Patent Application
20250041486
The present disclosure relates to a method of providing an implantable hydrogel material, an implantable hydrogel material and use thereof.
Injured and diseased tissues or organs have been traditionally treated or replaced by autologous grafts, allogenic grafts, or synthetic or natural-based biomaterials. However, there is a huge global shortage of tissue grafts. Donor tissue grafts may also cause donor site morbidity and loss of organ functionality and allogenic grafts are associated with the risk of disease transmission and often require the use of immunosuppressant drugs.
As for synthetic biomaterials, although many of them have achieved widespread clinical use, seamless integration and immunological response issues remain. These issues have led to the more recent paradigm shift to the development of tissue-engineered biomaterials that mimic the extracellular matrix (ECM) of the natural tissue.
For these biomaterials to be successful, they need to be mechanically robust and elastic to support and maintain tissue structure, and cell friendly and bio-interactive to allow seamless host-biomaterial integration that helps restore tissue functionality and avoid rejection.
Collagen is a family of ECM macromolecules within the body that contribute to both mechanical properties and biological function of various types of tissues such as cornea, skin, bone, tendons, ligaments, blood vessels, and the heart. Although native collagen is very robust and elastic due to the presence of natural crosslinks and proteoglycans, extracted collagen is rapidly degraded and lacks the mechanical toughness and elasticity, due to the dissociation of natural cross-links and removal of proteoglycans during isolation and purification process.
Chemical crosslinking techniques are often used to enhance the mechanical properties of collagen hydrogel biomaterials. This, however, may result in collagen-based scaffolds that are either too soft or too brittle and that do not actively interact with the body cells and tissues.
To improve the quality of collagen hydrogels, other materials may be incorporated in the collagen material. Lu et al., Composite aerogels based on dialdehyde nanocellulose and collagen for potential applications as wound dressing and tissue engineering scaffold, Composites Science and Technology 94 (2014) 132-138, reported preparation and characterization of a non-transparent aerogel based on collagen and aldehyde activated nanocellulose fibers (NCFs). Harsh chemicals such as sodium periodate were used to functionalize NCF aldehyde groups for crosslinking with collagen. Lohrasbi et al., Collagen/cellulose nanofiber hydrogel scaffold: physical, mechanical and cell biocompatibility properties, Cellulose 27(11), January 2020, discuss preparation and characterization of a non-transparent composite hydrogel based on collagen and cellulose nanofibers, wherein the mixture was partially stabilized by increasing pH and temperature, thereby forming a physically stabilised hydrogel. In U.S. Pat. No. 8,691,974 is shown three-dimensional bioprinting of biosynthetic cellulose implants and scaffolds for tissue engineering. Fermentation techniques are used for making biosynthetic cellulose (BC) using bacterial fermentation process controlling the 3D shape. In these implants there is no collagen.
Even though cellulose is a natural biopolymer, it is not one of the components of the ECM. Therefore, implantable scaffolds that are mainly comprised of cellulose and its derivatives alone may not be the best options for implantable devices or products.
Scaffolds made from high ratios of nanocellulose may not be permeable to nutrients and oxygen, why not ideal for tissue engineering applications at high ratios. Furthermore, the nanocellulose form (i.e., crystallinity, hydration and swelling) of cellulose may affect the degree of degradation, absorption and immune response (Lin et al., Nanocellulose in biomedicine: Current status and future prospects, European Polymer Journal 59 (2014) 302-325).
In view of the above, there is a need for implantable collagen-based materials that are mechanical robust, while not disrupting the collagen backbone and therefore not compromising biological properties for implantation. The collagen-based material should also be elastic and biocompatible to support and maintain tissue structure, and further cell friendly and bio-interactive, allowing seamless host-biomaterial integration that helps restore tissue functionality and avoid rejection.
It is an object of the present disclosure to provide an implantable collagen-based hydrogel material that alleviates or overcomes at least some of the disadvantages with known collagen-based hydrogel materials.
The invention is defined by the appended independent patent claims. Non-limiting embodiments emerge from the dependent claims, the appended drawings and the following description.
According to a first aspect, there is provided a method of providing an implantable hydrogel material. The method comprises to provide a solution having a collagen concentration of 0.1 to 30 wt. % and a cellulose nanofiber concentration of 0 to 30 wt. %. To the solution is added a non-polymeric short-range carbodiimide crosslinking agent to a concentration of 1-30 wt. % and riboflavin to a concentration of 0.1 to 10 wt. %. Thereafter the formed intermediate hydrogel material is exposed to ultraviolet A light, thereby forming an implantable hydrogel material.
In the above-described method, an implantable collagen-based hydrogel material is formed, which comprises collagen and which optionally also is nano-reinforced with an amount of cellulose nanofibers.
With “hydrogel material” is here meant a collagen network, which exhibits the ability to swell in water or aqueous solution without dissolution, retaining a significant portion of water or aqueous solution within its structure.
The use of non-polymeric short-range carbodiimide crosslinking agent and riboflavin (vitamin B2) with ultraviolet A light (UVA) exposure results in a double chemical-photochemical crosslinking of the material, i.e., both a chemical and physical crosslinking. The double-crosslinking stabilizes collagen in a viscous form. The crosslinking agents do not become integrated within the final implantable hydrogel material, resulting in an entirely natural, chemical-free, transparent hydrogel material, stabilized collagen.
A collagen molecule cross-links with itself, also called intra-molecular cross-links, which is generally done via the short-range small carbodiimide crosslinking agent. Riboflavin is not expected to result in such crosslinks due to their larger molecular size. Riboflavin has bulky cyclic groups (e.g., isoalloxazine) that may limit their penetration depth into the molecules.
A collagen molecule cross-links with another collagen molecule in its vicinity, also called inter-molecular cross-link. This is generally done via short-range small carbodiimide crosslinkers and larger size crosslinkers such as riboflavin in the presence of UVA. The larger spacing between collagen molecules allows riboflavin to penetrate in and creates crosslinks. A collagen fiber cross-links with another collagen fiber in its vicinity, also called inter-fibrillar crosslink. Short-range small carbodiimide crosslinkers cannot produce such crosslinks because of the large spacing between the collagen fibers. However, riboflavin in the presence of UVA can result in crosslinks between two collagen fibers.
Such double-crosslinked collagen material is mechanically robust, tough, and stress tolerant, translating into an ability to withstand surgical implantation, mechanical forces and enzymatic degradation in vivo. The formed hydrogel material is not only robust, elastic, and suturable but also interact with cells and the surrounding tissues when implanted into animal and human models.
The hydrogel material may optionally be nano-reinforced with an amount of cellulose nanofibers (CNF). CNF is added to the collagen to reinforce the collagen material even more than what may be obtained by the double cross-linking of pure collagen. The CNF added to the material may enhance collagen mechanical properties without changing biocompatibility, transparency and other biological properties of the implantable hydrogel material. Even though cellulose is a natural biopolymer, it is not one of the components of the extracellular matrix (ECM). Therefore, implantable scaffolds that are mainly comprised of cellulose and its derivatives alone may not be the best options for implantable devices or products.
The CNFs can be derived from various sources such as Ciona intestinals, wood pulp, and bacterial sources. A CNF molecule cross-links with itself through intra-molecular crosslinking. This is generally done via short-range small molecule carbodiimide crosslinkers and not the larger riboflavin.
A CNF molecule cross-links with another CNF molecule, also called inter-molecular or inter-fibrillary crosslinking. This is generally done via short-range small molecule carbodiimide crosslinkers when the CNFs are in each other vicinity and/or by larger size crosslinkers such as riboflavin in the presence of UVA when CNFs are more distanced. The larger spacing between CNF molecules allows riboflavin to penetrate in and creates crosslinks. This type of crosslinks can also occur between two CNF fibers that are separately and covalently attached to two different collagen fibers creating a bridge between two collagen fibers. This is one of the main advantages of mixing collagen and CNF and using riboflavin/UVA crosslinking.
A CNF molecule cross-links with a collagen molecule in its vicinity, also called inter-fibril-molecular, this is generally done via short-range small molecule carbodiimide crosslinkers and larger size crosslinker, riboflavin in the presence of UVA. The larger spacing between CNF molecules allows riboflavin to penetrate in and creates crosslinks.
With UVA light exposure, or near UV, the intermediate hydrogel material is exposed to light having a wavelength of 315-400 nm. The solution is allowed to react in UVA for at least 1 min, or at least 5 min, or at least 10 min, such as 1-60 min, 5-60 min, 10-60 min or 15-60 min.
The solution in which the collagen molecules, optionally the CNFs, the short-range small molecule carbodiimide crosslinkers and the riboflavin are mixed may be water or a buffer such as PBS. The collagen molecules may be dissolved in one solution and the CNFs in another solution, which then are mixed and thereafter the short-range carbodiimide crosslinkers and the riboflavin are added. Alternatively, the collagen molecules and the CNFs may be dissolved in the same solution.
The concentration of collagen in the formed mixture is 0.1-30 wt. %, or 0.5-30 wt. %, or 1-30 wt. %, or 5-30 wt. %, or 10-30 wt. %, or 15-30 wt. % or 20-30 wt. %, or 25-30 wt. %, or 0.1-25 wt. %, or 0.1-20 wt. %, or 0.1-15 wt. %, or 0.1-10 wt. %, 0.1-5 wt. %, or 0.1-1 wt. %, or 1-5 wt. %, or 5-10 wt. %, or 10-20 wt. %, or 20-25 wt. %.
The concentration of cellulose nanofiber in the solution may be 0-30 wt. %. When the solution comprises 0 wt. % this means that there is no cellulose nanofiber in the solution, only collagen molecules. At the higher concentrations of cellulose nanofibers, the hydrogel material might not be transparent or fully transparent (i.e., having a transmission of at least 80%). Depending on application area, different degrees of transparency might be needed. If used for example in the eye, a transmission of 80% or more would be desirable, and a content of cellulose nanofibers of about 10 wt. % or lower would give the desired transparency.
The concentration of the carbodiimide crosslinking-agent in the mixture may be 1-30 wt. %, or 5-30 wt. %, or 10-30 wt. %, or 15-30 wt. %, or 20-30 wt. %, or 25-30 wt. %, or 1-25 wt. %, or 1-20 wt. %, or 1-15 wt. %, or 1-10 wt. %, or 1-5 wt. %, or 5-10 wt. %, or 10-20 wt. %. 5
The concentration of riboflavin in the mixture may be 0.1 to 10 wt. %, 0.5 to 10 wt. %, 1 to 10 wt. %, 2 to 10 wt. %, 3 to 10 wt. %, 4 to 10 wt. %, 5 to 10 wt. %, 6 to 10 wt. %, 7 to 10 wt. %, 8 to 10 wt. %, 9 to 10 wt. %, 2 to 8 wt. %, 2 to 6 wt. %, 2 to 4 wt. %, 0.1 to 1.0 wt. %, or 0.2 to 1.0 wt. %, or 0.3 to 1.0 wt. %, or 0.4 to 1.0 wt. %, or 0.5 to 1.0 wt. %, or 0.6 to 1.0 wt. %, or 0.7 to 1.0 wt. %, or 0.8 to 1.0 wt. %, or 0.9 to 1.0 wt. %, or 0.1 to 0.9 wt. %, or 0.1 to 0.8 wt. %, or 0.1 to 0.7 wt. %, or 0.1 to 0.6 wt. %, or 0.1 to 0.5 wt. %, or 0.1 to 0.4 wt. %, or 0.1 to 0.3 wt. %, 0.1 to 0.2 wt. %, or 0.2 to 0.4 wt. %, or 0.4 to 0.6 wt. %, 0.6 to 0.8 wt. %.
In one embodiment, the concentration of cellulose nanofibers in the solution is 0.1-30 wt. %.
The concentration of cellulose nanofibers in the solution may 0.1-30 wt. %, or 0.5-30 wt. %, or 1-30 wt. %, or 5-30 wt. %, or 10-30 wt. %, or 15-30 wt. % or 20-30 wt. %, or 25-30 wt. %, or 0.1-25 wt. %, or 0.1-20 wt. %, or 0.1-15 wt. %, or 0.1-10 wt. %, or 1-5 wt. %, or 5-10 wt. %, or 0.2-10 wt. %, or 0.5-10 wt. %, or 0.8-10 wt. %, or 1-10 wt. %, or 2-10 wt. %, or 3-10 wt. %, or 4-10 wt. %, or 5-10 wt. %, or 6-10 wt. %, or 7-10 wt. %, or 8-10 wt. %, or 9-10 wt. %, or 0. 1-9 wt. %, or 0.1-8 wt. %, or 0.1-7 wt. %, or 0.1-6 wt. %, or 0.1-5 wt. %, or 0.1-4 wt. %, or 0.1-3 wt. %, or 0.1-2 wt. %, or 0.1-1 wt. %, or 0.1-0.5 wt. %, or 0.5-5 wt. %, or 5-10 wt. %, or 2-8 wt. %.
A ratio of collagen to cellulose nanofibers in the solution may be 95:5 to 99.9:0.1.
The ratio may be 95:5 to 99.9:0.1, or 96:4-99.9:0.1, or 97:3-99.9:0.1, or 98:2-99.9:0.1, or 99:1-99.9:0.1, or 99.5:0.5-99.9:0.1, or 95:5-99.5:0.5, or 95:5-99:1.0, or 95:5-98:2.0, or 95:5-97:3.0, or 95:5-96:4.
The non-polymeric short-range carbodiimide crosslinking agent may be selected from a group consisting of ethyl-N′-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC), 1-[3-(dimethylamino)propyl]-3-ethylcarbodiimide methiodide (EDCM), dicyclohexyl-carbodiimide (DCC), N-hydroxy-succinimide (NHS) and combinations thereof.
The collagen may be selected from a group consisting of Type I collagen, Type II collagen, Type III collagen, Type IV collagen, Type V collagen, Type VI collagen, denatured collagen from animal sources, and/or human recombinant collagens.
Twenty-nine types of collagens have been identified, but types I, II and Ill are the most abundant and make up the majority of the extracellular matrix macromolecules and have several roles. For instance, over 90% of the collagen in the human body is type I. All fibril-forming collagens (e.g., types I, II, III, V, XI, etc.) may be used in the hydrogel material. This is because many of the crosslinking sites that are needed for formation of crosslink networks are common among these collagens. In addition, these collagens are protein complexes whose basic units consist of the same triple helices (tropocollagen) in which three polypeptide chains are wound around each other like a piece of rope.
The collagen can be derived from different sources such as animals' tissue (e.g., pig skin, jelly fish, fish scale, human, etc.) or synthetically produced collagens such as human recombinant collagen.
Collagen molecules can self-assemble into micro-fibrils and then fibrils. For tissue engineering applications, one collagen type or a combination of various types can be used depending on the target tissue or organ that need to be replaced or repaired. For example, collagen type I is abundant in the human cornea, skin, tendon, and bone while collagen type II is abundant in cartilage and type III is abundant in veins Collagen type III is more abundant in skin, lung, cornea, and the vascular system, frequently in association with type I collagen.
The cellulose nanofibers may be derived from biomass, plants, and/or bacteria.
Nanocellulose can be sourced from biomass, plants, or bacteria. Nanocellulose derived from pulp and medical grade nanocellulose derived from Ciona intestinalis may be used.
Ciona tunicates are the only animals known to produce cellulose. It is believed that they acquired this ability through lateral gene transfer from cyanobacteria. Ciona is also an environmentally friendly source of nanocellulose as instead of cutting trees, it is produced from Ciona species which are aggressively growing in the ocean and costal sea areas resulting in depletion of nutrients for other sea creators and plants as well as blockage of the water inlet to factories that relay on sea water for their processing and causing trouble for boat and ship loading docks.
The cellulose nanofibers (CNFs) are composed of nano-sized cellulose fibrils with a high aspect ratio (length to width ratio). CNF from Ciona has an aspect ratio of about 396. Typical fibril widths are 5-20 nanometers with a wide range of lengths, typically several micrometers.
Ciona nanocellulose is longer and wider and thus stronger than CNFs from e.g., wood. For example, Ciona CNF has a length of 1000-3000 nm which is more than 10-folds longer than woody CNF at 100-150 nm. Ciona CNF has a diameter range of 15-30 nm which is higher than 4-10 nm for woody CNFs.
Ciona tunics are rich in cellulose (44%) and do not contain lignin, facilitating cellulose extraction and purification up to 99.5% especially for medical application. The source of Ciona CNF used is of a medical grade quality, which is a must for applications in tissue engineering while there is no woody medical grade CNF to our knowledge.
A molar ratio of the non-polymeric short-range carbodiimide crosslinking agent to riboflavin in the solution may be from 10:1-500:1.
The molar ratio of the non-polymeric short-range carbodiimide crosslinking agent to riboflavin may be from 10:1-500:1, or 50:1-500:1, or 100:1-500:1, or 200:1-500:1, or 300:1-500:1, or 400:1-500:1, or 10:1-400:1, or 10:1-300:1, or 10:1-200:1, or 10:1-100:1, or 100:1-300:1.
The pH of the solution may be 3-6.
pH plays a key role in the crosslinking efficiency and therefore is a tool to control degradation gradient within the hydrogel material. The pH of the solution may be 3-6, or 4-6, or 5-6, or 3-4, or 3-5, or 4-5.
The method may further comprise a step of rinsing the formed intermediate hydrogel material before exposing the intermediate hydrogel material to ultraviolet A light.
The intermediate hydrogel material may be rinsed with water or a buffer such as PBS.
According to a second aspect, there is provided an implantable hydrogel material comprising collagen in a concentration of 0.1-30 wt. %, wherein the collagen molecules are crosslinked.
The collagen molecules are both intra- and inter-molecularly crosslinked. Collagen molecules are both chemically and physically crosslinked.
The degree of crosslinking may be as high as 100%, or 80-100% crosslinked.
The content of collagen in the hydrogel material may be 0.1-30 wt. %, or 0.5-30 wt. %, or 1-30 wt. %, or 5-30 wt. %, or 10-30 wt. %, or 15-30 wt. % or 20-30 wt. %, or 25-30 wt. %, or 0.1-25 wt. %, or 0.1-20 wt. %, or 0.1-15 wt. %, or 0.1-10 wt. %, 0.1-5 wt. %, or 0.1-1 wt. %, or 1-5 wt. %, or 5-10 wt. %, or 10-20 wt. %, or 20-25 wt. %.
Such a hydrogel material is highly transparent, biocompatible, and mechanically strong and can tolerate surgical manipulations, sutures, and remain stable and intact when implanted in vivo.
The implantable hydrogel material may further comprise cellulose nanofibers in a concentration of 0.1-30 wt. %, wherein the cellulose nanofibers are crosslinked, and wherein there are crosslinks between the cellulose nanofibers and the collagen molecules.
There are both intra- and inter-fibrillary crosslinks and inter fibrillary-molecular crosslinks. The collagen molecules and the cellulose nanofibers are chemically and physically crosslinked. The degree of crosslinking may be as high as 100%, or 80-100% crosslinked.
The content of cellulose nanofibers in the formed hydrogel may be 0.1-30 wt. %, or 0.5-30 wt. %, or 1-30 wt. %, or 5-30 wt. %, or 10-30 wt. %, or 15-30 wt. % or 20-30 wt. %, or 25-30 wt. %, or 0.1-25 wt. %, or 0.1-20 wt. %, or 0.1-15 wt. %, or 0.1-10 wt. %, or 1-5 wt. %, or 5-10 wt. %, or 0.1-10 wt. %, or 0.2-10 wt. %, or 0.5-10 wt. %, or 0.8-10 wt. %, or 1-10 wt. %, or 2-10 wt. %, or 3-10 wt. %, or 4-10 wt. %, or 5-10 wt. %, or 6-10 wt. %, or 7-10 wt. %, or 8-10 wt. %, or 9-10 wt. %, or 0.1-9 wt. %, or 0.1-8 wt. %, or 0.1-7 wt. %, or 0.1-6 wt. %, or 0.1-5 wt. %, or 0.1-4 wt. %, or 0.1-3 wt. %, or 0.1-2 wt. %, or 0.1-1 wt. %, or 0.1-1 wt. %, or 0.5-5 wt. %, or 5-10 wt. %, or 2-8 wt. %.
A ratio of collagen molecules to cellulose nanofibers in the hydrogel material may be 95:5 to 99.9:0.1, or 96:4-99.9:0.1, or 97:3-99.9:0.1, or 98:2-99.9:0.1, or 99:1-99.9:0.1, or 99.5:0.5-99.9:0.1, or 95:5-99.5:0.5, or 95:5-99:1.0, or 95:5-98:2.0, or 95:5-97:3.0, or 95:5-96:4.
The hydrogel material may exhibit a light transmission of at least 80%.
The hydrogel material may be essentially transparent having a light transmission of 80% or more. At higher concentrations of cellulose nanofibers than 10 wt. % in the hydrogel material, the light transmission through the hydrogel material is lower than 80% and, hence not of interest in ophthalmic applications, which require high transmission through the material.
The implantable hydrogel material may be loadable with cells, tissue factors, growth factors, bioactive agents and/or drugs.
The hydrogel material may be loaded with cells, tissue factors, growth factors, bioactive agents and drugs, which may be delivered to the area of implantation improving and increasing tissue regeneration.
According to a third aspect, there is provided a use of the implantable hydrogel material described above in ophthalmic devices, skin replacement, cardiac wall repairs, or cardiac patch applications.
The ophthalmic device may be a corneal implant, a corneal lens, or a contact lens.
The corneal implant may be a full or partial corneal implant, a corneal inlay or a corneal onlay.
The hydrogel materials disclosed in the embodiments of the present invention can be bio-printed into implantable 3D architectures and morphologies for tissue engineering applications.
Below is described an implantable collagen-based material and its production. The material is both physically and chemically stabilized, thereby forming a mechanically robust, elastic and biocompatible implantable material, which support and maintain tissue structure. The material is further transparent, cell friendly and bio-interactive, allowing seamless host-biomaterial integration that helps restore tissue functionality and avoid rejection of the material from the implant site.
Such an implantable collagen-based hydrogel material may be formed by mixing a collagen molecule solution, having a concentration of collagen molecules in water or buffer of 0.1 to 30 wt. %, with a non-polymeric short-range carbodiimide crosslinking agent and riboflavin. The formed mixture having a concentration of the non-polymeric short-range carbodiimide crosslinking-agent of 1-30 wt. %, and a concentration of riboflavin of 0.1 to 10 wt. %. The mixture is allowed to react, forming an intermediate hydrogel material. The step of allowing the mixture to react, forming an intermediate hydrogel material may take place in a mould shaping the hydrogel material into its wanted shape, such as a lens or cornea etc. The formed intermediate hydrogel material may optionally be rinsed with water of a buffer such as PBS. The intermediate hydrogel material is exposed to ultraviolet A light, thereby forming an implantable hydrogel material. The formed material is a double-crosslinked collagen material, physically and chemically stabilized.
The hydrogel material may be reinforced by adding an amount of cellulose nanofibers in the solution during the production of the hydrogel material. A concentration of cellulose nanofibers added in the solution may be 0.1 to 30 wt. %. Such a hydrogen reinforced material has an increased mechanical robustness as compared to the hydrogel material without cellulose nanofibers, while elasticity, biocompatibility and transparency are maintained.
Depending on the amount of nanocellulose fibers in the mixture, the hydrogel material formed may be a pure double crosslinked (i.e., crosslinked by short range carbodiimide crosslinker and crosslinked using riboflavin and UVA) collagen material or a collagen material re-inforced with an amount of nanocellulose.
Double chemical-photochemical crosslinking of collagen and collagen reinforced with cellulose nanofibers (CNF) is shown in the schematics in
Double-crosslinked collagen, double crosslinked collagen-CNF and single crosslinked collagen (using EDC/NHS) formulations are associated with certain types of crosslinks described above. For example, it is believed that the crosslinks in single crosslinked collagen hydrogel materials are of type 1 and 3, in double crosslinked collagen hydrogel material crosslinks of type 1, 3, and 6 are present, while the ones in double crosslinked collagen-CNF are of all types from type 1 to type 6.
The implantable hydrogel material described above may be used in ophthalmic devices, skin replacement, cardiac wall repairs, or cardiac patch applications. The ophthalmic device may be a corneal implant, a corneal lens, or a contact lens.
In the experimental section below is a described non-limiting examples of production of such implantable hydrogel materials, and different tests thereof.
For this work purified medical-grade type I porcine collagen from porcine skin was used.
In the experiments described here, nanocellulose derived from pulp and medical grade nanocellulose derived from Ciona intestinalis were used.
MES buffer (2-(N-morpholino) ethanesulfonic acid) at a concentration range of 0.5 to 1 Molar was used for fabrication of the hydrogel material/scaffolds. Sterile phosphate buffered saline (PBS) was used as storage buffer.
1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), 1-[3-(Dimethylamino) propyl]-3-ethylcarbodiimide methiodide (EDCM), N-hydroxysuccinimide (NHS), and riboflavin (known as vitamin B2) 7,8-dimethyl-10-(D-ribo-2,3,4,5-tetrahydroxypentyl) isoalloxazine and 7,8-dimethyl-10-ribitylisoalloxazine were used as crosslinking agents
Medical grade purified freeze-dried type I collagen was dissolved in water or PBS at 20 wt. % solution concentration. The collagen solution was then mixed with crosslinkers 1-[3-(Dimethylamino) propyl]-3-ethylcarbodiimide methiodide (EDCM), 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), N-hydroxysuccinimide (NHS) at 20%, and riboflavin (vitamin B2) at 0.5% concentration. The solution was mixed thoroughly, dispensed into custom moulds, clamped and then cured. Removal from moulds was achieved by immersion in PBS for one hour at room temperature. Scaffolds were immersed in riboflavin and then exposed to UVA. Finally, scaffolds were rinsed with sterile PBS to extract any reaction residues.
Medical grade purified freeze-dried type I collagen was dissolved in water or PBS at 20% solution concentration. Nanocellulose solution at 2.5% was added to the collagen solution and then mixed with crosslinkers 1-[3-(Dimethylamino) propyl]-3-ethylcarbodiimide methiodide (EDCM), 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC), N-hydroxysuccinimide (NHS) at 20%, and riboflavin (vitamin B2) at 0.5% concentration. The solution was mixed thoroughly, dispensed into custom moulds, clamped and then cured. Removal from moulds was achieved by immersion in PBS for one hour at room temperature. Scaffolds were immersed in riboflavin and then exposed to UVA. Finally, scaffolds were rinsed with sterile PBS to extract any reaction residues.
Medical grade purified freeze-dried type I collagen was dissolved in water or PBS at 20% solution concentration. The collagen solution was then mixed with crosslinker solutions of 1-[3-(Dimethylamino) propyl]-3-ethylcarbodiimide methiodide (EDCM), 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), and N-hydroxysuccinimide (NHS) at 20%. The solution was mixed thoroughly and then dispensed into custom moulds. Moulds were clamped and samples were cured. Removal from moulds was achieved by immersion in PBS for one hour at room temperature. Finally, scaffolds were rinsed with sterile PBS to extract any reaction residues.
The solution was mixed thoroughly and dispensed into custom moulds with spacers of various thicknesses to delineate the final device thickness. Moulds were clamped and samples were cured at room temperature followed by curing at 37° C. Removal from moulds was achieved by immersion in PBS for one hour at room temperature. Finally, the moulded material was rinsed with sterile PBS to extract any reaction residues.
The solution was mixed thoroughly with chemical crosslinkers and bio-printed into 3D custom shapes of various sizes to delineate the final device architecture followed by curing at 37° C. Moulds maybe used to further shape the implants if necessary.
Following the chemical crosslinking, a photochemical crosslinking of samples was performed by immersion in riboflavin solution and exposure to ultraviolet A (UVA) light for 15-60 minutes at a wavelength of 365 nm and intensity of 1300μW/cm2.
Double crosslinked collagen and double crosslinked collagen-CNF, prepared as discussed above, were tested and compared to single crosslinked collagen formulations for major physical and biological properties including mechanical properties, transparency (% light transmission), and cell biocompatibility.
Tensile strength, elongation at break (elasticity), elastic modulus (stiffness), and energy at break (toughness) were measured with an Instron Automated Materials Testing System (Model 5943 Single Column Table Frame) equipped with BluHill software, a load cell of 50 N capacity and pneumatic metal grips at a crosshead speed of 10 mm/min. Test specimens were made by molding and curing the samples into dumbbell-shaped Teflon molds followed by equilibration in PBS. Specimens were attached to the grips and tensile force was applied until the sample break point. Data were automatically recorded by the software, and 66 dumbbell-shaped samples were used for each mechanical test.
Double crosslinked collagen (Double Xlinked), and double crosslinked collagen-CNF (CNF-double-Xlinked), were tested and compared to single crosslinked collagen (Single Xlinked) formulations. In
The results suggest significant increases in stress (strength of the scaffold), energy at maximum load (robustness), and elastic modulus (stiffness) for the double crosslinked collagen and double crosslinked collagen-CNF compared to single crosslinked collagen, without sacrificing the strain (elasticity).
Light transmission through the hydrogel material was measured across the UV and visible light spectrum (200-700 nm) at room temperature, using a High Performance USB4000 UV-Vis Spectrophotometer (Mettler Toledo, Stockholm). To enable direct comparison with light transmission through the native human cornea, the hydrogel material samples were 550 μm thick. Samples were immersed in PBS during measurement, and light transmission was compared with published data from healthy human corneal tissue (Meek, K. M., et. al. Transparency, swelling and scarring in the corneal stroma. Eye (Lond) 17, 927-936 (2003). Measurements were made for three independent samples per spectrum.
The light transmission results shown in
Culture of human corneal epithelial cells (HCE-2 50.B1 cell line, Lot No. 70015331, ATCC, USA) was established according to the manufacturer's instructions. Briefly, serum-free keratinocyte growth medium (1×, Gibco) was supplemented with L-Glutamine, 5 ng/ml epidermal growth factor (EGF) and 0.05 mg/ml bovine pituitary extract (BPE), 500 ng/ml hydrocortisone and 0.005 mg/ml insulin (Gibco). A T-75 cell culture flask was precoated with a mixture of 0.01 mg/ml fibronectin, 0.03 mg/mL bovine collagen type 1 and 0.01 mg/mL bovine serum albumin, and was incubated overnight at 37° C. The next day, the excess coating was aspirated, and the flask was allowed to stand for 15 minutes before seeding the cells. The HCE-2 cell vial was thawed, and cells were seeded at a density of 104 cells/cm2 on the precoated flask. Cells were incubated at 37° C. at 5% CO2, and growth media was changed every other day.
Hydrogel material, double crosslinked collagen and double crosslinked collagen-cellulose nanofibers, (300 μm thick) precut to 8 mm diameter were rinsed in PBS and equilibrated in the complete keratinocyte serum-free medium for 2 hours in a humidified cell culture incubator. The hydrogel samples were then laid down (concave side down) onto the bottom of a 48-well cell culture plate and allowed to adhere to the bottom of the plate for 2 hours in an incubator at 37° C. Three wells were used for each hydrogel sample, and three wells were used as controls (i.e., no biomaterial attached to the bottom of the well).
Upon confluence of the seeded HCE-2 cells in the culture flask, the cells were trypsinized with Trypsin-EDTA, then trypsinization was stopped with a complete growth medium and cells were harvested, counted, and seeded into the six prepared wells. Cells were seeded at a density of 105 cells/well for the control and the two hydrogel materials and incubated for 1 hour prior to adding additional media to reach a final volume of 200 μl of growth media per well. Growth media was changed every other day, and cells were maintained in culture for 16 days. On day 16, the cells were washed and covered with fresh media. Cells were stained with NucBlue Live cell stain (Hoechst 33342, Thermo Fisher Scientific) according to the manufacturer's instructions. Images of stained cells were captured using a Leica DMi8 inverted live-cell microscope under ultraviolet light excitation (385 nm) to detect live cells with fluorescent blue nuclear stain. Brightfield images were additionally obtained to observe cell morphology.
In
The capability of the double crosslinked collagen-CNF for drug loading and slow release was also tested. Three model drugs, e.g.,, dexamethasone, Nerve Growth factor (NGF), and Artemin were separately loaded into double crosslinked collagen-CNF bioengineered corneal implants and tested in vitro and in vivo in rabbit corneas for drug release and key physical properties of the corneas including transparency and mechanical properties. Dexamethasone is a steroid medicine that is used to relieve the redness, itching, and swelling caused by eye infections and other conditions or procedures (e.g.,, eye surgery). Dexamethasone eye drops are used to treat inflammation of the eyes caused by allergies and certain conditions, including damage caused by chemical and thermal burns as well as treating eye pain and inflammation after eye surgery.
A 1-3 mg/ml solution of dexamethasone was prepared. 30 microliters of the solution was then injected into the collagen solution before adding the crosslinkers as explained above. Similar protocols were used for NGF and Artemin. Upon crosslinking, the drugs will be covalently bound to collagen molecules making it possible for a long-term slow and continuous release of the drugs (up to 60 days) from the scaffolds.
No prior work to our knowledge has demonstrated a commercially and clinically viable, GMP-grade bioengineered scaffold from a sustainably sourced, cost-effective, widely available, and FDA-approved raw material, and no other technology has achieved manufacturability, packaging, sterility, or long shelf life that are ISO-compliant and independently third-party certified. Yet these often-overlooked aspects are critical for addressing the lack of donor tissue.
The optimum scaffold (double crosslinked collagen material (Double-Xlinked)) was tested in human corneas of 20 patients and data collected over 12 months post implantation. From a safety perspective, the scaffold did not result in thinning, loss of transparency, neovascularization, rejection, or other adverse event observed to varying degrees in most preclinical and clinical studies conducted by other groups.
The scaffolds are mechanically more robust than earlier versions we previously developed, with superior toughness, elastic modulus, and stress tolerance translating into an ability to withstand surgical implantation, mechanical forces and enzymes in vivo. Our results suggest it may not be necessary for bioengineered tissue to match the mechanical properties of the native cornea to be therapeutically effective. Standard corneal transplantation with mechanically tough human donor tissue often results in scar tissue formation at host-to-implant interfaces, a phenomenon we did not observe with the mechanically softer scaffolds.
| Number | Date | Country | Kind |
|---|---|---|---|
| 2151478-1 | Dec 2021 | SE | national |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/SE2022/051130 | 12/1/2022 | WO |
