Heterogeneous hybrid hydrogels, their manufacturing methods, and their use as non-degradable in situ filler implants.

Abstract

The filler implant comprises a scaffold made from a heterogeneous hybrid hydrogel formed from a copolymer comprising an ethylene radical-functionalized dendrimer monomer, an acrylamide compound selected from N-substituted methacrylamides and N-substituted acrylamides, a crosslinker, and a copolymerizable bioactive material. The scaffold is formed from microbeads with diameters between 1.5 microns and 10 microns, with the acrylamide compound accounting for a majority of the weight. The microbeads aggregate to form aggregates containing between 5 and 50 microbeads. The aggregates are connected by crosslinks to define a through-porous array that defines a three-dimensional percolation pathway. The through-porous array is formed from pores, the diameter of the majority of the pore volume of which is between 10 and 30 microns. The scaffold has viscoelastic properties with an elastic modulus of between 1 and 200 kPa.

Description

Technical Field

[0001] The present invention relates to a hybrid heterogeneous hydrogel, and more specifically, to the use of a hybrid heterogeneous hydrogel as a filler implant.

Background Art

[0002] Damage to the spinal cord parenchyma, whether it is due to local trauma, ischemia, tumor or removal surgery of vascular malformation, or other causes, will cut off the connection of nerve fibers, interrupt the transmission of nerve impulses that the brain commands motor functions, and as a result, prevent the treatment of sensory functions. This will result in complete paralysis or incomplete paralysis.

[0003] In the case of spinal cord injury, the injury caused by compression or any other impact on the spinal cord parenchyma progresses outward from the center through a neurodegenerative phenomenon that gradually transitions to the chronic phase of the initial injury. In the chronic phase, the injury is represented by a cystic cavity, and is separated and surrounded by heterogeneous scar tissue composed of glial cells, fibroblasts, pericytes, meningothelial cells, and extracellular matrix molecules, particularly proteoglycan and collagen. Such a form of healing is the reason why the nerve fibers of the inward and outward axonal pathways cannot be repaired, just like the spinal canal lumen.

[0004] Regenerative therapy aims to repair cells in the cystic cavity in the chronic phase of medullary injury and to re-vascularize the injured area, which may represent a significant volume reduction with vertebral body height varying from 2 to 6. Regenerative therapy also aims to grow nerve fibers through this cystic cavity to facilitate the reconnection of spinal nerve circuits and re-establish motor and sensory functions. This strategy takes into account the post-injury plasticity of the nervous system, which allows for the re-formation of relay circuits between the regrowing nerve fibers and intact neurons below the injury site. The strategy used in tissue engineering involves introducing a porous and permeable matrix to the site of injury (either acute or chronic), allowing this matrix to act as a physical, chemical, and mechanical scaffold in the repair processes of endogenous cells, blood vessels, and nerves, resulting in histological reconstruction of nerve tissue.

[0005] Scaffolds commonly used in experimental and clinical trials are represented by a specific class of biomaterials, namely hydrogels. These scaffolds are polymer matrices that form water-saturated and cross-linked polymer arrays. This polymer matrix is ​​used in tissue engineering, particularly when the nervous system is involved. The polymer matrix is ​​prepared from biodegradable or bioabsorbable polymers to construct a time-unstable porous structure. These polymers are either natural polymers such as alginates, agarose, chitosan, collagen, hyaluronic acid, fibrin, or peptides, or synthetic polymers such as polycaprolactone, poly(hydroxybutyrate), poly(orthoester), poly(α-hydroxyester), or polyanhydride. For this reason, biodegradable and / or bioabsorbable hydrogels have been proposed to promote nerve regeneration in spinal cord injury in experimental models or clinical trials in humans.

[0006] Biodegradable hydrogels are hydrogels that, when injected into a living organism, can be broken down by spontaneous chemical hydrolysis of polyanhydrides, poly(orthoesters), and poly(α-hydroxyesters). Thioether esters are broken down in the presence of water, while biopolymers (oligopeptides, proteins, or polysaccharides) are broken down by enzymes or other proteins produced by cells.

[0007] For example, U.S. Patent No. 7,163,545 discloses a poly(lactic-co-glycolic acid) matrix containing induction channels in axonal regeneration in combination with therapeutic agents. U.S. Patent No. 8,377,463 reports a device formed from poly(lactic-co-glycolic acid) for treating acute spinal cord injury that degrades in situ between 30 and 60 days, and which can be combined with therapeutic agents and / or stem cells. U.S. Patent Application Publication No. 2018 / 0037865 describes a composite hydrogel comprising a degradable matrix of hyaluronic acid, collagen, fibrin, chitosan, methylcellulose, polyoxyethylene, or a combination thereof, which degrades in situ at varying rates depending on the properties of the polymer, in combination with stem cells and therapeutic molecules. U.S. Patent Publication No. 20060002978 discloses a tubular matrix containing a porous polymer material formed from a homopolymer or copolymer of lactic acid and / or glycolic acid and / or poly(caprolactone). This polymer matrix may comprise aliphatic polyester, polyanhydride, polyphosphatine, polyvinyl alcohol, polypeptide, or alginate. U.S. Patent No. 8,877,498 reports a matrix presenting a composition comprising chitosan, chitin, cellulose, alginate, gelatin, hyaluronic acid, collagen, elastin, or a combination thereof, with a hierarchical structure having highly aligned channels and ridges along the wall, enabling the induction and regeneration of nerve fibers. International Publication No. 2014013188 describes a microgel suspension or 2-3 mm in size for the treatment of spinal cord injury. 3Acetylated chitosan biomaterials in gel form are disclosed. U.S. Patent Publication No. 2015 / 0166786 and International Publication No. 2013010087 describe a hydrogel comprising poly(lactic acid-co-glycolic acid) or poly(ε-caprolactone) in combination with poly(L-lysine) for the treatment of acute spinal cord injury in humans. U.S. Patent Publication No. 2015 / 0044259 describes a matrix composed of poly-D-lysine and peptidoglycan for promoting nerve fiber growth. European Patent Publication No. 2,347,763 and U.S. Patent Publication No. 2011 / 0177170 describe a matrix comprising collagen microparticles contained in a homogeneous gel composed of peptides, uronic acid, and hexosamine, which is implanted integrally with cell grafts in the treatment of central nervous system injury, with matrix degradation occurring over several weeks to months. International Publication No. 2013 / 084137 discloses a degradable implant using calcium sulfate hemihydrate containing parallel geometrically shaped channels combined with growth factors for the treatment of complete spinal cord injury in humans.

[0008] Biodegradable hydrogels, such as those presented in U.S. Patent No. 8,815,277 or International Publication No. 2011 / 002249, are also proposed. These systems offer the advantage of being non-invasive, meaning they do not require open spinal cord surgery for implantation. Generally, these systems degrade in situ within two weeks, which is significantly faster than regenerative processes.

[0009] A hydrogel made from a poly(lactic-co-glycolic acid) substrate has also been proposed for repairing nerve channels in the spinal cord. While this hydrogel degrades in 30-60 days in rats, if these results are applied to humans, this period is far too short to allow for successful tissue regeneration. Comparative studies have shown that isolated axons grow longer in humans than in rats, and the regeneration rate is three times slower in humans (Gordon, 2007, "The potential of electrical stimulation to promote functional recovery after peripheral nerve injury - comparisons between rats and humans," Acta Neurochir Suppl., 2007, Vol. 100, pp. 3-11). In the case of the spinal cord, the growth time is even longer because numerous nerve fibers need to regrow to achieve functional motor recovery. Axonal growth through the surgical repair site is slow and does not occur simultaneously until it reaches the appropriate target in the spinal segment below the site of injury.

[0010] These methods, which attempt to repair spinal cord injuries, cannot repair the anatomical structure of the damaged spinal cord segment simultaneously with achieving functional recovery of nerve function. Therefore, when applied to humans, the efficiency of central nervous system regeneration is severely limited and incomplete.

[0011] Once implanted in an organ, a degradable hydrogel matrix degrades as polymer chains are cleaved during tissue remodeling, coinciding with the migration and colonization of cells, blood vessels, and regenerating nerve fibers. Thus, as spinal cord tissue reconstruction occurs and the formation of new nerve tissue is promoted, the degradable hydrogel matrix exhibits a high degradation rate, rapidly diminishing its initial mechanical support properties. To optimize tissue regeneration, i.e., achieve complete regeneration in terms of time, it is essential to maintain a certain level of structural integrity in terms of space and time throughout the entire tissue remodeling process within the hydrogel scaffold.

[0012] Hydrolysis of ester bonds in poly(α-hydroxy acids) releases acidic compounds, and it is clear that these acidic compounds, when accumulated at the implantation site, lower the pH. As illustrated in Bostman OM and Pihlajamaki HK (2000), "Adverse tissue reactions to bioabsorbable fixation devices," Clin Orthop Rel Res, Vol. 371, pp. 216-227, a decrease in pH accelerates the hydrolysis rate in the center of the implant compared to the surface, rapidly reducing the initial mechanical properties of the implant and leading to local reactions to foreign bodies.

[0013] International Publication No. 2010 / 097524 describes heterogeneous hybrid hydrogels comprising the following monomers: A dendrimer monomer comprising a central core A and macromolecular dendritic branches of polyoxyethylene, wherein at least one of the macromolecular dendritic molecules is functionalized with an acrylate radical or a methacrylate radical. N-substituted methacrylamide or N-substituted acrylamide, and Copolymerizable bioactive materials selected from the group consisting of derivatives of complex sugars, derivatives of tissue adhesion peptides, and derivatives of polymer conjugates bound to antibodies against lipid derivatives. A heterogeneous hybrid hydrogel copolymer obtained from is disclosed. [Prior art documents] [Patent Documents]

[0014] [Patent Document 1] U.S. Patent No. 7,163,545 [Patent Document 2] U.S. Patent No. 8,377,463 [Patent Document 3] U.S. Patent Application Publication No. 2018 / 0037865 [Patent Document 4] U.S. Patent Application Publication No. 2006 / 0002978 [Patent Document 5] U.S. Patent No. 8,877,498 [Patent Document 6] International Publication No. 2014 / 013188 [Patent Document 7] U.S. Patent Application Publication No. 2015 / 0166786 [Patent Document 8] International Publication No. 2013 / 010087 [Patent Document 9] U.S. Patent Application Publication No. 2015 / 0044259 [Patent Document 10] European Patent Application Publication No. 2,347,763 [Patent Document 11] U.S. Patent Application Publication No. 2011 / 0177170 [Patent Document 12] International Publication No. 2013 / 084137 [Patent Document 13] U.S. Patent No.  8,815,277 [Patent Document 14] International Publication No. 2011 / 002249 [Patent Document 15] International Publication No. 2010 / 097524 [Non-Patent Document]

[0015] [Non-Patent Document 1] "The potential of electrical stimulation to promote functional recovery after peripheral nerve injury - comparisons between rats and humans" by Gordon in 2007, Acta Neurochir Suppl., 2007, Vol. 100, pp. 3 - 11 [Non-Patent Document 2] Bostman OM and Pihlajamaki HK (2000), "Adverse tissue reactions to bioabsorbable fixation devices," Clin Orthop Rel Res, Vol. 371, pp. 216-227. [Overview of the project] [Problems that the invention aims to solve]

[0016] One of the objectives of the present invention is to correct the aforementioned drawbacks, and more specifically, to provide an implant made from a hydrogel that has a lower degradation rate compared to prior art hydrogels and exhibits good conformity to mechanical stresses related to cell colony formation. [Means for solving the problem]

[0017] According to one feature of the present invention, the following monomers: A dendrimer monomer functionalized by a single branch containing an unsaturated ethylene radical. Acrylamide compounds selected from N-substituted methacrylamides and N-substituted acrylamides, as well as Crosslinking agent A heterogeneous hybrid hydrogel formed from copolymers obtained from at least three of these is proposed.

[0018] This heterogeneous hybrid hydrogel is mainly formed by multiple microbeads having a diameter greater than 1.5 microns but less than 10 microns, with N-substituted methacrylamide and N-substituted acrylamide accounting for the majority of its weight. The microbeads aggregate together to form aggregates containing between 5 and 50 microbeads, and these aggregates are connected to each other by crosslinking points, defining a through-porous array that defines a three-dimensional percolating path. The through-porous array defines pores, and the majority of these pores are formed by pores with a diameter between 10 and 30 microns. The heterogeneous hybrid hydrogel is characterized by having viscoelastic properties and an elastic modulus between 1 and 200 kPa.

[0019] According to one analysis, the diameter of the microbeads is between 2 and 5 microns.

[0020] Preferably, the microbeads comprise at least 90% by weight of an acrylamide compound. More preferably, the microbeads are composed of a crosslinked acrylamide compound.

[0021] In certain embodiments, the aggregate contains between 10 and 30 microbeads. It is advantageous to provide a compressible aggregate.

[0022] Advantageously, the functionalized dendrimer monomer comprises one or more polyoxyethylene polymeric dendritic branches, at least one of which is functionalized with one or more copolymerizable bioactive materials selected from the group consisting of derivatives of complex sugars, derivatives of tissue adhesion peptides, and derivatives of polymer conjugates conjugated to antibodies against lipid derivatives, the one or more copolymerizable bioactive materials coating the walls of a three-dimensional percolating array.

[0023] Preferably, the walls of the three-dimensional percolation array described above are functionalized with several different copolymerizable bioactive materials.

[0024] In a favorable configuration, the one or more active molecules described above are selected from the group consisting of derivatives of complex sugars, derivatives of tissue adhesion peptides or peptides having angiogenesis properties, derivatives of peptides that stimulate nerve regrowth, derivatives of peptides that stimulate cell proliferation and differentiation, derivatives of polymer conjugates conjugated to antibodies against lipid derivatives, and chemokines of the stroma-derived factor-1 (SDF-1) class.

[0025] The functionalized dendrimer monomer is further preferable to be provided as comprising a central core A, polymeric dendritic branches of polyoxyethylene, at least one of which is functionalized with a polymerizable acrylate radical or methacrylate radical.

[0026] According to another configuration, the acrylamide compound is N-(2-hydroxypropyl)methacrylamide (HPMA), and the functionalized dendrimer monomer comprises dendritic branches of polyoxyethylene functionalized with a bioactive agent that confers bioactive properties for tissue regeneration.

[0027] Another object of the present invention is to use a heterogeneous hybrid hydrogel with one of the above configurations as a filler implant designed to be inserted between the edges of an anatomical defect in an organ or tissue.

[0028] Prioritizing this, heterogeneous hybrid hydrogels are used as filler implants to fill the parenchymal lumen of the central nervous system.

[0029] Advantageously, heterogeneous hybrid hydrogels are used as filler implants to correct congenital malformations of the central nervous system or spina bifida.

[0030] According to one feature of the present invention, a method is proposed for producing heterogeneous hybrid hydrogels that are easy to implement in industrial production, and this method can produce heterogeneous hybrid hydrogels that are more suitable for the formation of filler implants, particularly in terms of physicochemical specifications.

[0031] The method for fabricating heterogeneous hybrid hydrogels involves the following monomers: A dendrimer monomer functionalized by one branch containing an unsaturated ethylene radical, wherein the other branches lack an ethylene radical, Acrylamide compounds selected from N-substituted methacrylamides and N-substituted acrylamides, as well as At least one bifunctional unsaturated ethylene crosslinking agent having two reactive vinyl bonds At least three of them, Free radical initiators and The reaction mixture comprises a reaction mixture comprising the following: formation of microbeads by phase separation induced by copolymerization and free radical copolymerization at temperatures between 45°C and 55°C. Microbeads have a diameter greater than 1.5 microns but less than 10 microns, and the majority of their weight consists of N-substituted methacrylamide and N-substituted acrylamide. Microbeads aggregate to form aggregates containing between 5 and 50 microbeads, and these aggregates are linked together by crosslinking points to define a heterogeneous hybrid hydrogel that defines a through-porous array that defines a three-dimensional percolation pathway. The through-porous array defines pores, the majority of which are formed by pores with a diameter between 10 and 30 microns. The above heterogeneous hybrid hydrogel has viscoelastic properties and an elastic modulus between 1 and 200 kPa. It is noteworthy that the above reaction mixture is poured into a tight, cylindrical heat-conducting mold.

[0032] According to one theory, the molar ratio of the functionalized dendrimer monomer to the crosslinking agent is between 0.1 and 0.8. Preferably, the molecular weight of the functionalized dendrimer monomer is between 6,220 g / mol and 23,280 g / mol.

[0033] Preferably, the reaction mixture for forming microbeads contains a copolymerizable bioactive material, which is selected from the group consisting of derivatives of complex sugars, derivatives of tissue adhesion peptides, and derivatives of polymer conjugates bound to antibodies against lipid derivatives.

[0034] In a favorable configuration, the reaction mixture is poured into a metal mold whose inner wall is coated with polytetrafluoroethylene. This mold is preferably heated in a water bath. Preferably, polymerization of the reaction mixture is carried out at a first temperature for at least 80 minutes to form microbeads, after which the temperature of the metal mold and the reaction mixture is raised by at least 5°C.

[0035] Advantageously, a method for forming aggregates of heterogeneous hybrid hydrogel microbeads using one of the above-described configurations is carried out in a method for fabricating filler implants. The method for fabricating filler implants comprises forming aggregates of heterogeneous hybrid hydrogel microbeads and assembling the aggregates together by a three-dimensional printing method to form a filler implant.

[0036] Other advantages and features will become more apparent from the following description relating to specific embodiments and implementations of the invention, which are given for non-limiting illustrative purposes only and are shown in the accompanying drawings. [Brief explanation of the drawing]

[0037] [Figure 1] This diagram schematically illustrates an implant equipped with a heterogeneous hybrid hydrogel that defines a three-dimensional penetrating channel. [Figure 2]This diagram illustrates a metal body defining four cylindrical wells, the walls of which are covered with a PTFE surface. [Modes for carrying out the invention]

[0038] Implants made from hydrogel substrates are designed to be implanted into cavities that represent tissue volume reduction, such as the medullary cavity. The degradation rate of the hydrogel (V d ) is the cell regeneration rate (V r If the regeneration is faster than the specified time, all support for tissue reconstruction is rapidly lost. Tissue reconstruction is limited to the periphery of the implantation site, and the cell regeneration process becomes incomplete. In addition, if biodegradation occurs in the polymer scaffold implanted in the intramedullary injury, a physical separation is created between the implant body and the spinal cord tissue. The space separating the implant from the tissue is filled with cerebrospinal fluid, which prevents the hydrogel from integrating into the nerve tissue. This space also prevents axons from reaching the scaffold body formed by the implant during the regeneration process.

[0039] The production of acid-derived degradation products can trigger acute inflammatory responses. Other degradation products have cellular-level toxicity and can interfere with organ homeostasis. These degradation products are produced over time until the polymer support is completely destroyed. These degradation products can be transported through systemic circulation and accumulate sequentially in target organs, and long-term accumulation of these products can damage organs.

[0040] For the regeneration process to be optimal and complete, the degradation rate of the implant's polymer array must be matched to the tissue repair rate. In this way, the decrease in total gel volume remains constant even as the volume of cellular bioburden increases, preventing mechanical compressive stress at the interface with the host organ. Such stress can lead to ischemic compression injury. While it is possible to control the degradation rate of polymer hydrogels in vitro under strictly controlled experimental conditions, it is impossible to control the degradation rate of the hydrogel matrix in vivo, nor to monitor the degradation rate of the polymer matrix relative to the cellular bioburden rate in vivo. Therefore, it is impossible to efficiently use biodegradable implants in the regeneration process.

[0041] Overall, the cellular regeneration and degradation processes of the implant begin at the interface in contact with the spinal cord and extend into the lumen. Furthermore, the core of the hydrogel matrix does not completely degrade, remaining in the form of oligomeric chains that have lost their structure and function as a growth substrate. Without any mechanical support, which ultimately hinders reconstruction, the implant becomes unusable.

[0042] When the matrix breaks down, the connections between cells that have migrated into the hydrogel structure may be further detached. These connections are essential for the formation and aggregation of functional tissues.

[0043] Therefore, using biodegradable polymer hydrogels as implants in tissue engineering to optimally repair tissue volume loss, particularly in the nervous system, requires considering both the progression of the hydrogel matrix's mechanical behavior during use to maintain structural integrity until new biological tissue is formed, and the actual in vivo degradation rate of the hydrogel matrix. This use is not possible because, as the matrix degrades, it gradually loses its structural topography and thus the mechanical properties defined by its initial structure. The structural integrity of an implant cannot be maintained over time. The structural integrity of an implant is an essential feature for maintaining the formation of tissue reconstruction throughout the entire tissue reconstruction process.

[0044] Nerve fibers have the ability to regenerate spontaneously after damage, either through elongation or lateral branch sprouting. This ability is significantly reduced if the ends of the regenerating fibers, i.e., the growth cones, cannot find any substrate to adhere to and elongate.

[0045] However, if the matrix is ​​degraded, the substrate through which the growth cone can proceed in the axon regeneration process may not be continuous. Therefore, the degradation rate V d The regeneration rate of the organization is V r It is advantageous to use implants made from the following hydrogels.

[0046] In particular, it is advantageous to use heterogeneous hybrid hydrogels that are non-degradable, i.e., have a degradation rate lower than their regeneration rate. Preferably, non-degradable heterogeneous hybrid hydrogels are polymer compositions that do not undergo in situ degradation by chemical hydrolysis or enzymatic hydrolysis, or cleavage by photodegradation, under physiological conditions equivalent to those of the human body. For example, chemical degradation is low or zero over a reference period equivalent to at least one or two years.

[0047] The degradability test of heterogeneous hybrid hydrogels is favorably carried out at 40°C in an acidic solution with a pH equal to 1 and an alkaline solution with a pH equal to 14. A piece of hydrogel is placed in each of the above solutions, and each solution is maintained at 40°C, for example, by a hot plate. The solutions are stirred. The heterogeneous hybrid hydrogel sample and the solution are observed periodically. For example, after one week, it is observed that the appearance of the non-degradable heterogeneous hybrid hydrogel has not changed and the solution remains clear. No suspended residue of the sample is observed in the solution. Furthermore, no decrease in the hydrogel mass is observed.

[0048] The acidic solution is preferably a 0.1 mol / L hydrochloric acid solution. The alkaline solution is preferably a 1 mol / L soda solution.

[0049] Heterogeneous hybrid hydrogels are also analyzed by high-performance liquid chromatography. A mixture containing 25% volume methanol and 75% volume water is preferred as the eluent. The flow rate of the eluent is preferably 1 mL / min. The column used may be a Nova-Pack C18 150 mm 3.9 mm reversed-phase column. Hydrogel degradation can be detected by chromatographic analysis, by searching for hydrogel components.

[0050] In the non-degradable heterogeneous hybrid hydrogel, chromatographic analysis revealed new peaks appearing very slowly over time. For example, chromatographic analysis was performed to monitor the main components of the hydrogel according to the present invention, such as the hydrogel containing HPMA (N-(2-hydroxypropyl)methacrylamide). High-performance liquid chromatography analysis showed no degradation of any of the HPMA-based hydrogels. These observations are supported by "in vivo" studies in which HPMA-based gels are implanted in the spinal cord. Analysis of cerebrospinal fluid by electrophoresis showed no oligomeric degradation products in the cerebrospinal fluid. Thus, the absence of degradation products indicates that the hydrogel is not degraded.

[0051] It is also advantageous that the hydrogel is compressible rather than bioabsorbable. Such hydrogels can be used as filler implants and can be inserted into organs, such as between the edges of anatomical defects in the nervous system, specifically into intramedullary cystic spaces after trauma.

[0052] However, if the implant does not degrade much or at all, its volume should not hinder cell remodeling. It is particularly advantageous to form a filler implant with a scaffold made from a porous heterogeneous hybrid hydrogel with perforations, thereby enabling cell regeneration within the filler implant. As illustrated in Figure 1, it is particularly advantageous that the heterogeneous hybrid hydrogel 1 has a porous structure, and the pores of this porous structure communicate with each other to create a percolation array 2 in the three-dimensional volume of the gel. This perforation array formed by the pores allows cells to migrate to the core of the hydrogel and proliferate, enabling them to survive by connecting through the transport and diffusion of nutrients necessary for cell metabolism.

[0053] If the hydrogel is non-degradable, the pore array will not induce the degradation of the hydrogel, which would otherwise lead to the accumulation of toxic or irritating molecules within the pores.

[0054] The above-described implant is a porous structure having pores that define percolation pathways within the three-dimensional hydrogel through which the implant passes. These pores facilitate the colonization of cells and blood vessels from the tissue to the implant core, as well as the circulation of bodily fluids, cell growth factors, cell repair factors, and physiological nutrients through the implant. The flow of various fluids through the implant promotes the vascularization of tissues generated within the implant. The implant forms a three-dimensional support matrix, and the pores passing through it induce the growth of cells, nerve fibers, and blood vessels. This cell growth is of higher quality when in contact with the implant than when not in contact with it. The implant must be formed solely from hydrogel.

[0055] Preferably, the porous portion of the hydrogel forming the implant should account for more than 85% of the volume, more preferably at least 90%, and still preferably at least 92%. This porous portion can be calculated using mercury porosimetry techniques.

[0056] It is preferable that the heterogeneous hybrid hydrogel is formed such that the majority of its porous portion consists of pores between 10 and 30 microns in diameter. Preferably, at least 60% of the porous portion consists of pores between 10 and 30 microns in diameter. In short, at least 60% of the porous volume consists of pores between 10 and 30 microns in diameter.

[0057] In heterogeneous hybrid hydrogels, it is also advantageous for pores between 30 and 300 microns in diameter to constitute a larger proportion than pores less than 10 microns in diameter. Preferably, the proportion of pores between 30 and 300 microns in diameter should exceed 20%, and more preferably, exceed 30%. This proportion of pores between 30 and 300 microns in diameter is particularly advantageous for accommodating large organisms such as multicellular tissues. Preferably, the proportion of pores less than 10 microns in diameter should be less than 15%, and even more preferably less than 10%. In heterogeneous hybrid hydrogels, it is even more advantageous for less than 2% of the pore volume to be less than 1 micron in size within the through-porous array.

[0058] This distribution of pore dimensions ensures that biocompounds can reliably pass through the implant, which covers the entire dimensional spectrum of biocompounds in living tissue. As a result, it becomes easier to use the implant as a tissue equivalent of the organ being repaired.

[0059] The pores of the implant described above are adapted to the circulation of chemokines secreted by inflammatory cells. The structure of this implant allows for the infiltration of glial cells, mesenchymal cells, pia mater-associated stem cells / progenitor cells that secrete SDF-1 factor, and stem cells / progenitor cells associated with the epithelium of central channels with neuronal differentiation potential, as well as vascular infiltration and the growth of regenerating nerve fibers.

[0060] It is also advantageous to form heterogeneous hybrid hydrogels from materials that exhibit properties that gradually change the composition of their porous array or the amount of cellular bioburden as cell colony formation progresses. Heterogeneous hybrid hydrogels are deformable under the stresses imposed by cell colony formation and / or cellular bioburden. The implant gradually deforms under the pressures imposed by the cell tissue as regeneration occurs. This hydrogel must be deformable under the mechanical stresses of cellular bioburden, thereby enabling the maintenance of a structure with percolation channels.

[0061] It is particularly advantageous to form a filler implant with a scaffold made from a porous heterogeneous hybrid hydrogel, which is viscoelastically deformable at a sufficiently low degradation rate to be considered a non-degradable heterogeneous hybrid hydrogel. The volume occupied by the porous implant changes as regeneration occurs and the regeneration rates around and inside the hydrogel partially coincide.

[0062] As cell and nerve regeneration occurs, the polymer backbone of the heterogeneous hybrid hydrogel deforms by a constant or nearly constant volume. During cell accumulation in the implant and during the expansion of the new tissue being formed, the polymer array deforms under the mechanical pressure exerted by the cell accumulation. The implant structure is fabricated from a material with an elastic modulus between 1 and 200 kPa to accommodate the stress exerted by the cellular bioburden. The value of the elastic modulus can be measured at 50% of its deformation before cutting.

[0063] As cellular bioburden gradually increases, the implant deforms in response to the mechanical stress introduced by the cellular bioburden within the implant. The implant deforms elastically, and then viscoelastically, particularly in the perforating channels, so as not to hinder cell growth. Even as deformation occurs, the implant maintains a three-dimensional percolation array that stretches to ensure complete tissue regeneration. Once the tissue is reconstructed, the residual polymer array of the implant acts as an intercellular support matrix that stabilizes the new tissue formed by mechanical strain.

[0064] The surface roughness of the implant is increased to improve adhesion to the host tissue by increasing the contact surface area between the implant and the tissue cells in contact, thereby promoting adhesion between the two environments. The pores on the surface of the implant connected to the percolation array allow for the circulation of bodily fluids, cell growth factors, and physiological nutrients necessary for cells, while simultaneously promoting the formation of cell and blood vessel colonies from the contact tissue to the core of the implant.

[0065] The use of heterogeneous hybrid hydrogels capable of forming matrices with elastic properties and porous structures suitable for the therapeutic and regenerative regeneration of organs and / or tissues, specifically as implantable biomaterials, is known from French Patent No. 2,942,408. It has been observed that the degradation rate of such materials is slower than the regeneration rate of cell tissues.

[0066] It is particularly advantageous to improve the heterogeneous hybrid hydrogel, which is known from French Patent No. 2,942,408, to form a non-degradable scaffold having a porous structure that specifically defines percolation channels in three dimensions and exhibits viscoelastic behavior with an elastic modulus between 1 and 200 kPa.

[0067] It is advantageous to use heterogeneous hybrid hydrogels, which are copolymers obtained from at least three monomers: a dendrimer monomer functionalized with an ethylene radical, an acrylamide monomer, and a crosslinking agent. In certain configurations, the heterogeneous hybrid hydrogel contains copolymerizable bioactive materials. In other configurations, the heterogeneous hybrid hydrogel does not contain any copolymerizable bioactive materials. Bioactivation can then be performed on the hydrogel that has been subsequently formed.

[0068] Copolymerizable bioactive materials are selected from the group comprising derivatives of complex sugars, derivatives of tissue adhesion peptides, and derivatives of polymer conjugates conjugated to antibodies against lipid derivatives. Preferably, copolymerizable bioactive materials are methacryloyl or methacrylamide derivatives of each of the complex sugars, tissue adhesion peptides, and polymer conjugates conjugated to antibodies against lipid derivatives.

[0069] The copolymerizable bioactive material can preferably be a methacryloyl derivative or methacrylamide derivative of a complex sugar selected from, for example, glucosamine, N-acetyl-glucosamine, N-diglycidyl-glucosamine, N-acetylgalactosamine, N-acetylneuraminic acid (sialic acid), and polysialic acid.

[0070] The copolymerizable bioactive material can preferably be a methacryloyl derivative or methacrylamide derivative of a tissue adhesion peptide selected from tissue adhesion peptides containing amino acid sequences such as Arg-Gly-Asp, Ile-Lys-Val-Ala-Val, Ala-His-Ala-Val-Ser-Glu, and Tyr-Ile-Gly-Ser-Arg; oligopeptide derivatives of tissue differentiation molecules, such as bone morphogenetic proteins or proteins of the SDF-1 (stromal cell-derived factor-1) family; or chemokines that have the ability to recruit and attract endogenous stem cells expressing the CXCR4 receptor and to stimulate axonal growth when tissue regeneration occurs.

[0071] The copolymerizable bioactive material can preferably be a methacryloyl derivative or methacrylamide derivative of a polymer conjugate bound to an antibody against myelin and its axon-associated lipid derivative.

[0072] The acrylamide monomer is preferably an N-substituted methacrylamide or an N-substituted acrylamide. The value of the modulus is determined in part by the crosslinking density of the hydrogel, i.e., by the number of covalent bonds between polymer chains formed from the acrylamide monomer, preferably by monitoring the number of covalent bonds between polymer HPMA chains.

[0073] N-substituted methacrylamides are preferably N-monoalkylmethacrylamides, N,N-dialkylmethacrylamides, N-hydroxyalkyl methacrylamide, In preference, N-(2-hydroxypropyl)methacrylamide (HPMA), N-alkyl, N-hydroxyalkyl methacrylamide, The following are selected from the group consisting of N,N-dihydroxyalkylmethacrylamide.

[0074] The N-substituted acrylamide is preferably an N-monoalkylacrylamide, an N-hydroxyalkylacrylamide, N,N-dialkylacrylamide, N-alkyl, N-hydroxyalkylacrylamide, and Selected from the group consisting of N,N-dihydroxyalkylacrylamides.

[0075] The dendrimer monomer preferably comprises a central core A and polymeric dendritic branches of polyethylene oxide (PEO). Only one of the dendritic branches is functionalized, preferably at its terminal position, by an ethylene radical. The ethylene radical is unsaturated in order to react with monomers having at least one reactive vinyl double bond. The other dendritic branches are preferably terminated by a hydroxyl functional group and do not contain an ethylene radical. The other dendritic branches can further be functionalized with other functional groups, such as ester or amide functional groups, prior to copolymerization. The ethylene radical is preferably an acrylate radical or a methacrylate radical. Because the dendrimer structure is star-shaped, it is possible to introduce a polyfunctional, variable geometric shape into the hydrogel, allowing it to respond specifically to multiple interactions expected in other materials. An index indicating that only one branch is functionalized generally corresponds to a statistical result indicating that the dendrimer monomer is functionalized by one branch containing an unsaturated ethylene radical.

[0076] Functionalized hydroxyl groups in heterogeneous hybrid hydrogels can be modified by attaching entities such as polypeptides, active ingredients, ligands, polymerizable groups, or bioactive agents such as oligosaccharides.

[0077] According to a particular embodiment, the central core A is a carbosilane, a polycarbosilane, a star-shaped polycarbosilane, or the following formula (1): [ka] Selected from the bases by, In the formula, n is an integer between 1 and 20, preferably equal to 6.

[0078] In one variant, the central core A is poly(divinylbenzene).

[0079] According to the preferred embodiment, the dendrimer monomer is given by the following formula (2): [ka] It is equivalent to the above, and in the formula, R is either H or CH3. X and Y are integers between 1 and 100, and the sum of X + Y is a multiple of 4. Z and Z' are either the same or different, and their values ​​are between 1 and 100. A is preferably a carbosilane, a polycarbosilane, a star-shaped polycarbosilane, or a group according to formula (1).

[0080] According to another preferred embodiment, the dendrimer monomer is given by the following formula (3): [ka] It is equivalent to the above, and in the formula, R is either H or CH3. n is an integer between 1 and 20, preferably equal to 6. Z and Z' are either the same or different, and their values ​​are between 1 and 100. X is equal to 1, 2, or 3. Y is equivalent to the equation Y = 4 - X.

[0081] Advantageously, the dendrimer monomer is functionalized with a methyl methacrylate radical and has four dendritic branches of PEO and a silane center core A having a hexanol crosslink, i.e., R=CH3, n=6, X=3, Y=1, and Z=Z', the structure of formula (3). For clarity, this dendrimer monomer is identified by the notation Si-PEO4-MMA.

[0082] The mechanical and chemical properties of hydrogels can also be tuned by grafting functional groups with specific properties, such as hydrophilicity, hydrophobicity, and / or tensile activity, onto the free hydroxyl terminals of dendritic branches. Hydroxyl functional groups can be readily functionalized due to the nucleophilicity of their oxygen atoms. These functional groups are known to be readily activated, for example, by basic treatment. The unique properties of heterogeneous hybrid hydrogels can also be modified by varying the number of dendrimer monomers incorporated into the hydrogel skeleton and the properties of the central core A. The solubility of polymers in water and organic solvents is tuned by the number of dendritic branches.

[0083] Heterogeneous hybrid hydrogels are primarily formed or composed of multiple microbeads that aggregate to define a transpermeable porous array. The microbeads are spherical or nearly spherical in shape, primarily with a diameter between 1.5 microns and less than 10 microns. Preferably, the microbeads have a diameter primarily between 2 and 5 microns to better regulate channel deformation in response to stresses imposed by cellular bioburden.

[0084] To better control the three-dimensional structure of a three-dimensional percolation array, it is preferable to aggregate hydrogel microbeads in the form of aggregates comprising at least 5 to less than 50 microbeads, preferably at least 10 microbeads. Advantageously, the microbeads aggregate with each other to form aggregates of 10 to 30 microbeads. It is also particularly advantageous to aggregate the hydrogel microbeads in a configuration called a "grape cluster," which facilitates the efficient formation of a porous array while ensuring good deformability of the scaffold. The aggregates are preferably formed of microbeads with a diameter between 5 and 10 microns. The aggregates preferentially form the main array of the hydrogel matrix. The aggregates are fixed to each other by attachment points, allowing one aggregate to move relative to another. In the grape cluster configuration, the cross-section becomes larger from one end to the other and is approximately circular perpendicular to the length of the aggregate. This grape-like three-dimensional structure is advantageous over a linear structure for promoting the formation of multicellular tissue structures and / or nerve fiber growth by facilitating cell-implant interaction. This grape-like structure also improves adhesion.

[0085] Within the aggregate, microbeads are fixed to each other by deformable crosslinking points, allowing them to move relative to one another and adapt the three-dimensional structure of the heterogeneous hybrid hydrogel scaffold in response to applied mechanical stress. The aggregate can be compressed. Polymer chains deform between crosslinking points by stretching. As the crosslinking points move away from each other, the porous portion of the hydrogel increases. As cell and nerve regeneration occurs, the overall porous volume of the implant increases at a constant rate, forming an anisotropic structure.

[0086] The microbead aggregates are constructed in a manner that allows them to move relative to each other, enabling contraction of the microbead aggregates in response to capillary pressure exerted by migrating cells in contact with the surface of the microbeads. The microbead aggregates are configured such that their contraction expands the porous array, thereby increasing the porous volume available for expansion of the cortex.

[0087] The aggregates, by being fixed to one another, form a porous array comprising mesoporous and macroporous regions. The multiple pores defined by the aggregates connect to one another, forming a percolation array that exhibits a certain degree of flexibility and passes through the hydrogel in its initial configuration. At a microscopic level, this percolation array allows cells, blood vessels, and nerve fibers to infiltrate in three-dimensional space through the hydrogel when the tissue remodeling process occurs. In parallel, the microbeads define micropores with a size of less than 20 nm, preferably varying in size within the range of 1.5 to 11 nanometers, and an average diameter of 6 nanometers. These micropores are closed and not interconnected. The micropores form craters on the surface of the microbeads, thereby creating a honeycomb-like surface on the hydrogel microbeads. These micropores create surfaces, contributing to an increase in the internal surface area of ​​the polymer hydrogel array formed from these hydrogel microbeads. As the surface topography of the polymer array and microbeads increases in this way, interactions with the cell membrane, particularly the focal adhesion of migrating cells, are promoted.

[0088] In preference, the above percolation array should be at least 25m 2 / g, advantageously, at least 50m 2 Define a large internal ratio surface equal to / g.

[0089] In prior art, it was observed that the mass of cellular bioburden increased over time due to the degradation of the hydrogel matrix, partially filling the volume reduced by the hydrogel matrix. The rate of hydrogel degradation was greater than the rate of cell colony formation, which clearly complicates the distribution of bioburden within the initial volume occupied by the implant. In contrast, with porous, viscoelastically deformable, non-degradable hybrid hydrogels, bioburden increases per volume unit of tissue defect by utilizing the compression and deformation of the hydrogel. Cell colony formation is better controlled. The deformable polymer array autonomously organizes, migrates, and differentiates to form new functional tissues, and is modified over time as cells grow.

[0090] Forming implants with non-uniform crosslinking densities is particularly advantageous. The degree of crosslinking can be determined by the fabrication method used. While copolymerization of hydrogels by a method that induces phase separation of the reaction mixture is preferred, copolymerization methods in which phase separation occurs from polymerization induced by thermal means are also advantageous. Controlling phase separation by performing crosslinking within a defined temperature range is particularly advantageous. For example, good results were obtained with phase separation between 40°C and 60°C. Crosslinking between 45°C and 55°C yielded scaffolds with better performance. Crosslinking between 49°C and 51°C yielded scaffolds with even better performance. Using such temperature ranges makes it possible to better determine the dimensions of the microbeads.

[0091] Limiting the polymerization rate is also advantageous to ensure the formation of a hydrogel with good mechanical performance and permeable channels. Preferably, the crosslinking time for implant formation should exceed 6 hours, or even 12 hours.

[0092] It is particularly advantageous to fabricate implants in which the aggregate is a region of a hydrogel array with a high crosslinking density. The hydrogel is formed by aggregates that are strongly crosslinked by regions with weak crosslinking. The strongly crosslinked aggregates form mechanically movable support regions that are connected to each other by regions with weak crosslinking.

[0093] The crosslinking points between aggregates are formed by covalent bonds, i.e., chemical bonds that can be considered non-degradable, thereby resulting in good mechanical strength between aggregates. It is advantageous to provide aggregates that are statistically distributed within a hydrogel volume that defines regions with high polymer chain concentration and regions with low polymer chain concentration that form voids containing free water.

[0094] The average distance between two bridge points in a region of strong bridge bridging is less than 20% of the average distance between two bridge points in a region of weak bridge bridging, preferably less than 10% of the average distance between two bridge points in a region of weak bridge bridging.

[0095] Preferably, the proportion of strongly crosslinked regions is at least 60% of the total volume into which the hydrogel penetrates. It is also advantageous for the proportion of strongly crosslinked regions to be less than 80% of the total volume into which the hydrogel penetrates.

[0096] The surface of the implant is not flat. This surface exhibits protrusions resulting from defects, such as aggregates of microbeads.

[0097] To impart elastic properties to the microbeads and ensure deformation of the implant due to cell load, it is particularly advantageous to fabricate a heterogeneous hybrid hydrogel in which the microbeads comprise or are composed of at least 90% by weight, or at least 95% by weight, of an acrylamide compound, such as N-substituted methacrylamide or N-substituted acrylamide. Specifically, the diameter of microbeads made mostly of an acrylamide compound, or made solely of an acrylamide compound, is between 1.5 and 10 microns, preferably between 2 and 5 microns.

[0098] Preferably, the microbeads consist mostly of HPMA by weight, preferably strongly crosslinked HPMA. Advantageously, strongly crosslinked microbeads have a crosslinking rate exceeding 1 mol% of the crosslinker. Using such a crosslinking rate ensures that a sufficient number of lateral chemical bonds exist between the linear polymer chains for each volume unit of the hydrogel. This configuration allows the hydrogel to achieve sufficient aggregation of the polymer array, e.g., HPMA, with a final swelling ratio of 96% (1g water / 1g dry material) at equilibrium. In this case, this swelling ratio is preferably 0.95 mol%. Furthermore, using a weakly crosslinked material (e.g., HPMA) can form bonds between the microbeads, defining the microbead aggregates.

[0099] It is particularly advantageous to provide perforation channels for implants defined by microbeads, the functionalized dendrimer monomers of the microbeads comprising one or more polyoxyethylene polymeric dendritic branches. The polyoxyethylene polymeric dendritic branches are functionalized with one or more active molecules, such as one of the copolymerizable bioactive materials described above. Subsequently, functionalizing the surface of the percolation array can facilitate cell remodeling within the implant.

[0100] The active molecules are selected from the group consisting of derivatives of complex sugars, derivatives of tissue adhesion peptides or peptides with angiogenesis properties, derivatives of peptides that stimulate nerve regrowth, derivatives of peptides that stimulate cell proliferation and differentiation, derivatives of polymer conjugates bound to antibodies against lipid derivatives, and chemokines of the interstitial factor-1 (SDF-1) class.

[0101] It is also advantageous to offer the use of microbeads in which the acrylamide compound is N-(2-hydroxypropyl)methacrylamide (HPMA), and the dendrimer monomer functionalized with an ethylene radical comprises a polyoxyethylene dendrimer functionalized with a bioactive agent that provides bioactive properties for tissue regeneration.

[0102] Similar to French Patent No. 2,942,408, heterogeneous hybrid hydrogels can be fabricated by radical copolymerization. The preferred polymerization method is carried out in a polar organic medium, preferably a dipolar organic medium, such as an acetone / DMSO mixture.

[0103] Advantageously, in addition to radical copolymerization, polymerization-induced phase separation (PIPS) accompanies the copolymerization. In this way, the initial homogeneous solution consisting of monomers and a polar solvent separates during copolymerization to form microbeads. Advantageously, the molar ratio of functionalized dendrimer monomer to crosslinking agent is between 0.1 and 0.8, for example, the PEO-MMA / MbisAA ratio is between 0.1 and 0.8. At this ratio, phase separation defines twisted balls called coils, forming microbeads. In the temperature range of 45°C to 55°C, it becomes even easier to obtain coils mainly made from acrylamide compounds with diameters between 1.5 and 10 microns. At this ratio, the formation of three-dimensional percolation arrays also becomes possible. By selecting functionalized dendrimer monomers with molecular masses between 6,220 g / mol and 23,280 g / mol, it becomes even easier to control the number of coils produced and the size distribution of the coils. Within this specific range, increasing the molar mass of the functionalized dendrimer monomer leads to an increase in the number of coils and a decrease in the size distribution.

[0104] A method for forming a heterogeneous hybrid hydrogel comprises a first reaction step of a reaction mixture, in which coiled polymer chains form insoluble nuclei. As illustrated in Figure 2, the reaction mixture is poured into a tight, cylindrical, thermally conductive mold 3 having a PTFE wall 4 in contact with the reaction mixture. The reaction mixture may be degassed with argon.

[0105] Microbeads aggregate by associating, and these aggregates connect with each other to form a heterogeneous hybrid hydrogel designed to create filler implants.

[0106] Furthermore, it is possible to form filler implants by supplying microbead aggregates along with aggregates of these aggregates using a three-dimensional printing method. The shape of the implant is directly defined by the three-dimensional printing device.

[0107] Copolymerization is carried out using a crosslinking agent. This crosslinking agent can be an acrylamide such as methylenebisacrylamide (MbisAA), its precursor, or a divinyl compound such as divinylbenzene (DVB). The free radical polymerization initiator is selected from known initiators such as azobisisobutyronitrile (AIBN) or benzoyl peroxide.

[0108] The preferred crosslinking agent is methylenebisacrylamide (MbisAA), which has two vinyl groups. Methylenebisacrylamide allows for greater reactivity compared to N-(2-hydroxypropyl)methacrylamide (HPMA), which has only one vinyl group. This enables more rapid incorporation of methylenebisacrylamide into the growth chain that forms the nucleus before microbead formation.

[0109] Polyoxyethylene exhibits neuroprotective effects on nerve membranes and reduces oxidative stress after injury, making the use of polyoxyethylene hydrogels particularly advantageous (Luo et al., "Polyethylene glycol immediately repairs neuronal membranes and inhibits free radical production after spinal cord injury," J. Neurochem, 2002, Vol. 83, p. 471). It is preferable to use hydrogels formed from one or more polymers in a star-branched structure with polyoxyethylene branches to improve the biocompatibility of the hydrogel, particularly by preventing nonspecific absorption of blood proteins, thereby regulating the local inflammatory response by reducing complement activation through cleavage of C3 proteins into peptides responsible for phagocytic cell recruitment (Nilsson, B. et al., "The role of complement in biomaterial-induced inflammation," Mol Immunol, 2007, Vol. 44, p. 82).

[0110] These branched molecules can be modified by bioconjugation at the terminal positions of the polyoxyethylene branches with functional groups of the classes peptides, bioactive sugars, and chemokines. It is particularly advantageous to form hydrogel materials in which a porous structure is formed on a scaffold comprising a structure composed of N-(2-hydroxypropyl)methacrylamide and a branched structure formed by a branched polyoxyethylene polymer vector of functional groups. The functional groups are preferably short peptides such as oligopeptides that interact with cellular integrin receptors, for example, the following sequences: Arg-Gly-Asp(RGD), Arg-Gly-Asp-Ser(RGDS), Ile-Lys-Val-Ala-Val(IKVAV), but are not limited to these. Other functional groups can be oligomeric conjugates with one or more sialic acids, such as sialyl lactose (Neu5Ac-α2,3-Gal-β1,4-Glc), as well as sulfate oligosaccharides such as HNK1 (SO4-3-GlcAβ1-4Galβ1-4GlcNac-R), or fucosylated oligosaccharides, such as Fucα(1-2)Gal. The bioactive agents may also be selected from growth factors that stimulate axonal regeneration, including but not limited to BDNF ("brain-derived neurotrophic factor"), IGF-1 ("insulin-like growth factor"), NT-3 ("neurotrophin"), and GDNF ("glial neurotrophic factor"), or from growth factors that stimulate the proliferation of neural precursors, such as FGF ("fibroblast growth factor"), EGF ("epidermal growth factor"), PDGF ("platelet-derived growth factor"), VEGF ("vascular endothelial growth factor"), PIGF ("placental growth factor"), NGF ("nerve growth factor"), and TGF ("transforming growth factor"), and may form conjugates with POE.

[0111] Bioactive agents that stimulate the recruitment of progenitor stem cells, such as SDF-1 (stromal-derived factor-1) chemokine, G-CSF (granulocyte colony-stimulating factor), GM-CSF (granulocyte-macrophage colony-stimulating factor), or SCF (stem cell factor) cytokines and interleukins (IT-8), which have the ability to attract endogenous stem cells expressing the receptor CXCR4 and stimulate axonal growth when tissue regeneration occurs.

[0112] One embodiment of the hydrogel is advantageously prepared in an inert atmosphere and comprises a hydrogel in which N-(2-hydroxypropyl)methacrylamide is combined with a crosslinking agent containing two vinyl groups, namely N,N'-methylenebisacrylamide or N,N'-methylenebismethacrylamide, in a molar ratio of 100:1 with the dipolar organic solvent acetone / DMSO (93 / 7 v / v) for 30.4% by weight of the total weight of the reaction mixture.

[0113] The reaction mixture is degassed with argon and poured into a tight, cylindrical, heat-conductive mold having a PTFE wall in contact with the reaction mixture. Advantageously, the initial reaction mixture is placed in a metal mold, preferably made from stainless steel with an inner wall coated with polytetrafluoroethylene (PTFE). This reaction mixture is purged directly with an inert gas inside the metal mold. Preferably, the metal mold has a circular cross-section with a height of at least twice its diameter. The use of a metal mold makes it easier to use the reaction mixture with a water bath that maintains the temperature of the metal mold and the reaction mixture during polymerization. This configuration allows for better control of the microbead and pore dimensions.

[0114] Radical polymerization is preferentially carried out at 50°C in the presence of an azobisisobutyronitrile initiator. During polymerization, it has been observed that it is particularly advantageous to raise the temperature of the metal mold and reaction mixture by at least 5°C, preferably 10°C or more. A temperature increase of less than 20°C is advantageous. The maximum polymerization temperature is also advantageous to be less than 70°C, and more preferably less than 65°C, so that the xerogel, i.e., the hydrogel before saturation with water, does not decompose. This temperature increase makes it possible to obtain a more uniform distribution of polymer chains in the hydrogel and better efficiency. Preferably, the temperature increase is carried out after the microbeads or the majority of them have been formed. The temperature increase can be carried out at least 80 minutes after polymerization, and more preferably 90 minutes after polymerization. Polymerization can be carried out on a first temperature plateau used for microbead formation, for example, between 45°C and 55°C, preferably equal to 50°C. Subsequently, a second plateau follows the first temperature plateau, or optionally, a temperature gradient or other form of annealing to a temperature at least 5°C higher than the temperature of the first temperature plateau. The use of two different temperature ranges makes it possible to achieve better control over the dimensions of the microbeads and the aggregates. Preferably, during the first temperature plateau, polymerization is carried out until the oligomer concentration reaches a threshold that allows for the formation of longer chains by oligomer condensation. As a result of oligomer condensation, at least two phases with different densities appear. It is advantageous to perform phase separation by altering the polymerization temperature.

[0115] To detect when the oligomer concentration has reached a threshold, it is advantageous to monitor the absorbance signal of the reaction. For example, the absorbance signal is monitored by absorbance analysis (optical density) in the ultraviolet-visible range. The detection of a sufficient amount of long chains obtained by oligomer condensation can correspond to the absorbance threshold, or the progress of absorbance to reach the threshold. When the threshold is detected, it is understood that the amount of long chains is sufficient. The second polymerization step is carried out at an even higher temperature, preferably at a second plateau at least 5°C higher than the first plateau. In the second plateau, copolymerization can be achieved by accelerating nucleation and continuing polymerization. Crosslinked oligomer coils that form microbeads are formed. The microbeads aggregate randomly, and crosslinking points are formed. During the second polymerization step at an even higher temperature, the absorbance value decreases. Alternatively, the second plateau can be replaced with a temperature gradient where the lowest temperature is at least 5°C higher than the temperature of the first plateau, i.e., a more complex temperature transition.

[0116] During polymerization, it is preferable to tightly close the mold 3, for example, with a polytetrafluoroethylene cover, which is advantageously accompanied by seal 5. Using this cover limits the evaporation of the solvent from the reaction mixture, making it possible to achieve better reproducibility.

[0117] Preferably, the dimensions of the mold are selected such that the hydrogel formed from the mold has a diameter equal to 175 mm and a height equal to 400 mm relative to the hydrated gel. It is also advantageous to use one identical metal part to define several molds in the form of several wells 6. The mold 3 may have openings 7 designed to close the wells 6 in cooperation with screws.

[0118] At the end of polymerization, the xerogel is in a dry or anhydrous state and is removed from the mold. Coating the inner wall of the mold with polytetrafluoroethylene is particularly advantageous because it facilitates the removal of the hydrogel in its dry state and prevents damage to the hydrogel. The xerogel present in the mold is brittle, and therefore becomes a brittle material that is easily damaged when extracted from vials used in prior art methods.

[0119] Washing the xerogel in ethanol / exothermic water allows it to reach a swelling ratio of 96% at equilibrium. Alternatively, ethanol can be replaced with methanol. Advantageously, the xerogel is transferred to a first receptor in the form of a perforated cage, which is preferably made from polytetrafluoroethylene. Preferably, the first cage is housed within a second cage containing a liquid. This liquid may be water, ethanol, methanol, or a mixture of water and ethanol or methanol. The liquid in the second receptor washes the xerogel by passing through the holes in the first receptor.

[0120] The second receptor is particularly advantageous if it is impermeable to visible radiation. The second receptor can be made from polycarbonate.

[0121] The xerogel attached to the first receptor is subjected to a washing cycle, which advantageously involves the application of several successive washing baths. The water content in the various washing baths is increased in order to wash the xerogel and saturate it with water to form a hydrogel.

[0122] The hydrogel is advantageously crosslinked with methylenebisacrylamide in a ratio equal to or substantially equal to 1 mol%, which gives the hydrogel advantageous physical properties as a filler implant.

[0123] For hydrogel to be used surgically as a filler implant handled under sterile conditions, it is necessary to ensure the sterility of the product. The gel is advantageously placed in a cylindrical container made from high-grade PTFE, i.e., a material that does not interact with the gel's chemical properties, and filled with injectable-grade water from which pyrogenic substances have been removed. The PTFE container containing the gel is sterilized by autoclaving at 121°C for 30 minutes and then tightly closed again with a screw-on cover, thereby maintaining the gel's saturation with water and sterility. This container is then placed inside a second polystyrene container with a "snap-safe cap" facing the sterile surgical field. Thus, the PTFE container can be handled in a sterile state. The second polystyrene container may have a height equal to 54 mm, an inner diameter equal to 34 mm, and a thickness equal to 1.5 mm.

[0124] Heterogeneous hybrid hydrogels are advantageously used in filling strategies to modify the natural healing phenomenon by regenerating and filling in the presence of a viscoelastic matrix that has properties to alter the configuration of porous geometric spaces and polymer arrays. The elastic properties of the hydrogel offer the advantage of modifying the mechanical elastic properties of the substrate and reconstructing a mechanical environment similar to that of in vivo cells. Such hydrogels can accept and induce bundles of cells, nerve fibers, and blood vessels.

[0125] Heterogeneous hybrid hydrogels define an elastically deformable, continuous porous medium that is non-degradable and non-bioresorbable, with a geometric shape that conforms to the regenerative dynamics of cells, blood vessels, and nerves. The hydrogel matrix, once implanted in the incision area, adheres to the spinal cord, thus maintaining its attachment to the transplant site by following the movement of the spinal column and the movement of the spinal cord caused by the pulsation of the arteries that irrigate this organ. The implant forms a support structure because its open porous structure allows cellular flow to migrate into the polymer matrix via percolation pathways in the porous array.

[0126] In certain embodiments, heterogeneous hybrid hydrogels are implanted in a graft cavity, such as the medullary cavity. This cavity can be formed by cutting and removing non-viable scar tissue from the inner edge of the injury. The edges of the cavity are formed by healthy nerve tissue. In advantageous embodiments, cerebrospinal fluid is drained after the cavity is formed. This makes it possible to create a "de novo" injury equivalent to an acute injury that reactivates the inflammatory processes of endogenous cell repair (stem cell recruitment, nerve terminal sprouting, angiogenesis). This implantation method further comprises a step of filling the post-traumatic cavity by injecting the hydrogel into the medullary cavity. The hydrogel can be cut to conform to the shape and geometric form of the cavity.

[0127] It is particularly advantageous to partially dehydrate the hydrogel before inserting it into the lumen. After being placed in the lumen, the hydrogel swells upon contact with the fluid circulating in the lumen until the hydrogel implant is in contact with the entire surface of the lumen, thus forming an integral (100%) interface with the white matter of intact nerve tissue surrounding the injury. The hydrogel can absorb large amounts of water and swell in the presence of water and water-containing biological fluids. It is advantageous for the hydrogel to contain at least 80% water by volume at equilibrium. It is advantageous to provide a hydrogel with a water content of 75% or less by volume. Preferably, the implant is dehydrated so that its volume is reduced between 10% and 30%. It is also advantageous to dehydrate the surface of the lumen, e.g., the parenchymal surface. It is preferable to dehydrate the wall with an ophthalmic sponge. The implant is then inserted into the lumen, preferably the parenchymal lumen, and then rehydrated to reach at least 95% of the initial volume, preferably 100% of the initial volume. Hydration to bring the implant to its initial volume can be done in less than one minute. Ideally, an implant volume with a swelling ratio equal to 100% should represent between 80% and 100% of the volume of the cavity to be filled. As the implant is rehydrated, its porous surface comes into contact with the surface of the cavity, thereby improving adhesion. A dehydration gradient that allows for greater hydration in the center than in the periphery is advantageous. Furthermore, performing a large amount of dehydration on the surface improves the quality of subsequent contact with the cavity wall.

[0128] To improve the adhesion between the implant and the lumen wall, it is preferable to dehydrate and then hydrate the wall designed to make contact. The implant may also be in the shape of multiple grape clusters formed by the microbeads. The three-dimensional structure of the implant having the microbeads as described above defines the surface roughness having protrusions and projections. The roughness is preferably 30 micrometers or less, preferably 15 micrometers or less, and more preferably 5 micrometers or less. A roughness of 0.1 micrometers or more is also advantageous. This range of roughness promotes the circulation of interstitial fluid between the hydrogel surface and the stroma. This surface texture is created by the bundle structure of hydrogel microbeads. Adhesion is further improved on a nanometer scale by surface pores on the projections.

[0129] It is particularly advantageous to cover the surface of the hydrogel with biowhite matter surrounding the injury, which contains inverting nerve fibers, exverting nerve fibers, and association nerve fibers. The implant is designed to closely conform to the geometric shape of the surface of the lumen in particular. These steps result in an anatomical reconstruction of the spinal cord at the location of the traumatic area of ​​the spinal cord. To take advantage of the bioadhesion properties of the implant to biological tissue, it is particularly advantageous to cover the implant with biowhite matter surrounding the injury, which contains inverting nerve fibers, exverting nerve fibers, and association nerve fibers.

[0130] To improve the quality of reconstruction, it is advantageous to give the implant bioadhesive properties or to improve the bioadhesive properties of the implant. By matching the surface properties of the implant, the adhesion between the implant and the biological tissue is improved.

[0131] Improving the bioadhesion properties of implants allows them to be inserted into the cavity, enabling high-quality cell reconstruction without the need for surgical sutures on the biological tissue.

[0132] The implant is particularly advantageous because it actively stimulates hemostasis, more specifically the first stage of hemostasis. It is also advantageous that the implant actively stimulates platelet aggregation. In this way, the implant induces blood coagulation, facilitating the achievement of good tissue reconstruction. This specific implant configuration allows for the control of microbleeding as close as possible to the interface between the implant and the lumen, thereby preventing degradation at the interface between the implant and the tissue. In addition to in vitro studies on blood coagulation and platelet aggregation, in vivo studies have also demonstrated the sufficient hemostatic control provided by the implant.

Claims

1. The following monomers: Dendrimer monomers functionalized by a single branch containing an unsaturated ethylene radical, Acrylamide compounds selected from N-substituted methacrylamides and N-substituted acrylamides, and Crosslinking agent An implant comprising a heterogeneous hybrid hydrogel formed by a copolymer obtained from the three of the following: The crosslinking agent is methylenebisacrylamide, and the molar ratio of the dendrimer monomer to the crosslinking agent is between 0.1 and 0.

8. The heterogeneous hybrid hydrogel is mainly formed by a plurality of microbeads having a diameter greater than 1.5 microns and less than 10 microns, with N-substituted methacrylamide and N-substituted acrylamide accounting for the majority of the weight. The microbeads contain at least 90% by weight of the acrylamide compound, The aforementioned microbeads have a crosslinking ratio exceeding 1 mol% of the crosslinking agent. The aforementioned microbeads aggregate to define a through-porous array that defines a three-dimensional percolation pathway. The porous portion of the aforementioned heterogeneous hybrid hydrogel accounts for more than 85% of the volume. In the aforementioned through-porous array, the proportion of pores with a diameter between 30 and 300 microns exceeds 20%, and the proportion of pores with a diameter between 10 and 30 microns exceeds 60%. The microbeads define closed, non-communicating micropores less than 20 nm in size, the micropores form craters on the surface of the microbeads, the microbeads aggregate to form aggregates containing between 5 and 50 microbeads, the aggregates exhibit a higher crosslinking density than in areas with weaker crosslinking, thereby enabling the aggregates to move in relation to each other. The aggregate exhibits a "grape cluster" three-dimensional structure, and each aggregate has a cross-section that increases in size from one end to the other along the length of the aggregate, and the cross-section is substantially circular perpendicular to the length of the aggregate. The proportion of pores with a diameter between 30 and 300 microns exceeds 20%. The heterogeneous hybrid hydrogel has viscoelastic properties and an elastic modulus between 1 and 200 kPa. N-substituted methacrylamides are selected from the group consisting of N-monoalkylmethacrylamides, N,N-dialkylmethacrylamides, N-hydroxyalkylmethacrylamides, N-(2-hydroxypropyl)methacrylamides, N-alkyl, N-hydroxyalkylmethacrylamides, and N,N-dihydroxyalkylmethacrylamides. N-substituted acrylamides are selected from the group consisting of N-monoalkylacrylamides, N-hydroxyalkylacrylamides, N,N-dialkylacrylamides, N-alkyl, N-hydroxyalkylacrylamides, and N,N-dihydroxyalkylacrylamides. The dendrimer monomer comprises a central core A and polymeric dendritic branches of polyethylene oxide, wherein the central core A is a carbosilane, polycarbosilane, star-shaped polycarbosilane, or the following formula (1): 【Chemistry 1】 (In the formula, n is between 1 and 20.) A group selected from the groups by, or a poly(divinylbenzene), or The dendrimer monomer is given by the following formula (2): 【Chemistry 2】 It is equivalent to the above, and in the formula, R is H or CH ３ And, X and Y are integers between 1 and 100, and the sum of X + Y is a multiple of 4. Z and Z' are either the same or different, and their values ​​are between 1 and 100. A is a carbosilane, polycarbosilane, star-shaped polycarbosilane, or a group according to formula (1), or The dendrimer monomer is given by the following formula (3): 【Transformation 3】 It is equivalent to the above, and in the formula, R is H or CH ３ And, n is between 1 and 20. Z and Z' are either the same or different, and their values ​​are between 1 and 100. X is equal to 1, 2, or 3, Y is equivalent to the equation Y = 4 - X. An implant characterized by the following features.

2. The implant according to claim 1, wherein the aggregate contains between 10 and 30 microbeads.

3. The functionalized dendrimer monomer comprises one polymeric dendritic branch of polyoxyethylene, and the polymeric dendritic branch of polyoxyethylene is copolymerizable to one or more derivatives selected from the group consisting of derivatives of complex sugars, derivatives of tissue adhesion peptides, and derivatives of polymer conjugates conjugated to antibodies against lipid derivatives. Methacryloyl derivatives or methacrylamide derivatives of complex sugars, Methacryloyl derivatives or methacrylamide derivatives of tissue adhesion peptides, Glucosamine methacryloyl derivative or methacrylamide derivative, Methacryloyl derivatives or methacrylamide derivatives of N-acetyl-glucosamine, Methacryloyl derivatives or methacrylamide derivatives of N-diglycidyl-glucosamine, Methacryloyl derivatives or methacrylamide derivatives of N-acetylgalactosamine, Methacryloyl derivatives or methacrylamide derivatives of N-acetylneuraminic acid (sialic acid), Polysialic acid methacryloyl derivative or methacrylamide derivative, A methacryloyl derivative or methacrylamide derivative of a tissue adhesion oligopeptide containing an amino acid sequence selected from Arg-Gly-Asp, Ile-Lys-Val-Ala-Val, Ala-His-Ala-Val-Ser-Glu, and Tyr-Ile-Gly-Ser-Arg, and Methacryloyl or methacrylamide derivatives of polymer conjugates bound to antibodies against myelin The implant according to claim 1, functionalized with copolymerizable bioactive materials selected from, wherein the one or more copolymerizable bioactive materials coat the walls of a three-dimensional percolation array.

4. The implant according to claim 3, wherein the walls of the three-dimensional percolation array are functionalized with several different copolymerizable bioactive materials.

5. The implant according to claim 1, wherein the functionalized dendrimer monomer comprises a central core A and polyoxyethylene polymeric dendritic branches, and one of the dendritic branches is functionalized with a polymerizable acrylate radical or methacrylate radical.

6. The implant according to claim 1, wherein the acrylamide compound is N-(2-hydroxypropyl)methacrylamide (HPMA), and the functionalized dendrimer monomer comprises a polyoxyethylene dendritic branch surrounded by a bioactive agent that imparts bioactive properties for tissue regeneration.

7. A permanent filler implant manufactured from the implant described in claim 1.

8. A method for producing a heterogeneous hybrid hydrogel for implants according to Claim 1, comprising the following monomers: A dendrimer monomer functionalized by one branch containing an unsaturated ethylene radical, wherein the other branches lack an ethylene radical, Acrylamide compounds selected from N-substituted methacrylamides and N-substituted acrylamides, and Three components: at least one bifunctional unsaturated ethylene crosslinking agent having two reactive vinyl bonds, Free radical initiators and The reaction mixture comprises a reaction mixture comprising the formation of microbeads by copolymerization at a temperature between 45°C and 55°C and phase separation induced by free radical copolymerization. The aforementioned microbeads define closed, non-communicating micropores with a size of less than 20 nm. The aforementioned microbeads have a diameter greater than 1.5 microns but less than 10 microns, and the majority of their weight consists of N-substituted methacrylamide and N-substituted acrylamide. The microbeads aggregate to form aggregates containing between 5 and 50 microbeads, and these aggregates are linked to each other by crosslinking points, thereby defining a heterogeneous hybrid hydrogel that defines a through-porous array that defines a three-dimensional percolation pathway, and the through-porous array defines pores, the majority of which are formed by pores with a diameter between 10 and 30 microns, and the proportion of pores with a diameter between 30 and 300 microns exceeds 20%. The heterogeneous hybrid hydrogel has viscoelastic properties and an elastic modulus of 1 to 200 kPa. N-substituted methacrylamides are selected from the group consisting of N-monoalkylmethacrylamides, N,N-dialkylmethacrylamides, N-hydroxyalkylmethacrylamides, N-(2-hydroxypropyl)methacrylamides, N-alkyl, N-hydroxyalkylmethacrylamides, and N,N-dihydroxyalkylmethacrylamides. N-substituted acrylamides are selected from the group consisting of N-monoalkylacrylamides, N-hydroxyalkylacrylamides, N,N-dialkylacrylamides, N-alkyl, N-hydroxyalkylacrylamides, and N,N-dihydroxyalkylacrylamides. The dendrimer monomer comprises a central core A and polymeric dendritic branches of polyethylene oxide, wherein the central core A is a carbosilane, polycarbosilane, star-shaped polycarbosilane, or the following formula (1): 【Chemistry 4】 (In the formula, n is between 1 and 20.) A group selected from the groups by, or a poly(divinylbenzene), or The dendrimer monomer is given by the following formula (2): 【Transformation 5】 It is equivalent to the above, and in the formula, R is H or CH ３ And, X and Y are integers between 1 and 100, and the sum of X + Y is a multiple of 4. Z and Z' are either the same or different, and their values ​​are between 1 and 100. A is a carbosilane, polycarbosilane, star-shaped polycarbosilane, or a group according to formula (1), or The dendrimer monomer is given by the following formula (3): 【Transformation 6】 It is equivalent to the above, and in the formula, R is H or CH ３ And, n is between 1 and 20. Z and Z' are either the same or different, and their values ​​are between 1 and 100. X is equal to 1, 2, or 3, Y is equivalent to the equation Y = 4 - X, Polymerization is carried out in a first temperature plateau, followed by a second plateau or temperature gradient, thereby forming the microbeads, wherein the temperature of the first temperature plateau is between 45°C and 55°C, and the temperature of the second plateau or temperature gradient is at least 5°C higher than the temperature of the first temperature plateau. The method involves pouring the reaction mixture into a tight, cylindrical heat-conducting mold.

9. A method for producing a heterogeneous hybrid hydrogel according to claim 8, wherein the molar ratio of the functionalized dendrimer monomer to the crosslinking agent is between 0.1 and 0.

8.

10. A method for producing a heterogeneous hybrid hydrogel according to claim 9, wherein the molecular weight of the functionalized dendrimer monomer is between 6,220 g / mol and 23,280 g / mol.

11. The reaction mixture that forms the microbeads contains a copolymerizable bioactive material, and the copolymerizable bioactive material is Methacryloyl derivatives or methacrylamide derivatives of complex sugars, Methacryloyl derivatives or methacrylamide derivatives of tissue adhesion peptides, Glucosamine methacryloyl derivative or methacrylamide derivative, Methacryloyl derivatives or methacrylamide derivatives of N-acetyl-glucosamine, Methacryloyl derivatives or methacrylamide derivatives of N-diglycidyl-glucosamine, Methacryloyl derivatives or methacrylamide derivatives of N-acetylgalactosamine, Methacryloyl derivatives or methacrylamide derivatives of N-acetylneuraminic acid (sialic acid), Polysialic acid methacryloyl derivative or methacrylamide derivative, A methacryloyl derivative or methacrylamide derivative of a tissue adhesion oligopeptide containing an amino acid sequence selected from Arg-Gly-Asp, Ile-Lys-Val-Ala-Val, Ala-His-Ala-Val-Ser-Glu, and Tyr-Ile-Gly-Ser-Arg, and A polymer conjugate bound to an antibody against myelin, consisting of a methacryloyl derivative or a methacrylamide derivative. A method for producing a heterogeneous hybrid hydrogel according to claim 8, selected from the group consisting of the following.

12. A method for producing a heterogeneous hybrid hydrogel according to claim 8, wherein the reaction mixture is poured into a mold whose inner wall is coated with polytetrafluoroethylene, and the mold is heated in a water bath.

13. A method for producing a heterogeneous hybrid hydrogel according to claim 8, comprising polymerization of the reaction mixture carried out at a first temperature for at least 80 minutes to form the microbeads, wherein after polymerization the temperature of the mold and the reaction mixture rises by at least 5°C.

14. A method for producing a filler implant, comprising forming a microbead aggregate of heterogeneous hybrid hydrogel as described in claim 8, and assembling the aggregates together by a three-dimensional printing method to form a filler implant.

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