canine pancreatic islets were suspended in a 1% slurry containing 0, 0.025, or 0.05% (w / v) photoinitiator. 2959 in Dulbecco's phosphate buffered saline (DPBS). Approximately 50 IEQ were loaded into the wells of a 24-well plate and irradiated with long-wave ultraviolet (UV) light for 3, 5, or 10 minutes. The plates were irradiated using a PortaRay 400 UV lamp in low power mode at a distance of 6 ", corresponding to approximately 40 mW / cm according to the manufacturer's data (Uvitron International). 2 After exposure, islets were washed twice with supplemented CMRL 1066 islet culture medium and incubated for 3 hours at 37°C and 5% CO2 before evaluation. Cell death was assessed by propidium iodide staining and quantified as described in the previous section.

[0093] Fabrication of MeHA hydrogel microparticles

[0094] Once non-toxic levels of cross-linker are established, microparticle production begins. Figure 2 A general schematic diagram of the CSS method for hydrogel microparticle fabrication is provided. MeHA was first synthesized by reacting HA (MW 1 MDa, Lifecore Biomedical) with a 50-fold molar excess of glycidyl methacrylate (Sigma) in a 50:50 water:DMSO mixture in the presence of triethylamine and tert-butylammonium bromide (Sigma) for 5 days. The MeHA was then dialyzed against deionized (DI) water for 2 days and then lyophilized. 1 The degree of methacrylation was determined to be 54-72% by H NMR (Avance AV-III 500, Bruker) by calculating the ratio of the relative peak areas of methacrylate protons to methyl protons.

[0095] In a solution containing 100 mM calcium chloride, 15 mM HEPES and 0.05% (w / v) OptiPrep A 2.5%:1% (w / w) methacrylated HA:PEGDA polymer blend (MW 3.4 kDa, Laysan Bio, Inc.) was prepared in a custom buffer (CosmoBio USA, Inc.) with 2959 HCl. In this embodiment, a low concentration of PEGDA was added to the MeHA solution as a crosslinker to increase the crosslink density of the hydrogel. The solution was passed through a 0.22 μm filter. The viscosity of the precursor solution was measured at room temperature using a Cannon-Manning semi-micro calibrated glass capillary viscometer.

[0096] The precursors were extruded by an automated droplet generator into a stirred bath containing 300 mM mannitol, 0.1% Tween 20, 0.15% (w / v) sodium alginate (Protanal LF 10 / 60, FMC Corp.), and adjusted to pH 7.6 using a custom 15 mM HEPES buffer. For the MeHA precursor, 0.05% (w / v) 2959 Initiator is added to the alginate bath to maintain balance between the droplet and the bath, and minimize Once the droplets are dropped into the bath, leaching (which will reduce cross-linking) is performed. A Buchi 395-Pro encapsulator (Buchi Corporation, Newcastle, Delaware (DE)) equipped with an air jet nozzle system and an internal fluid nozzle of 400 microns in diameter in a 1.5 mm concentric air nozzle is used for droplet generation. Compressed nitrogen is used to extrude the droplets. The MeHA droplets form a core-shell construct with a respective alginate shell encapsulating the hydrogel precursor solution / droplet core. The hydrogel precursor solution / droplet core is then irradiated with long-wave UV light to initiate free radical photocrosslinking of MeHA and PEGDA. The irradiance at the bath center is approximately 40 mW / cm 2 The alginate shell was then removed from the core-shell structure by rinsing the core-shell structure in 25 mM citrate buffer (DPBS) for 5 minutes under gentle stirring. The resulting microparticles were collected using a steel mesh sieve and suspended in DPBS. A second rinse in 50 mM citrate was continued for 5 minutes to ensure complete dissolution and removal of the alginate shell.

[0097] Fabrication of ThHA hydrogel microparticles

[0098] The ThHA precursor was prepared by dissolving 1.2% (w / w) of thiolated HA (HyStem, Biotime Inc) in a custom buffer (CosmoBio USA, Inc) containing 100mM calcium chloride and OptiPrep to adjust the solution density. The pH of the precursor solution was adjusted to approximately 7.0 using a custom 15mM HEPES buffer to reduce the disulfide bridging of the ThHA macromer. The solution viscosity was measured as described above. The core-shell construct was produced as described above. However, cross-linking was achieved by chemical cross-linking (compared to free radical photocrosslinking using MeHA microparticles). The terminal acrylate group of the PEGDA molecule was bonded to the thiol group on the ThHA by Michael-type addition. In short, the precursor solution produced a core-shell construct with a ThHA precursor core and an alginate shell after contact with an alginate bath. The alginate bath also contained 0.4% PEGDA as a cross-linking agent (MW 3.4kDa, Laysan Bio, Inc) and was gently stirred for 5 minutes. Bath and hydrogel precursor both can comprise cross-linking agent (for example, PEGDA), but in this scheme, it is preferably only included in the bath. Cross-linking is carried out by diffusing into the core by smaller 3.4kDa PEGDA molecules through alginate shell. The bath is then diluted half with DPBS, which reduces the solution pH to 7.4. Continue stirring for an additional 30 minutes with the HA precursor in the cross-linked core. Collect the core-shell construct, and remove the alginate shell as described above to produce microparticles.

[0099] Fabrication of PEGDA hydrogel microparticles

[0100] PEGDA hydrogel precursor solution was prepared by dissolving PEGDA 3.4kDA and 20kDa (Laysan Bio, Inc.) at 18% and 12% (w / w), respectively, in a 100 mM solution containing 10 mM calcium chloride, 10 mM HEPES, and 0.025% (w / v) The solution was filtered using a 0.22 micron syringe and the viscosity was measured as described above. For MeHA microparticles, the precursor was extruded into a stirred alginate bath, but with 0.025% (w / v) All solutions used for MeHA and PEGDA microparticle production were prepared in water degassed by sonication to eliminate excess oxygen (a known inhibitor of photocrosslinking). The core-shell constructs were irradiated as described above to achieve crosslinking of the core. The constructs were further processed as described above to dissolve the alginate shell and produce hydrogel microparticles.

[0101] Fabrication of AHA hydrogel microparticles

[0102] First, AHA was synthesized by reacting HA (MW 1 MDa or 200 kDa, Lifecore Biomedical) with acryloyl chloride and glycidol (TCI) in the presence of triethylamine and N-dimethylformamide (Sigma) for 5 days. AHA was then dialyzed against deionized (DI) water for 5 days and then lyophilized. 1 The degree of acrylation was determined to be 53-72% by H NMR (Avance AV-III 500, Bruker) by calculating the relative peak area ratio of acylated protons to methyl protons. AHA precursor solution was then prepared by dissolving 4% acrylated HA in a solution containing 100 mM calcium chloride, 15 mM HEPES buffer and 0.05% (w / v) The precursor solution was filtered using a 0.22 micron syringe and the viscosity was measured as described above. For MeHA particles, the precursor was extruded into a stirred alginate bath. The core-shell construct was irradiated as described above to achieve core crosslinking. The alginate shell was removed using citrate buffer as described above to collect the AHA hydrogel microparticles.

[0103] Physical properties and size distribution of hydrogel microparticles

[0104] Images of 100 individual microparticles were captured using a Biotek Cytation 5 multimode imaging reader. The microparticles were imaged in a multiwell plate in PBS, and the images were analyzed with Adobe Photoshop to determine the average microparticle diameter and size distribution of each microparticle type (N = 100 microparticles). The hydrogel microparticles were further characterized by determining the swelling ratio "Q" (i.e., the ratio of the hydrated mass of the swelling to the dry mass) of each microparticle type. The microparticles were immersed in excess DI water for 24 hours before measurement to remove dissolved salts and ensure that equilibrium swelling was reached. Excess surface moisture was removed from the microparticles, and then weighed on a pre-weighed watch glass. Subsequently, the microparticles were dried overnight at 60°C and reweighed to obtain dry mass (N = 3).

[0105] Diffusion characteristics of particles

[0106] Following previously published procedures, hydrogel microparticles were incubated overnight in 0.1 mg / mL FITC-labeled dextran in DPBS with molecular weights of 10, 40, 70, and 500 kDa (Invitrogen Molecular Probes). The microparticles were rinsed with DPBS and imaged by laser scanning confocal microscopy (Olympus Fluoview 300) to monitor the outflow of the probe. Micrographs were captured at set time points between 3 and 30 minutes after removal of the microparticles from the FITC-dextran incubation solution.

[0107] Aperture calculation

[0108] Empty PEGDA and MeHA microparticles were examined using environmental scanning electron microscopy (ESEM) performed on a FEIQuanta 600F ESEM instrument. The samples were immersed in water on a cooling stage and the temperature in the SEM chamber was set to 6°C. The chamber was then evacuated to 1000 Pa and the relative humidity was maintained at 100%. Images were captured at a magnification of 10,000 times. Standard SEM images were also collected. For these images, the samples were critical point dried or freeze-dried and coated with 10nm of platinum and imaged at 5kV in high vacuum mode on the same ESEM instrument. SEM images were captured at a magnification of 1000 times.

[0109] result

[0110] Core-shell spheroidization procedure

[0111] Hydrogel microparticles are Figure 2 The CSS method summarized in

[15] starts with a slowly hardening hydrogel precursor solution (functionalized PEG or HA) that is embedded in the alginate shell as a droplet when it hits the alginate bath, giving the functionalized PEG or HA time to harden. The alginate shell formed around the core microparticle is characterized by a concentric ring-linked morphology ( Figure 3 A). Subsequently, the alginate shell was removed, leaving only the inner hydrogel core ( Figure 3 B). Figure 3 B shows a predominantly spherical shape of the microparticles. However, other shapes are acceptable for the final product. Figure 4 Additional examples of acceptable end-product characteristics are provided. Figure 4 A provides an example of a more durable (aka long-lasting) PEGDA-based microparticle with a slower degradation profile, loaded with canine islet cells, and having a completely spherical shape. Figure 4 B shows PEGDA-based microparticles of similar durability loaded with islets but with an elongated shape; a 3D ellipsoid shape. Figure 4 C illustrates faster degrading (aka short-lived) MeHA-based microparticles with a faster degradation profile, loaded with stem cells, and with a flattened edge on the left. Figure 4 D shows the same type of particle, but with a 3D ellipsoidal shape. All of these particles, as well as other related shapes such as teardrop or egg shapes, are functional examples of CSS particle end products.

[0112] Cytotoxicity studies

[0113] During microparticle fabrication, while the alginate shell remains, gelation of the core requires crosslinking so that it maintains a 3-dimensional, self-sustaining structure after shell removal. To optimize the CSS method, the cytotoxic effects of the crosslinking component of the microparticle fabrication process on freshly dispersed cells (canine pancreatic islets) were evaluated. Figure 5 A depicts the impact of calcium concentration and exposure time on canine islet viability. The three conditions tested were 100 and 200mM calcium with PEGDA or only 100mM calcium. Under the shortest measurement exposure time of 5 minutes, no significant differences were observed in the cytotoxicity of any group. Compared with the untreated control (islets from the same batch in the culture medium) of the matching, the viability of all groups exceeded 97.5%. However, compared with the 5-minute measurement, at each subsequent time point, the viability of the 200mM calcium group was seen to decline significantly. At both 10 minutes and 15 minutes, the viability of the 200mM calcium group was significantly lower than that of the two 100mM groups. Within 10 minutes, no difference was observed in any 100mM group. However, at 15 minutes, compared with 5 minutes and 10 minutes time points and with the 100mM group without PEGDA at 15 minutes, the viability of the 100mM group comprising PEGDA was shown to decrease significantly. Interestingly, no significant changes in viability were observed at any time point within the 100 mM group without PEGDA, with viability at 97.6% at 15 minutes compared to the control.

[0114] Effects of photoinitiator concentration and UV light exposure time on cell viability Figure 5 B. Canine pancreatic islets were incubated at concentrations of 0, 0.025, and 0.05% (w / v) 2959 ~ 40mW / cm 2 Long-wave UV exposure was performed for 3, 5, and 10 minutes, and cytotoxicity was assessed. Interestingly, the absence of a photoinitiator ( The group containing 2959) showed the highest degree of cytotoxicity. Compared with the 3-minute exposure time, the viability of the 0% group was significantly reduced at 5 and 10 minutes. The 0% group had significantly lower viability at 5 and 10 minutes compared to the 0.025% group. The only significant difference between the 0.025% and 0.05% groups was at 10 minutes, with viable cell fractions of 94.5% and 87.2%, respectively.

[0115] Physical properties and size distribution of hydrogel microparticles

[0116] To fully test the new CSS platform, microparticles were produced from two different starting materials: HA and PEGDA. HA production was further divided into thiolated HA (ThHA) and methacrylated HA (MeHA) to evaluate both chemically cross-linked and photocross-linked hydrogels using the CSS platform. Table 2 provides the physical properties of the hydrogel formulations before and after cross-linking, as well as the average microparticle diameter and diameter range.

[0117] Table 2. Microparticle fabrication characteristics and physical properties

[0118]

[0119]

[0120] 1. The swelling ratio "Q" is the ratio of the hydrated equilibrium mass of the gel to the dry mass.

[0121] 2. The diameters of HA and PEGDA beads were significantly different (p<0.001).

[0122] For PEGDA, the precursor mass fraction of the gel precursor was 30%, while ThHA and MeHA were made from 1.2% and 2.5% solutions, respectively. High PEGDA precursor concentrations are necessary to achieve solution viscosities suitable for forming well-shaped microparticles. The gel formulations and corresponding precursor viscosities are shown in Table 2. ThHA was the only chemically cross-linked hydrogel, and it required significantly more time to cross-link (35 minutes) compared to the photo-cross-linked gels (MeHA and PEGDA). The average diameter of the MeHA microparticles was the largest, followed by PEGDA, and then the ThHA microparticles, which were smaller than the other two groups. The PEGDA microparticles exhibited the lowest equilibrium swelling ratio "Q", indicating that the overall hydrogel was more compact. This is in Figure 6 Further illustrated in Figure 1, which plots the diameters of 100 representative microparticles composed of ThHA, MeHA, PEGDA, or AHA, and illustrates the good monodispersity of the microparticles, especially for the HA material. Despite being manufactured using the same droplet generation equipment and parameters before cross-linking, the ThHA microparticles were much smaller in size than PEGDA or MeHA and had a narrower diameter range than PEGDA. The average MeHA and AHA microparticle diameters were the highest among the four groups. Interestingly, the PEGDA particles containing islets exhibited a slightly larger average diameter but improved monodispersity compared to the empty PEGDA microparticles (938 μm [CV: 8.4%] for PEGDA with islets and 904 μm [CV: 15.7%] for empty PEGDA microparticles).

[0123] Diffusion characteristics of particles

[0124] Figure 7The diffusion of fluorescently labeled dextran into PEGDA, ThHA, and MeHA hydrogel microparticles is shown. The microparticles were incubated overnight in fluorescent dextran of different molecular weights, rinsed, and immediately examined by confocal microscopy to assess the extent of dextran penetration and the efflux rate as a measure of diffusion properties. The 10kDa probe was able to penetrate all microparticles. However, the fluorescence in the PEGDA microparticles was weak, especially near the center of the construct, indicating a lower probe penetration rate compared to the HA microparticles (both of which showed strong fluorescence). In addition, the fluorescence of the 10kDa probe rapidly weakened in the HA group and was absent or almost absent 30 minutes after rinsing. In contrast, localized fluorescence was still observed in the PEGDA group at 30 minutes, again indicating a higher diffusion barrier for the 10kDa probe compared to the two HA gels.

[0125] Low levels of fluorescence from the 40 kDa probe were detected at the periphery of the PEGDA microparticles, but almost no signal was observed in the center, indicating a diffusion barrier to the probe ( Figure 7 Larger dextrans (70 and 500 kDa) showed almost no fluorescent signal in the PEGDA microparticles, indicating that the penetration of these probes into the gel matrix during the overnight incubation period was negligible.

[0126] Both ThHA and MeHA microparticles showed strong fluorescence from all probes evaluated, indicating that the diffusion barrier is minimal even for large molecules. In addition, the relative fluorescence signal in MeHA appears significantly higher for all probes immediately after washing compared to ThHA, indicating higher probe concentrations and, therefore, even lower diffusion barriers. This is best exemplified by the 70 kDa probe, which has moderate infusion into ThHA microparticles but appears to be saturated in MeHA microparticles. All additional studies focused on the hydrogel formulations with the largest and smallest diffusion barriers, PEGDA and MeHA, respectively.

[0127] discuss

[0128] We have developed and evaluated a new method for manufacturing hydrogel microparticles for cell encapsulation and delivery. Current methods for cell microencapsulation are primarily based on alginate spheres due to their rapid gelation rates. Other hydrogels used for microencapsulation involve demanding oil emulsion techniques. In contrast, the CSS method can be used with a variety of hydrogel materials, is compatible with standard, commercially available, and GMP-ready equipment, and exhibits minimal cytotoxicity to pancreatic islet cells.

[0129] Low cytotoxicity is a key feature of cell encapsulation strategies. Therefore, we evaluated the effects of some key enabling components of the CSS approach, specifically calcium content and UV exposure, on islet cytotoxicity. Not surprisingly, 200 mM calcium was rapidly cytotoxic. However, we found that even in highly concentrated and viscous PEGDA solutions, islets tolerated 100 mM calcium for at least 10 minutes. Exposure to 40 mW / cm 2 Long-wavelength (~365 nm) UV light for up to 10 minutes was the least toxic to canine islet cells. Unexpectedly, the group without a photoinitiator showed the most cell death after UV exposure, which is in contrast to other similar studies. However, these other studies used much lower UV intensities, between 4-10 mW / cm 2 and the photoinitiator concentration used in this paper is relatively low. Therefore, in our experiments, the photoinitiator can provide protection by preferentially absorbing UV photons. A concentration-dependent increase in cytotoxicity was still observed in the 2959 group, which is consistent with previously published works.

[0130] Despite being manufactured using the same 400-micron nozzle system, the sizes of the microparticles varied significantly, with the diameter of the MeHA beads being almost twice that of the ThHA beads, corresponding to an approximately 8-fold difference in the total volume of each microparticle. The size differences between the groups are most likely the result of differences in volume swelling kinetics after fabrication. Hydrogel swelling is controlled by many variables but is strongly correlated with the initial (i.e., pre-crosslinked) concentration of the polymer and the crosslink density within the gel after fabrication. More specifically, solvent molecules (e.g., water) adsorbed onto the gel polymer matrix exert outward pressure, causing the gel to expand (swell), while crosslinks within the gel matrix resist this outward expansion. For example, hydrogels with a high initial polymer concentration and a low crosslink density would be expected to expand significantly in volume after fabrication. Conversely, gels with a high ratio of crosslinks to polymer concentration tend to resist expansion after fabrication. The smaller final size of the ThHA microparticles in our study is likely a result of the low initial polymer fraction (1.2%) combined with the relatively high crosslink density. It is important to note that volume swelling after fabrication should not be confused with the swelling ratio "Q" (hydrated mass / dry mass). In particular, Q describes the final gel structure after volume swelling has reached equilibrium, where lower Q values generally indicate stronger and more compact gels. Thus, a gel can undergo significant volume swelling and still have a relatively low Q value, which appears to be the case for the PEGDA microparticles in this study. In contrast, ThHA microparticles undergo relatively little volume swelling (if any) but have higher Q values compared to PEGDA microparticles. Furthermore, MeHA microparticles, despite having a low initial polymer concentration (3.5%), swell significantly and have a high Q value, indicating a very low relative crosslink density.

[0131] The diffusion characteristics differed significantly between the microparticle groups. ThHA and MeHA microparticles were more permeable to dextran probes up to 500 kDa in size, with MeHA microparticles appearing to have the highest diffusion rate of all three groups. In contrast, PEGDA particles appeared to strongly restrict the diffusion of dextran with a MW of 40 kDa and larger. Although various factors can affect the diffusivity of the gel, the swelling ratio "Q" (the ratio of the equilibrium hydrated mass of the gel to the dry mass) is strongly correlated with the permeability of the hydrogel. The Q values of PEGDA, ThHA, and MeHA microparticles were 17.3, 27.7, and 105.7, respectively. Therefore, the diffusion behavior of the dextran probe observed in the microparticles was consistent with its corresponding Q value and the measured surface pore size.

[0132] Example 2

[0133] Durability and stability of the final product of rapidly degradable hydrogel microparticles

[0134] introduce

[0135] The description of durable versus rapidly degrading hydrogel microparticles depends on the stability of the physical composition of the microparticles. Both in vitro and in vivo stability studies were conducted to determine whether there were differences between the three formulations (PEGDA, MeHA, and AHA). While the CSS method provides detailed adjustments for hydrogels that degrade at different rates, two examples of durable hydrogels are provided below that can last for months to years compared to hydrogels that degrade more quickly.

[0136] method

[0137] In vitro durability testing

[0138] For MeHA, PEGDA, and AHA, 1.000 grams of hydrated microparticles were collected and placed in a degradation buffer consisting of 10 mM HEPES-buffered phosphate-buffered saline and 1% gentamicin. The degradation solution was maintained at a pH of 7.3 + / - 0.10. The microparticles in solution were stored at 37°C for 1, 3, and 6 weeks, then collected and dried at 60°C overnight to obtain the dry mass after degradation. The dry mass after degradation was compared to the dry mass of the microparticles at day 0 to determine the percent mass loss. In addition, the swelling ratio was measured at each time point to assess gel strength and degradation.

[0139] In vivo durability testing

[0140] Canine pancreatic islet isolation

[0141] Canine islets were isolated from pancreases obtained locally from euthanized donors at a local veterinary clinic with consent from the organ donors. Procurement and digestion protocols have been published in detail previously. After organ removal and cleaning, collagenase digestion was performed using standard published protocols, followed by density gradient purification. Isolated islets were converted to islet equivalents or "IEQs" for quantification purposes by disulfide hydrazone staining according to standard published protocols. Canine islets were cultured in CMRL 1066 supplemented with 10% fetal bovine serum, 2 mM glutamine, 10 mM nicotinamide, and 1% antibiotic-antimycotic solution at 37°C and 5% CO2.

[0142] Encapsulated islet xenotransplantation

[0143] Diabetes was induced in immunodeficient NOD / SCID mice (NOD.CB17-Prkdc) treated with streptozotocin (STZ, 220-250 mg / kg). scid , Jackson Laboratory) to evaluate the functional capacity of canine islets encapsulated by CSS. Six mice received islets encapsulated in PEGDA, two mice received islets in PEGMAL (synthesis procedure is as follows), and four mice received the same dose of islets encapsulated in MeHA. Two mice were transplanted (IP) with an equal dose of unencapsulated islets as a control. At 10 weeks into the study, all mice receiving MeHA-encapsulated islets, along with the unencapsulated islet group, developed hyperglycemia and were terminated. Four mice from the PEGDA group were also terminated at 10 weeks to serve as comparison for subsequent histological studies. The remaining two PEGDA mice were terminated after another 6 weeks. Monitoring of the two PEGMAL mice is ongoing (over 16 weeks after transplantation).

[0144] As described above, the durable PEGDA and fast-degrading MeHA preparations of the microparticles encapsulating pancreatic islets were manufactured, except that canine pancreatic islet cells were mixed into the precursor. According to a scheme similar to the above-mentioned ThHA microparticles, a durable PEGMAL preparation of chemical crosslinking (i.e., non-UV) was prepared. In brief, the hydrogel precursor consisting of 25% (w / w) 8-arm 40kDa PEGMAL (JenKemUSA) and canine pancreatic islets was extruded into a pH 6.6 alginate bath containing 0.25% (m / v) 1kDa SH-PEG-SH crosslinking agent (BiochemPEG) and stirred for 5 minutes. After 5 minutes, the cross-linked alginate bath solution was then diluted to a final pH of ~7.2 with HEPES buffered islet culture medium and stirred for another 30 minutes to fully crosslink the hydrogel core. The alginate shell was then removed as described.

[0145] Microparticles containing islets were administered into the intraperitoneal (IP) space of mice via an 18G catheter. The islet dose was based on a previous ascending dose study showing that approximately 4,000 islet equivalents per mouse resulted in normoglycemia and was consistent with previous publications on canine islet transplantation into mice. Blood glucose levels were measured daily for 14 days after transplantation and every two weeks thereafter in all mice.

[0146] After animal termination (10 and 16 weeks), microparticles containing islets were recovered from the IP space and placed in culture medium for evaluation. Microparticles containing islets were stained with disulfide hydrazone to detect the presence of insulin, and stained with calcein (live cells) and propidium iodide (apoptotic / necrotic cells) to assess cell viability of explanted tissue. Color, bright field, and fluorescence images of the stained microparticles were captured using a Cytation 5 imaging multimode reader (Biotek Instruments, Inc.).

[0147] Before being transferred to PBS, pancreas and tissue samples were removed from the peritoneum with attached microparticles and fixed overnight in 10% neutral buffered formalin. The tissue was embedded in paraffin blocks and sliced appropriately for staining. For hematoxylin / eosin (H&E) staining, the sections were deparaffinized by placing the sections in Clear Rite 3 solution 3 times for 3 minutes each, followed by subsequent dehydration and rehydration steps before staining with hematoxylin and eosin. The sections were rinsed in water and then further rinsed with tap water using Bluing Reagent. Finally, the sections were immersed in 100% anhydrous alcohol and Clear Rite 3 to improve color stability and coverslips were applied accordingly. Slides were examined and images were captured on a BioTekCytation 5 cell imaging multimode reader.

[0148] Encapsulated islet allotransplantation

[0149] Two adult male dogs recently diagnosed with diabetes received canine islet transplants encapsulated with either durable PEDGA or rapidly degrading MeHA (N=1 / group). Canine islets were encapsulated according to the previously described protocol. Each dog received a transplant consisting of approximately 60 mL of loaded microparticles delivered directly into the peritoneum via catheter infusion. Body weight and blood glucose were monitored daily for the first 20 days after transplantation.

[0150] result

[0151] In vitro degradation

[0152] MeHA, PEGDA, and AHA microparticles were placed in HEPES-buffered degradation solution at 37°C for 1, 3, or 6 weeks and then dried to determine the dry polymer mass. After 6 weeks of degradation, it was found that the MeHA microparticles retained only 24% of their starting mass, while PEGDA and AHA retained 37% and 53%, respectively. Figure 8A Although PEGDA loses a lot of mass ( Figure 8C ), but it maintained its gel strength during 6 weeks as evidenced by the unchanged swelling ratio ( Figure 8B ). In contrast, MeHA and AHA microparticles showed a sharp increase in their swelling ratios, indicating a significant decrease in gel strength as the microparticles degraded. The variation in gel strength between PEGDA- and HA-based microparticles can be attributed to both crosslink density and polymer chemistry. For PEGDA microparticles, the polymer chains are tightly held together by multiple covalent bonds, resulting in a higher crosslink density than for HA microparticles. Furthermore, PEGDA polymer chains contain regions of hydrophobic moieties. Both the crosslink density and the hydrophobic regions slow the rate of water diffusion into the microparticles compared to the hydrolytic degradation occurring on the microparticle surface. Naturally, if degradation were limited to the surface, the internal gel network would remain intact and maintain gel strength. However, in MeHA and AHA microparticles, the crosslink density is much lower, and the polymer chemistry results in a more hydrophilic structure than that of PEGDA. The lower crosslink density and hydrophilicity of HA microparticles both promote a higher rate of water diffusion into the microparticles, leading to a bulk degradation mechanism. In bulk degradation, water penetrates deeply into the hydrogel network and causes hydrolytic degradation, which compromises the integrity of the gel network. This type of degradation may explain why the remaining AHA microparticles after 6 weeks of degradation have a higher polymer mass than those of PEGDA. The loss of gel strength over time in HA microparticles indicates that hydrolytic degradation is occurring, but the degradation products have not yet diffused out of the gel network. In this case, the remaining polymer mass does not accurately assess the extent of degradation because, although the polymer is within the gel network, it no longer contributes to hydrogel integrity.

[0153] Function of canine islet-loaded microparticles in diabetic mice

[0154] Canine islets encapsulated in PEGDA or MeHA were transplanted into diabetic NOD / SCID mice. A control group received unencapsulated islets. Figure 9The mean daily blood glucose levels of each group during the initial study group after transplantation are shown. All mice in MeHA, PEGMAL and PEGDA groups were normoglycemic in the second week after transplantation. However, despite reintroducing exogenous insulin treatment, the animals receiving rapidly degraded MeHA microparticles gradually recovered to hyperglycemic levels (starting from about 3 weeks after transplantation). These mice were euthanized in the time of about 10-11 weeks. The mice receiving the islets encapsulated with durable PEGDA and PEGMAL maintained normoglycemia throughout the study period (PEGMAL ongoing, 16 weeks) and had excellent glycemic control in the absence of exogenous insulin.

[0155] In contrast, all mice that received unencapsulated islets failed to achieve sustained normoglycemia at any time point. These mice had an initial average blood glucose of 386 ± 23 mg / dL before transplantation and 494 ± 61 mg / dL on the day of transplantation. After transplantation with the same volume of islets as the other two groups, one animal had only a normal blood glucose reading on day 2, while another mouse showed fluctuations in blood glucose values from day 19 to day 30, but then returned to persistently high blood glucose levels for the remainder of the study.

[0156] After necropsy, microparticles were recovered from the IP space of the mice. The results are shown in Figure 10 In AD, durable PEGDA microparticles containing islets were easily recovered at approximately 10 weeks, but rapidly degrading MeHA microparticles were few in number and difficult to identify. Samples of recovered PEGDA microparticles were found to be randomly dispersed throughout the IP cavity and remained transparent and structurally intact. The recovered PEDGA microparticles showed little sign of degradation (i.e., smooth edges, no significant change in size, and mechanically stable), and the transplanted cells remained contained within the microparticles ( Figure 10 A). In contrast, rapidly degrading MeHA microparticles were misshapen and smaller in diameter, indicating significant hydrogel degradation. The MeHA microparticles did not appear to contain intact islets ( Figure 10 B). At a later time point (16 weeks), durable PEDGA microparticles continued to be readily recoverable at necropsy. Islets within PEGDA microparticles were stained with disulfide hydrazone, calcein, and propidium iodide (PI) to confirm the presence of viable, functional islets.

[0157] PEGDA microparticles recovered approximately 16 weeks after transplantation contained healthy, viable islets identified by deep red disulfide hydrazone staining ( Figure 10 C). The same islets were co-labeled with green calcein, a common indicator of viable cells ( Figure 10D). Some positive PI staining (red; apoptotic / necrotic cells) was observed, but was generally localized in non-islet (i.e., disulfide-negative) tissue that appeared to be adhered to the surface of the recovered microparticles. In contrast, islets (disulfide-positive cells) were almost completely devoid of PI (dead cells) staining ( Figure 10 D).

[0158] Durable PEGDA microparticles were found to adhere to tissues such as the liver and abdominal wall, but generally did not adhere to the lumen and floated freely within the lumen. Histology of the implant site confirmed the location of the PEGDA microparticles. Although the histological processing method destroyed the integrity of the PEGDA microparticles, there was no obvious fibrotic tissue at the site of attachment. Figure 11 The healthy fat cells at the intersection of the two microparticles are shown, with the positions of the microparticles represented by black lines. The blue-grey membrane at the center of each microparticle position is the hydrogel remaining after treatment. In contrast, no gel material was identified in the histological sections of the MeHA group. For the PEGDA group, there was generally no fibrosis at the microparticle-omentum interface. To ensure that the normalization of blood glucose levels in the mice treated with PEGDA was due to transplantation of the graft, rather than endogenous islets remaining in the pancreas after streptozotocin treatment, the presence of islets in the histological sections of the pancreas was examined. Only one islet-like structure was identified in six pancreatic sections of each mouse (results not shown).

[0159] A similar study was completed using canine islets transplanted into two diabetic dogs. Both animals were diabetic before transplantation and did not receive regular exogenous insulin. One dog received canine islets encapsulated in durable PEGDA microparticles, while the other received canine islets encapsulated in a biodegradable MeHA hydrogel. Figure 12 The mean non-fasting blood glucose from measurements taken each morning and afternoon is shown. Dogs that received islets in durable hydrogels slowly returned to lower blood glucose levels within the normal range (solid line) and maintained these normal levels for the duration of the experiment (8 weeks). In contrast, animals that received rapidly degrading hydrogels slowly returned to a fully diabetic state within 20 days after transplantation (dashed line).

[0160] discuss

[0161] Interestingly, while both mouse groups achieved normoglycemia after IP injection of microparticles, unencapsulated islets at the same dose and transplant site did not achieve normoglycemia. Since the animal model included immunocompromised mice, the results suggest that the hydrogel matrix provides a physical support environment that promotes better cell function and survival, especially when transplanted by injection into the IP cavity. This conclusion was further confirmed by the observation that MeHA microparticles degraded over time and simultaneously returned to hyperglycemia. It is further noteworthy that PEGDA and PEGMAL islet microparticles (which are more structurally robust) resulted in superior control of blood glucose levels in the peri-transplant period before the recurrence of hyperglycemia in the MeHA group compared to MeHA transplants.

[0162] Although the dog study included only two animals, the results are important because they contrast the responses of immunocompromised mice with those of immunocompetent dogs. In the presence of an intact canine immune system, the weaker hydrogel (MeHA) degraded faster, leading to a more rapid return to hyperglycemic blood sugar levels. The results do indicate that this particular formulation of degradable hydrogel microparticles has an in vivo persistence of approximately days to weeks, depending on the substance. However, the durable hydrogel microparticles exhibited an in vivo persistence of at least two months.

[0163] Example 3

[0164] Durability versus intra-bead proliferation of rapidly degradable microparticle final products

[0165] introduce

[0166] Manipulation of the physical and chemical properties of cell encapsulants has become a powerful tool for in vitro tissue engineering research. It is known that the mechanical and chemical properties of the microenvironment change the behavior of various cell types, but most notably, stem cells are affected by the mechanical properties of the cell surroundings. The ability of stem cells to continue to proliferate within a hydrogel is an important feature for therapeutic treatment of soft tissue defects caused by trauma or secondary to surgical intervention. For example, muscle or skin wounds would benefit from local delivery microbeads with proliferating stem cells. This would result in a smaller dose of microparticles because the number of cells would continue to increase when in the microparticles. However, not all hydrogels allow cells to proliferate within the microparticles. Here, we show that rapidly degrading hydrogel microparticles allow for robust cell proliferation within the microparticles.

[0167] method

[0168] Human embryonic kidney cells (HEK293) were cultured in DMEM medium with high glucose and fetal bovine serum and maintained in an incubator at 37°C and 5% CO2. When the cells reached an appropriate density, they were separated from the culture flask by adding 0.5% trypsin-EDTA for 2 minutes. After separation, the cells were transferred to a growth medium. In order to test the ability of cells to proliferate in hydrogels, two types of hydrogel polymers, rapidly degradable MeHA and durable PEGDA, were manufactured into microparticles using the above-mentioned CSS manufacturing method. The same amount of polymer precursor (1 mL) was mixed with the same number of proliferating human embryonic kidney cells (HEK293; 4 million cells), followed by forming microparticles using the method described previously. The microparticles were kept for 3 weeks under culture conditions. In order to assess the proliferation of cells in the microparticles, images were captured within 34 days using a BioTek Cytation 5 cell imaging multimode reader.

[0169] result

[0170] While formulating the microparticles, it was noted that the cell morphology and its distribution within the microparticles was unique, depending on the hydrogel. Initially, HEK293 cells aggregated within the durable PEGDA microparticles ( Figure 13 A). In contrast, the same cells loaded at the same density remained as single cells in the degradable (MeHA) microparticles ( Figure 13 C).

[0171] Over time, there was little change in the morphology or number of cells in the PEGDA microparticles ( Figure 13 B). In contrast, cells in the degradable MeHA microparticles proliferated within the hydrogel, producing large clusters within the microparticles, as can be seen after 34 days of culture ( Figure 13 D) The clusters do not appear to protrude from or attach to the surface of the MeHA beads, but remain encapsulated within the bead matrix. However, the same cells within the durable PEGDA microparticles do not proliferate rapidly or form clusters of proliferating cells within the PEGDA beads.

[0172] discuss

[0173] The microparticles in their final product form either prevent or allow cell proliferation within the microparticles. Using the data presented herein, we show that durable hydrogels prevent cell proliferation, while softer, rapidly degrading MeHA hydrogels allow for significant proliferation of blast cells. Furthermore, the morphology of the cells changes depending on the formulation used. This can be important when encapsulating proliferating versus non-proliferating cells and can be adjusted using CSS technology.

[0174] Example 4

[0175] Durability versus cell migration in degradable microparticle final products

[0176] introduce

[0177] In some therapeutic cases, the migration of cells or their products out of the microparticles is an important aspect of the treatment. For example, endogenous stem or progenitor cells have been shown to migrate to the brain during normal development and also under pathological conditions (such as brain injury or disease). Some studies have shown that natural stem cells actually migrate or "hone" to the damaged area to help it repair. However, when controls are in place to measure the number of stem cells in non-damaged tissue, most studies have failed to confirm this claim.

[0178] Therefore, in order for cells to help repair any tissue, they must be localized to that area for an extended period of time. Placing unencapsulated stem cells in a specific area is challenging because the cells are small and would quickly diffuse out of the area or be attacked by the immune system. Therefore, a localized approach that allows for the controlled release of cells to a specific area would have enormous therapeutic benefits. Here, we demonstrate these properties of degradable microparticles.

[0179] method

[0180] The three types of hydrogel polymers MeHA, AHA and PEGDA were manufactured into microparticles using the above-mentioned CSS manufacturing method. Human embryonic kidney cells (HEK293) were cultured in DMEM culture medium with high glucose and fetal bovine serum and maintained at 37°C and 5% CO2 in an incubator. Rat bone marrow stem cells (rBMSC) were cultured in high glucose DMEM, fetal bovine serum and Glutagro. When the cells reached an appropriate density, they were separated from the culture flask by adding 0.5% trypsin-EDTA for 2 minutes. After separation, the cells were transferred to a growth medium. The same amount of polymer precursor (1 mL) was mixed with the same number of HEK293 cells or rBMSC (4 million) and then formed into beads.

[0181] To assess the amount of cells that migrated from the microparticles, the microparticles were removed from the culture medium at defined time points and washed on a mesh. The wash medium and the medium surrounding the microparticles were collected. Cells were counted using a Cytation 5 imaging multimode reader (Biotek Instruments, Inc.) and an EVE Cell Counter.

[0182] result

[0183] Figure 14Representative images of wells with collected spent medium and wash medium with single cells are provided at day 4 of encapsulation. In the culture medium of HEK293 cells encapsulated in PEGDA hydrogels, few cells were counted and many of them were dead (red) ( Figure 14 A). In contrast, degradable hydrogels allow the release of high numbers of viable cells, e.g. Figure 14 The high number of viable cells (green) remaining in the wells is shown in B. However, these cell numbers are small compared to the cumulative cell release at week 2.

[0184] Figure 15 Average data from studies on cumulative cell release are summarized. During the first 6 days of culture, only a few thousand cells were released from the beads. By day 14, over 20 million cells had been released from the rapidly degrading MeHA microparticles, while only 220,000 cells had been released from the durable PEGDA microbeads. Cumulatively, by day 21, 30.8 million cells had been released from the MeHA beads, while less than 5% of that had been released from the PEGDA beads.

[0185] discuss

[0186] The results demonstrated a clear difference in the number of proliferating cells migrating from the durable versus degradable microparticles. The final microparticle product allowed cells to migrate from the microparticles into the surrounding tissue. The rapidly degrading formulation achieved this at a faster and larger scale compared to the durable hydrogel. The results demonstrate the utility of rapidly degrading hydrogels in allowing cells to migrate to localized injury sites, even before the microparticles physically degrade.

[0187] Example 5

[0188] Durability versus Viability, Morphology, and Functionality of Rapidly Degrading Microparticle Final Products

[0189] introduce

[0190] Durability or fast degradation / short-term hydrogel formulations have many applications. It has been demonstrated in the literature that microencapsulated mesenchymal stem cells (MSCs) significantly enhanced paracrine-mediated healing in a mouse hind limb injury model, compared with unencapsulated MSCs, showing increased survival and pro-angiogenic activity. Once encapsulated, cells must maintain viability and must have therapeutic function. It has been demonstrated that, compared with the cartilage formation path using softer materials, increasing the rigidity of the encapsulating matrix guides MSC differentiation toward more osteogenic paths. In addition, compared with cells grown with 2D on the same material, the rat neuron cells cultured in 3D collagen hydrogels show better survival and behave more like natural neuronal networks.

[0191] method

[0192] Viability and morphology

[0193] Two types of hydrogel polymers (rapidly degradable (MeHA and AHA) and durable (PEGDA)) were fabricated into microparticles using the CSS fabrication method described above. The same amount of polymer precursor (1 mL) was mixed with the same number of mouse MSCs (16 million) or insulinoma cells (30 million cells) and then formed into beads. The beads were maintained under culture conditions for 2 weeks. For each formulation, the viability and function of the cells were high when cultured within the microparticles. After encapsulation, the viability of MSCs and insulinoma cells within the microparticles was measured using a fluorescent viability stain. Calcein was added to the culture medium as a measure of viable cells because calcein can only enter cells with intact cell membranes. Propidium iodide was added to the culture medium to identify dead cells. After 30-60 minutes or incubation in the fluorophore, images of the microparticles were collected on a Cytation 5 imaging multimode reader (Biotek Instruments, Inc.).

[0194] Similar experiments were also conducted using a non-proliferating somatic cell source (canine pancreatic islets) encapsulated in two different hydrogels. Microparticles were produced as described above, where cells (islets or cultured cells) were mixed into the precursor. For canine islets, a 1.1-fold concentration of precursor was mixed with canine islet slurry at a 10:1 volume ratio just before droplet generation. Islet microparticle production was performed aseptically in a closed, sterile bioreactor system.

[0195] The finished microparticles were stained with disulfide hydrazone to detect the presence of insulin, as well as calcein (live cells) and propidium iodide (apoptotic / necrotic cells) to assess cell viability of the explanted tissue. Color, bright field, and fluorescence images of the stained microparticles were captured using a Cytation 5 imaging multimode reader (Biotek Instruments, Inc.).

[0196] result

[0197] MSCs encapsulated in PEGDA or MeHA showed high viability, as shown in Figure 2. Figure 16 A and B. Similar results were seen in the AHA ( Figure 16 E). Green fluorophore indicates live cells stained with calcein. Red / yellow cells indicate dead cells stained with propidium iodide. Interestingly, these cells were loaded into the microparticles as single cells, but they rapidly aggregated in the durable PEGDA gel but not in the degradable MeHA hydrogel, where they remained primarily as single cells. Monitoring cell morphology during the manufacturing process indicated that the cells aggregated during, rather than after, microparticle fabrication.

[0198] In studies using canine pancreatic islets, cells had high viability in both hydrogel microparticles ( Figure 16 C-PEDGA and 16D-MeHA). They continued to function in either hydrogel, as demonstrated by staining for their insulin content. Islets were loaded into the microparticles as cell clusters. There did not appear to be any changes in aggregation frequency or cluster size in either formulation.

[0199] Similar results were obtained using a rat insulinoma cell line. Figure 17 The high viability of cells within one day after encapsulation was shown, particularly when encapsulated in AHA. Over the next 7 days, the viability of cells in AHA dropped to approximately 50%. In contrast, insulinoma cells in PEGDA started with a lower viability but maintained close to this average throughout the study.

[0200] Islet cell function is tested by staining the cells for insulin. Disulfide hydrazone staining is a standard indicator of insulin production. It stains insulin-positive cells in a red / brown hue. Figure 18 The images in Figure 3 illustrate positive insulin staining of islets in both durable PEGDA and degradable MeHA microparticles approximately one week after encapsulation, demonstrating their continued ability to produce insulin after encapsulation in either microparticle. Unlike mouse MSCs, the morphology of these somatic islet cells did not change significantly in either hydrogel formulation.

[0201] Example 6

[0202] Durability versus rapid degradation of microparticle final products for local cellular delivery

[0203] introduce

[0204] Durable or rapidly degradable hydrogel formulations have many applications. It has been demonstrated in the literature that microencapsulated MSCs significantly enhanced paracrine-mediated healing in a mouse hindlimb injury model, showing increased survival and pro-angiogenic activity compared to unencapsulated MSCs. One of the problems with using MSC paracrine release is that unencapsulated MSCs do not stay in the injured area and are quickly removed from the body. Therefore, encapsulation with a sticky hydrogel would trap cells in the implanted area, where they can secrete beneficial compounds. We conducted experiments to determine whether the durability of microparticles or degradable formulations showed differences in their ability to adhere to surrounding tissues and remain in the area.

[0205] method

[0206] Adhesion to tissue

[0207] Immediate tissue adhesion was tested during the transplantation procedure into the omentum of dogs. Four dogs were transplanted with microparticles containing canine islets. On the day of transplantation, the animals were anesthetized using standard veterinary procedures, which included a cephalic vein catheter for infusion of fluids before, during, and after surgery. The surgeon injected the microparticles into the omentum veil. Care was taken not to enter or damage any intestinal tissue, blood vessels, nodes, or lymphatic vessels. A total of 10,000 islet equivalents were contained in 1 mL of hydrogel microparticles. The manufacture of PEGDA microparticles containing islets has been described previously. Photos were obtained before the omentum was gently replaced on the abdomen and the abdominal wall was closed followed by intradermal sutures to close the skin.

[0208] Another example of tissue adhesion was demonstrated by injecting microbeads into the rat knee. Briefly, 100 μL of 300-micron PEGDA microparticles in PBS were extruded through a 25G needle into the joint capsule of the Sprague-Dawley rat knee joint. Prior to extrusion, the solution and beads were dyed blue to confirm the presence of the microparticles within the joint capsule.

[0209] Histology of adherent tissue

[0210] Healthy Sprague-Dawley rats were used to assess the acute biocompatibility and location of microparticles within the peritoneal cavity (N=2 for each microparticle formulation). Empty microparticles were implanted around the omentum according to the previously described procedure. The microparticles were delivered via syringe as a suspension in DPBS. Approximately 1 mL of loosely packed microparticles were deposited around and within the omentum of the rats. Pain markers and activity levels of the rats were monitored and scored daily for 10 days following microparticle implantation.

[0211] Rats were sacrificed 14 days after implantation to assess the implant site to determine the location of the microparticles within the peritoneal cavity and to examine for possible tissue abnormalities. Tissue samples were collected and preserved in neutral buffered formalin, embedded in paraffin, and sectioned at 7 micrometers. Sections were stained with H&E and evaluated under a microscope.

[0212] result

[0213] PEDGA microparticles were infused into and around the abdominal omentum, e.g. Figure 19 Shown in Figure A and described in the Methods section, images were collected immediately after microparticles in culture medium were infused into and around the omentum, showing that the microparticles adhered to the tissue even when the omentum was manually manipulated. Figure 19 B and 19C illustrate results from two different dogs, both showing the sticky nature of the microparticles.

[0214] The dyed PEGDA microparticles were injected into the knee of a rat. After manual manipulation of the joint, the joint was dissected. Even through the muscle, the blue color from the microparticles was visible ( Figure 20A). Upon further dissection, the microparticle was visible and found to remain within the knee joint ( Figure 20 B and C).

[0215] Durable PEGDA and degradable MeHA and AHA microparticles (cell-free) were implanted into healthy Sprague-Dawley rats to evaluate the safety and biocompatibility of microparticles produced using the CSS method, as well as their location and attachment to surrounding tissues. Necropsies were performed after 2 and 10 weeks, and no signs of acute inflammation, excess fluid, or tissue abnormalities were observed throughout the abdominal cavity, although some tissue adhesion was identified at the suture site of the omentum to the gastric wall. Both microparticle formulations were found intact in the omentum and were visually translucent and appeared to be surrounded by a very thin, membranous layer. A small number of microparticles were found to adhere to the liver surface. However, the vast majority of microparticles were attached to the omentum, with some free-floating microparticles.

[0216] Hematoxylin and eosin staining of explanted tissues for the three microparticle groups is shown in Figure 21 No significant differences were noted between the groups. Microparticles were generally intact, but appeared to have shrunk in size and in some cases appeared crushed / deformed as a result of tissue processing for histology. Microparticles were generally found to have a thin layer of cells around their periphery, typically two to three cells thick, and were generally devoid of cells within, suggesting that empty microparticles were injected into the animals. Most individual microparticles were surrounded by normal omental tissue, suggesting a lack of inflammatory response to the microparticles from various formulations.

[0217] Figure 22 Quantification of the foreign body reaction to microparticles in the omentum is shown. The width of the fibrotic zone, determined by low cellularity and higher collagen content, was smallest in AHA and higher around PEGDA microparticles. Nevertheless, the thickness of the fibrotic ring around PEGDA microparticles was still less than that reported for other hydrogel formulations such as alginate. Similarly, the lymphocyte zone was less in AHA microparticles. The lymphocyte zone is defined as cellular immune cells that migrate around the microparticles. They were also higher for PEGDA microparticles, but still considered minimal ( Figure 22 ).

[0218] discuss

[0219] A unique and advantageous feature of the final product is that, regardless of whether it is constructed from a durable or degradable hydrogel matrix, it is adhesive. This adhesive property allows it to quickly attach to surrounding tissue and remain in that area until it degrades. In addition to attaching islets to surrounding tissue in the peritoneal space, the use of adhesive hydrogels for the administration of MSCs or other cells that release anti-inflammatory or pro-healing cytokines could be injected into joints or other inflamed tissues for prolonged delivery of cytokines to a localized area without concern for redistribution to other areas of the body.

[0220] Example 7

[0221] Study of parameters affecting hydrogel properties

[0222] In this example, various parameters were studied to better understand their relationship to each other and their impact on the properties and characteristics of the resulting hydrogel microparticles. As discussed in more detail below, various microparticle characteristics can be fine-tuned and adjusted according to the specific desired treatment regimen and results. It is worth noting that each of these parameters is a dynamic factor in the manufacture of hydrogel microparticles, which means that its impact will be related to the amount or ratio of other components used in the formulation and / or other processing parameters. Further, modifications to certain parameters will not necessarily produce proportional changes in the resulting hydrogel. Likewise, however, the subtle differences in different microparticle formulations and processing parameters allow for many different customizable characteristics, depending on the cell type to be encapsulated, the implantation site, the condition to be treated or prevented, and other desired results.

[0223] To explore these different parameters, the swelling ratio (Q) was used as a proxy for the relative degradation rate of the microparticles.

[0224] A. Mass fraction of core polymer material

[0225] The mass fraction refers to the % composition by mass of the reactive polymer species in the hydrogel precursor solution used for the polymer backbone (e.g., 2.5 g PEGMAL in 10 g of precursor solution = 25% mass fraction). Adjusting the mass fraction of the reactive polymer in the precursor solution has a consistent effect on the resulting Q value. The lower the mass fraction, the higher the Q value and the softer the gel. This is illustrated by cross-linking various multi-arm PEGMAL species with 500 Da PEG dithiol. The cross-linker concentration in this example was kept constant at a mass fraction of 0.25% in the alginate bath. A non-reactive PEG species (PEG 8000) was included in the lower mass fraction group to maintain a constant solution viscosity. The results are shown in Figure 23 middle.

[0226] B. Concentration of cross-linking substances in the alginate bath

[0227] The concentration of the crosslinker (e.g., PEG dithiol, DTT) in the alginate bath will affect the resulting hydrogel microparticles. However, the relationship between the crosslinker and the hydrogel matrix is not straightforward, as there can be too little or too much crosslinker in the bath. Insufficient crosslinker results in a weak gel with insufficient crosslinking; however, too much crosslinker will also saturate the reactive sites on the polymer backbone, similarly preventing crosslinking and producing a weak gel. Therefore, this parameter must also be balanced with the amount and type of polymer precursor used for the backbone.

[0228] 1. Increasing the cross-linker concentration leads to lower Q values.

[0229] In this example (core polymer = 4-arm 10 kDa PEGMAL), increasing the cross-linker concentration (DTT) resulted in lower Q values. This indicates that the cross-linker concentration never reached a level where saturation of the core polymer occurred. The results are shown in Figure 24A middle.

[0230] 2. Increasing the cross-linker concentration leads to higher Q values.

[0231] In this example (core polymer = 8-arm 40 kDa PEG-vinyl sulfone), increasing the crosslinker (BMA, CAS 123-81-9) concentration initially resulted in lower Q values, but further increases resulted in higher Q values, in contrast to those observed in Example 1. It is noteworthy that the relative molar concentrations of the crosslinker (thiol) to the core reactive group (vinyl sulfone) were different from the previous examples. Therefore, the conditions in Example 2 may have led to saturation of the core polymer molecules and, therefore, inefficient crosslinking, resulting in a weaker gel network and a higher Q. The results are shown in Figure 24B middle.

[0232] C. Alginate bath pH

[0233] Thiol-ene reactions are strongly influenced by the pH of the environment. Higher pH levels significantly increase the kinetics of thiol-ene reactions. Therefore, adjusting the pH of the alginate bath can greatly affect the kinetics of gel crosslinking and the resulting Q value.

[0234] 1. Increasing pH leads to higher Q values.

[0235] In this example (core polymer = 8-arm 40 kDa PEG-vinyl sulfone; cross-linker = BMA), increasing the bath pH resulted in higher Q values. The results are shown in Figure 25A The increased reaction rate can lead to faster reaction with the core polymer and eventual partial saturation of the core polymer, thus a weaker gel network and a higher Q value.

[0236] 2. Increasing pH results in lower Q values.

[0237] In this example (core polymer = 8-arm 40 kDa PEGMAL; cross-linker = 500 Da PEG dithiol), increasing pH results in lower Q values. The polymer and cross-linker are balanced to avoid saturation. Therefore, in this case, increasing pH results in faster reaction kinetics and more efficient cross-linking. The results are shown in Figure 25B Note that the increase in Q with respect to pH in Example 2 is quite small compared to the very steep increase in Example 1. Thus, both the direction and magnitude of the effect of pH on Q appear to be highly influenced by the relative amounts of crosslinker and core polymer reactive sites.

[0238] D. Molecular weight of cross-linked substances

[0239] The molecular weight of the cross-linking species affects both the rate of diffusion into the core-shell construct and the distance between cross-links in the gel network. Therefore, differences in the MW of the cross-linking species can affect the final Q value of the resulting gel microparticles, depending on the other parameters used in the process.

[0240] 1. Increasing MW leads to higher Q value.

[0241] In this example (core polymer = 8-arm 40 kDa PEG-vinyl sulfone), increasing the MW of the cross-linking species from 154 Da to 500 Da resulted in a significant increase in the Q value. The molar concentration of the cross-linking species was 10 mM for both formulations. The results are shown in Figure 26A middle.

[0242] 2. Increasing MW results in no change in Q value.

[0243] In this example (core polymer = 8-arm 40 kDa PEGMAL), increasing the MW of the cross-linker resulted in no significant difference in the Q value of the resulting gel. The effect of the change in the MW of the cross-linker compared to the polymer reactive sites may be offset by adjusting the relative amount of cross-linker used. The results are shown in Figure 26B middle.

[0244] E. Molecular weight of core polymer material

[0245] Assuming that the reactive sites per gram of polymer are maintained, the molecular weight of the core polymer material appears to affect the Q value more predictably than the molecular weight of the crosslinker. For example, 8-arm 40 kDa PEGMAL and 4-arm 20 kDa PEGMAL have the same number of reactive sites (maleimides) per gram. However, as shown below, when prepared under otherwise identical conditions, the 40 kDa material produces a stronger gel (lower Q value).

[0246] 1. Effect of MW of core polymer at constant reactive sites / g.

[0247] This example demonstrates the effect of core polymer molecular weight on the Q value of gel microparticles prepared by CSS using three side-by-side comparisons. In all formulations, despite having the same mass fraction and reactive sites / gram, the Q value decreases as the core MW increases. Note that the effect of core MW becomes more pronounced as the crosslinker concentration decreases (and the size increases). The results are shown in Figure 27A middle.

[0248] 2. Effect of MW of core polymer at unequal reactive sites / gram.

[0249] In this example, microparticles made from either an 8-arm 40 kDa or a 4-arm 10 kDa PEGMAL core material were compared. In all cases, the mass fraction of the core polymer solution was 10%. The 4-arm 10 kDa material had twice as many reactive sites per gram as the 8-arm 40 kDa, yet despite the lower potential crosslink density, the 8-arm 40 kDa microparticles had lower Q values (i.e., stronger gels) than the 4-arm 10 kDa microparticles. The results are shown in Figure 27B middle.

[0250] F. Effect of UV cross-linking time

[0251] In this example, AHA and MeHA microspheres of different molecular weights were produced at different UV exposure times. Briefly, 2.5% 1 MDa or 200 kDa AHA or 1 MDa MeHA solutions (each with 1% 3.4 kDa PEGDA) were extruded dropwise into a stirred alginate bath as described above and exposed to UV light for 2.5, 5, or 7.5 minutes. The resulting diameter and hydrogel swelling ratio for each polymer type were quantified. The results are shown in Figure 28A and Figure 28B Figure 3 shows the changes in microsphere diameter and swelling ratio as a function of UV exposure time. These results show that both diameter and swelling ratio decrease with increasing UV exposure time. Notably, the swelling ratio of 1 MDa MeHA hydrogel microspheres decreased significantly after 7.5 minutes of UV exposure compared to 5 minutes, with swelling ratios of 509.1 and 113.4, respectively. This trend was consistent for each polymer type and molecular weight and demonstrates the control of hydrogel durability by varying UV exposure time.

[0252] G. Potential control of hydrogel degradation via hydrolytic stability of cross-linking species, regardless of Q value.

[0253] 1. BMA acts as a more hydrolytically unstable cross-linker.

[0254] Certain crosslinkers can be used to influence the degradation profile of microparticles. For example, the dithiol ester crosslinker molecule BMA could potentially be used effectively as a means to accelerate the hydrolytic degradation of thiol-ene type hydrogels.

[0255] A series of hydrogel microparticles composed of PEG-vinyl sulfone cross-linked with either DTT or BMA were prepared and stored in 50 mL polypropylene tubes in DPBS at room temperature (~27°C) for at least 6 months. Table 3 below lists the composition and corresponding swelling ratio of each formulation for reference. After 6 months of storage, the microparticles were qualitatively re-examined for signs of degradation. Notably, all microparticles cross-linked with BMA completely degraded within 6 months. Those cross-linked with DTT appeared relatively unchanged after 6 months, although formal characterization had not been performed at this time. Furthermore, two of the completely degraded BMA-cross-linked microparticles, which previously exhibited the lowest Q values of all the microparticles included in the table, still completely degraded compared to DTT-cross-linked microparticles, which had higher Q values. This result suggests that the BMA cross-linker may be a useful tool for generating microparticles that are initially very strong (low Q) but degrade more rapidly than other microparticles cross-linked with DTT or PEG dithiol of similar initial strength.

[0256] Table 3. Long-term durability of microparticles cross-linked with DTT or BMA.

[0257]

[0258] Table 4 below provides a list of additional dithiol species and the corresponding pKa of one or more thiol groups. The hydrolytic instability of the bond generated by the thiol-ene reaction is related to the pKa of the reactive thiol. Therefore, these species can be used as chemical crosslinkers in thiol-ene reaction schemes to control the hydrolytic instability and, therefore, the degradation rate of the microparticles.

[0259] Table 4. Examples of dithiol cross-linker species with varying degrees of hydrolytic instability.

[0260]

[0261]

Claims

1. A spherical or ellipsoidal non-alginate hydrogel microparticle for local delivery and sustained release of therapeutic cells and / or tissues at an implantation site, the microparticle comprising a 3-dimensional matrix of a covalently cross-linked non-alginate polymer compound and a therapeutically effective amount of cells and / or tissues embedded therein, wherein the cells have a viability of at least 50%, and wherein the microparticle has a size greater than 30 μm, wherein the non-alginate hydrogel microparticle is formed without using an oil emulsion, wherein the polymer compound is polyethylene glycol maleimide, and the matrix further comprises a cross-linking agent cross-linked with the non-alginate polymer compound, the cross-linking agent being selected from the group consisting of dithiothreitol, polyethylene glycol dithiol, and ethylene glycol bis-thioglycolate.

2. The microparticles according to claim 1, wherein The microparticles have a size greater than 300 μm.

3. The microparticles according to claim 1, wherein The matrix consists essentially of the polyethylene glycol maleimide cross-linked with the dithiothreitol, polyethylene glycol dithiol, or ethylene glycol bis-mercaptoacetate.

4. The microparticles according to claim 1, wherein The matrix consists essentially of covalently cross-linked polyethylene glycol maleimide and polyethylene glycol dithiol precursor compounds.

5. The microparticles according to claim 1, wherein Up to 50% of the volume of the microparticles comprises the cells or tissue. The microparticles according to claim 1 , wherein The microparticles have a storage stability in PBS at 37°C of less than 6 months.

7. The microparticles according to claim 1, wherein The microparticles have a storage stability in PBS at 37°C of greater than 6 months.

8. The microparticles according to claim 1, wherein The microparticles are degradable microparticles that will break down in less than 3 months when implanted in the body.

9. The microparticles according to claim 1, wherein The microparticles are durable microparticles that will not decompose for at least 3 months when implanted in the body.

10. The microparticles of claim 1, further comprising a cell culture medium component.

11. The microparticles according to claim 1, wherein The cells are non-proliferative or proliferative.

12. The microparticle of claim 1, further comprising cellular products secreted with the cells, including signaling molecules, cytokines, chemokines, therapeutic proteins, hormones, vesicles, antibodies, viruses, exosomes.

13. The microparticles according to claim 1, wherein The cells are engineered or transgenic cells.

14. The microparticles according to claim 1, wherein The cells are clusters of multiple aggregated cells.

15. The microparticles according to claim 14, wherein The clusters include two or more different types of cells.

16. The microparticles according to claim 1, wherein The microparticles have a size of 50 μm to 5 mm.

17. A composition for local delivery and sustained release of therapeutic cells and / or tissues and / or cell products at an implantation site, comprising a plurality of hydrogel microparticles according to any one of claims 1 to 16 dispersed or suspended in a biocompatible delivery vehicle.

18. The composition of claim 17, further comprising a plurality of hydrogel microparticles comprising a 3-dimensional matrix of covalently cross-linked non-alginate polymer compounds having a size greater than 30 μm and being substantially free of any cells or tissues.

19. The composition according to claim 17, wherein The plurality of hydrogel microparticles includes a mixture of degradable and durable hydrogel microparticles dispersed or suspended in the biocompatible delivery vehicle.

20. The composition according to claim 19, wherein The durable hydrogel microparticles include insulin-producing cells and / or tissues, and wherein the degradable hydrogel microparticles include stem cells.

21. The composition according to claim 17, wherein The plurality of hydrogel particles include a plurality of first hydrogel particles comprising a first cell and / or tissue type and a plurality of second hydrogel particles comprising a second cell and / or tissue type different from the first cell and / or tissue type.

22. The composition according to claim 21, wherein The first cell and / or tissue type comprises insulin-producing cells and / or tissues, and the second cell and / or tissue type comprises stem cells or glucagon-producing cells and / or tissues.

23. The composition according to claim 17, wherein The plurality of hydrogel particles include a plurality of first hydrogel particles having a first size and a plurality of second hydrogel particles having a second size different from the first size.

24. The composition according to claim 17, wherein The carrier is a biocompatible liquid suspension or gel.

25. The composition according to claim 17, wherein The carrier is a cell culture medium solution or a gel.

26. The composition according to claim 17, wherein The composition may be injectable or implantable via a catheter.

27. The composition of claim 17, comprising a therapeutically effective amount of the microparticles.

28. The composition according to claim 17, wherein The composition is in the form of a putty, paste or gel.

29. The composition according to claim 17, wherein The composition is in the form of a liquid suspension.

30. Use of a plurality of microparticles according to any one of claims 1 to 16 or a composition according to any one of claims 17 to 29 for the preparation of a medicament for local delivery and sustained release of therapeutic cells and / or tissues at an implantation site.

31. The use according to claim 30, wherein The medicament is formulated so that the composition is injected into or near the implantation site.

32. The use according to claim 31, wherein The injection is intra-articular.

33. The use according to claim 30, wherein The implantation site is an area of inflammation, injury, arthritis, or degeneration in the subject.

34. The use according to claim 30, wherein The implantation site is a joint.

35. The use according to claim 30, wherein The implantation site is the area near the gland where the microparticles release the cellular product.

36. The use according to claim 35, wherein The cellular products released by the microparticles in the vicinity of the glands are hormones.

37. The use according to claim 30, wherein The implantation site is in a capsular space or a physiologically contained space.

38. The use according to claim 37, wherein The implantation site is in the cardiac capsule or the renal capsule.

39. The use according to claim 30, wherein The implantation site is the omentum or peritoneum of the subject.

40. The use according to claim 30, wherein The implantation site is in or adjacent to a tumor.

41. The use according to claim 30, wherein The implantation site is in or adjacent to the site of tumor resection.

42. The use according to claim 30, wherein the implantation site is a sepsis site.

43. The use according to claim 30, wherein The microparticles biodegraded within 3 months after administration.

44. The use according to claim 30, wherein The microparticles do not biodegrade for at least 6 months after administration.

45. The use according to claim 30, wherein The implantation site was a joint, where the microparticles mechanically degraded within 3 months after administration.

46. The use according to claim 30, wherein The implantation site is a joint, wherein the microparticles do not mechanically degrade for at least 6 months after administration.

47. The use according to claim 30, wherein The hydrogel microparticles adhere to the tissue at the implantation site, thereby maintaining the implanted cells at the implantation site in therapeutic amounts and for a period of time.

48. The use according to claim 30, wherein The cells and / or tissues and / or cell products are slowly released from the microparticles through the pores in the microparticles and / or biodegradation of the microparticles to the local area of the implantation site to alleviate or reduce the severity of the condition in the subject.

49. The use according to claim 30, wherein The microparticles biodegrade within 5 days or less after administration.

50. The use according to claim 30, wherein The microparticles biodegraded within 24 hours after administration.

51. The use according to claim 30, wherein The microparticles biodegraded within 1 week after administration.

52. The use according to claim 30, wherein The microparticles biodegraded within 1 month after administration.

53. The use according to claim 30, wherein The microparticles biodegrade in less than 6 months after administration.

54. The use according to claim 30, wherein The microparticles biodegrade after administration, wherein the biodegradation includes hydrolysis of the polymer backbone.

55. The use according to claim 30, wherein The microparticles biodegrade after administration, wherein the biodegradation includes hydrolysis of the cross-linking agent.

56. The use according to claim 30, wherein The microparticles biodegrade after administration, wherein the degradation rate can be controlled by varying a parameter selected from the group consisting of polymer precursor molecular weight, crosslinker molecular weight, crosslinker to polymer precursor ratio, crosslinker hydrolysis, crosslinking kinetics, UV exposure time, and combinations thereof.

Images