Patent Application
20220047856
The present invention relates to a hydrogel for in-vivo directional release of medication. In particular it concerns a controlled and local release of medication. More in particular, the present invention relates to a hydrogel for close contact to organs and skeletal structures.
Hydrogels are three-dimensional, physically or chemically cross-linked networks of water-soluble polymers. Their hydrophilic nature, water content similar to living tissue and elasticity, make them excellent candidates for biomedical applications. There is therefore quite some prior art on biodegradable hydrogels that are designed to release medication in the (human or animal) body in a sustained way.
For instance, in the Journal of Advanced Research, volume 8, Issue 3, May 2017, pages 217-233, a thorough review by E. A. Kamoun et al may be found on hydrogels and their medical application. As indicated in the introduction of this article, a further overview may be found in European Polymer Journal, volume 65, April 2015, pages 252-267 by E. Caló et al, “Biomedical applications of hydrogels: A review of patents and commercial products”.
Q. Feng et al describes “Mechanically resilient, injectable, and bioadhesive supramolecular gelatin hydrogels crosslinked by weak host-guest interactions assist cell infiltration and in situ tissue regeneration” in Biomaterials, Volume 101, September 2016, Pages 217-228.
In RSC Adv., 2017, 7, 34053, T. T. H. Thi et al describe injectable hydrogels as a novel platform for the release of hydrophobic drugs. An additional Schiff base reaction was introduced into a phenol-phenol crosslinked gelatin hydrogel to increase adhesiveness. β-cyclodextrin possessing a hydrophobic cavity and oxidized to present aldehyde groups (hereinafter “oβ-CD”) was grafted to the gelatin backbone via Schiff base reaction, with the cavity providing encapsulation for hydrophobic drugs. Simply blending gelatin-tyramine (hereinafter “GTA”) and oβ-CD in the presence of horseradish peroxidase and hydrogen peroxide (hereinafter “HRP/H2O2”) rapidly and controllably formed GTA-oβ-CD hydrogels in situ. The optimal composition of GTA-oβ-CD hydrogels was found to be 5 wt % GTA with 1 wt % oβ-CD. Their elastic modulus and degradation rate were 1.8- and 1.5-fold higher than those of GTA hydrogels owing to additional imine bonds. Hydrophobic drugs (e.g., dexamethasone and curcumin) could be dissolved homogeneously in GTA-oβ-CD matrices with greater loading efficiencies than in GTA matrices. An in vitro test of cell viability using human dermal fibroblasts demonstrated that GTA-oβ-CD hydrogels were cytocompatible. In summary, dual-functional injectable GTA-oβ-CD hydrogels can be used as a promising platform to improve tissue adhesion and hydrophobic drug delivery.
Important factors to consider during the design of these hydrogels include 1) duration of delivery, and 2) location of delivery with respect to its working mechanism. For example, for effective local pain relief it is essential that an anaesthetic is delivered and remains in situ for a period in close proximity to the origin of pain. The problem of sustained release is particularly challenging for small molecules, such as Bupivacaine (hereafter “Bupi”).
Bupi is a very effective and relatively inexpensive local anaesthetic. However, the duration of its effect is limited to approx. 8 hours. Increasing the dose or concentration of conventional bupivacaine solutions to obtain prolonged durations of effect can lead to both systemic and local toxicity, cf., Gitman M, Barrington M J “Local Anesthetic Systemic Toxicity: A Review of Recent Case Reports and Registries” in Regional Anesthesia & Pain Medicine 2018; 43:124-130. Cardio- and central nervous system toxicity are well-known systemic toxic effect of bupivacaine. It is therefore of interest to find a way of releasing Bupi locally and in a delayed fashion, whereby it may work longer and with a decreased incidence of local and systemic cytotoxicity compared to conventional bupivacaine applications such as local bolus injection.
In the yet unpublished NL patent application 2020071 by the present applicant a deformable body and combination of such deformable body and a surgical screw element is described. The deformable body may be made of a visco-elastic material, a degradable felt material, a sponge-like material, a gelatine material, a gel, in particular a hydrogel, a polymer or any combination thereof. The deformable body may comprise an anaesthetic and/or another pharmaceutical compound. It has a surface through which the anaesthetic may be released, for example, the bone contact surface. The release in the surrounding area may be avoided, by use of a further substantially non-pervious wall of the deformable body.
As further improvement on the deformable body, the present inventors set out to design a biocompatible, biodegradable hydrogel with controlled, sustained and directional release of medication. Moreover, the inventors set out to design a hydrogel that is versatile and easy to produce on a large scale, is easy to cross-link and can be cross-linked in a controlled manner to produce a hydrogel that is both flexible and strong. In this regard it should be understood that the hydrogel must be both sufficiently flexible and strong as to allow it to be implanted and to withstand local circumstances and forces so as to stay at the location of implantation for sufficient time to release the medication and not break or otherwise be damaged. This means that the hydrogel can adapt to a shape of a surface of a skeletal structure against which it is pressed, whereby intimate contact with the outer bone surface of the skeletal structure is achieved. Typically this requires a hydrogel with an elastic/compressive modulus of between 50 and 1000 kPa, more preferably between 100 and 600 kPa.
The present 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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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:
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.
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.
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.
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 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.
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.
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).
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.
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).
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.
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
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.
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,
| Number | Date | Country | Kind |
|---|---|---|---|
| PCT/NL2018/050832 | Dec 2018 | NL | national |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/NL2019/050352 | 6/11/2019 | WO | 00 |
