Preparation and application of biocompatible injectable and in-situ gelling hydrogels and cellulose nanofibril-based biocompatible injectable and in-situ gelling hydrogels for tissue and organ repair.

Abstract

Biocompatible in situ gelling and injectable hydrogels composed of cellulose nanofibrils and their preparation and use for tissue and organ repair and regeneration. This invention demonstrates cellulose nanofibrils with tailored size (fibril length and distribution), high crystallinity and purity, variable aqueous concentration, and controlled charge for injectable hydrogels. This application discovers that in situ gelation can be achieved by controlling the degree of substitution of cellulose nanofibrils through carboxymethylation. After injection, they form three-dimensional gels and can provide attachment sites and guidance for cells to promote tissue and organ repair. They can also be loaded with drugs, growth factors, or signaling molecules that can be delivered to target locations within the body and released in a controlled manner. They can be mixed with animal or human cells, including stem cells, and provide transport of stem cells, cells, signaling molecules, and / or active ingredients to the site of injury.

Description

Technical Field

[0001] Cross - Reference to Related Applications This application claims priority and the benefit of its filing date to the disclosure of U.S. Provisional Patent Application No. 63 / 059,342, filed Jul. 31, 2020. The disclosure of that application is hereby incorporated by reference in its entirety.

[0002] Background of the Invention Field of the Invention The present invention relates to a biocompatible cellulose nanofibril aqueous dispersion and its use as an injectable hydrogel. These hydrogels are particularly suitable for injection in animals or humans, with or without drugs, growth factors, extracellular vesicles, or cells, for the purpose of repairing or replacing tissues or organs. Hydrogels, for example, materials composed of hydrophilic polymer networks capable of retaining large amounts of water while maintaining structural integrity in a preferred embodiment, can be biocompatible and can be tuned to deliver therapeutic agents to repair tissues and organs, thus making them attractive biomaterials. In embodiments, the present invention may closely resemble or include properties of the natural extracellular matrix that provide sufficient structural support to cells for the regeneration of tissues or organs. The cellulose nanofibrils described herein are preferred candidates for application as injectable hydrogels for tissue and organ repair and regeneration due to their properties, which include a unique fibril structure exhibiting similar dimensions to collagen in the extracellular matrix, high purity and crystallinity, extremely large water retention capacity, and shear-thinning properties. Not all biomaterials or hydrogels are suitable for injection. Regarding injectability, biomaterials require specific viscosity properties sufficient for optimal or preferred flow. Most hydrogels based on synthetic polymers such as polyethylene oxide, polypropylene oxide, or copolymers of the two have relatively low viscosity and do not have a suitable or preferred effect of shear rate on viscosity. The same is generally true for hydrogels based on natural polymers such as alginate or chitosan. The consequence of low viscosity is a lack of control over injectability, thereby negatively impacting the accuracy of reaching the target tissue or organ site and the effectiveness of the hydrogel in permeating and shaping within the tissue or organ. Cellulose nanofibril aqueous dispersions, such as those described in this invention, possess unique rheological properties (flowability) characterized by solid-like behavior without shear and very strong shear-thinning properties. This is an ideal property for satisfying good injectability.The rheological properties (flowability) according to the present invention can be varied by, in aspect, controlling the concentration of the nanocellulose dispersion and / or by, in aspect, controlling the fibril length and / or fibril distribution during the mechanical homogenization process. Post-injection gelation (e.g., in-situ gelation) can, in aspect, be achieved by controlling the degree of substitution of cellulose nanofibrils through carboxymethylation, as demonstrated and discovered by the present invention. In addition, surface interactions with drugs, growth factors, extracellular vesicles ("EVs"), or cells can be regulated by surface modification of cellulose nanofibrils as described herein.

[0003] Such injectable hydrogels can be used, for example, for the repair and reconstruction of soft tissues. Examples include the repair of skin, adipose tissue, or cartilage. Injectable hydrogels can also be used, for example, for wound healing and / or reconstruction of breast, bone, or cartilage, as well as for ligament repair. In some embodiments, injectable hydrogels may be loaded with growth factors such as bone morphogenetic proteins (BMPs) or TGF-beta, and can be used for the repair of hard tissues such as bone in both medical and dental applications. Injectable hydrogels can also be used, for example, to deliver drugs and growth factors for spinal repair or cancer treatment. [Background technology]

[0004] Description of related technical fields The shortage of transplantable organs is a serious global problem. In the United States, approximately 20 people die every day while waiting for a transplant, and a new patient is added to the organ transplant list every 10 minutes (HRSA, USA). The severe organ shortage situation can lead to high mortality rates among those on the waiting list or to people seeking organs through illegal and unethical channels (WHO). The total cost of treating endo-stage organ failure is estimated at $400 billion annually in the United States alone (1). Tissue engineering, for example, which combines biomaterials and cells, is a key alternative that could help solve this global healthcare problem. Hydrogels, materials composed of hydrophilic polymer networks that can retain large amounts of water while maintaining structural integrity, are attractive biomaterials because most are biocompatible and can be modified to deliver therapeutic agents or cells to repair tissues and organs.

[0005] Biopolymers, including proteins such as collagen and polysaccharides such as alginates, have been used as hydrogels for wound dressings, tissue engineering scaffolds, and drug delivery media (2-5). Nanocellulose hydrogels have been shown to function as tissue engineering scaffolds because they offer a combination of biocompatibility, fibril morphology, and water retention capacity (6-8). Basu et al. proposed the use of wood-derived nanofibrillated cellulose in NFC hydrogels for wound healing applications (9). To achieve self-supporting hydrogels, Basu et al. applied ion-induced crosslinking of nanofibers. However, Basu et al. only described carboxylated TEMPO-oxidized NFCs. In the present invention described herein, it has been found that carboxymethylated cellulose nanofibers have increased divalent ion-induced crosslinking ability compared to TEMPO-oxidized NFCs. Therefore, carboxymethylation according to the present invention is a method that can modify the surface of nanofibrils by introducing carboxymethyl groups.

[0006] The rheological properties (viscosity and flow behavior) of cellulose nanofibril dispersions are influenced by fibril length and fibril length distribution. Paakko et al. described enzymatic hydrolysis combined with mechanical shearing and high-pressure homogenization for the preparation of nanoscale cellulose fibrils from wood (10). However, Paakko et al. do not appear to have investigated the effect of the homogenization cycle on fibril length and rheological properties.

[0007] Most injectable hydrogels reported in the literature have relatively low stiffness / robustness, which is a limitation for long-term application for tissue and organ repair. Yang et al. used cellulose nanocrystals to reinforce carboxymethylcellulose hydrogel (11). De France et al. described the use of cellulose nanocrystals to reinforce poly(oligoethylene glycol methacrylate) injectable hydrogel (12).

[0008] Recent reviews describe the potential use of cellulose nanofibril hydrogels in biomedical applications, particularly for novel uses as bioinks for 3D bioprinting and cell culture support materials (13-15). U.S. Patent No. 10,675,379B2, issued June 9, 2020, proposes cellulose nanofibril hydrogels for use as bioinks for 3D bioprinting, cell culture, tissue engineering, and regenerative medicine applications (16). The reference describes various mechanical, enzymatic, and chemical steps for producing cellulose nanofibril dispersions in a liquid medium, in which case the cellulose nanofibrils have a length of approximately 1–100 microns and a width of approximately 10–20 microns, and possess desirable morphological and rheological properties for use as bioinks in 3D bioprinter applications. 3D bioprint samples must be crosslinked before use, and the constructs then need to be surgically implanted. This adds complexity and cost to the intervention, as it requires the presence of a skilled worker in addition to a bioprinter in the operating room. The present invention improves upon more invasive surgical procedures for implantation and, in its embodiment, describes direct injection of hydrogels without the need to use bioinks and 3D bioprinting.

[0009] Other references describe a 3D discontinuous entity for cell culture comprising an aqueous medium and a hydrogel body containing cellulose nanofibrils and / or derivatives suspended in the aqueous medium (17). Another U.S. patent, No. 10,612,003, describes a plant-derived cell culture material for cell culture or cell delivery comprising sterile mechanically disintegrating cellulose nanofibers and / or derivatives in the form of a 3D hydrogel matrix having a nanofiber concentration in the range of about 0.01 to 1.7 wt%, in which case the cellulose nanofibers and / or derivatives are structurally type I cellulose, and multiple cells are uniformly distributed within the three-dimensional matrix (18). Both patents only describe plant nanocellulose hydrogels for in vitro cell culture. [Overview of the project] [Problems that the invention aims to solve]

[0010] In summary, there is a growing need for biocompatible injectable and in-situ gelling hydrogels with robust mechanical properties for tissue and organ repair. The tunicate cellulose nanofibril hydrogels prepared as described in this invention would be a preferred candidate for injectable hydrogels for tissue and organ repair. [Means for solving the problem]

[0011] Therefore, the present invention improves upon the technology and, in aspect, describes the preparation of biocompatible injectable, in-situ gelling hydrogel formulations based on tunicate cellulose nanofibrils for the repair and regeneration of animal and human tissues and organs. Injectable hydrogels are attractive for tissue and organ repair because, in contrast to surgical implantation materials, they exhibit minimally invasive delivery procedures, reduce healing time, lower hospital costs, cause less pain to patients, reduce scarring, and decrease the risk of postoperative infection.

[0012] Summary of the Invention The present invention teaches the preparation of biocompatible injectable, in-situ gelling hydrogels, preferably in embodiments, composed of cellulose nanofibrils derived from tunicates, for applications such as tissue and / or organ repair. After injection, in embodiments, hydrogels composed of cellulose nanofibrils having adjusted size (fibril length and distribution), crystallinity, concentration, surface chemical composition, and charge form a three-dimensional ("3D") matrix that provides attachment sites and guidance to cells to promote tissue and / or organ repair. In embodiments, they may also be preloaded with drugs, growth factors, or signaling molecules, enabling delivery to target sites in the body at release rates adjusted by the properties of the matrix to optimize efficacy. They may be mixed with animal or human cells, including stem cells, to provide transport of stem cells to the site of injury. Instead of using cells or stem cells, it may also be possible to stimulate repair carried out by endogenous cells using extracellular vesicles (EVs) or autologous tissue aspirates. In this case, a biocompatible cellulose nanofibril dispersion may be used as the main component of the injectable hydrogel. The present invention describes the adjustment of the length and length distribution of cellulose nanofibrils by mechanical treatment in a homogenizer, and viscosity adjustment by selecting a range of fibril concentrations to provide the viscosity and shear reduction desired for controlled injectability. The surface of the cellulose nanofibrils may be modified by carboxymethylation to control the surface charge and enable in-situ gelling. For example, the advantages of using cellulose nanofibrils derived from tunicates as a component of injectable hydrogels include, for example, a combination of beneficial rheological properties such as shear reduction and rapid recovery (flow through various tissues after injection), linked to a 3D architecture favorable for cell attachment and growth, in-situ gelling by crosslinking with endogenous calcium or other divalent cations in a physiological environment, and / or biocompatibility. These features indicate that injectable hydrogels based on cellulose nanofibril dispersions are functional injectable hydrogels for tissue and / or organ repair.

[0013] The embodiments of the present invention are not limited to the present invention, 1. Biocompatibility, 2. Water retention capacity and holdings, 3. Rheological properties (suitable viscosity, shear viscosity reduction, and cavity filling capacity), 4. Controlled in-situ gelation, 5. Controlled targeted interactions with other materials such as drugs, growth factors, conductive components, and cells, and 6. Controlled in vivo function It has the ability to confer desirable properties to hydrogels, including one or more of the following.

[0014] Brief explanation of the drawing The accompanying drawings illustrate some specific aspects of embodiments of the present invention and are not intended to limit or define the invention. Together with this specification, the drawings illustrate specific principles of the present invention. [Brief explanation of the drawing]

[0015] [Figure 1A-C] Figure 1A shows AFM images of fibrils homogenized in 9 cycles. Figure 1B shows the normal distribution of fibril length at different homogenization cycle degrees. Figure 1C shows the normal distribution of fibril length at different homogenization cycle degrees. [Figure 2] This shows the effect of shear rate on the viscosity of homogenized hydrogels at various cycle counts. [Figure 3] The effect of mechanical treatment on injectability in 2% and 10% gelatin is shown as a model of different tissue densities. [Figure 4] This shows the effect of cellulose nanofibril concentration on rheological properties. [Figure 5] This study demonstrates the effect of cellulose nanofibril concentration on injectability in soft tissue models composed of 2% and 10% gelatin. [Figure 6] This paper demonstrates gelation experiments using alginate and various tunicate cellulose nanofibril hydrogels. [Figure 7]Shows the experimental configuration for testing the gelation process. [Figure 8] Shows the effect of the addition of a cross-linking agent (100 mM calcium chloride in this example) on the gelation of the selected hydrogel. The cross-linking agent was added 60 seconds later in this example. [Figure 9] Shows the hemoglobin release profiles from cross-linked TUNICELL-alginate hydrogels of various compositions.

Modes for Carrying Out the Invention

[0016] Detailed description of various embodiments of the present invention The present invention is described with reference to specific embodiments having various features. It will be apparent to those skilled in the art that various modifications and variations can be made to the present invention without departing from the scope or spirit of the present invention. It will be understood by those skilled in the art that these features can be used alone or in any combination based on the requirements and specifications of a given application or design. Embodiments including various features can consist of or essentially consist of such various features. Other embodiments of the present invention will be apparent to those skilled in the art from consideration of the present disclosure. The description of the present invention provided is merely exemplary in nature, and accordingly, variations that do not depart from the essence of the present invention are intended to be within the scope of the present invention. All references cited herein are hereby incorporated by reference in their entirety.

[0017] This invention demonstrates that cellulose nanofibrils, in aspects, possess a size (fibril length and distribution), high crystallinity, high purity, variable aqueous concentration, and controlled charge, which are adapted for injection hydrogels. This application shows that in-situ gelation can be achieved by controlling the degree of substitution of cellulose nanofibrils through carboxymethylation. After injection, it can form a three-dimensional gel that provides attachment sites and guidance to cells to promote tissue and organ repair and / or development and / or regeneration. It may also include one or more pharmaceuticals, growth factors, or signaling molecules that are delivered to a target site in the body and released under controlled conditions. It may also be mixed with animal or human cells, including stem cells, to provide, for example, the transport of stem cells or cellular signals to the site of injury.

[0018] The rheological properties (flowability) can be varied in embodiments by controlling the concentration of the nanocellulose dispersion and / or by controlling the fibril length and / or fibril distribution during the mechanical homogenization process. In embodiments, the fibril length and / or distribution can be adjusted by altering the number of cycles in the mechanical homogenization process. This alters the rheological properties of the hydrogel to promote injectability, in-situ gelation, and / or cell viability. Additional cycles in mechanical homogenization will reduce viscosity by reducing the fiber length in embodiments, resulting in improved injectability of the hydrogel.

[0019] In some embodiments, differences in crystallinity and / or purity can be achieved by using different cellulose sources. Tuctor nanocellulose has a preferred crystallinity and purity. In some embodiments, a higher crystallinity can enhance one or more mechanical properties of the hydrogel, such as possessing a specific desirable shape. In some embodiments, a higher purity can enhance biocompatibility and tissue integration in animals and / or humans for tissue and organ repair applications.

[0020] In embodiments, the present invention may enable variable concentrations of cellulose nanofibrils in water or other liquids or solutions. Higher concentrations of cellulose nanofibrils will, in embodiments, result in higher viscosity of the hydrogel and further alter the porosity to accommodate the diffusion of preloaded cell signals or active compounds into the hydrogel. Thus, the concentration of the cellulose nanofibril dispersion can be used to adjust the preferred fluidity of the hydrogel to obtain enhanced cell viability, specific diffusion rates, and / or tissue requirements.

[0021] In the embodiment, the charge of the cellulose fibril can be controlled by varying the degree of substitution. This substitution is a surface modification of the cellulose fibril, in this case, for example, by adding carboxyl groups or carboxymethyl groups to the fibril. By using different numbers or amounts of carboxyl groups or carboxymethyl groups, it is possible to influence the crosslinking kinetics of the cellulose fibril by altering its surface properties, such as charge and its ability to crosslink molecules and fibers in the hydrogel. The addition of carboxyl groups or carboxymethyl groups may enable further modification by grafting functional molecules onto cellulose fibers in the hydrogel, and may enhance functionality in specific applications.

[0022] In some embodiments, this application demonstrates that in-situ gelation can be achieved by controlling the degree of substitution of cellulose nanofibrils through carboxymethylation. In other embodiments, carboxymethylated cellulose nanofibrils can exhibit increased divalent ion-induced crosslinking ability compared to, for example, TEMPO-oxidized cellulose nanofibrils. Therefore, the carboxymethylation according to the present invention is a method or mechanism that can modify the surface of nanofibrils by introducing carboxymethyl groups.

[0023] In embodiments, the hydrogel according to the present invention may also be preloaded with drugs, growth factors, or signaling molecules to enable delivery to target sites within the body. In embodiments, the release rate is adjusted by the properties of the matrix to optimize efficacy and / or safety. In embodiments, surface interactions with drugs, growth factors, extracellular vesicles, cells, and autotissue aspirates can be adjusted by surface modification of the cellulose nanofibrils described herein. In embodiments, variable porosity can be achieved by varying the concentration of nanofibrils in the hydrogel, thereby modifying the diffusion of active molecules from the hydrogel into adjacent tissues, organs, and / or body parts after injection. For example, but not limited to, chemical modification of nanofibers by adding carboxyl groups and / or carboxymethyl groups may enable grafting of functional components into the hydrogel. By modifying nanocellulose via periodate oxidation, etc., oxidized cellulose nanofibrils can be bioconjugated to other biopolymers, such as, but not limited to, fibronectin, laminin, and collagen, for enhanced cell interactions.

[0024] In embodiments, the hydrogel is injected into a wound, tissue, body part, or organ and crosslinked in situ by adding one or more crosslinkable molecules, such as divalent cations, to the hydrogel simultaneously with, nearly simultaneously with, substantially simultaneously with, approximately simultaneously with, or separately over time. The hydrogel may also be crosslinked by the physiological conditions or properties of the wound, tissue, body part, or organ, such as the endogenous calcium concentration. In embodiments, in-situ gelling enables the injection of a hydrogel containing or not containing active ingredients, cells, or materials, e.g., a smooth, comfortable, effective, medically effective, and / or medically safe injection, and subsequently allows the injected hydrogel to be molded or shaped into a desired form within the wound, tissue, body part, or organ. In embodiments, the hydrogel includes mechanical strength, such as high mechanical strength to facilitate in-situ molding or shaping so that the hydrogel can retain a desired shape indefinitely, permanently, temporarily, or over a period of time. Cellulose nanofibers according to the present invention may have mechanical properties that enable in-situ molding and shaping. In-situ crosslinking can allow the injected hydrogel to retain its shape.

[0025] In some embodiments, in-situ gelation includes hydrogel formation or gelation at the injection site and / or immediately after injection of the hydrogel. In some embodiments, gelation may be autogelation or may be caused by the addition of crosslinking molecules, such as, but not limited to, divalent cations such as calcium. For example, in some embodiments, gelation may occur, for example, at or after injection of the hydrogel into human or animal tissue, body part, organ or wound. In some embodiments, this makes it possible to make the hydrogel functional without having to perform more invasive surgical procedures. In some embodiments, the gel becomes hard, rigid, stiff, semi-rigid, more viscous, elastic, semi-elastic, soft or semi-soft at or after injection of the hydrogel. In some embodiments, in-situ gelation includes a hydrogel with lower viscosity before injection, which becomes more viscous or gel in vivo at or inside the injection site, for example, but is not limited to. Therefore, in other applications, injecting a hydrogel may enable the formation of a scaffold. In some embodiments, in-situ gelation includes a hydrogel of lower viscosity for the purpose of injection, and then a hydrogel of higher viscosity at or after injection. In some embodiments, the hydrogel can gel slowly, moderately slowly, quickly, moderately quickly, spontaneously, or almost spontaneously immediately after, at, or after injection. In some embodiments, hydrogel formation occurs in situ with, or without, a crosslinking agent. In some embodiments, the crosslinking agent may be included in or added to the hydrogel before, during, or after injection. In some embodiments, the gel is macroscopic or microscopic. In some embodiments, gelation is reversible or irreversible. In some embodiments, the gel consists solely of a cellulose nanofibril dispersion or also includes one or more polymers, biopolymers, active ingredients, cells, stem cells, cell signals, or EVs. In some embodiments, cellulose nanofibrils are chemically modified, for example, by carboxymethylation, TEMPO oxidation, periodate oxidation, or enzymatic treatment. In some embodiments, cellulose fibers are bioconjugated with other biopolymers, for example, but not limited to collagen, laminin, and / or fibronectin.

[0026] To further the understanding of the present invention, the following examples of specific aspects of several embodiments are given. The following examples should not be construed as limiting the scope of the present invention in any way. [Examples]

[0027] Example 1 The effect of mechanical pretreatment on injectability Dispersions of enzymatically pre-treated tunicate cellulose nanofibrils TUNICELL ETC were homogenized using a high-pressure fluidizer (Microfluidizer M-110EH, Microfluidics Corp. USA) at various cycle counts (number of passes to the homogenizer). The ETC dispersions were then evaluated for fibril length distribution using atomic force microscopy (AFM) (8). Freshly cut mica sheets were treated with a poly-L-lysine (0.01%) solution for 5 minutes and then air-dried. Droplets of diluted ETC dispersion (0.02% dry content) were then deposited onto the mica and incubated for 5 minutes, followed by rinsing with DI water. The dried samples were examined using an AFM NanoScope III scanning probe microscope with a Type G scanner equipped with Nanoscope software (v.4.43; Digital Instruments, Santa Barbara, CA, USA). Measurements were performed in tapping mode using a standard silicon tip (height: 15 μm, radius of curvature: 8 nm) to determine the length and width of nanocellulose fibrils. The calculations were performed using ImageJ (National Institutes of Health, Bethesda, MD, USA) and the average of 5-10 fibrils. The rheological properties of the ETF were assessed using a TA Discovery HR2 rheometer (TA Instruments, New Castle, DE, USA) with a Peltier aluminum plate (diameter 20 mm, gap = 300 μm). To determine the linear viscoelastic region (LVR), the vibration amplitude was set to a frequency of 1 Hz in the range of 0.1 Pa to 1000 Pa. From the LVR, 10 -3 Hz~10 3A force of 10 Pa was selected for vibration frequency measurements in the Hz range. The shear rate was 0.1 s at 25°C. -1 from 1000s -1 The shear viscosity was evaluated by increasing the concentration.

[0028] Table 1 summarizes the results from AFM fibril length determination. The average fibril length after 6 cycles of mechanical treatment was over 3 μm. The average fibril length decreased to 2.64 μm after 9 cycles of run and further decreased to 2.43 μm after 12 cycles of homogenization. The fibril dimensions analyzed by AFM after 9 cycles are shown in Figure 1a, and the fibril size distribution is shown in Figure 1b. It is important to note that mechanical treatment by homogenization reduced the average fibril length. Increasing the mechanical treatment to 20 cycles reduced the average fibril length to 2.27 micrometers, but the fibril size distribution became wider again, probably due to the formation of finer material (see Figure 1c).

[0029] [Table 1]

[0030] Figure 2 shows the effect of homogenization with various cycle counts on the rheological properties of cellulose nanofibril hydrogels. Hydrogels exhibit shear-thinning properties, which generally means that increasing shear rate results in lower viscosity. The lower panel shows an enlarged region at a shear rate of 1 × 1 / s. At this magnification, it is more readily apparent that 20 cycles of homogenization result in lower viscosity at all shear rates, as can be seen in Table 1. The injectability of different hydrogels was compared by observing the shape of the hydrogels as they emerged from the needle after coloring them with an oil-based red dye. In the upper panel of Figure 3, it can be seen that the hydrogel homogenized with 6 cycles formed droplets when dispensed. Increased homogenization resulted in relatively better flow when dispensed through the needle. At 20 cycles, in this example, the dispersion flowed relatively more smoothly through the needle. The injectability of these hydrogels was studied and tested by soft tissue simulations. Gelatin gels were cast at two different concentrations, 2% and 10%, to simulate tissues of different densities. The hydrogels were then dispensed using a needle inserted into the top of a vial containing the gelatin matrix (see lower panel of Figure 3). Evaluation was performed by qualitative comparison of the length, width, and shape of the injected hydrogels after they entered the gelatin matrix. A tendency was observed for improved injectability with increasing homogenization. Hydrogels homogenized for 20 cycles exhibited the narrowest width and longest linear tracks compared to hydrogels homogenized for fewer cycles.

[0031] Example 2 Effect of concentration on injectability Dispersions of enzymatically pretreated tunicate cellulose nanofibrils TUNICELL ETC were prepared by 9 cycles of homogenization and post-processed to increase the concentration. Post-processing included vacuum filtration. The concentration was increased from 2.5% to 3.25% and then to 4%. The effect of concentration on the viscosity-shear rate relationship was investigated using a rheometer under the conditions described in Example 1. Figure 4 shows that all three hydrogels exhibited shear reduction with higher viscosity at higher nanocellulose concentrations across the entire range of shear rates investigated. Injectability was compared by visual inspection of the hydrogels dispensed from a 20-gauge syringe needle and by comparing the shape of the hydrogels dispensed in 2% and 10% gelatin soft tissue models (Figure 5). Through the examples, it was concluded that hydrogels with higher or higher concentrations have preferable injectability.

[0032] Example 3 In situgelization The in-situ gelling ability of cellulose nanofibril dispersions was investigated by crosslinking dispensed hydrogels with different surface charges using calcium chloride and comparing them to alginates. Three different cellulose nanofibril hydrogels were selected for testing: enzymatic tunicate cellulose (ETC), carboxymethylated tunicate cellulose (CTC), and TEMPO-oxidized tunicate cellulose (TTC). Surface charges were determined by zeta potential measurement (ζ potential, i.e., the average charge of the fibrils) using DLS (Nano ZS-ZEN3600; Malvern Instruments, Malvern, UK).

[0033] When the CTC sample was subjected to conductivity titration, the charge density was 367 μmol / g. This corresponds to a degree of substitution (DS) of 0.062. The charge density of the TEMPO-oxidized TTC sample was 664 μmol / g. Alginate (Nova Matrix, Pronova SLG100, Norway) was used for comparison. Table 2 summarizes the zeta potential measurement results.

[0034] Table 2

[0035] The crosslinking solution was 100 mM calcium chloride in DI water. Gelation was screened by dropping the calcium chloride solution onto a matrix dispensed via a 20 gauge needle. Alginate samples were evaluated as a 3% solution in DI water. Gelation was then tested by performing vibration-time measurements for 10 minutes at a 1.5% strain and a frequency of 1 Hz using a Discovery HR-2 rheometer (TA Instruments, Crawley, UK). All measurements were performed at 25°C using a 20 mm plate-plate geometry (gap: 500 μm). 60 seconds after the start of measurement, 1 ml of 0.1 M CaCl2 was dispensed around the sample while collecting data on storage modulus and loss modulus. Figure 7 shows the experimental setup for testing hydrogel gelation. Figure 8 shows the effect of adding the crosslinking agent, 100 mM calcium chloride solution, on the gelation of selected hydrogels. The crosslinking agent was added after 60 seconds. The storage modulus in shear mode, which describes the stiffness of the hydrogel, is presented as a function of time. At time 0, it is possible to observe the difference between an alginate solution that was not a hydrogel before crosslinking and different TUNICELL hydrogels. Unmodified ETC materials have a higher or highest storage modulus, followed by carboxymethylated TUNICELL (CTC) and then TEMPO-oxidized TUNICELL (TTC). After the addition of a 100 mM calcium chloride solution, the storage modulus of the alginate increases due to rapid crosslinking, sometimes immediately. This crosslinked hydrogel sometimes had a higher or highest stiffness than the analyzed hydrogel, sometimes immediately after the addition of the crosslinking agent. Sometimes, unmodified TUNICELL ETC was substantially unaffected by the addition of the crosslinking agent. TTC hydrogels reach equilibrium with a moderate increase in storage modulus, sometimes rapidly or quickly. The CTC hydrogel showed a relatively slower rate of increase in storage modulus, exhibiting a more favorable storage modulus than alginate after 500 seconds. This indicates that carboxymethylation is a suitable modification method for cellulose nanofibril dispersions to provide crosslinking ability and in-situ gelling properties.Table 3 summarizes the storage moduli of the materials analyzed 540 seconds after the addition of a 100 mM calcium chloride solution. CTC had the highest storage moduli among the materials analyzed.

[0036] [Table 3]

[0037] Example 4. Biocompatibility TUNICELL hydrogels were further refined by processing with a high-pressure fluidizer in a validated cleanroom facility. The processing protocol resulted in high crystallinity, high aspect ratio, >99% pure cellulose, and free from contaminated hemicellulose and lignin. Table 4 summarizes the carbohydrate composition of TUNICELL. The released carbohydrates after complete acid hydrolysis of TUNICELL were examined by high-performance anion exchange chromatography using pulsed amperometry detection (HPAEC-PAD) with a Carbopac PA1 column (Dionex, Sunnyvale, CA, United States) on an ICS3000 system (Dionex, Sunnyvale, CA, United States). Medical-grade ultra-high purity TUNICELL produced under cleanroom standards was electron beam sterilized and had bioload levels of <10 CFU / ml and endotoxin levels of ≤0.5 EU / ml in compliance with FDA regulations for implantable devices. The endotoxin levels of TUNICELL were tested using Lonza's PyroGene® Recombinant Factor C Assay. Bioload tests were performed in accordance with Section 2.6.12 of the European Pharmacopoeia.

[0038] [Table 4]

[0039] Example 5. Hemoglobin delivery TUNICELL hydrogels were also evaluated for drug and growth factor delivery. As an example, hemoglobin delivery was evaluated to enhance tissue oxygenation and accelerate the wound healing process. Two different hydrogels were prepared using 80:20 and 40:60 TUNICELL-alginate mixtures. Hemoglobin was loaded onto the crosslinked hydrogels using a 100 mM calcium chloride solution by immersing the crosslinked hydrogels in hemoglobin solution. The hemoglobin release rate was evaluated by determining the hemoglobin concentration using UV spectroscopy when the crosslinked hydrogels were placed in HBSS solution. By varying the ratio of TUNICELL to alginate, the composition of the hydrogels was altered, allowing for differential diffusion rates of hemoglobin that could be adjusted to suit wound and wound healing requirements (Figure 9). Faster hemoglobin release was observed with hydrogels using higher concentrations of TUNICELL. Therefore, this was found to be an applicable method for dispensing and delivering hemoglobin to wound tissue under controlled conditions.

[0040] Example 6. Enhancement of cell interactions To control cell adhesion, TUNICELL hydrogels ETC were modified by periodate oxidation followed by bioconjugation of selected extracellular proteins. A solution of 0.73 g of sodium periodate (1.5 mol periodate / anhydrous glucose units) in 5 mL of DI water was added to a glass bottle covered with aluminum foil containing 15 g of ETC (2.4% concentration). The reaction was stirred at room temperature for 24 hours, after which the hydrogel was centrifuged and rinsed with DI water. Next, fibronectin, collagen I, and laminin were bioconjugated to the oxidized construct by adding 1–15 mL of 100 μg protein / mL solution to the oxidized hydrogel and incubating at 37°C for 24 hours. The bioconjugated hydrogel was then briefly rinsed in deionized water and centrifuged to the desired concentration. The bioconjugated hydrogel was used with and without cells for injection into soft tissue to repair defects. Bioconjugated hydrogels exhibited enhanced cell adhesion, contributing to and improving tissue repair. Bioconjugated hydrogels were mixed with human chondrocytes and injected into the joints for cartilage repair. After 28 days, human cartilage developed within the area where the hydrogels were implanted.

[0041] Those skilled in the art will see that the disclosed features may be used individually, in any combination, or omitted, based on the requirements and specifications of a given application or design. When an embodiment refers to a particular feature as "including," it should be understood that the embodiment may alternatively "consist of" or "essentially consist of" one or more of those features. Other embodiments of the invention will become apparent to those skilled in the art by practicing the invention in consideration of this specification. The algae used herein are not limited to macroalgae and microalgae, and all forms of algae.

[0042] Where a range of values ​​is provided herein, it should be noted that each value between the upper and lower limits of that range is also specifically disclosed. These smaller upper and lower limits may likewise be independently included in or excluded from that range. The singular forms “a,” “an,” and “it” include multiple references unless specifically defined in context. This specification and the examples are considered illustrative in nature, and any variations that do not depart from the essence of the invention are intended to be within the scope of the invention. All references cited herein are individually and entirely incorporated herein by reference, and are intended to provide an efficient way to supplement the available disclosure of the invention and to provide background detailing the level of skill of those skilled in the art.

[0043] References As stated above, the following references are incorporated herein by reference in their entirety. 1. WHO, organ failure. And E. Calo, VV Khutoryanskiy, Biomedical applications of hydrogels: A review of patents and commercial products, European Polymer Journal 65(2015)252-267. 2. JA Rowley, G. Madlambayan, DJ Mooney, Alginate hydrogels as synthetic extracellular matrix materials, Biomaterials 20 (1) (1999) 45-53. 3. KY Lee, DJ Mooney, Alginate: properties and biomedical applications, Prog. Polym. Sci. 37 (1) (2012) 106-126. 4. D. J. Overstreet, D. Dutta, S. E. Stabenfeldt, B. L. Vernon, Injectable hydrogels, J. Polym. Sci. B: Polym. Phys. 50 (13) (2012) 881-903. 5. S. Van Vlierberghe, P. Dubruel, E. Schacht, Biopolymer-based hydrogels as scaffolds for tissue engineering applications: a review, Biomacromolecules 12 (5) (2011) 1387-1408. 6. G. Helenius, H. Backdahl, A. Bodin, U. Nannmark, P. Gatenholm, B. Risberg, In vivo biocompatibility of bacterial cellulose, J. Biomed. Mater. Res. A76 (2006) 431-438. 7. Backdahl, H.; Helenius, G.; Bodin, A.; Nannmark, U.; Johansson, B. R.; Risberg, B.; Gatenholm, P. Mechanical properties of bacterial cellulose and interactions with smooth muscle cells. Biomaterials 2006, 27 (9), 2141-9. 8. Apelgren, P., Karabulut, E., Amoroso, M., Mantas, A., Martinez Avila, H., Kolby, L., Kondo, K., Toriz, G., and Gatenholm, P., 2019. In Vivo Human Cartilage Formation in Three Dimensional Bioprinted Constructs with a Novel Bacterial Nanocellulose Bioink, ACS Biomater. Sci. Eng. 5, 2482-2490. 9. Basu A, Lindh J, Alander E, Stromme M, Ferraz N. On the use of ion-crosslinked nanocellulose hydrogels for wound healing solutions: physicochemical properties and application-oriented biocompatibility studies. Carbohydr Polym 2017; 174: 299-308. 10. Paakko M, Ankefors M, Kosonen H, et ak. Enzymatic hydrolysis combined with mechanical shearing and high pressure homogenization for nanoscale cellulose fibrils and strong gels. Biomacromolecules 2007; 8(6):1934-1941. 11. Yang, X.; Bakaic, E.; Hoare, T.; Cranston, E. D. Injectable Polysaccharide Hydrogels Reinforced with Cellulose Nanocrystals: Morphology, Rheology, Degradation, and Cytotoxicity. Biomacromolecules 2013, 14 (12), 4447-4455. 12. De France, K. J.; Chan, K. J. W. W.; Cranston, E. D.; Hoare, T. R. Enhanced Mechanical Properties in Cellulose Nanocrystal-Poly (oligo Ethylene Glycol Methacrylate) Injectable Nanocomposite Hydrogels through Control of Physical and Chemical Cross-Linking. Biomacromolecules 2016, acs.biomac.5b01598. 13. Du H, Liu W, Zhang M, Si C, Zhang X, Li B., Cellulose nanocrystals and cellulose nanofibrils based hydrogels for biomedical applications. Carbohydr Polym. 2019 Apr 1; 209: 130-144. doi: 10.1016 / j.carbpol.2019.01.020. 14. del Valle JL, Diaz A, Puiggali J. Hydrogels for biomedical applications: cellulose, chitosan, and protein / peptide derivatives. Gels 2017; 3: 1-28. 15. Curvello R, Raghuwanshi VS, Gamier G, Engineering nanocellulose hydrogels for biomedical applications. Advances in Colloid and Interface Science 267 (2019) 47-61. 16. Patent US10, 675, 379B2 Cellulose nanofibrillar bioink for 3D bioprinting for cell culturing, tissue engineering and regenerative medicine applications. 17. Lou Y. R. et ah, The use of Nanofibrillar cellulose hydrogel as a flexible three-dimensional model to culture human pluripotent stem cells, Stem cell and development 2014; Vol.23: 4, pp. 380-392. 18. U.S. Patent No.US10, 612, 003A Plant derived cell culture material

Claims

1. An aqueous dispersion for injection containing cellulose nanofibrils, wherein the cellulose nanofibrils are biocompatible and derived from one or more tunicated animals, the aqueous dispersion exhibits shear reduction during injection, and the aqueous dispersion contains cellulose nanofibrils that form a cross-linked hydrogel in situ.

2. The aqueous dispersion for injection according to claim 1, wherein the cellulose nanofibril is modified by a chemical modification selected from one or more of the following: carboxymethylation, TEMPO oxidation, periodate oxidation, enzymatic treatment, or a combination thereof.

3. The aqueous dispersion for injection according to claim 1 or 2, wherein the cellulose nanofibrils are derived from wood, one or more plants, bacteria, algae, or a combination thereof.

4. The aqueous injectable dispersion according to claim 1 or 2, wherein the aqueous injectable dispersion is used for tissue repair, organ repair, tissue regeneration, organ regeneration, cell therapy, cancer treatment, or a combination thereof in humans or animals.

5. The aqueous dispersion for injection according to claim 1 or 2, further comprising one or more biopolymers, one or more synthetic polymers, or a combination thereof.

6. The aqueous dispersion for injection according to claim 5, wherein the one or more biopolymers are selected from one or more of alginate, hyaluronic acid, collagen, laminin, fibrin, dextran, gellan, and chitosan.

7. The aqueous dispersion for injection according to claim 1 or 2, further comprising one or more pharmaceuticals, pharmaceutical compounds, pharmaceutical agents, therapeutic compounds, therapeutic agents, or drugs.

8. The aqueous dispersion for injection according to claim 7, which is injected into a human or animal to deliver one or more of the pharmaceutical, pharmaceutical compound, pharmaceutical agent, therapeutic compound, therapeutic agent, or drug to a target area, body part, tissue, organ, or combination thereof.

9. The aqueous dispersion for injection according to claim 8, having a controlled release rate.

10. The aqueous dispersion for injection according to claim 1 or 2, further comprising one or more growth factors, one or more signaling molecules, or a combination thereof.

11. The aqueous dispersion for injection according to claim 10, which is injected into a human or animal to deliver one or more growth factors, one or more signaling molecules, or a combination thereof to a target region, body part, tissue, organ, or a combination thereof.

12. The aqueous dispersion for injection according to claim 11, having a controlled release rate.

13. The aqueous dispersion for injection according to claim 1 or 2, further comprising human or animal cells.

14. The aqueous dispersion for injection according to claim 13, which is injected in a human or animal to deliver the human or animal cells to a target region, body part, tissue, organ or combination thereof, and which is injected in a human or animal to act as a scaffold.

15. The aqueous dispersion for injection according to claim 1 or 2, which is bioconjugated with one or more adhesion proteins, one or more other molecules that affect cell adhesion, or a combination thereof.

16. The aqueous dispersion for injection according to claim 15, which is injected into a target region, body part, tissue, organ or combination thereof in a human or animal for attracting, binding, acting as a scaffold or a combination thereof.

17. The aqueous dispersion for injection according to claim 1 or 2, further comprising a dispersion containing fibrils having a length of 0.05 to 20 μm.

18. An aqueous dispersion for injection according to claim 1 or 2, comprising a dispersion having a solid content of more than 0.5% by weight and less than 5% by weight.

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