HYDROGEL FOR IN-VIVO DIRECTIONAL RELEASE OF MEDICATION

Abstract

The invention provides a hydrogel for in-vivo release of medication comprising at least one medication, wherein the surface of the hydrogel comprises a coating such that the surface has one or more sub-surfaces with permeability that is at least 2× higher than the average permeability of the entire surface, wherein the hydrogel has an elastic modulus of between 50 and 1000 kPa.

Claims

1. A hydrogel for in-vivo release of medication comprising at least one medication in the form of a small molecule, wherein the hydrogel has a surface and the surface of the hydrogel comprises a coating that is composed of a material that is less permeable to the medication than the material of the hydrogel itself such that the surface has one or more sub-surfaces with permeability that is at least 2× higher than the average permeability of the entire surface, wherein the hydrogel has an elastic modulus of between 50 and 1000 kPa.

2. The hydrogel of claim 1, the hydrogel has an elastic modulus of between 100 and 600 kPa.

3. The hydrogel of claim 1, having a degree of swelling in the range of 2-20 calculated as (swollen weight−dry weight)/dry weight.

4. The hydrogel of claim 1, comprising a cross-linked biopolymer.

5. The hydrogel of claim 4, wherein the cross-linked biopolymer is a protein-based and/or polysaccharide-based polymer.

6. The hydrogel of claim 1, wherein between 10 and 90% of the surface is covered by the coating.

7. The hydrogel of claim 1, wherein the coating has a thickness between 10 nm to 200 μm.

8. The hydrogel of claim 1, wherein the coating is based on a biodegradable polymer.

9. A method for the preparation of the hydrogel of claim 1, wherein the coating is comprised of a precursor material and some of the precursor material to the coating is allowed to partially diffuse into the hydrogel.

10. The hydrogel according to claim 1 for use in the treatment of musculoskeletal disorders.

11. The hydrogel of claim 2, having a degree of swelling in the range of 2-20 calculated as (swollen weight−dry weight)/dry weight.

12. The hydrogel of claim 2, comprising a cross-linked biopolymer.

13. The hydrogel of claim 3, comprising a cross-linked biopolymer.

14. The hydrogel of claim 11, comprising a cross-linked biopolymer.

15. The hydrogel of claim 12, wherein the cross-linked biopolymer is a protein-based and/or polysaccharide-based polymer selected from the group consisting of hyaluronic acid, chitosan, cellulose, gelatin, and combinations thereof.

16. The hydrogel of claim 13, wherein the cross-linked biopolymer is a protein-based and/or polysaccharide-based polymer selected from the group consisting of hyaluronic acid, chitosan, cellulose, gelatin, and combinations thereof.

17. The hydrogel of claim 4, wherein the cross-linked biopolymer is gelatin.

18. The hydrogel of claim 1, wherein between 20 and 80% of the surface is covered by the coating, optionally wherein between 30 and 70% of the surface is covered by the coating.

19. The hydrogel of claim 1, wherein the coating is selected from the group consisting of PLGA, PCL, gelatin, alginate, and combinations thereof.

20. The hydrogel according to claim 1 for use in the treatment of musculoskeletal disorders for treatment of infection, inflammation, malignant processes, growth disorders, degenerative disorders, treatment of pain arising from said disorders, or treatment of pain arising from surgical treatment of said disorders.

Description

DRAWINGS

[0011] FIG. 1 is a series of images of a hydrogel having its top part coated and containing methylthioninium chloride (methylene blue). As can be seen, methylthioninium chloride (which is both a medication and dye) is only released in the opposite direction.

DETAILED DESCRIPTION OF THE INVENTION

[0012] Hydrogels may be synthesized by cross-linking water-soluble polymers. Water-soluble polymers such as poly(acrylic acid), poly(vinyl alcohol), poly(vinylpyrrolidone), poly(ethylene glycol), polyacrylamide and polysaccharides (e.g. hyaluronic acid) are the most common systems used to form hydrogels. These water-soluble polymers are non-toxic and widely used in various pharmaceutical and biomedical applications. Although there are many different hydrogels, the present invention focusses on medical hydrogels that are biocompatible and can be implanted and used in-vivo. Moreover, they must be biodegradable. For instance, protein-based and/or polysaccharide-based polymers may be used, such as, hyaluronic acid, chitosan, and cellulose. Preferably, the hydrogel is based on gelatin. In addition to, or instead of the protein-based and/or polysaccharide based polymers, the hydrogel may also comprise other non-toxic water-soluble synthetic or natural polymers. The other polymers may compose up to 50% by weight of the entire polymer content. Given its availability, biocompatibility and cost, the use of gelatin as sole polymer component is preferred. Of particular interest is a hydrogel based on gelatin that is functionalized with a cyclodextrin.

[0013] Although hydrogels for release in-vivo of medication are known, the present inventors found that existing hydrogels could be improved in terms of their directional release. As a result, the new hydrogels of the present invention can be implanted and fixated to specific locations where medication, in particular to achieve pain relief, is required. This may be a hydrogel in the form of e.g. deformable body, whereby the hydrogel conforms to the shape of a skeletal structure or surgical implant or even organ to which it is fixated. Of relevance in this respect is that a hydrogel with a specific elastic modulus in the aforementioned range is used. Moreover, the hydrogel preferably has a degree of swelling in the range of 2-20, preferably in the range of 2-6, calculated as swollen weight (at equilibrium swelling)−dry weight/dry weight.

[0014] The direction of release of medication is achieved by partly covering the surface of the hydrogel with a coating. As a result, the hydrogel will have a sub-surface or sub-surfaces with little or no coating and hence unrestricted permeability of the medication, and a subsurface or surfaces with coating and therefore a reduced permeability for the medication. Preferably the nature and thickness of the coating is selected such as that the permeability at the desired contact surface, e.g., the bone or organ contact surface is at least 2× higher than the average permeability of the entire surface. Having the implanted hydrogel affixed adjacent to the body part that is to be treated, and moreover with the uncovered surface of the hydrogel adjacent to the body part that is to be treated, release of medication in other directions is reduced or even avoided. This has the advantage of reduced-side effects and the possibility to work with lower concentrations of medication or, alternatively, with a longer working time due to a slower release of the regular amount of medication.

[0015] The coating may be composed of the material of the hydrogel, provided that it contains no medication and is sufficiently thick. Suitably it is between 10 nm and 200 μm thick. Preferably, however, the coating is composed of a material that is less permeable to the medication than the material of the hydrogel itself. The coating may be flexible or shell-like. Similar to the hydrogel, the coating must be composed of biocompatible biopolymers. The biodegradability may be the same or prolonged compared to the hydrogel. Suitable materials include, but are not limited to polycaprolactone (hereinafter “PCL”), poly(lactic-co-glycolic acid) (hereinafter “PLGA”), gelatin, or alginate. The permeability of the coating may be adjusted, such that even very small molecules cannot get through. Moreover, the coating can be made hydrophobic, or hydrophilic, depending on its intended use.

[0016] The hydrogel may take any particular shape. In a co-pending application, the use of a hydrogel as carrier for local release of medication in the form of a ring is described (PCT/NL2018/050832, incorporated herein by reference) where it is used in combination with a screw. In another co-pending application the use of a hydrogel as carrier for local release of medication in the form of a sleeve, e.g. for a joint prosthesis is described (NL2023208, incorporated herein by reference). The hydrogel may also be shaped in the form of a (board) thumb pin for attachment to bone or any other solid tissue. Finally, the hydrogel may also be shaped to provide a tight fit in crevices in organs and similar body structures. In each of these embodiments, the hydrogel is coated such as to ensure that those parts of its surface that are not in contact with the body part that is to be treated by direct release are covered by the coating.

[0017] The coatings may be applied onto the hydrogel by any common coating process, including dip coating, brush coating, spray coating and the like. Alternatively, the entire surface of the hydrogel may be coated, whereas the relevant sub-surfaces intended for contact with the body part that is to be treated are freed from coating. Moreover, the coating may be formed and shaped first, as a shell, whereupon the hydrogel in introduced e.g., as an non-crosslinked solution. In this case the shell of coating acts as a mould during the cross-linking and formation of the hydrogel. Alternative methods include overmolding and the like.

[0018] Using a coating material and method that allows some of the precursor material to the coating to partially diffuse into the hydrogel may be particularly beneficial, in particular if this material is water-soluble. After polymerization/crosslinking/setting, the coating will be physically entangled with hydrogel directly underneath the interface, ensuring a good bond. This method is of particular interest, as it reduces chances of coating material breaking off, which is detrimental as it affects the directional release, but which is also detrimental as it might cause migration of particles of coating that may create their own problems.

[0019] Preferably between 10 and 90% of the surface of the hydrogel is covered by a coating. For instance, between 20 and 80% of the surface is covered by a coating, more preferably between 30 and 70% of the surface is covered by a coating.

[0020] The present hydrogel is particularly suitable for treatment of musculoskeletal disorders. These disorders include infection, inflammation, malignant processes, growth disorders, degenerative disorders or treatment of pain arising from (surgical treatment of) these disorders.

[0021] In addition to the medication one or more further ingredients may be included, preferably further ingredients selected from co-medication, glycerol and other co-solvents, colorants, and buffers.

[0022] Methods for making the feedstock for the hydrogel are known. Thus, it is known to functionalize gelatin and related biopolymers with tyramine. See Thi et al, 2017 RSC Adv, which has been cited above, and which is included herein by reference. Of importance, but common in the field of medical application is to remove all forms of contamination. By way of example, the hydrogel may be prepared by the following method: [0023] 1. Solutions of a suitable cross-linking water-soluble (bio)polymer(s), cross-linker and medication are prepared. [0024] 2. Solutions of biopolymer(s) and cross-linker are mixed at pre-determined concentrations to achieve a cross-linked hydrogel with an elastic modulus in the range of 50 to 1000 kPa. [0025] 3. The obtained hydrogel is then submerged in a solution of medication to allow for diffusion of the medication into the hydrogel. Glycerol or similar co-solvent can be added to the medication solution. Glycerol then also diffuses into the hydrogel where it acts as a plasticizer, providing additional robustness and flexibility to the hydrogel. Alternatively, the drug (e.g. in a nano-/microparticle formulation) can be mixed in with the polymer solution prior to crosslinking. [0026] 4. The gel is then dried. [0027] 5. Next, the hydrogel is coated in part, e.g. with a solution of a biopolymer with a different permeability for the medication compared to the hydrogel, to ensure directional release of the encapsulated medication. The coating may also enhance the mechanical properties of the hydrogel. Alternatively, it is also possible to form a shell of the coating in a pre-defined shape, and introduce the solution of step 2, together with the medication, into this shell, whereby the coating acts as a mould for the hydrogel.

EXAMPLES

Materials

[0028] Gelatin (porcine skin, type A, 300 g bloom strength), 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide (EDC), N-hydroxysuccinimide (NHS), tyramine hydrochloride, 2-morpholinoethanesulfonic acid monohydrate (MES), sodium persulfate (SPS), sodium periodate, 3-cyclodextrin, phosphate-buffered saline (PBS), riboflavin (RB), ethylene glycol and glycerol were purchased from Sigma-Aldrich. Cellulose dialysis membranes (Spectra/Por™, 0.5 kDa; 12 kDa molecular weight cut-off) were purchased from Spectrum Laboratories. Bupivacaine was obtained from Siegfried, Switzerland.

Synthesis of Gelatin-Tyramine (GTA)

[0029] Gelatin type A (5 g) was dissolved in a 50 mM MES buffer (300 ml) at 50° C. After dissolution of the gelatin, EDC (13.7 mmol), NHS (6.85 mmol) and tyramine (15 mmol) were added to the gelatin solution. The reaction mixture was left to react for 24 h at 40° C. with stirring. After 24 the mixture was dialyzed against water for 72 h and the product was then obtained by lyophilization.

Tyramine Content Measurement

[0030] The degree of functionalization of gelatin was determined by measuring the absorbance of the polymer solution (0.1%, w/v) at 275 nm and calculated from a calibration curve obtained by measuring the absorbance of known percentages of tyramine in distilled water.

Oxidation of β-Cyclodextrin

[0031] Oxidized β-cyclodextrin was prepared by reaction with sodium periodate. Briefly, β-cyclodextrin (5 g) was dispersed in distilled water followed by addition of sodium periodate (3.77 g) and stirred at room temperature in the dark, overnight. The reaction was terminated by the addition of ethylene glycol. The mixture was dialyzed against deionized water using a dialysis membrane with an MWCO of 500 Da (Spectrum Labs) for 3 days and the product was collected by lyophilization. The degree of oxidation was determined by .sup.1H NMR, using either deuterated dimethyl sulfoxide (DMSO-d6) or deuterium oxide (D2O) as solvent. Whereas β-cyclodextrin has a ratio of protons at 4.8-4.9 ppm versus 4 ppm of about 2.04, progress of the reaction can be seen by a change in the ratio, to about 1.49.

Fabrication of GTA/β-Cyclodextrin Hydrogels

[0032] Prior to hydrogel crosslinking, solutions of GTA, op-CD, SPS and Riboflavin were prepared. Unless indicated otherwise, GTA had a degree of functionalization of 10-25%, whereas oβ-CD with an oxidation degree of the secondary hydroxyl groups of 15-30% was used. These solutions were mixed so that final concentrations of 20 wt % GTA, 0-10 wt % oβ-CD, 20 mM SPS and 2 mM Riboflavin were obtained. The obtained solution was exposed to visible light for 30 minutes to enable hydrogel formation. The cross-linked hydrogel had a degree of swelling of 3-6. Moreover, it had an elastic modulus of 100-600 kPa.

[0033] The obtained hydrogel was then submerged overnight in a bupivacaine solution with a concentration of bupivacaine of 50 mg/mL to allow for diffusion of bupivacaine into the gel. The bupivacaine solution contained a concentration glycerol of 30 vol %. As a result, the concentration of bupivacaine in the hydrogel was 50 mg/mL (±20).

[0034] Next, the hydrogel was coated with a coating solution comprising 10% PCL in dichloromethane (DCM) In this case the hydrogel was dipped into the solution for a number of times to achieve a coating of about 180 μm. The coating was found to provide additional strength to the hydrogel.

Drug Loading and In Vitro Drug Release Assay

[0035] For the investigation of drug release properties, the obtained hydrogels were loaded with bupivacaine by immersion in an aqueous solution of bupivacaine at 50 mg/mL for 24 hours. The bupivacaine solution contained a concentration glycerol of 30% vol. As a result, the concentration of bupivacaine in the hydrogel was ±50 mg/mL (±20).

[0036] The release of bupivacaine from the hydrogels was measured by placing the hydrogels in a vial containing 1 mL of 0.1M citrate buffer, pH 6 at 37° C.

[0037] At predetermined time points, aliquots of 100 uL samples were taken from the release solution and replaced with fresh buffer. The samples were diluted 1:10. Bupivacaine release was determined by UPLC using ammonium formate (10 mM, pH 2.4) and a mixture of acetonitrile/water/formic acid (96:5:0.2, v:v:v) as mobile phase. This control experiment proves that the hydrogel may be used for sustained release of medication

Drug Loading and In Vitro Drug Release Assay

[0038] For the investigation of directional release, the obtained hydrogels were now loaded with methylthioninium chloride. A PCL shell was acquired by dip-coating of a metal mould. The mould was dipped twice in 10% PCL solution to obtain a 180 um thick film. A photocrosslinkable pre-gel solution was then prepared, methylthioninium chloride was added by mixing a 1 wt % solution in the pre-gel solution to obtain a final concentration of 0.1 wt % methylthioninium chloride in the hydrogel. The gel was then cross-linked on top of the PCL film using exposure to a visible light-source.

[0039] Release of methylthioninium chloride from the gel was simulated in a 3% alginate gel, cross-linked with calcium chloride to obtain a tissue-like consistency. In the images, FIG. 1, the PCL film is on top of the gel. As shown, release was only visible in the non-PCL-covered direction. Hydrogels were positioned vertically to eliminate any effect of gravity on the direction of release. This experiment proofs that the hydrogel with coating may be used for sustained directional release of medication.