Hydrogels for in vivo drug release

Abstract

The present invention relates to a hydrogel for the in vivo release of a drug, comprising at least one drug, wherein the hydrogel comprises: (i) a proteinaceous biopolymer functionalized with a functionalizing agent, preferably a primary aminoalkylphenol, capable of forming a guest-host interaction with oxidized β-cyclodextrin, more preferably gelatin (GTA) functionalized with tyramine, and (ii) oxidized β-cyclodextrin (oβ-CD), wherein the hydrogel is crosslinked by exposure to visible light in the presence of a biocompatible photoinitiator, such that the degree of swelling calculated as (swollen weight-dry weight) / dry weight is in the range of 2 to 20. The present invention further relates to a process for the preparation of the hydrogel, and to a medicament for the treatment of musculoskeletal diseases, preferably for the treatment of infections, inflammations, malignant processes, growth disorders, degenerative diseases, trauma, autoimmune diseases or for the treatment of pain caused by these diseases (surgical treatment of these diseases).

Description

Technical Field

[0001] The present application relates to hydrogels for in vivo release of drugs. In particular, it relates to controlled and local release of drugs, such as bupivacaine and / or other local anesthetics. More particularly, the present invention relates to hydrogels for close contact with skeleton structures. The present invention further relates to methods for preparing the same in vitro and in vivo. Background Art

[0002] Hydrogels are three-dimensional physically or chemically cross-linked networks of water-soluble polymers. Their hydrophilicity, water content similar to that of living tissue, and elasticity make them excellent candidates for biomedical applications. Therefore, there is a considerable prior art of biodegradable hydrogels designed to release drugs in a sustained manner in the (human or animal) body.

[0003] For example, in "J. Adv. Res. 2017, 8, 217-233", EA Kamoun et al. comprehensively reviewed hydrogels and their medical applications. As pointed out in the introduction of the article, a further overview can be found in Eur. Polym. J. 2015, 65, 252-267, "Biomedical applications of hydrogels: A review of patents and commercial products" by E. Caló et al.

[0004] Q.Feng et al. described "Mechanically resilient, injectable, and bioadhesive supramolecular gelatin hydrogels crosslinked by weak host-guest interactions assist cell infiltration and insitu tissue regeneration" in Biomaterials 2016, 101, 217-228.

[0005] In RSC Adv., 2017, 734053, TTHThi et al. described injectable hydrogels as a new platform for the release of hydrophobic drugs. An additional Schiff base reaction was introduced into phenol-phenol cross-linked gelatin hydrogel to increase adhesion. β-cyclodextrin (hereinafter "oβ-CD") having a hydrophobic cavity and oxidized to present an aldehyde group was grafted onto the gelatin backbone by the Schiff base reaction, and the cavity provided encapsulation for the hydrophobic drug. In the presence of horseradish peroxidase and hydrogen peroxide (hereinafter "HRP / H 2 O 2”), gelatin-tyramine (hereinafter “GTA”) and oβ-CD were simply mixed to quickly and controllably form GTA-oβ-CD hydrogels in situ. The optimal composition of GTA-oβ-CD hydrogels was found to be 5wt% GTA and 1wt% oβ-CD. Due to the additional imine bonds, their elastic modulus and degradation rate were 1.8 and 1.5 times higher than those of GTA hydrogels. Hydrophobic drugs (e.g., dexamethasone and curcumin) were uniformly dissolved in the GTA-oβ-CD matrix, and the loading efficiency was much higher than that of the GTA matrix. In vitro cell viability assays using human dermal fibroblasts showed that the GTA-oβ-CD hydrogels were cell compatible. In summary, the bifunctional injectable GTA-oβ-CD hydrogels can be used as a promising platform to improve tissue adhesion and hydrophobic drug release.

[0006] Important factors to consider during the design of these hydrogels include: 1) the duration of delivery, and 2) the location of delivery relative to its mechanism of action. For example, in order to effectively relieve local pain, the anesthetic must be administered close to the source of pain and remain in place for a period of time. For small molecules, such as bupivacaine (hereinafter "Bupi"), the problem of sustained release is particularly challenging.

[0007] Bupi is a very effective and relatively inexpensive local anesthetic. However, its duration of action is limited to about 8 hours. Increasing the dose or concentration of conventional bupivacaine solutions to extend the duration of action can lead to systemic and local toxicity, see "Gitman M, Barrington MJ "Local Anesthetic Systemic Toxicity: A Review of Recent Case Reports and Registries", Regional Anesthesia & Pain Medicine 2018; 43: 124-130". Cardiac and central nervous system toxicity are well-known systemic toxic effects of bupivacaine. Therefore, it is of interest to find a method for local and delayed release of Bupi, whereby the duration of action of bupivacaine is longer and the incidence of local and systemic cytotoxicity is reduced compared to traditional bupivacaine applications (such as local bolus injections).

[0008] In addition, few hydrogels meet the strict requirements of clinical transformation in terms of preparation, application, mechanical properties and biocompatibility of hydrogels. The inventors set out to design biocompatible, biodegradable hydrogels with controlled release and sustained release of drugs. In addition, the inventors set out to design hydrogels, which are widely used, easy to large-scale production, easy to crosslink, and can be crosslinked in a controlled manner, thereby producing both flexible and strong hydrogels. In this regard, it should be understood that the hydrogel must have enough flexibility and strength to allow it to be implanted, and withstand the local environment and force, so as to stay at the implanted position for enough time to release the drug without breaking or being otherwise damaged. This means that the hydrogel can adapt to the surface shape of the skeletal structure to which it is applied, so as to achieve close contact with the exoskeleton surface of the skeletal structure. In a similar manner, it can be pressed on the surface of other musculoskeletal tissues or surgical implants. Typically, this requires a hydrogel with an elastic / compression modulus (Young's modulus) between 100kPa and 600kPa. Summary of the invention

[0009] The present invention provides a hydrogel for in vivo release of a drug, comprising at least one drug, wherein the hydrogel comprises:

[0010] (i) a biopolymer functionalized with a primary aminoalkylphenol, preferably gelatin functionalized with tyramine (GTA), and

[0011] (ii) oxidized β-cyclodextrin (oβ-cd),

[0012] Wherein the hydrogel is crosslinked by exposure to visible light in the presence of a biocompatible photoinitiator, such that the degree of swelling calculated as (swollen weight-dry weight) / dry weight is in the range of 2 to 20, preferably 2 to 6. The swollen weight is the equilibrium weight of the hydrogel in vivo. The swollen weight can be determined experimentally after swelling in vitro for 24 hours (or when equilibrium is reached) in a simulated body fluid, such as PBS, at body temperature, such as 37°C. BRIEF DESCRIPTION OF THE DRAWINGS

[0013] Figure 1 Schematic diagram of the formation of GTA and cross-linking of GTA. Here, cross-linking of type (a) is shown under the influence of riboflavin / sodium persulfate (SPS) and light.

[0014] Figure 2 Schematic diagram of the cross-linking of the amine group of gelatin with the aldehyde group of oβ-CD in the Schiff reaction (b) and the cross-linking of the type (c) showing the guest-host interaction between the tyramine group of GTA and the cavity in cyclodextrin.

[0015] Figure 3a shows a top view of a ring element according to an embodiment of the present invention, Figure 3b Shows Figure 3a The cross section AA of the ring element;

[0016] Figure 4 The arrangement on the rod of the screw element is shown. Figure 3a A combination of a screw element and a ring element;

[0017] Figure 5 The screws screwed into the patient's bone Figure 4 combination.

[0018] Figure 6a An embodiment of an artificial hip implant is shown; Figure 6b Shows Figure 6a An embodiment of an artificial hip joint implant comprises a first embodiment of a sleeve element arranged on a neck of a femoral component of the artificial hip joint implant.

[0019] Figure 7a and Figure 7b An embodiment of a "thumbnail" having a ring-like or screw-like fastening means is shown. Figure 7c Embodiment 7a is shown arranged for attachment to bone.

[0020] Figure 8 A graph showing the dynamic amplitude test of the GTA-oβCD mixture is shown.

[0021] Fig. 9 Graph showing the compressive force of hydrogels with a GTA content of 20 wt% before and after glycerol treatment.

[0022] Fig.10 Graph showing the elongation at break (% of original inner diameter) of annular hydrogels with a GTA content of 20 wt% before and after glycerol treatment.

[0023] Fig.11 Graph showing the compression modulus of 20 wt % GTA hydrogels before and after glycerol treatment.

[0024] Fig.12 A graph showing the cumulative release of bupivacaine crystals from a hydrogel matrix in 0.01 M, pH 6 citrate buffer. The hydrogel was partially coated with a PLGA membrane. Release was expressed as a percentage of the total drug content. DETAILED DESCRIPTION

[0025] Hydrogels can be synthesized by cross-linking water-soluble polymers. Water-soluble polymers such as poly(acrylic acid), poly(vinyl alcohol), poly(vinyl pyrrolidone), poly(ethylene glycol), polyacrylamide and polysaccharides (e.g., hyaluronic acid) are the most common systems for forming hydrogels. These water-soluble polymers are biocompatible and widely used in various pharmaceutical and biomedical applications. It has been found that these common hydrogels can be used as carriers for Bupi, but lack important properties, such as being flexible / deformable but firm, biocompatible and biodegradable, and allowing for extensive adjustment of their properties. Interestingly, it has been found that cross-linking of biopolymers such as gelatin and similar polymers can provide an improved carrier.

[0026] Although there are many different hydrogels, the present invention focuses on medical hydrogels that are biocompatible and can be implanted and used in the body. In addition, they must be biodegradable. Since the hydrogels include biopolymers, i.e., natural polymers produced by living organisms, the hydrogels are biodegradable, i.e., they can naturally break down in the human body.

[0027] Although the present invention is described with particular reference to the use of gelatin, any water-soluble biocompatible biopolymer comprising amino and hydroxyl groups may be used. For example, proteinaceous and / or polysaccharide polymers such as hyaluronic acid, chitosan and cellulose may be used. Preferably, the biopolymer is a proteinaceous polymer such as elastin, silk, collagen, fibrin or gelatin. More preferably, the proteinaceous biopolymer is silk, collagen, fibrin or gelatin. These polymers contain tyrosine groups that can be used for riboflavin-mediated cross-linking. Most preferably, the hydrogel is based on gelatin. The hydrogel may also include other biocompatible water-soluble synthetic or natural polymers. By weight, other polymers may account for up to 50% of the total polymer content. In view of the availability, biocompatibility and cost of gelatin, it is preferred to use gelatin as the only polymer component.

[0028] The amount of oxidized β-cyclodextrin and biopolymer, preferably gelatin, may vary within a wide range. Preferably, the amount of oβ-CD may be in the range of 0.1% to 10% by weight of the hydrogel, preferably in the range of 2% to 6% by weight of the hydrogel. Using higher amounts of oxidized β-cyclodextrin may interfere with the chemical cross-linking of gelatin due to increased interactions between the tyramine functional groups and the cavities of the oxidized β-cyclodextrin.

[0029] The biopolymer, preferably gelatin, is preferably functionalized using tyramine, 4-(2-aminoethyl)phenol as a functionalizing agent. In addition to or instead of tyramine, a polymer of the formula NH 2-Other primary aminoalkylphenols of R-PhOH and substituted forms thereof. Tyramine is the most commonly used compound to introduce phenolic hydroxyl groups on the gelatin backbone via functionalization with gelatin carboxylic acid groups. Alternatively, phenolic hydroxyl groups can be introduced via reaction with gelatin amino groups using a functionalizing agent such as hydroxyphenylpropionic acid. What is important is the biocompatibility of the functionalizing agent and its potential to form guest-host interactions with cyclodextrins. Tyramine is preferably used as the sole reagent to functionalize biopolymers in view of its availability, biocompatibility and cost.

[0030] The degree of functionalization may vary within wide limits. To ensure adequate cross-linking, preferably 5% to 50%, preferably 20% to 25%, of the carboxylic acid groups in the gelatin are reacted with tyramine or a similar functionalizing agent. If alternative biopolymers are used, a similar degree of functionalization is required.

[0031] The use of β-cyclodextrin in hydrogels is known. In the present invention, β-cyclodextrin is oxidized. In order to ensure grafting to gelatin, oxidation of β-cyclodextrin is required. The degree of oxidation can be 5% to 30%, preferably 20% to 30% of the secondary hydroxyl groups. Oxidation converts the secondary hydroxyl groups in the molecule into aldehyde groups. The preferred degree of oxidation allows maximum grafting of oβ-CD to the gelatin backbone while limiting the possible cytotoxic effects of any unreacted aldehyde groups and ensuring sufficient solubility of oβ-CD in water.

[0032] Although hydrogels based on GTA and cyclodextrin are known, the inventors have found that the physical and chemical properties of existing hydrogels can be improved. Therefore, the new hydrogels of the present invention can be implanted and fixed to a specific location where a drug is needed, in particular a location where pain relief is needed. This can be, for example, a hydrogel in the form of a deformable body, whereby the hydrogel conforms to the shape of the bone structure or surgical implant to which it is fixed. Therefore, the new hydrogels always release their contents, such as drugs like Bupi, in the appropriate position. It is important in this regard to achieve a specific crosslinking density so that the swelling degree calculated as swollen weight (at equilibrium swelling)-dry weight / dry weight is in the range of 2 to 20, preferably in the range of 2 to 6. The crosslinking density is achieved by using the following types of crosslinking:

[0033] (a) phenol-phenol cross-linking in functionalized biopolymers using primary aminoalkylphenols or similar functionalizing agents,

[0034] (b) Schiff base crosslinking between the amino groups on the functionalized biopolymer and the aldehyde groups of oβ-CD, and

[0035] (c) Guest-host interaction between the phenolic moiety of the functionalizing agent grafted on the biopolymer and the cavity of oβ-CD.

[0036] The present invention provides, inter alia, excellent control and adjustability of the formation of phenol crosslinks. Thus, hydrogels can be prepared with a variety of ratios between crosslink types (a), (b) and (c). Furthermore, by adjusting the crosslink density, the elasticity can also be varied. The importance of this is discussed below in the text, where various embodiments of the hydrogels of the present invention are discussed.

[0037] The direction of drug release can be further enhanced by partially coating the surface of the hydrogel with a coating. Attaching the implanted hydrogel near the body part to be treated and attaching the exposed surface of the hydrogel near the body part to be treated can reduce or even avoid drug release in other directions. This has the advantage of reducing side effects and making it possible to use lower concentrations of drug, optionally with a longer working time due to the slower release of conventional doses of drug.

[0038] The coating can be made up of the material of hydrogel, provided that it does not contain medicine and is thick enough. However, preferably, the coating is made up of a material that is less permeable to medicine than the material of hydrogel itself. The coating can be flexible or shell-like. Similar to hydrogel, the coating must be made up of biocompatible biopolymers. Compared with hydrogel, biodegradability can be the same or extended. Suitable materials include but are not limited to polycaprolactone (hereinafter "PCL"), poly (lactic acid-co-glycolic acid) (hereinafter "PLGA"), gelatin or alginate.

[0039] Therefore, the nature of the photoinitiator is important. Crosslinking systems for type (a) crosslinking are known in the art and are based on HRP / H 2 O 2 . The use of a combination of riboflavin, sodium persulfate (SPS), and visible light to achieve crosslinking is new. Riboflavin, also known as vitamin B2, circulates naturally in the body, is biocompatible, and is currently used in clinical applications for corneal collagen crosslinking (Belin, Michael W., et al. Cornea 2018, 37, 1218-1225). Riboflavin is exposed to visible light in the presence of SPS to produce reactive intermediates. By visible light, it is meant the portion of the electromagnetic spectrum that is visible to the human eye. Typically the human eye will respond to wavelengths between 380 nanometers and 740 nanometers, or even 780 nanometers. In particular, the present invention has been tested with wavelengths between 400 nanometers and 700 nanometers. Other useful photoinitiators are ferrocene and anthraquinone.

[0040] In addition, compared with the HRP / H 2 O 2The use of light-induced cross-linking provides better control and adjustability than with conventional systems. Preferably, riboflavin and SPS are used in a molar ratio (riboflavin:SPS) of 1:5 to 20, preferably about 1:10. For example, riboflavin and SPS may be used at 0.1 mM to 10 mM riboflavin and 1 mM to 100 mM SPS. Preferably, riboflavin is a flavin mononucleotide, which is a water-soluble form of riboflavin.

[0041] As discussed above, by varying the amount of biopolymer to cyclodextrin, by varying the functionalization of the biopolymer, by varying the amount of oxidation of the cyclodextrin, and by varying the amount of photoinitiator, the crosslink density, and therefore the degree of swelling, and the elastic modulus, can be adjusted to the intended use, e.g., a toroidal hydrogel, a sleeve-shaped hydrogel, a thumbnail-shaped hydrogel, a wedge-shaped hydrogel (e.g., for osteotomy), or other appropriately shaped hydrogels, such as a shape suitable for use between a plate and a bone. The following properties are desirable for the ability to implant and provide sustained release of a drug.

[0042] ● High degree of cross-linking, resulting in mechanically strong hydrogels and slow degradation. Therefore, the effect of degradation on drug release is minimized;

[0043] ● Targeted release of the drug towards a preferred location, for example through the use of a coating, and

[0044] • The ability to adapt to the shape of the structure to which it is applied, such as a bone structure or the surface of a surgical implant, thereby achieving intimate contact with the structure.

[0045] Turning to Figure 3, hydrogels can find different applications. In a co-pending application, the use of hydrogels as local release carriers of annular drugs is described (PCT / NL2018 / 050832, incorporated herein by reference), where the hydrogels are used in conjunction with screws. Figure 3a A hydrogel according to a previous invention is shown in FIG. Figure 3a A top view of a ring element (generally indicated by reference numeral 1) according to an embodiment of the present invention is shown. The ring element 1 comprises a deformable ring body 2 having a central opening 3. The ring element 1 is designed to be placed on the shank of a surgical screw element, such as a pedicle screw of a spinal fixation system. The ring body 2 is made of a hydrogel. Figure 3bA cross section AA of a ring element 1 is shown. The outer surface wall of the ring element 1 is formed by a bone contact surface 4 and a compartment wall 5. The bone contact surface 4 and the compartment wall 5 define a chamber 6. The chamber 6, which is substantially formed by the ring body 2, contains a drug. In a preferred embodiment, the bone contact surface 4 has a first release rate of the drug to be released after implantation, and the compartment wall 5 has a second release rate of the drug to be released. Preferably, the first release rate is substantially greater than the second release rate, for example at least 2 times, preferably at least 10 times, the second release rate. According to a preferred embodiment, the drug will be released from the chamber 6 mainly through the bone contact surface 4 according to a desired release curve over time. Figure 4 The surgical screw element 20 and the screw assembly 21 mounted on the shank 21 of the screw element 20 are shown. Figure 3a The screw element 20 is a combination of a ring element 1 and a shank 21. The shank 21 comprises a thread 22 for screwing the screw element 20 into a hole provided in the patient's bone. The screw element 20 further comprises a proximal portion 23 having a screw head 24 integrated with the shank 21 and a separate connector portion 25. It is worth noting that the screw element 20 is a standard threaded element, that is, not specifically suitable for use in combination with the ring element 1. The screw element 20 is, for example, a screw element of the spinal fixation system disclosed in US 2010 / 031228. Figure 5 The combination of the surgical screw element 20 and the ring element 1 is shown after being screwed into a hole 50 provided in a bone 51 of a patient to be treated. The bone 51 is, for example, a pedicle of the patient's spine. The hydrogel according to the invention, having a swelling degree of about 4 and an elastic modulus of 400 kPa, is very suitable for said application.

[0046] In another co-pending application, the use of hydrogels as carriers in the form of sleeves for the local release of drugs, for example for joint prostheses, is described (NL2023208, incorporated herein by reference). Figure 6a A hydrogel according to a previous invention is shown in FIG. Figure 6a An artificial hip implant, generally represented by reference numeral 60, is shown. The artificial hip implant 60 comprises a femoral component 61 attached to the patient's femur and an acetabular component 65 attached to the patient's acetabulum (hip socket). The femoral component 61 comprises a rod 62 to be placed in the femur, a neck 63 and a head 64 supported on the neck 63. The acetabular component 65 comprises an acetabular cup 66, which may have a sleeve 67 arranged in the acetabular cup 66. Artificial hip implant 60 is well known in the art. Using these types of artificial hip implants 60 to carry out hip replacement is one of the most common orthopedic surgeries at present, but short-term and long-term patient satisfaction vary greatly. In view of the possible negative effects of hip replacement, such as pain and infection risks, it is necessary to apply one or more medical active agents to the patient after implanting the artificial hip implant 60.

[0047] Figure 6b Shows Figure 6a The artificial hip joint implant 60, wherein the sleeve element 68 according to an embodiment of the present invention is arranged on the neck 63 of the femoral component 61. The sleeve element 68 includes the hydrogel according to the present invention and further includes a drug, such as a pain treatment drug, such as an anesthetic or analgesic. The hydrogel according to the present invention has a swelling degree of about 4 and an elastic modulus of 400kPa, which is very suitable for the application.

[0048] In another application, a hydrogel can be used if a portion of the hydrogel is rigid enough to act as the shaft of a "thumbnail" while the remainder of the hydrogel forms the head of the thumbnail. Figure 7a As illustrated in the figure, wherein the needle shaft is (72) and the head of the nail is (71). The drug may be contained in the needle shaft or the head of the nail. In this embodiment, it is attractive to coat the outer surface of the head (73) of the needle shaft away from the needle shaft with a biopolymer coating having a different permeability than the hydrogel. This ensures a directional release of the drug to the direction of the bone. The hydrogel according to the present invention, having a swelling degree of about 4 and an elastic modulus of 400 kPa, is very suitable for the application. Preferably, the head of the thumb nail is coated externally. In addition, the needle shaft can be made of a biodegradable material other than the hydrogel alone.

[0049] Each of these embodiments is well suited for treating musculoskeletal diseases, and is particularly useful for treating bone diseases, as the hydrogel is able to conform to the shape of the bone or implant to which it is applied. These diseases include infection, inflammation, malignant processes, growth disorders, degenerative diseases, trauma, autoimmune diseases, or the treatment of pain caused by these diseases (surgical treatment). Preferably, these diseases include infection, inflammation, malignant processes, growth disorders, degenerative diseases, or the treatment of pain caused by these diseases (surgical treatment of these diseases).

[0050] The invention is described with reference to the use of Bupi, but any (local) anesthetic may be used. Local anesthetics are generally divided into amides and esters; amides are more commonly used. The anesthetic is preferably an amino-amide local anesthetic, such as articaine, procaine, chloroprocaine, etidocaine, prilocaine, bupivacaine, levobupivacaine, ropivacaine, mepivacaine, lidocaine, dibuvacaine or other aminocaines, but may also be an ester, such as tetracaine, procaine or chloroprocaine. The anesthetic may also comprise a combination of two or more anesthetics. Preferably, the anesthetic is bupivacaine, liposomal bupivacaine or levobupivacaine, lidocaine or an anesthetic combination comprising bupivacaine, liposomal bupivacaine and / or levobupivacaine. The drug may also be or comprise an antibiotic or anticancer agent, a growth factor, an immunomodulatory drug, etc. The drug may also be hydrophilic or hydrophobic. Due to the hydrophilicity of the hydrogel, hydrophilic drugs are easily incorporated into the hydrogel. The hydrophobic cavity of oβ-CD provides encapsulation for hydrophobic drugs. Therefore, with respect to hydrophobic drugs, the hydrogel of the present invention has advantages over hydrogels without oβ-CD. Preferably, the drug is hydrophobic. An index for measuring the hydrophobicity of a drug is the octanol-water partition coefficient P, which is the ratio of the concentration of the drug in a mixture of octanol and water at equilibrium. For hydrophobic drugs, log P>0, preferably log P>2.

[0051] The hydrogel may include additional components, such as colorants, stabilizers, cosolvents, buffers and similar common additives. If and to a certain extent, bupivacaine is used as a drug, it is preferably used in an amount of 0.01 mg / mL to 200 mg / mL volume. In addition, the drug itself can be encapsulated in nanoparticles or microparticles in the size range of 50 nm to 200 μm, which are then included in the hydrogel. It can be encapsulated in PLGA, PCL, gelatin, alginate or liposomes.

[0052] The release of bupivacaine from the hydrogel can be further prolonged by in situ crystallization of the drug. Specifically, an alkaline solution is used as the swelling medium of the hydrogel to induce the formation of bupivacaine crystals in the hydrogel matrix. The slow dissolution of the crystals in the surrounding medium ensures the controlled release of bupivacaine.

[0053] In addition to the drug, one or more other ingredients may be included, preferably selected from the group consisting of auxiliary drugs, cosolvents, colorants, and buffers. An auxiliary drug may be considered to be any additional drug added to the hydrogel, preferably a drug that enhances the effect of at least one drug present in the hydrogel. Cosolvents include, but are not limited to, plasticizers. One such plasticizer is glycerol. Adding glycerol to the hydrogel matrix produces higher elasticity, but does not affect the hardness of the sample. The elongation at break of the resulting hydrogel may be between 100% and 300%, preferably between 120% and 250%.

[0054] Methods for preparing hydrogel raw materials are known. Thus, it is known to functionalize gelatin and related biopolymers with tyramine and related primary aminoalkylphenols. Likewise, oxidized cyclodextrins are known. See Thi et al., RSC Adv. 2017, cited above and incorporated herein by reference. However, it is important to remove all forms of contamination, which is common in the field of medical applications. For example, hydrogels can be prepared by the following method:

[0055] 1. Prepare solutions of GTA, oβ-CD, SPS, riboflavin, and bupivacaine.

[0056] 2. The solutions are mixed to obtain predetermined concentrations of GTA, oβ-CD, SPS, and riboflavin. These concentrations can be varied depending on the desired mechanical and release properties, with higher GTA, oβ-CD (within a certain range), and SPS concentrations producing more tightly cross-linked hydrogels.

[0057] 3. The resulting solution is then exposed to visible light for a predetermined amount of time. The time can vary depending on the desired mechanical properties, with low exposure times producing less densely cross-linked hydrogels.

[0058] 4. The resulting hydrogel is then immersed in a bupivacaine solution to allow the bupivacaine (and, if necessary, a cosolvent such as glycerol) to diffuse into the hydrogel.

[0059] 5. The hydrogel is then allowed to dry and set aside.

[0060] 6. Optionally, the hydrogel may be partially coated with a biopolymer solution having a different permeability to bupivacaine than the hydrogel to ensure directional release of the encapsulated drug. The coating may also enhance the mechanical properties of the hydrogel. After step 6, the coating may be applied to the hydrogel. Alternatively, a coating shell of a predefined shape may also be formed and the solution of step 2 introduced into the shell, so that the coating acts as a mold for the hydrogel.

[0061] For example, a coated hydrogel can be prepared as follows:

[0062] 1. Prepare a solution of biopolymer (e.g. PCL) in an appropriate solvent (in the case of PCL this is dichloromethane). The concentration range of PCL solution used was 0.5 wt% to 25 wt%.

[0063] 2. Two options are available:

[0064] a. Dip the metal mold into the biopolymer solution. Apply the polymer solution to the mold. Then dry the polymer solution and remove it from the mold. A coating in the shape of the mold is produced.

[0065] b. Immerse the hydrogel in a solution of the biopolymer. Coat the hydrogel with the solution. Then allow the polymer solution to dry on the hydrogel to form a coating around the hydrogel. Using another method, the PCL solution can also be sprayed on the hydrogel to achieve a coating by spraying.

[0066] 3. Depending on the viscosity of the polymer solution, the duration of the dipping step, the speed of the dipping step (immersion and retraction) and the number of dipping repetitions (eg, between 1 and 10), the desired coating thickness can be obtained.

[0067] 4. Then, the coating provides directional release, slow release and degradation from / to the hydrogel due to the reduction of diffusion surface and some mechanical support to the hydrogel.

[0068] Bupivacaine encapsulated in nanoparticles / microparticles can be added to the solution in step 2 to load the hydrogel with encapsulated bupivacaine prior to exposure to light in step 3. In this case, additional bupivacaine can be introduced into the hydrogel in step 4, but step 4 can also be skipped.

[0069] If tyramine-functionalized gelatin (GTA) is used in the above method, biopolymers functionalized with other functionalizing agents may be used in place of GTA or in addition to GTA. Similarly, in the above method, when bupivacaine is used as a drug, other drugs may be used, as well as other ingredients selected from auxiliary drugs, cosolvents, colorants, and buffers.

[0070] This method can be used to crosslink hydrogels outside the body where the hydrogels are used. As a modification of the above method, step 6 is omitted, and the drug can also be mixed with a solution of GTA, oβ-CD, riboflavin, SPS and bupivacaine and injected into the body at the location where the drug is to be administered, such as by perfusion or injection. Step 3 is then performed in vivo.

[0071] Example

[0072] Use the method described in the instruction manual.

[0073] Material:

[0074] Gelatin (porcine skin, type A, 300 g gel strength), 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide (EDC), N-hydroxysuccinimide (NHS), tyramine hydrochloride, 2-morpholineethanesulfonic acid monohydrate (MES), sodium persulfate (SPS), sodium bicarbonate (NaHCO 3), sodium periodate, β-cyclodextrin, glycerol, phosphate buffered saline (PBS), riboflavin (RB), and ethylene glycol were purchased from Sigma Aldrich. Cellulose dialysis membrane (Spectrum / Por TM , 0.5 kDa; 12 kDa molecular weight cut-off) was purchased from Spectrum Laboratories. Bupivacaine was obtained from Siegfried, Switzerland.

[0075] Synthesis of Gelatin-Tyramine (GTA):

[0076] Type A gelatin was dissolved in MES buffer at 50°C, and then tyramine, EDC, and NHS were added. The reaction mixture was stirred overnight. The mixture was then dialyzed with water, and the product was obtained by freeze drying.

[0077] Tyramine content measurement

[0078] The degree of functionalization of gelatin was determined by measuring the absorbance of polymer solutions (0.1%, w / v) at 275 nm and calculating from a calibration curve obtained by measuring the absorbance of known percentages of tyramine in distilled water.

[0079] Oxidation of β-cyclodextrin:

[0080] Oxidized β-cyclodextrin was prepared by reaction with sodium periodate. Briefly, β-cyclodextrin was dispersed in distilled water, then sodium periodate was added and stirred overnight at room temperature in the dark. The reaction was terminated by adding ethylene glycol. The mixture was dialyzed against deionized water for 3 days using a dialysis membrane with a MWCO of 500 Da (Spectrum Labs), and the product was collected by freeze drying. Deuterated dimethyl sulfoxide (DMSO-d6) or deuterium oxide (D2O) was used as solvent by 1 The degree of oxidation is determined by H NMR. Although the proton ratio of β-cyclodextrin is 4.8 to 4.9 ppm, and the proton ratio of 4 ppm is about 2.04, the progress of the reaction can be observed by changing the ratio to about 1.49.

[0081] Preparation of GTA / β-cyclodextrin hydrogel:

[0082] Prior to cross-linking of the hydrogel, solutions of GTA, oβ-CD, SPS, and riboflavin were prepared.

[0083] Unless otherwise stated, the degree of functionalization of GTA is between 10% and 25%, while the degree of oxidation of the subhydroxyl groups of oβ-CD is between 15% and 30%. These solutions are mixed to obtain final concentrations of 20 wt% of GTA, 0 wt% to 10 wt% of oβ-CD, 0 mM to 100 mM of SPS and 0 mM to 10 mM of riboflavin. Samples without oβ-CD serve as controls. The obtained solutions are exposed to visible light, in particular to visible light using a white light lamp in the wavelength range of 400 nm to 700 nm, to ensure hydrogel formation.

[0084] Determination of swelling ratio:

[0085] Swelling studies were performed on disc-shaped GTA hydrogels prepared using custom molds. After synthesis, the hydrogel samples were immersed in PBS at pH 7.4 and 37 °C for 24 h or until equilibrium swelling was reached. At predetermined time points, excess water was removed and the samples were weighed (W s The hydrogel was then freeze-dried to obtain the dry weight (W d The swelling ratio is defined as (W s -W d ) / W d .

[0086] To determine the gel composition, the samples were freeze-dried after synthesis (W d1 ), and then soaked in water for 24 hours to remove the soluble components. The hydrogel was then dried (W d2 ), and the gel fraction (%) was calculated as (W d1 / W d2 )×100%.

[0087] crystallization

[0088] To induce the formation of bupivacaine crystals, after synthesis, the hydrogel was incubated in 0.1 M NaHCO 3 The hydrogel was then immersed in an aqueous buffer solution (pH 8.5) (optionally containing 30% (v / v) glycerol) for 2 hours to induce drug crystallization (optionally allowing glycerol to diffuse) into the hydrogel. The hydrogel was then dried at 37°C.

[0089] Mechanical testing:

[0090] The compression tests on the disc-shaped swollen hydrogels were performed using a dynamic mechanical analyzer (DMA Q800, TA Instruments, UK). The experiments were performed at room temperature in controlled force mode with a force ramp rate of 3 N min -1 to 18 N. The elastic modulus is determined from the slope of the stress-strain curve at 5% to 10% strain.

[0091] The force required to induce plastic deformation (i.e. not return to original form / shape / state) of the tested material was measured at room temperature using a Mark-10 ES10 manual force test stand with a maximum capacity of 500 N. A parallel plate design was used on a hydrogel with a diameter of 5 mm and a cross section of 1 cm. The applied compressive force was measured using a digital force gauge in the Mark-10 ES10.

[0092] The elongation of the ring-shaped hydrogel was measured using a custom method by mounting the sample on a caliper tool. The initial inner diameter of the ring was first measured using a caliper and then gradually extended. The inner diameter when the ring broke was recorded. The elongation was then calculated as ((ID at break - ID at rest) / ID at rest) * 100%.

[0093] Rheological behavior:

[0094] In order to study the interaction between GTA and oβCD, rheological experiments were performed.

[0095] Rheology - Gel Time

[0096] Immediately after mixing the solutions of functionalized gelatin and oβCD (final concentrations of 20% and 2% (w / v), respectively), the solutions were deposited on a rheometer plate and the gelation was monitored over time at 37°C (strain set to 1% constant and angular frequency set to 1 rad / s). At approximately 8 minutes, a crossover between the storage modulus (G') and the viscous modulus (G") was observed, with G' being higher than G", indicating a gel point. The gelation was monitored for a total of 90 minutes, after which the experiment was stopped. This experiment clearly demonstrated that in addition to the phenol-phenol coupling, there were additional crosslinks (Schiff base reaction; host-guest interactions) between oβCD and the modified polymer in the final hydrogel.

[0097] Rheology - Dynamic Amplitude Testing (High Strain Deformation)

[0098] After the gelation experiment, dynamic oscillation amplitude tests were performed to evaluate the reversibility of the cross-linking interactions in the network. Alternating cycles were performed at low strain (1%) and high strain (2000%) at a constant angular frequency of 1 rad / s, and G' and G" were monitored. Each cycle was performed for 200 s ( Figure 8 ). At low strain (1%), G' (open circles) is higher than G" (grey circles), indicating solid-like behavior. Then, in the next cycle, the strain is increased to 2000% and G" is higher than G', indicating viscous-like behavior. Next, when the strain is reduced, the material is able to recover its original stiffness. This experiment shows the self-healing ability of the hydrogel material, which is due to the interaction between oβCD and tyramine-functionalized gelatin, which acts as a sacrificial bond in addition to the dityrosine cross-linking, helping to improve the mechanical properties of the hydrogel material.

[0099] Drug loading and in vitro drug release assay

[0100] To investigate the drug release properties, the obtained hydrogels were loaded with bupivacaine by immersing them in a 50 mg / mL aqueous solution of bupivacaine for 24 h, followed by pH-induced crystallization of bupivacaine within the matrix.

[0101] Bupivacaine release from the hydrogels was measured at 37°C by placing the hydrogels in vials containing 1 mL of 0.1 M citrate buffer, pH 6.

[0102] At predetermined time points, 100 uL aliquots were taken from the release solution and replaced with fresh buffer. The samples were diluted 1:10. The release of bupivacaine was determined by UPLC using a mixture of ammonium formate (10 mM, pH 2.4) and acetonitrile / water / formic acid (96:5:0.2, v:v:v) as the mobile phase.

[0103] Analyze the results

[0104] The swelling ratio and sol fraction of the hydrogels were analyzed and increased with the increase of photoinitiator concentration. The swelling ratio was similar at all tested concentrations; slightly above 4. The sol fraction (non-crosslinked part of the gel) was lowest at 2mM concentration, close to 0%, and about 1% for the other hydrogels. 2mM is the preferred concentration for the hydrogels.

[0105] Hydrogels (control) were prepared based on 10 wt%, 15 wt% and 20 wt% GTA using riboflavin / SPS as a photoinitiator by exposure to visible light for a fixed time. The maximum compressive strength of the hydrogel with 20 wt% GTA was 52 ± 13 N. After glycerol treatment, the hydrogel with 10 wt% GTA showed a compression force of 200 N to 400 N, and the hydrogel with 15% GTA showed a compression force greater than 400 N. The hydrogel with 20% GTA was the most resistant to compression, with a score well above 400 N ( Fig. 9 ).

[0106] Furthermore, the addition of glycerol to the hydrogel matrix improved the elasticity, as the elongation of the annular hydrogels significantly increased from 80.3 ± 6.5% to 189.3 ± 56.8% ( Fig.10 ).

[0107] Similarly, hydrogels (control) were prepared, but now with a fixed GTA content of 20 wt%, using increasing SPS concentrations and then exposed to visible light. At 0 mM SPS, no hydrogel was formed. At low concentrations, the hydrogel did not retain its form. Higher concentrations of SPS caused the hydrogel to exceed the maximum capacity of the mechanical tester (500 N). Based on the present invention, the selected SPS concentration provides the best mechanical properties and the lowest cytotoxic effects.

[0108] Hydrogels were prepared again, but now with a fixed GTA content of 20 wt%, a fixed RB / SPS concentration and different amounts of oβ-CD, 0 wt% (control), 2 wt% (invention), 4 wt% (invention) and 6 wt% (invention). The higher the oβ-CD concentration, the less swelling and the higher the degree of cross-linking. The swelling ratio was about 6 to 3.

[0109] The effect of irradiation time on the compression modulus was evaluated on hydrogels with a GTA content of 20 wt%, a fixed RB / SPS ratio, and a fixed oβ-CD concentration. The compression modulus of the hydrogel increased with increasing irradiation time. The compression modulus was from 167.5 (±20) kPa (5 minutes), from 328.3 (±29) kPa (10 minutes); from 455.7 (±63) kPa (20 minutes), from 552.3 (±138) kPa (30 minutes). With increasing irradiation time, the swelling degree of the hydrogel decreased, which is consistent with the increase in cross-linking degree. The swelling reached about 4 within an irradiation time between 10 and 30 minutes. The addition of glycerol to the hydrogel did not result in a change in the compression modulus ( Fig.11 ).

[0110] Similarly, hydrogels were prepared with different concentrations of oβ-CD. The effect of adding bupivacaine to the hydrogel on the compression modulus was tested. It was found that it had no effect on the compression modulus. As mentioned before, the increase of oβ-CD did increase the compression modulus of the hydrogel.

[0111] The hydrogels were prepared according to Thi et al., RSC Adv. 2017 cited above, where the GTA content was 20 wt%, the HRP content was 0.5 EU / mL, and the H 2 O 2 The hydrogels prepared according to the present invention have similar contents of GTA and oβ-CD but are prepared with RB and SPS as photoinitiators. At the same polymer concentration, the hydrogels according to the present invention show higher compression modulus.

[0112] The cytotoxicity of the individual components of the hydrogels of the present invention was tested using a live / dead assay using human MSC cells. Cells were exposed to the compounds (dissolved in PBS) at the specified concentrations for 1 hour. Cell viability was compared to control wells where the hydrogels were exposed to PBS only for 1 hour. After 48 hours, a live / dead assay was performed according to the manufacturer's protocol. For all concentrations of riboflavin and oβ-CD tested, cell viability was similar to the control wells.

[0113] Using alamar Blue TM The metabolic activity of human MSC cells was evaluated. The cells were exposed to hydrogels according to Thi et al., RSC Adv. 2017 cited above, or to hydrogels with equivalent GTA content according to the present invention for 48 hours. Subsequently, the analysis was performed according to the protocol. The metabolic activity of the tested hydrogels was almost similar to that of the control group, confirming the cytocompatibility of the current photoinitiator. No significant differences were observed in cell viability or metabolic activity in the hydrogels before and after glycerol treatment.

[0114] Furthermore, the use of an alkaline buffer to induce bupivacaine crystallization within the hydrogel did not result in any cytotoxic effects upon contact with cells.

[0115] The release of crystallized bupivacaine from partially coated PLGA hydrogels in 0.01 M citrate buffer at pH 6.0 was over 168 hours. The release profile was characterized by an initial burst release (approximately 20% of the total drug content in the first 8 hours) followed by a nearly linear release phase of bupivacaine. After 168 hours of drug release, 20% of the total drug content remained within the hydrogel as determined by completion of the hydrogel degradation experiment ( Fig.12 ).

[0116] To determine the effect of coating on directional release, hydrogels containing methylene blue were cast on 180 μm thick PCL films obtained by casting. The release of methylene blue in hydrogels was simulated in 3% alginate gel, cross-linked with calcium chloride to obtain a tissue-like consistency. Only the top of the hydrogel was coated with PCL. The samples were examined at 0, 1, 2 and 3 hours. The hydrogels were placed vertically to eliminate any effect of gravity on the direction of release. Visual inspection showed that methylene blue was released only in the non-PCL coated direction. The same experiment was performed for bupivacaine. The cumulative release was evaluated and it was found that the cumulative release reached about 2 mg after 8 hours from the PCL coated direction, while in the non-PCL coated direction, this cumulative release was reached within 1 hour. This confirms that the coating can be used to provide directional release of drugs within the hydrogel.

Claims

1. A hydrogel for in vivo release of a drug, comprising at least one drug, wherein the hydrogel include: (i) a proteinaceous biopolymer functionalized with a functionalizing agent that introduces phenolic hydroxyl groups capable of forming guest-host interactions with oxidized β-cyclodextrin, and (ii) oxidized β-cyclodextrin (oβ-CD), wherein the hydrogel is crosslinked upon exposure to visible light in the presence of a biocompatible photoinitiator such that the degree of swelling calculated as (swollen weight - dry weight) / dry weight is in the range of 2 to 20, and wherein the biocompatible photoinitiator is a combination of riboflavin and sodium persulfate, wherein the proteinaceous biopolymer (i) is gelatin, wherein the hydrogel comprises an elastic modulus between 100 kPa and 600 kPa, the elastic modulus being measured from the slope of a stress-strain curve obtained by subjecting the hydrogel to a controlled force mode at 3 N min -1 The force ramp-up rate up to 18N was obtained by dynamic mechanical analysis at strains between 5% and 10%, wherein the hydrogel has an elongation at break between 100% and 300%, the elongation at break being measured from a ring-shaped hydrogel by mounting the sample on a caliper tool, measuring the initial inner diameter (ID) of the ring, and then gradually extending the ring, and calculating the elongation as ((ID at break - ID at rest) / ID at rest)*100%.

2. The hydrogel according to claim 1, wherein the functionalizing agent is a primary aminoalkylphenol.

3. The hydrogel according to claim 2, wherein the functionalized proteinaceous biopolymer is tyramine-functionalized gelatin (GTA). The hydrogel according to claim 1 , further comprising a co-solvent. The hydrogel according to claim 4 , wherein the cosolvent is a plasticizer.

6. The hydrogel according to claim 5, wherein the plasticizer is glycerol.

7. The hydrogel according to claim 1 or 2, wherein 10% to 30% of the secondary hydroxyl groups in oβ-CD have been converted into aldehyde groups.

8. The hydrogel according to claim 7, wherein 15% to 25% of the secondary hydroxyl groups in oβ-CD have been converted into aldehyde groups.

9. The hydrogel according to claim 1 or 2, wherein the amount of oβ-CD is 0.1% to 10% by weight of the hydrogel.

10. The hydrogel according to claim 9, wherein the amount of oβ-CD is in the range of 2% to 6% by weight of the hydrogel.

11. The hydrogel according to claim 1 or 2, comprising bupivacaine as a drug in an amount of 0.01 mg / mL to 200 mg / mL by volume.

12. The hydrogel according to claim 11, comprising bupivacaine as a drug in crystalline form.

13. The hydrogel according to claim 1 or 2, wherein the drug is encapsulated in the proteinaceous biopolymer.

14. The hydrogel of claim 13, wherein the drug is encapsulated in PLGA, PCL, gelatin, alginate or liposomes.

15. The hydrogel according to claim 1 or 2, comprising bupivacaine as a drug and one or more other ingredients.

16. The hydrogel of claim 15, wherein the one or more other ingredients are selected from the group consisting of auxiliary drugs, colorants and buffers.

17. The hydrogel according to claim 1 or 2, comprising the following types of crosslinks: (a) phenol-phenol cross-linking in the proteinaceous biopolymer functionalized with the functionalizing agent; (b) Schiff base crosslinking between the amino groups present on the functionalized proteinaceous biopolymer and the aldehyde groups of oβ-CD, and (c) Guest-host interaction between the phenolic moiety of the functionalizing agent grafted on the proteinaceous biopolymer and the cavity of oβ-CD.

18. The hydrogel of claim 17, wherein the functionalizing agent is a primary aminoalkylphenol.

19. The hydrogel according to claim 1 or 2, partially covered by a coating.

20. The hydrogel of claim 19, wherein the coating is a biodegradable polymer coating.

21. The hydrogel of claim 1 or 2, wherein the functionalizing agent is tyramine.

22. The hydrogel according to claim 1 or 2, wherein the drug is a hydrophobic drug.

23. The hydrogel of claim 22, wherein the proteinaceous biopolymer functionalized with the functionalizing agent is gelatin functionalized with tyramine (GTA).

24. The hydrogel according to claim 1 or 2, having a swelling degree in the range of 4 and an elastic modulus of 400 kPa, the elastic modulus being measured from the slope of a stress-strain curve obtained by subjecting the hydrogel to a controlled force mode at 3 N min -1 The force ramp rate up to 18N was obtained by dynamic mechanical analysis at strains between 5% and 10%, and (a) have holes for attachment to bone, or (b) is shaped as a sleeve to surround a portion of a bone, or (c) is shaped like a thumb nail and includes a rigid portion for fixation to bone.

25. The hydrogel of claim 1 having an elongation at break between 120% and 250%.

26. A method for preparing a hydrogel according to any one of claims 1 to 25, said method include: preparing a mixed solution of a protein biopolymer functionalized with the functionalizing agent, oβ-CD, a biocompatible photoinitiator and a drug, exposing the solution to visible light to form a hydrogel, and If the drug is not included in the mixed solution, contacting the hydrogel with a solution of the drug, thereby allowing the drug to diffuse into the hydrogel, The hydrogel is allowed to dry.

27. The hydrogel according to claim 1 or 2, for use in treating musculoskeletal diseases.

28. The hydrogel according to claim 27, for use in the treatment of musculoskeletal infections, inflammations, malignant processes, growth disorders, degenerative diseases, trauma, autoimmune diseases or the treatment of pain caused by these diseases.

29. The hydrogel according to claim 27 or 28, which is prepared in vivo by administering a combination of: (a) a proteinaceous biopolymer functionalized with a functionalizing agent mixed with a drug in liquid form or encapsulated in microparticles, and (b) oxidized oβ-CD, and wherein the combination is exposed to visible light in vivo to generate the hydrogel in vivo.

30. The hydrogel of claim 29, wherein the administration is injection.

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