Title of Invention: Composite Material for Drug Delivery

Technical Field

The present invention relates to a composite material for delivery of active ingredients. The present invention also relates to a method of fabricating such a composite material.

Background Art

Polysaccharides are ubiquitous in in the extracellular matrix (ECM), where they interact with specific proteins and growth factors to enhance their activities. Heparan sulfate, as one kind of polysaccharide, has been reported to enhance the activity of BMP-2, thus promoting bone regeneration. Polysaccharides can withstand heat and pressure, retaining their functionality and bioactivity to a greater extent than proteins or growth factors. The incorporation of polysaccharides into a matrix, destined for their controlled release, is expected to find applications in enhancing the activity of endogenous growth factors. Such a composite material may be utilised in the form of a scaffold, thus enabling their long-term sustained release and engendering therapeutic benefit in tissue regeneration applications.

Polycaprolactone (PCL) is a biodegradable polyester that has been approved by Food and Drug Administration (FDA) for specific applications in the human body. It has been used to engineer scaffolds, which possess load bearing properties and are considered suitable as being for bone tissue engineering applications. Its implementation is generally favoured over other materials, such as poly(lactic-co-glycolic acid) (PLGA) and poly(lactic acid) (PLA), as it avoids forming acidic by-products during its degradation. Moreover, PCL's degradation kinetics is on a time- scale that is compatible with bone regeneration applications. However, PCL's hydrophobic nature and its slow degradation, over several months, generally gives rise to the release profile of incorporated hydrophilic molecules with limited tenability.

Poly(ethylene glycol) (PEG) and poly(ethylene oxide) (PEO) are non-toxic, water-soluble polymers that have been extensively used as drug carriers and drug release modifier, due to their capability of improving the wettability of the matrix. Although the release of incorporated molecules can be extended by introducing PEG additive to PCL matrix, phase separation between hydrophobic PCL and hydrophilic PEG may occur during the molten phase of a blending process. The intrinsic phase separation alters the composite's microstructure, thus affecting its performance. The release of hydrophilic drugs is particularly sensitive to the microstructure of the composite, compared to hydrophobic drugs, since hydrophobic drug's release is more dependent on its diffusion within matrix while hydrophilic drug relies on its solubility in water. Therefore, it will be meaningful to modulate the potential phase separation within the composite during the process of drug release controlling.

There is therefore a need to provide a composite material that overcomes or at least ameliorates, one or more of the disadvantages described above. Summary

In first aspect, there is provided a composite material for delivering an active ingredient, the composite material comprising a substantially homogeneous mixture of a biodegradable polyester; a ceramic; and a porogen.

In another aspect, there is provided a composite material for delivering an active ingredient, the composite material comprising a substantially homogeneous mixture of a biodegradable polyester and a ceramic; wherein the composite material comprises pores.

The delivery of drug molecules is generally regulated by their diffusion and the degradation of the polymer matrix. Advantageously, the hydrophobic nature of polyesters such as polycaprolactone (PCL) and its slow degradation, over several months, enables the long-term, sustained release of small molecules. More advantageously, the composite material of the present disclosure relies on the hydrophobicity of the biodegradable polyester such as polycaprolactone and their diffusion within a matrix.

In an embodiment, the active ingredient is heparin. Heparin is a model polysaccharide with a molecular weight of about 12-15 kDa: a comparatively large hydrophilic molecule, which may exhibit a burst release from pure PCL, followed by negligible sustained release, or over time- scales release far longer than therapeutically-relevant time scales for tissue engineering. At low loading, poor delivery efficiency may observed which may be associated with heparin release from the surface of the material, while high loading of heparin may result in delivery of most of the incorporated heparin which may be associated with heparin release from the bulk material. Advantageously, the addition of ceramics such as tricalcium phosphate (TCP) particles, uniformly distributed within the PCL matrix may facilitate the controlled release of heparin, resulting in greater delivery efficiency.

TCP is a hydrophilic and water-insoluble particle. Advantageously, this biocompatible, osteoconductive ceramic may be incorporated into PCL matrix to enhance its mechanical strength. Further advantageously, TCP may modulate the formation of phase-separated pore domains such as poly(ethylene glycol) (PEG) domains within the PCL matrix, regulating the release profile of hydrophilic active ingredients. Advantageously, hydrophilic and water- insoluble inorganic particles, may be used to promote the compatibility of immiscible polymer blends, which may prevent phase separation. Advantageously, this may result in a composite structure having increased porosity in order to provide a sustained, tuneable release profile for active ingredients such as heparin. Scaffolds fabricated from the composite material of the present disclosure may find applications in tissue engineering, such as bone and cartilage regeneration, or surgical patches and wound/skin care applications. While this improvement may be ascribed to the more hydrophilic nature of the composite material created by the hydrophilicity of TCP, the sustained release of active ingredients such as heparin may not be substantially improved and its delivery efficiency may remain poor at low loading. This may be attributed to the insolubility of the TCP in aqueous medium, which may not be directly contributing to the formation of an interconnected channel network and therefore enhancement in deliver of the active ingredient such as heparin.

Advantageously, the introduction of porogens such as soluble poly(ethylene glycol) (PEG) additive to the PCL/TCP composite may give rise to a sustained, long-term release of the active ingredient such as heparin, with enhanced delivery efficiency. This is in contrast to the addition of PEG to only PCL, which may result in lower delivery efficiency and low sustained release, after two weeks. Advantageously, the aqueous dissolution of PEG may contribute to generating an interconnected porous network within the composite material. Further advantageously, TCP may play a role in regulating the structure of the interconnected pores generated by the PEG. During heat pressing, TCP may act as a solid, hydrophilic intermediary, thus nucleating the formation of individual PEG domains within PCL matrix. In the absence of TCP particles, phase-separation between PCL and PEG may induce aggregation of PEG domains, which may result in larger pores being formed and thus potentially hinder the formation of a network of interconnected channels in the composite material, after dissolution of PEG domains in aqueous medium. Thus, TCP may advantageously stabilise PEG domains, formed within the PCL matrix, altering the distribution of the PEG domains in the composite material. TCP may therefore advantageously play a key role on the sustained release of active ingredients such as heparin, by reducing the size of hydrophilic PEG domains and causing them to be more uniformly distributed throughout the composite material. The resulting network of interconnected pores, formed by aqueous dissolution of PEG and assisted by TCP, may enhance the delivery efficiency of active ingredients such as hydrophilic heparin molecules.

Advantageously, PEG may be completely removed from the PCL/PEG/TCP composite by dissolution in aqueous medium within a day, while only 75% of PEG may be removed from the PCL/PEG composite. This may further support the hypothesis that the TCP particle nucleation of PEG domains within the PCL matrix may generate a continuous network of porous domains. Further advantageously, despite the rapid dissolution of porogens such as PEG, the delivery of active ingredients such as heparin may persist for several weeks. Advantageously, adsorption of active ingredients such as heparin onto the surface of the ceramic such as the TCP particles may contribute to its observed long-term release. Further advantageously, a gradual heparin desorption may contribute significantly to its sustained release, over several weeks. Further advantageously, at modest loading of porogens such as PEG, the mechanical properties of the composite material may be comparable to those that do not contain PEG. This shows that the uniform distribution of PEG domains within the PCL matrix, induced by TCP particles, may result in uncompromised mechanical properties. Further advantageously, the bioactivity of the active ingredient such as heparin released from the composite material may not be affected by the fact it was previously incorporated into the composite material.

Advantageously, the innovation of a composite material, blended from a biodegradable polyester such as PCL, a ceramic such as TCP and a porogen such as PEG constituents that are individually FDA-approved for human implantation, may enable tunable and sustained release of heparin bioactive molecules. The composite material of the present disclosure may facilitate long-term release of hydrophilic molecules from a slowly degrading hydrophobic matrix. Tuning the structure's composition may regulate the release profile of active ingredients such as heparin, enabling its long-term delivery. The aqueous dissolution of the porogen such as PEG may ensure high heparin delivery efficiency, through an interconnected pore network. This may be enabled by the synergy of hydrophilic, insoluble TCP particles with the hydrophilic PEG additive, within a hydrophobic matrix. The intrinsic phase separation between hydrophobic PCL and hydrophilic PEG during heat processing may be effectively reduced by the presence of TCP particles, and thus may result in the formation of distinct and uniformly distributed PEG domains within the PCL matrix. Furthermore, the adsorption of heparin onto TCP particles may also help to slow down the leaching process, thus accounting for a significant fraction of the structure's long-term release profile of the active ingredient. An in vitro assessment of heparin bioactivity indicates that it remains unaltered, despite material processing and the molecule's release into aqueous medium. The mechanical properties of the composite may also remain uncompromised despite the addition of additives and as such, this material may be useful in the fabrication of tissue engineering scaffolds, as used in bone and cartilage regeneration, surgical patches for wound healing and skin care applications, and the implementation of localized drug release. In another aspect, there is provided a method for fabricating a composite material for delivering an active ingredient as defined above, comprising the steps of: contacting a biodegradable polyester, a ceramic, and a porogen to form a mixture; pulverizing the mixture; sintering the pulverized mixture into a mold to form the composite material; and removing the porogen.

In an embodiment, the method may not require the use of any sacrificial material or template. Advantageously, the method of the present disclosure comprises a porogen in the initial mixture, facilitating the formation of the pores without having to use any sacrificial material or template. For example, conventional techniques for fabricating porous materials may require the use of a sacrificial material which must be discarded during the fabrication process and therefore does not form part of the final product. In other conventional techniques, a template may be used, which again does not form part of the final product. The use of a sacrificial material or a template may result in materials being wasted and additional steps in the fabrication process for removing the sacrificial material and/or the template, causing the process to be less cost- efficient and time -efficient. As such, the method of the present disclosure has the advantage that it does not require the use of sacrificial materials or templates during the fabrication process. In the method of the present application, the pores are formed in situ by incorporating a porogen into the initial mixture for forming the composite material. In another aspect, there is provided the use of the composite material as defined above, for delivery of an active ingredient.

Definitions

The following words and terms used herein shall have the meaning indicated: The term "substantially homogeneous" as used herein refers to particles which are generally evenly dispersed within the mixture or dispersion without substantial agglomeration within the resin media or separation from the resin media.

The term "ceramic material", for the purposes of this application, refers to inorganic, non- metallic materials made from compounds of a metal and a non metal. Ceramic materials may be crystalline or partly crystalline.

The term "active ingredient" for the purposes of this application, refers to any component that provides pharmacological or biological activity or other direct effect in the diagnosis, cure, mitigation, treatment, or prevention of disease, or to affect the structure or any function of the body of man or animals. The term "porogen", for the purposes of this application, refers to any substance that is physically disposed within a settable composition to reserve space while the composition is being prepared but once the composition is at least partially set, the porogen is preferentially or selectively removed over time to result in porosity in the set or partially set composition to provide latent pores in the settable composition. The term "sacrificial material", for the purposes of this application, refers to any material that forms part of a substrate, which must be removed from the final product. The sacrificial material may be mixed into the substrate or deposited on the surface of the substrate. The sacrificial material may be part of the substrate from the beginning of the manufacturing process, or at any time during the manufacturing process, but may not be part of the final product.

The term "template", for the purposes of this application, refers to an object used to replicate shapes or designs. The template may be used during the manufacturing process to replicate a particular shaper or design, but may not be part of the final product.

The terms "pharmaceutically acceptable excipient", "pharmaceutically compatible excipient", and "excipient" are used interchangeably in this disclosure, and refer to substances such as disintegrators, binders, fillers, and lubricants used in formulating pharmaceutical products. They are generally safe for administering to humans according to established governmental standards, including those promulgated by the United States Food and Drug Administration.

The word "substantially" does not exclude "completely" e.g. a composition which is "substantially free" from Y may be completely free from Y. Where necessary, the word "substantially" may be omitted from the definition of the invention. Unless specified otherwise, the terms "comprising" and "comprise", and grammatical variants thereof, are intended to represent "open" or "inclusive" language such that they include recited elements but also permit inclusion of additional, unrecited elements.

As used herein, the term "about", in the context of concentrations of components of the formulations, typically means +/- 5% of the stated value, more typically +/- 4% of the stated value, more typically +/- 3% of the stated value, more typically, +/- 2% of the stated value, even more typically +/- 1% of the stated value, and even more typically +/- 0.5% of the stated value.

Throughout this disclosure, certain embodiments may be disclosed in a range format. It should be understood that the description in range format is merely for convenience and brevity and should not be construed as an inflexible limitation on the scope of the disclosed ranges. Accordingly, the description of a range should be considered to have specifically disclosed all the possible sub-ranges as well as individual numerical values within that range. For example, description of a range such as from 1 to 6 should be considered to have specifically disclosed sub-ranges such as from 1 to 3, from 1 to 4, from 1 to 5, from 2 to 4, from 2 to 6, from 3 to 6 etc., as well as individual numbers within that range, for example, 1, 2, 3, 4, 5, and 6. This applies regardless of the breadth of the range.

Certain embodiments may also be described broadly and generically herein. Each of the narrower species and subgeneric groupings falling within the generic disclosure also form part of the disclosure. This includes the generic description of the embodiments with a proviso or negative limitation removing any subject matter from the genus, regardless of whether or not the excised material is specifically recited herein.

Detailed Disclosure of Optional Embodiments

Exemplary, non-limiting embodiments of a polymer matrix will now be disclosed.

A composite material for delivering an active ingredient may comprise a substantially homogeneous mixture of a biodegradable polyester; a ceramic; and a porogen.

A composite material for delivering an active ingredient may comprise a substantially homogeneous mixture of a biodegradable polyester and a ceramic; wherein the composite material comprises pores.

The composite material may comprise a continuous matrix of said biodegradable polyester, wherein said matrix comprises a homogeneous or substantially uniform dispersion of ceramic particles throughout said matrix. The ceramic particles may be micro-sized or nano-sized. The ceramic particles may possess a substantially uniform size distribution.

The composite material may be both biocompatible and bioresorbable. The biodegradable polyester may confer suitable mechanical properties to the scaffold, the porogen may tune the release rate of the active ingredient and the ceramic may nucleate and uniformly distribute the porogen within the biodegradable polyester.

The porogen may be optionally present in the composite material. The porogen may be disposed within said polyester matrix to form a network of optionally interconnected channels distributed throughout the polyester matrix. The porogen may be polymeric or oligomeric. The porogen may be substantially non-reactive with said polyester matrix. The porogen may form discrete or nucleated domains within said polyester matrix. The polyester matrix structure may comprise a plurality of recesses or pocket domains ("pore" or "pore -like structure"), which may be connected or disjointed, wherein each pore may be occupied by a porogen polymer or oligomer, optionally nucleated or coupled to a ceramic particle.

A composite material for delivering an active ingredient, the composite material comprising a substantially homogeneous mixture of a biodegradable polyester; a ceramic; and optionally, a porogen, wherein the composite material comprises a plurality of pores which, when in use, facilitates the delivery of the active ingredient.

The biodegradable polyester may be selected from the group consisting of polycaprolactone (PCL), polyhydroxybutyrate, polylactic acid (PL A), poly(DL-lactide) (PDLLA), poly(DL-lactic acid); poly(D-lactide) (PDLA), poly(D-lactic acid), poly(L-lactide) (PLLA), poly(L -lactic acid), poly(L-lactide-co-D,L-lactide) (PLDLLA), poly(acid anhydride) (PAA), poly(cyanoacrylate) (PC A), poly(glycolide) (PGA), poly (gly colic acid), poly(glycolideco-lactide) (PGALA), poly(glycolic acid-co-lactic acid), poly(orthoester) (POE) and any mixture thereof.

Preferably, the biodegradable polymer does not produce acidic by-products during degradation. Advantageously, this may assist in bone regeneration. The biodegradable polyester may be polycaprolactone (PCL).

The ceramic may be replaced by an inorganic salt, in solid form.

The cation of the inorganic salt may be a Group 1 or Group 2 element of the Periodic Table of Elements. The cation of the inorganic salt may be selected from the group consisting of Na ⁺ , K ⁺ , Li ⁺ , Rb ⁺ , Cs ⁺ , Be ²⁺ , Mg ²⁺ , Ca ²⁺ , Sr ²⁺ , Ba ²⁺ and any mixture thereof. The cation of the inorganic salt may be Ca ²⁺ .

The cation of the inorganic salt may also be selected from the group consisting of Fe ²⁺ , Fe ³⁺ , Cr ²⁺ , Cr ³⁺ and any mixture thereof.

The anion of the inorganic salt may be selected from the group consisting of F , CI , Br-, Γ, O ² , S ² , OH , SO ₄ ² , NO ₃ ² , PO ₄ ³ , SO ₃ ² , PO ³ , PO ₃ ³ , HPO3 ² , HP(0) ₂ OH , H ₂ P ₂ O ₅ ² , H ₂ PO ₂ , and any mixture thereof. The anion of the inorganic salt may be P0 ₄ ³ . The inorganic salt may be selected from the group consisting of calcium-phosphate ceramics, hydroxyapatite, carbonated apatite, calcium phosphate, calcium-silicate based bioactive glass, anorgamnic bone and dental tooth enamel, monetite, brushite, calcium sulfate, coral, zeolite, silicate, Montmorillonite, hectorite, saponite, clays, alumina ceramic, zirconia ceramic, and any mixture thereof.

The inorganic salt may be tricalcium phosphate (Ca ₃ (P0 ₄ ) ₂ ) (TCP). The inorganic salt may be hydroxyapatite. The tricalcium phosphate may be beta-tricalcium phosphate.

The composite material may comprise an active ingredient.

The active ingredient may be a pharmaceutically active ingredient. The pharmaceutically active ingredient may have a molecular weight greater than 1 kDa, 2kDa, 5 kDa, 10 kDa, 20 kDa, 50 kDa, 100 kDa, 200 kDa or 500 kDa. The pharmaceutically active ingredient may have a molecular weight in the range of about 1 kDa to about 2 kDa, about 1 kDa to about 5 kDa, about 1 kDa to about 10 kDa, about 1 kDa to about 20 kDa, about 1 kDa to about 50 kDa, about 1 kDa to about 100 kDa, about 1 kDa to about 200 kDa, about 1 kDa to about 500 kDa, about 2 kDa to about 5 kDa, about 2 kDa to about 10 kDa, about 2 kDa to about 20 kDa, about 2 kDa to about 50 kDa, about 2 kDa to about 100 kDa, about 2 kDa to about 200 kDa, about 2 kDa to about 500 kDa, about 5 kDa to about 10 kDa, about 5 kDa to about 20 kDa, about 5 kDa to about 50 kDa, about 5 kDa to about 100 kDa, about 5 kDa to about 200 kDa, about 5 kDa to about 200 kDa, about 10 kDa to about 20 kDa, about 10 kDa to about 50 kDa or about 20 kDa to about 50 kDa, about 20 kDa to about 100 kDa, about 20 kDa to about 200 kDa, about 20 kDa to about 500 kDa, about 50 kDa to about 100 kDa, about 50 kDa to about 200 kDa, about 50 kDa to about 500 kDa, about 100 kDa to about 200 kDa, about 100 kDa to about 500 kDa or about 200 kDa to about 500 kDa. The pharmaceutically active ingredient may have a molecular weight greater than 1 kDa. Advantageously, the composite material as defined above may facilitate controlled release of molecules having a molecular weight greater than 1 kDa more effectively than for molecules with a smaller molecular weight, as molecules with a smaller molecular weight may diffuse through the polyester matrix and thus bypass the need for an interconnected pore network. Molecules with a smaller molecular weight may therefore not be able to undergo controlled, extended release, as they may be released faster.

The pharmaceutically active ingredient may be hydrophilic.

The pharmaceutically active ingredient may be selected from the group consisting of polysaccharide, polypeptide, polynucleotide, lipid, nanoparticle, hormone, antiresorptive agent, bone -forming agent, antibiotic, anti-inflammatory drug, anticancer drug and any mixture thereof. Lipids, which are amphiphilic, may benefit from the biphasic nature of the composite material, as the composite material may have hydrophilic and hydrophobic domains which may be compatible with the amphiphilic nature of the lipid. Nanoparticles, such as clay or small ceramic particles may also be a suitable pharmaceutically active ingredient. The nanoparticle may be less than 1 μm in size to be effectively released from the composite material.

The pharmaceutically active ingredient may be selected from the group consisting of starch, cyclodextrin, cellulose, chitin, fibrin, elastin, fibroin, alpha-melanocyte-stimulating hormone, growth factor, biosphosphonate, denosumab, cathepsin inhibitor, estrogen, selective estrogen receptor modulator, parathyroid hormone peptide, strontium ranelate, calcium-sensing receptor antagonist, gentamicin, ciprofloxacin, gatifloxacine, dexamethasone, ibuprofen, cisplatin and any mixture thereof. The pharmaceutically active ingredient may be a polysaccharide. The polysaccharide may be selected from the group consisting of heparan sulfate, heparin, hyaluronic acid, keratin sulfate, chondroitin/dermatan sulfate, dextran and any mixture thereof. The polysaccharide may comprise monomers selected from the group consisting of glucuronic acid (β-D -glucuronic acid) (GlcA), N-acetylglucosamine (GlcNAc), α-L-iduronic acid (IdoA), 2-O-sulfo-α-L-iduronic acid (IdoA(2S)) 2-deoxy-2-acetamido- α-D-glucopyranosyl (GlcNAc), 2-deoxy-2-sulfamido-α- D-glucopyranosyl (GlcNS), 2-deoxy-2-sulfamido-α-D-glucopyranosyl-3-0-sulfate (GlcNS(3S)), 2-deoxy-2-sulfamido-α-D-glucopyranosyl-6-0-sulfate (GlcNS(6S)) and any mixture thereof.

The polysaccharide may be heparin or heparan sulphate.

The backbone structure of heparan sulfate may be described generically as ΔUA,+/-2S 1-4 GlcNAc/S,+/-6S. The key component disaccharides (produced by the action of bacterial heparinase on high grade porcine HS, then isolated by high resolution gel filtration and ion exchange chromatography) may be best described as follows, and will be understood by persons skilled in the art:

The pharmaceutically active ingredient, the polysaccharide, heparin or heparin sulphate may have a molecular weight in the range of about 10 kDa to about 18 kDa, about 10 kDa to about 12 kDa, about 10 kDa to about 14 kDa, about 10 kDa to about 16 kDa, about 12 kDa to about 14 kDa, about 12 kDa to about 16 kDa, about 12 kDa to about 18 kDa, about 14 kDa to about 16 kDa, about 14 kDa to about 18 kDa or about 16 kDa to about 18 kDa.

The porogen may be at least partially water soluble.

The porogen may be selected from the group consisting of sodium chloride, paraffin beads, poly(ethylene glycol) (PEG), poly(ethylene oxide) (PEO), polysaccharides, poly oxazolines, poly acrylates, poly acrylamides, synthetic polymers, and any mixture thereof. The porogen may be poly(ethylene glycol) (PEG). Preferably, the porogen may be soluble in an aqueous solution. The porogen may advantageously be leached out of the composite material over time to form pores. The synthetic polymer may be copolymers and block copolymers.

The porogen may be a block copolymer including diblock copolymer and triblock copolymer. The copolymer may comprise PEG and/or PEO, and may also comprise hydrophilic and/or hydrophobic moieties.

Polysaccharides that may be used as a porogen may be selected from the group consisting of chondroitin sulfate, hyaluronic acid, sucrose, fructose and any mixture thereof.

Synthetic polymers that may be used as a porogen may be selected from the group consisting of poly (oxazolines) and, poly (acrylates), poly (acrylamides) and any mixture thereof.

The biodegradable polyester and porogen may be present in a ratio in the range of about 15:1 to about 2:1, about 15:1 to about 12:1, about 15:1, to about 9:1, about 15:1, to about 6:1, about 15:1 to about 4:1, about 15:1 to about 2:1, about 12:1, to about 9:1, about 12:1, to about 6:1, about 12:1, to about 4:1, about 12:1 to about 2:1, about 9:1 to about 6:1, about 9:1, to about 4:1, about 9:1, to about 2:1, about 6:1 to about 4:1, about 6:1, to about 2:1, about 4:1, to about 2:1. The biodegradable polyester and porogen may be present in a ratio in the range of about 9:1, to about 4:1, by weight. The biodegradable polyester and ceramic may be present in a ratio in the range of about 3:1 to about 4:1, about 3:1 to about 3.2:1, about 3:1 to about 3.4:1, about 3:1 to about 3.6:1, about 3:1 to about 3.8:1, about 3.2:1 to about 3.4:1, about 3.2:1 to about 3.6:1, about 3.2:1 to about 3.8:1, about 3.2:1 to about 4:1, about 3.4:1 to about 3.6:1, about 3.4:1 to about 3.8:1, about 3.4:1 to about 4:1, about 3.6:1 to about 3.8:1, about 3.6:1 to about 4:1 or about 3.8:1 to about 4:1, by weight.

The biodegradable polyester, porogen and ceramic may be present in a ratio in the range of about 72:8:20 to about 64:16:20, by weight.

When the active ingredient is a polysaccharide, the ratio of biodegradable polyester, porogen and ceramic in the range of about 72:8:20 to about 64:16:20 may achieve the highest sustained release of the active ingredient. The active ingredient may be present in amount in the range of about 0.1 wt% to about 90 wt%, about 0.1 wt% to about 0.2 wt%, about 0.1 wt% to about 0.5 wt%, about 0.1 wt% to about 1 wt%, about 0.1 wt% to about 2 wt%, about 0.1 wt% to about 5 wt%, about 0.1 wt% to about 10 wt%, about 0.1 wt% to about 20 wt%, about 0.1 wt% to about 50 wt%, about 0.2 wt% to about 0.5 wt%, about 0.2 wt% to about 1 wt%, about 0.2 wt% to about 2%, about 0.2 wt% to about 5%, about 0.2 wt% to about 10 wt%, about 0.2 wt% to about 20 wt%, about 0.2 wt% to about 50 wt%, about 0.2 wt% to about 90 wt%, about 0.5 wt% to about 1%, about 0.5 wt% to about 2%, about 0.5 wt% to about 5 wt%, about 0.5 wt% to about 10 wt%, about 0.5 wt% to about 20 wt%, about 0.5 wt% to about 50 wt%, about 0.5 wt% to about 90 wt%, about 1 wt% to about 2 wt%, about 1 wt% to about 5 wt%, about 1 wt% to about 10 wt%, about 1 wt% to about 20 wt%, about 1 wt% to about 50 wt%, about 1 wt% to about 90 wt%, about 2 wt% to about 5 wt%, about 2 wt% to about 10 wt%, about 2 wt% to about 20 wt%, about 2 wt% to about 50 wt%, about 2 wt% to about 90 wt%, about 5 wt% to about 10 wt%, about 5 wt% to about 20 wt%, about 5 wt% to about 50 wt%, about 5 wt% to about 90 wt%, about 10 wt% to about 20 wt%, about 10 wt% to about 50 wt%, about 10 wt% to about 90 wt%, about 20 wt% to about 50 wt%, about 20 wt% to about 90 wt%, or about 50 wt% to about 90 wt%.

The composite material may comprise pores having a size in the range of about 1 μm to about 10 urn, about 1 μm to about 2 μm, about 1 μm to about 5 μm, about 2 μm to about 5 μm, about 2 μm to about 10 μm, or about 5 μm to about 10 μm. The pores may be small and uniformly interconnected to facilitate the controlled release of the active ingredient.

The composite material may comprise a pharmaceutically acceptable excipient. The pharmaceutically acceptable excipient may be selected from the group consisting of disintegrators, binders, fillers, lubricants, glidants and any mixture thereof. The pharmaceutically acceptable excipient may enhance the properties of the composite material. Disintegrants or disintegrants, as used herein, may refer to one or more of agar-agar, algins, calcium carbonate, carboxmethylcellulose, cellulose, clays, colloid silicon dioxide, croscarmellose sodium, crospovidone, gums, magnesium aluminium silicate, methylcellulose, polacrilin potassium, sodium alginate, low substituted hydroxypropylcellulose, and cross-linked polyvinylpyrrolidone hydroxypropylcellulose, sodium starch glycolate, and starch. Binders, as used herein, may refer to one or more of microcrystalline cellulose, hydroxymethyl cellulose, hydroxypropylcellulose, and polyvinylpyrrolidone.

Fillers, as used herein, may refer to one or more of calcium carbonate, calcium phosphate, dibasic calcium phosphate, tribasic calcium sulfate, calcium carboxymethylcellulose, cellulose, dextrin derivatives, dextrin, dextrose, fructose, lactitol, lactose, magnesium carbonate, magnesium oxide, maltitol, maltodextrins, maltose, sorbitol, starch, sucrose, sugar, and xylitol.

Lubricants, as used herein, may refer to one or more of agar, calcium stearate, ethyl oleate, ethyl laureate, glycerin, glyceryl palmitostearate, hydrogenated vegetable oil, magnesium oxide, magnesium stearate, mannitol, poloxamer, glycols, sodium benzoate, sodium lauryl sulfate, sodium stearyl, sorbitol, stearic acid, talc, and zinc stearate.

Glidant, as used herein, may refer to one or more of magnesium stearate, colloidal silicon dioxide, starch and talc. A method for fabricating a composite material for delivering an active ingredient as defined above, may comprise the steps of: contacting a biodegradable polyester, a ceramic, and a porogen to form a mixture; pulverizing the mixture; sintering the pulverized mixture in a mold to form the composite material; and removing the porogen.

The contacting step may further comprise an active ingredient. Advantageously, incorporating the active ingredient at the contacting step may be more time and cost-efficient, as an excessive amount of the active ingredient may not be required to facilitate the release from the composite material.

The active ingredient may be added to the composite material after the porogen removing step.

Advantageously, for the active ingredient to be added after the sintering step, the water-soluble porogen may be removed, thus facilitating access of the active ingredient to the core of the composite material. The active ingredient may be introduced in the form of an aqueous solution or a quasi-solid medium that may allow the diffusion of the active ingredient into the composite material. Since active ingredients such as polypeptides may be denatured during material processing, in particular when exposed to high temperatures (e.g. during sintering), their introduction after the formation of the composite material may avoid the degradation of the active ingredient.

Advantageously, the composite material as defined above may be a solid comprising the active ingredient. That is, the composite material may not comprise any liquids, and hence may exhibit substantially longer shelf life than active ingredients stored in aqueous solution. The contacting step may be done by blending, extrusion or a combination thereof. The blending and/or extrusion may be done using a single screw mixer or a twin screw mixer.

Blending may be shear blending. Shear blending may be used to fabricate the composite material (from pellets, powder or a combination thereof of the starting material.

The composite material may be extruded into strands and also deposited by fused deposition modelling (3D printing) or deposited in a tubular form, such as part of a sheath/core cylinder.

The material may also be sectioned into smaller particles, from sub -millimeter to millimeter or several millimeter dimensions.

The pulverizing step may be done by milling. The milling may be cryomilling. The cryomilling may be performed at a vibrational frequency in the range of about 15 Hz to about 20 Hz, about 15 Hz to about 17.5 Hz or about 17.5 Hz to about 20 Hz. Advantageously, cryomilling may provide micro-sized or nano-sized polyester particles, which in turn reduces melt viscosity during subsequent a heating / sintering / molding step. Advantageously, this may achieve a substantially uniform distribution of the ceramic particles within the polyester matrix.

The cryomilling may be performed for a duration in the range of about 25 minutes to about 40 minutes, about 25 minutes to about 30 minutes, about 25 minutes to about 35 minutes, about 30 minutes to about 35 minutes, about 30 minutes to about 40 minutes, or about 35 minutes to about 40 minutes.

The cryomilling may be preceded by a pre -cooling at a vibrational frequency in the range of about 3 Hz to about 7 Hz, about 3 Hz to about 5 Hz or about 5 Hz to about 7 Hz. The pre-cooling may be performed for a duration in the range of 2 minutes to about 7 minutes, about 2 minutes to about 5 minutes or about 5 minutes to about 7 minutes.

The cryomilling may facilitate the grinding of the biodegradable polyester, ceramic, and porogen into a finely ground substantially homogeneous powder by cooling with a cryogen. By using a cryogen such as liquid nitrogen or liquid argon, milling of the mixture into a fine powder is possible, which is normally not possible at higher temperatures due to the low glass transition temperature/melting point of the components causing the material to soften.

The powder fragments following cryomilling may have a size in the range of about 100 μm to about 2 mm, about 100 μm to about 200 μm, about 100 μm to about 300 μm, about 100 μm to about 400 μm, about 100 μm to about 500 μm, about 100 μm to about 1 mm, about 100 μm to about 2 mm, about 100 μm to about 5 mm, about 200 μm to about 300 μm, about 200 μm to about 400 μm, about 200 μm to about 500 μm, about 200 μm to about 1 mm, about 200 μm to about 2 mm, about 200 μm to about 5 mm, about 300 μm to about 400 μm, about 300 μm to about 500 μm, about 300 μm to about 1 mm, about 300 μm to about 2 mm, about 300 μm to about 5 mm, about 400 μm to about 500 μm, about 400 μm to about 1 mm, or about 500 μm to about 1 mm, about 400 μm to about 2 mm, about 400 μm to about 5 mm, about 500 μm to about 1 mm, about 500 μm to about 2 mm, about 500 μm to about 5 mm, about 1 mm to about 2 mm, about 1 mm to about 5 mm, or about 2 mm to about 5 mm.

The powder fragments of the biodegradable polyester or polycaprolactone (PCL) following cryomilling may have a size in the range of about 100 μm to about 500 μm, about 100 μm to about 200 μm, about 100 μm to about 300 μm, about 100 μm to about 400 μm, about 200 μm to about 300 μm, about 200 μm to about 400 μm, about 200 μm to about 500 μm, about 300 μm to about 400 μm, about 300 μm to about 500 μm or about 400 μm to about 500 μm.

The powder fragments of the ceramic or tricalcium phosphate (TCP) following cryomilling may have a size in the range of about 100 μm to about2 mm, about 100 μm to about 200 μm, about 100 μm to about 300 μm, about 100 μm to about 400 μm, about 100 μm to about 500 μm, about 100 μm to about 1 mm, about 100 μm to about 2 mm, about 100 μm to about 5 mm, about 200 μm to about 300 μm, about 200 μm to about 400 μm, about 200 μm to about 500 μm, about 200 μm to about 1 mm, about 200 μm to about 2 mm, about 200 μm to about 5 mm, about 300 μm to about 400 μm, about 300 μm to about 500 μm, about 300 μm to about 1 mm, about 300 μm to about 2 mm, about 300 μm to about 5 mm, about 400 μm to about 500 μm, about 400 μm to about 1 mm, or about 500 μm to about 1 mm, about 400 μm to about 2 mm, about 400 μm to about 5 mm, about 500 μm to about 1 mm, about 500 μm to about 2 mm, about 500 μm to about 5 mm, about 1 mm to about 2 mm, about 1 mm to about 5 mm, or about 2 mm to about 5 mm. The powder fragments of the porogen or poly(ethylene glycol) (PEG) following cryomilling may have a size in the range of about 100 μm to about2 mm, about 100 μm to about 200 μm, about 100 μm to about 300 μm, about 100 μm to about 400 μm, about 100 μm to about 500 μm, about 100 μm to about 1 mm, about 100 μm to about 2 mm, about 100 μm to about 5 mm, about 200 μm to about 300 μm, about 200 μm to about 400 μm, about 200 μm to about 500 μm, about 200 μm to about 1 mm, about 200 μm to about 2 mm, about 200 μm to about 5 mm, about 300 μm to about 400 μm, about 300 μm to about 500 μm, about 300 μm to about 1 mm, about 300 μm to about 2 mm, about 300 μm to about 5 mm, about 400 μm to about 500 μm, about 400 μm to about 1 mm, or about 500 μm to about 1 mm, about 400 μm to about 2 mm, about 400 μm to about 5 mm, about 500 μm to about 1 mm, about 500 μm to about 2 mm, about 500 μm to about 5 mm, about 1 mm to about 2 mm, about 1 mm to about 5 mm, or about 2 mm to about 5 mm.

The sintering step may be hot pressing. The hot pressing step may be performed at a hot pressing temperature in the range of about 75 °C to about 115 °C, about 75 °C to about 90 °C, or about 90 °C to about 115 °C. The hot pressing step may be preceded by pre -heating the pulverized mixture for a duration in the range of about 4 minutes to 8 minutes, about 4 minutes to about 6 minutes or about 6 minutes to about 8 minutes at the hot pressing temperature.

The hot pressing step may be performed a pressure in the range of about 3000 psi to about 6000 psi, about 3000 psi to about 4500 psi, or about 4500 psi to about 6000 psi. The pressure may be applied for a duration in the range of about 1 minute to about 4 minutes, about 1 minute to about 2 minutes, about 1 minute to about 3 minutes, about 2 minutes to about 3 minutes, about 2 minutes to about 4 minutes or about 3 minutes to about 4 minutes.

The hot pressing step may result in the solidification of the pulverized mixture. The mold may be a film mold. The composite material may be extruded into a strand, pelletized, then hot -pressed into a mold, giving it its final shape.

The hot pressing step using a film mold may result in a film of the composite material. The film may have a thickness in the range of about 100 μm to about 300 μm, about 100 μm to about 150 μm, about 100 μm to about 200 μm, about 100 μm to about 250 μm, about 150 μm to about 200 μm, about 200 μm to about 250 μm, about 200 μm to about 300 μm, or about 250 μm to about 300 μm. Advantageously, the disclosed method provides flexibility in obtaining a desired shape, geometry or size of the composite material. This may be advantageous over methods which utilize a scaffold template for forming the polyester/ceramic composite material, wherein the size and shape of the composite would necessarily be limited by the scaffold template being used.

The removing step of the porogen may be done by leaching.

The leaching may be done in aqueous solution. The aqueous solution may be water or buffer. The aqueous solution may be selected as a solvent wherein the porogen material is selectively or preferentially more soluble in, relative to the composite material. Advantageously, the leaching step may result in the formation of pores in the composite material, which may in turn facilitate the controlled release of the active ingredient from the composite material.

The method may not require the use of any sacrificial material or template. The template may be polyurethane.

The contacting step, the pulverizing step and the sintering step of the method may not require the use of a solvent.

The solvent may be an organic solvent or an aqueous solvent. The solvent may be selected from the group consisting of water, buffer, methanol, ethanol, isopropanol, acetone, hexane, ethyl acetate, acetonitrile, chloroform, dichloromethane, dimethyl sulfoxide, dimethyl formamide, diethyl ether, toluene, tetrahydrofuran and any mixture thereof.

Use of the composite material as defined above, may be for delivery of an active ingredient.

Brief Description of Drawings

The accompanying drawings illustrate a disclosed embodiment and serves to explain the principles of the disclosed embodiment. It is to be understood, however, that the drawings are designed for purposes of illustration only, and not as a definition of the limits of the invention.

Fig.l

[Fig. 1] refers to SEM images of powdered components, (a) PCL powder, with (b) the same material at higher magnification; (c) and (d) similarly depict PCL powder in the presence of TCP; (e) cross-section of heat-pressed PCL TCP sample; (f) fluorescent optical microscopy of heat-pressed PCL/TCP composite, loaded with A488-heparin.

Fig.2

[Fig. 2] refers to images depicting heparin fluorescence, showing its uniform distribution within the material. Fig. 2 shows a fluorescence image, on the left, and an optical microscopy image, on the right, of (a) a composite film containing heparin and (b) a composite film containing heparin labeled with A-488 fluorophore.

Fig.3

[Fig. 3] refers to graphs showing the characterization of (a) heparin extracted from material by FTIR and (b) bioactivity of heparin released from scaffold material.

Fig. 4 [Fig. 4] refers to graphs showing heparin release from (a) PCL or (b) PCL/TCP films.

Fig. 5

[Fig. 5] refers to graphs showing (a) PEG and (b) TCP effect on the heparin release from PCL matrix respectively. Heparin loading, at 1% for all samples, is not included in the percentage composition. Fig. 6

[Fig. 6] refers to a graph showing film mass loss over time.

Fig. 7

[Fig. 7] refers to images showing heparin adsorbed onto TCP particles: (a) heparin solution concentration versus incubation time with 4 mg TCP particles, and (b) adsorbed heparin amount versus TCP mass following 3hours' incubation, fluorescence image of TCP particles after absorbing (c) as-received heparin and (d) A488-heparin. Inset in parts (c) and (d) (cd) are from white light microscopy.

Fig. 8

[Fig. 8] refers to an XRD spectrum of: (a) PEG, (b) PCL, (c) TCP, (d) PCL/PEG (80/20), and (e) PCL/PEG/TCP (64/16/20). Fig. 9

[Fig. 9] refers to images showing the surface and cross-section morphology of films: (a), (b), (c) and (d) are PCL/PEG (80/20) films, while (e), (f), (g) and (h) are PCL/PEG/TCP (64/16/20) films. Fig. 10

[Fig. 10] is a graph showing the tensile modulus of the composite materials.

Fig. 11

[Fig. 11] is a graph showing the in vitro activity of released heparin.

Examples

Non-limiting examples of the invention and a comparative example will be further described in greater detail by reference to specific Examples, which should not be construed as in any way limiting the scope of the invention.

Example 1: Materials and Methods

Materials

Medical grade polycaprolactone (PCL), with molecular weight (MW) -70 kDa and β-tricalcium phosphate (TCP) were purchased from Osteopore International Pte (Singapore). Alexa fluor® 488 hydrazide (A488), penicillin/streptomycin and p-nitrophenyl phosphate substrate were sourced from Life Technologies. Polyethylene glycol (PEG, with MW -10 kDa), sodium chloride (NaCl, >99.5%), l-Ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC), 4- morpholinoethanesulfonic acid (MES), heparin sodium salt from porcine intestinal mucosa, phosphate buffered saline (xlO) (PBS), toluidine blue O (TBO), ethylenediaminetetraacetic acid (EDTA), Igepal and NP40 were purchased from Sigma-Aldrich. Dichloromethane (DCM) was sourced from J.T.Baker. Dulbecco's Modified Eagle's Medium (DMEM) and fetal bovine serum (FBS) were purchased from Hyclone laboratories. Tris(hydroxymethyl)aminome thane (Tris) was purchased from Invitrogen. SDS was purchased from BDH laboratories. Triton X- 100 was purchased from Bio-rad. Protease inhibitor mixture was purchased from Calbiochem. Bicinchoninic acid (BCA) Protein Assay Kit was purchased from Pierce. P-nitrophenyl phosphate substrate was purchased from Zymed Laboratories Inc. Calf intestinal alkaline phosphatase (CIAP) was purchased from New England Biolabs. Fluorescent labeling of heparin

Heparin was labeled with A488, by EDC cross-linking. Briefly, 10% (w/v) EDC solution in 0.1 M MES buffer (50 μΐ, pH 4.5) was added to 300 μΐ of 1% (w/v) heparin in 0.1 M MES buffer. Following incubation at room temperature for 20 min, 50 μΐ of 1% (w/v) A488 solution in MES buffer was added to the heparin/EDC solution. The reaction tube, protected from light, was incubated overnight at room temperature. A488 labeled heparin (A488 -heparin) was purified through an Amersham PD10 desalting column and then freeze dried. The A488 -heparin distribution within the processed material was probed by fluorescence microscopy.

Fabrication of composite material A Retsch Cryomill was used to powder material components and generate their mixtures, before hot pressing. A 50 ml hardened steel grinding jar, containing two hardened 15 mm steel balls and the materials to be milled, was pre-cooled at 5 Hz for 5 min. This is followed by 35 min of milling at 17.5 Hz. The resulting powder mixture was subsequently heat-pressed into a mold, preheating it at 90°C for 6 min, then applying increasing pressure over 1.5 min and maintaining at 4500 psi for 2 min. A film mold was used to generate -200 μm thickness films. Heat-pressed samples were cooled in air and subsequently removed from the mold.

For heparin, the highest sustained release performance was obtained from the following material compositions: polycaprolacton (PCL) / poly(ethyleneglycol) (PEG) / tricalcium phosphate (TCP) ratios of 72:8:20 and 64: 16:20 by weight. Fourier transform infrared spectroscopy (FTIR)

Approximately 20 mg of PCL composite were dissolved in 10 ml DCM, followed by centrifuging at 10000 rpm for 10 min, to allow decanting of dissolved PCL. TCP and heparin precipitates were dispersed into 10 ml water, followed by shaking at 100 rpm for 10 min to ensure dissolution of heparin in water. This was followed by centrifuging at 10000 rpm, for 10 min, to precipitate TCP particles. Heparin was then lyophilized from the aqueous phase. Approximately 2 mg of lyophilized heparin was ground in a mortar with potassium bromide (KBr, Uvasol ^® Merck), yielding a -1% concentration, which was then compressed at -10 tons to obtain a transparent disc. The disc was probed with FTIR spectrometer (Perkin Elmer Spectrum 2000), running 8 scans, each with a resolution of 4 cm ^-1 , over a spectral range from 4000 to 400 cm ¹ .

Morphology study

A488-heparin and TCP distributions within the composite material were imaged by fluorescence microscopy (Olympus BX51) and scanning electron microscopy (SEM) (JEOL JSM-6360LA), respectively. For the latter, samples were fixed onto an aluminum support with conductive adhesive tape, and coated with gold for 30 seconds, using an SPI-Modulate™ sputter coater, to improve their electrical conductivity. SEM at lOkV was used to image the samples' morphology. In vitro release

Films, cut into 1 cm square samples, were used to quantify heparin release into PBS (pH 7.4). Individual samples were immersed in 4 ml of PBS, in a 5 ml amber glass bottle, placed in a 37°C incubator (Labnet 21 IDS) and shaken at 100 rpm. At predetermined time intervals, 0.4 ml aliquots of the release medium were removed to assay released heparin. The same volume of fresh buffer was replaced, thus maintaining a constant total volume. Heparin concentration in the released medium was measured by TBO staining. Briefly, 0.1 ml 1% (w/v) NaCl solution was added to 0.4 ml of released medium, supplemented with 0.5 ml of 0.005% (w/v) TBO, dissolved in 0.01 M aqueous HCl/0.2% (w/v) NaCl solution. This solution was shaken for 10 min and subsequently centrifuged at 12000 rpm for 10 min to separate the supernatant solution from the precipitates. The absorbance of 200 ml of the supernatant solution, placed in a 96-well plate, was measured at 631 nm, using a Spectra Max Plus 384 spectrophotometer. Calibration curves were obtained from known heparin concentrations ranging from 1 to 50 μg/ml in PBS, with fresh PBS used as a reference. Cumulative heparin release was calculated from these calibration curves.

X-ray diffraction

Phase structures of PCL, PEG, PCL/PEG and PCL/PEG/TCP were probed using an X-ray diffractometer (XRD, Brukers D8 GADDS), using a scanning speed of 3°/min to determine patterns over 2Θ angles ranging from 2.5° to 33.5°. Heparin adsorption onto TCP particles

Heparin was adsorbed from 1 ml of 40μg/ml solution onto 0.5-5 mg of TCP particles. The mixed solution was shaken for predetermined time to figure out sufficient adsorption time. Centrifugation at 12000 rpm for 10 min, was used to precipitate the TCP particles, followed by withdrawal from the supernatant of 0.4ml aliquots to quantify the heparin concentration by TBO staining, as described above. Depletion of the heparin solution concentration was attributed to its adsorption onto TCP. The latter was confirmed by fluorescence microscopy imaging of A488-heparin, remaining adsorbed onto TCP particles, after thrice rinsing in DI water.

Mechanical properties of the PCL composite

Tensile properties of composites were measured, using an Instron Model 5569 universal tester. 200 μm thickness dogbone-type films (3.14 mm wide, with 12 mm length) were fabricated for these tensile measurements. Tensile loading was applied at a constant 1 mm min, with a 100N load cell.

In vitro bioactivity of heparin released from the scaffold

Bioactivity of released heparin was determined by measuring its ability to induce alkaline phosphatase (ALP) in C2C12 cells, in the presence and absence of BMP -2. A medium with released heparin was obtained by incubating a composite rod loaded with 33μg of heparin in 250μ1 of PBS at 37°C for 7 days. The rod released an estimated 40% of its incorporated heparin, that is 13^g in 250 μΐ PBS, which was then diluted with the culture media to obtain the heparin concentration of 10.6 μg/ml for the ALP activity assay.

Cells were cultured in DMEM, containing 1000 mg/L of glucose, 10% FBS, and 1% penicillin/streptomycin. 40,000 cells were seeded into 24-well plates and incubated for 24 hours at 37°C. Following medium replacement with fresh medium containing 5% FBS, the cells were exposed to the medium containing released heparin (10.6 μg/ml), in the presence and absence of BMP-2 (100 ng/ml), for a period of 3 days. The basal ALP activity of C2C12 cells was assessed with control cultures: PBS as a negative control, BMP-2, and BMP-2 with as received heparin (10 μg/ml) as a positive control. After 3 days, the cells were lysed in radio-immunoprecipitation assay (RIPA) buffer (150 mM NaCl, 10 mM Tris, pH 7.4, 2 mM EDTA, 0.5% Igepal/NP40, 0.1% SDS, 1% Triton X-100, protease inhibitor mixture), and protein concentration was determined using BCA Protein Assay Kit. ALP activity was determined using equal amounts of total protein. Briefly, proteins (10 mg) were incubated with assay buffer (40 mL) including p-nitrophenyl phosphate substrate at 37°C for 1 h. CIAP was used as positive control. The absorbance at 405 nm was measured using a Victor ³ 1420 Multilabel Counter (PerkinElmer Life Sciences).

Example 2: Characterization of composite material

Conventionally, composite materials loaded with drug molecules are generally blended by either solvent casting or melt extrusion. The former involves organic solvents, engendering residues that may limit their therapeutic applications. Blending and extrusion equipment, used for the latter, typically requires relatively large sample size, in the tens or hundreds of grams. A twin screw extruder may also be used. Some twin screw extruders are capable of processing small sample sizes, of only a few grams, with adequate mixing efficiency. The high cost of bioactive ingredients, such as specific heparan sulfate that potentiates endogenous growth factors, may necessitate preparing smaller quantities of composite materials. The present study describes a solvent-free method of cryomilling followed by heat pressing, for preparing relatively small samples size of the composite material, in the order of a gram or less. Cryomilling can pulverise polymers and can also serve to mix the polymer with other components. Heat-pressing the resulting mixture into a mould allows the fabrication of laboratory-scale samples. Samples of larger sizes, such as tissue engineering scaffolds, may be prepared using fused deposition modelling (FDM) or 3D printing. These methods would either use strands, output from an extrusion (single or twin screw), or pellets fabricated by extrusion or heat pressing. FDM and 3D printing are capable of generating complex, three-dimensional architectures, including macroporous structures that are suitable for supporting the culture of anchorage -dependent cells and/or in tissue regeneration applications. Note that a controlled release device has the capability to deliver bioactive agents that regulate cell proliferation and evolution, by making advantageous use of the contact between cultured cells and the device, used as a cell culture substrate. Distribution

The uniform distribution of TCP particles and heparin within PCL matrix is expected to play a critical role, impacting the stability of the material and having a consequential influence on its release profile, therefore the distribution of TCP particles and heparin in the scaffold material were studied by SEM and fluorescence microscopy. Fig. 1 (a) and (b) depicts the PCL powder fragments following cryomilling, having a size range of 100-500 μm. Fig. 1 (c) shows PCL fragments cryomilled in the presence of TCP having a slightly different morphology. The surface of these fragments, imaged with higher magnification in Fig. 1 (d), appears to show features in the order of microns that do not appear in Fig. 1 (b), implying that they are likely to be TCP particles attached to the polymer surface. This would confirm the intimate mixing between the TCP particles within the PCL matrix which contributes to their uniform distribution. Fig. 1 (e) depicts the cross-section of a heat -pressed composite, showing the even distribution of TCP particles within the material.

The PCL and TCP component ratio, within the composite material, was quantified by thermogravimetric analysis (TGA). The results show that there is a mass reduction of approximately 79% in the composite material above 500 °C, leaving a residue of approximately 21%, which corresponds closely to the initial mass ratio for PCL to TCP of 80% to 20%, respectively.

Heparin labelled with A488 and incorporated within the scaffold material was detected by fluorescence microscopy, as shown in Fig. 1 (f). A high level of uniform fluorescence is observed for the scaffold material (102), in contrast to the absence of signal from a similar composite loaded with as-received (unlabelled) heparin (104). This indicates that heparin is reasonably uniformly distributed within the scaffold material.

Fig. 2 (a) and (b) depicts heparin fluorescence, showing its uniform distribution within the material. Fig. 2 (a) shows a fluorescence image, on the left, and an optical microscopy image, on the right, of a composite film containing heparin. The weak fluorescence signal (left hand image in Fig. 2 (a)) arises from auto-fluorescence of the film material. Fig. 2 (b) shows a fluorescence image, on the left, and an optical microscopy image, on the right, of a film containing heparin labeled with A-488 fluorophore. These images demonstrate the uniform distribution of fluorescent heparin within the composite film.

Molecular Structure

Given the potential for scission of the polymer molecular chains during the cryomilling process, the molecular weight (MW) of PCL after cryomilling was measured by gel permeation chromatography (GPC), as shown in Fig. 3 (a). The results reveal a minor reduction in the average PCL MW, thus indicating that modest chain scission leaves the polymer chain largely unaffected by cryomilling. Since the molecular structure of heparin plays a critical role in conferring its bioactivity, FTIR offers a preliminary assessment of how material processing may influence its molecular structure. FTIR spectrum for heparin, were usually presented between 2000 cm ¹ and 600 cm ^-1 . Fig. 3 (b) shows the spectrum generated by the as-received and heparin extracted from the composite material, exhibiting characteristic peaks at 1235 and 1045 cm ^-1 , which are related to the sulfonation of disaccharide units. The bands in the region of 750-950 cm ¹ are generated by sulphate half ester groups. Features at 1420 and 1630 cm ¹ arise from the weak symmetric and asymmetric stretching of COO- moieties. The spectra of the as-received and extracted heparin exhibit similar features, implying that the material fabrication process does not affect the functional moieties. However, this probe does not reveal the overall disaccharide sequence, which regulates its group-specific antigen binding to proteins (e.g. BMP-2).

Example 3: Heparin release from PCL or PCL/TCP composite films

The delivery of drug molecules is generally regulated by their diffusion and the degradation of the polymer matrix. The hydrophobic nature of PCL and its slow degradation, over several months, enables the long-term, sustained release of small molecules, as exemplified by the Capronor implantable drug delivery system for the sustained administration of hormonal contraceptives. This device relies on the hydrophobicity of these molecules and their diffusion within matrix. Hydrophilic drug molecules, such as gentamicin sulfate (Mw 1.5 kDa), exhibit a different release profile, either delivered within a short time frame or remaining entrapped within the PCL matrix, until it degrades.

The release of heparin from either PCL or PCL/TCP was quantified, as shown in Fig. 4. Fig. 4 (a) shows that pure PCL has a short-term (burst) release, followed by minimal sustained release. With moderately higher loading fractions such as at 5 and 10%, the delivery efficiency increases significantly, to 50% and over 90%, respectively. The small fraction of the incorporated heparin that is released even at low loading fractions may be due to heparin being released from the composite material surface. In contrast, high delivery efficiencies imply the release of heparin from the bulk material, presumably through an interconnected pore network, that is generated by its dissolution of PCL. Despite higher delivery efficiency with higher loading, pure PCL appears unsuitable for the sustained release of heparin delivery over a period of time. Although 2% heparin loading in pure PCL offered the compromise of a modest sustained release, it showed low efficiency (<l/3, after 25 days in PBS), which rendered it unsuitable for the delivery of high added value bioactive agents.

Heparin is a comparatively large hydrophilic molecule, which exhibits a similar burst release from pure PCL, followed by negligible sustained release, or over time-scales release far longer than therapeutically-relevant time scales for tissue engineering. At low loading, poor delivery efficiency may be associated with heparin release from the surface region of the material, while high loading delivers most of the incorporated heparin implying release from the bulk material. The latter presumably occurs through a network of interconnected channels, generated by aqueous dissolution of the heparin, acting as a porogen Heparin release from PCL/TCP films exhibits similar trends as from pure PCL, but with higher delivery efficiency, as shown in Fig. 4 (b). Tricalcium phosphate (TCP) is a hydrophilic and water-insoluble particle. This biocompatible, osteoconductive ceramic has been incorporation into PCL matrix to enhance its mechanical strength, for bone regeneration applications. Low heparin loading of 1% results in low delivery efficiency, while 5% loading offers a delivery efficiency approaching 85%. This is a substantial improvement over the 50% delivery observed from 5% loading in pure PCL. The addition of TCP particles, uniformly distributed within the PCL matrix, therefore facilitates the release of heparin, resulting in greater delivery efficiency. Similarly to PCL, 3% loading of heparin in PCL/TCP shows a modest sustained release, but also with a similarly low delivery efficiency, of 45% over 40 days. These data imply that TCP particles within the PCL matrix may contribute to the formation of pathways for diffusion, thus offering an improvement in heparin release efficiency.

While the improvement may be ascribed to the more hydrophilic nature of matrix created by the hydrophilicity of TCP, the sustained release of heparin is not substantially improved and its delivery efficiency remains poor at low loading. This is attributed to the insolubility of TCP in aqueous medium, which is not directly contributing to the formation of a network of interconnected channels and improving the heparin delivery.

Although the delivery efficiency of heparin is enhanced by adding TCP particles into the PCL matrix, its sustained release remains limited, especially for low heparin loading of 1 %. It was therefore necessary to create more channels or pores within the PCL matrix for heparin diffusion, to achieve an enhanced sustained release.

Example 4: Heparin release from PCL/PEG/TCP composite

PEG has been used to tune the release of drug molecules from hydrophobic matrix. As a water- soluble polymer, it offers hydrophilic channels for molecular diffusion. The influence of this additive on the release profiles of heparin from PCL/TCP composites was explored. As shown in Fig. 5, introducing PEG tunes the sustained release of heparin molecules and enhances delivery efficiency in a dose-dependent manner. Loaded at 8%, this additive modifies the release profile, giving rise to a sustained release of heparin over 2 months, with an enhanced efficiency of 50%. Higher PEG loading fractions transitioned into mainly burst release, with higher heparin delivery efficiency.

Although PEG addition contributes to creating channels for heparin diffusion, the addition of PEG to pure PCL is insufficient to generate a sustained, long-term release of heparin, as shown in Fig. 5 (b). There is limited sustained release from a PCL/PEG composite after an initial short term burst release. Without being bound to theory, TCP particles may serve to nucleate PEG domains to form small and evenly dispersed domains within the PCL matrix, which may enable tuning of the sustained release profile of heparin, over several weeks. Weight loss data, from exposure to aqueous medium, are shown in Fig. 6. Low solubility of TCP particles (0.02 mg/ml) is not expected to contribute to weight loss, thus attributing the film mass loss to PEG dissolution. The data shows that most of the additive elutes within the first day, with negligible additional weight loss over the subsequent months. Almost all of the incorporated PEG was removed from the PCL/PEG/TCP composite after 6 months, as compared with 75% removed from PCL/PEG composite. This suggests an enhanced interconnected pore structure in the PCL/PEG/TCP composite that enables a tuneable, sustained release profile of heparin.

The introduction of soluble PEG additive to PCL/TCP composite gives rise to a sustained, long- term heparin release, with enhanced delivery efficiency. This contrasts with the addition of PEG to PCL, which results in lower delivery efficiency and low sustained release, beyond two weeks. While aqueous dissolution of PEG may contribute to generating an interconnected porous network, TCP is postulated, without being bound to theory, to play a role in regulating the evolution of this structure. During heat pressing, TCP may act as a solid, hydrophilic intermediary, thus nucleating the formation of individual PEG domains within the PCL matrix. In the absence of TCP particles, phase -separation between PCL and PEG may induce the aggregation of PEG domains, which could be confirmed by the larger pores formed in the film after dissolution of PEG domains in aqueous medium. Thus, it appears that TCP stabilizes PEG domains, formed within the PCL matrix, altering the PEG domain distribution. Thus, TCP is believed to play a key role on the sustained release, by reducing the size of hydrophilic PEG domains and causing them to be more uniformly distributed throughout the matrix. The resulting network of interconnected pores, formed by aqueous dissolution of PEG and assisted by TCP, may enhance the delivery efficiency of hydrophilic heparin molecules. The incomplete heparin release, over a time scale approaching 2 months may be attributed to the slow degradation of PCL and the deeply buried of heparin within PCL matrix.

Prior studies have implied that the release of biomolecules from PCL/PEG is associated with PEG leaching. In the present system, most of the incorporated PEG is released within a day. Despite this rapid dissolution of PEG, heparin delivery persists for several weeks. Given that the molecular weight of heparin is comparable to that of PEG, molecular weight of heparin does not necessarily account entirely for its sustained release. Heparin adsorption onto the surface of TCP particles may be the cause of its observed long-term release. Heparin was depleted from aqueous solution when incubated with TCP particles and its surface adsorption was confirmed by fluorescence imaging of coated TCP. Most of the heparin incorporated within PCL/PEG/TCP composite material is expected to be segregated from the PEG phase, thus are expected to have ample opportunity to adsorb onto the TCP particles that nucleate these domains. A gradual heparin desorption may contribute significantly to its sustained release, over several weeks. Example 5: Adsorption of heparin on TCP particles

Given the rapid dissolution of the PEG additive from the composite material and the ensuing diffusion of heparin through the network of channels generated by PEG dissolution, a retarding mechanism was sought that could partially facilitate long-term, sustained release of heparin. Heparin adsorption on TCP was thought to retard heparin release from the composite material. The heparin would then be eventually released from the composite material via its subsequent desorption.

Fig. 7 provides evidence of heparin adsorbing onto TCP particles from an aqueous solution. Fig. 7 (a) shows the depletion of aqueous heparin solution concentration, when incubated with 4 mg of TCP particles. Fig. 7 (b) shows that this adsorption, which saturates after about 3 hours, increases linearly with the mass of TCP particles, for approximately 5 μg of heparin per milligram of TCP. Fluorescence images in Fig. 7 (c) and Fig. 7 (d) confirm the adsorption of A488-heparin onto TCP particles, which are known to have a hydrophilic surface. A subsequent release of the adsorbed heparin by desorption may contribute to its sustained release profile. A rough estimate of ~4 mg TCP in the film would adsorb -20 μg of heparin, which is -10% of its loading in the film.

EXAMPLE 6: Phase separation

Potential phase separation between hydrophobic PCL and hydrophilic PEG during heat processing might influence the material's mechanical properties. Thus, the miscibility between PCL and PEG in this system was determined using XRD.

Fig. 8 shows XRD scans of pure PCL, PEG and TCP, as well as their composites. The main features associated with PCL and PEG lie within an 18-25° angle range, where TCP does not generate any signal. PCL revealed two distinct peaks at 21.3° and 23.7°, while PEG revealed two peaks at 19.1° and 23.4°, indicating their crystalline nature. PCL/PEG composites clearly show three peaks, at 19.3°, 21.5° and 23.7°. The first and second peaks represent PEG and PCL phases, respectively. The attenuation of these features in the presence of TCP particles may reflect the formation of smaller PEG domains, nucleated by TCP particles.

Without being bound to theory but based on the hypothesis that TCP nucleates PEG domains within the PCL matrix, it was thought that it may be possible to see these smaller and more uniformly distributed domains within the structure of the composite material. SEM images of the material surface and cross -section morphologies, before and after exposure to aqueous medium, are shown in Fig. 9. These samples were heat-pressed for 40 min, to exacerbate PCL and PEG phase separation. Prior to exposure to aqueous solution, the surfaces of both PCL/PEG and PCL/TCP/PEG composites are smooth (Fig. 9 (a) and (e)). Following exposure to water, pores in the order of one micron in size were formed by PEG dissolution from the PCL/PEG composite (Fig. 9 (b)). The lack of visible pores at the surface of PEG/PCL/TCP composite film (Fig. 9 (f)) suggests that the PEG domains in the PEG/PCL/TCP composite film were of a smaller size, below the detectable resolution of these SEM images.

The film cross-section reveals significantly larger diameter pores, formed near its midplane (Fig. 9 (c), (d), (g) and (h)). After exposure to aqueous medium, pores formed in PCL/PEG/TCP are smaller and more uniformly distributed than those formed in PCL/PEG, indicating that the addition of TCP particles reduced the PEG domain size, from 10-30 μm to 2-10 μm. This reduction in domain size may improve the mechanical properties of the composite material.

Ancillary evidence of differences in phase separation is provided by the XRD data and SEM images, indicating smaller, well-distributed PEG domains in the presence of TCP, supporting the hypothesis that TCP particles may influence the composite structure. The PEG domains in the presence of TCP are in contrast to those in the absence of TCP, where the domains are larger in size and are thought to result from significant coalescence. Weight loss of the composite following aqueous dissolution indicates the complete removal of PEG from PCL/PEG/TCP composite, while only 75% of PEG is removed from PCL/PEG composite. This appears to support the theory that nucleation of PEG domains within the PCL matrix is caused by the TCP particles, which in turn generates an enhanced pore network within the composite.

EXAMPLE 6: Tensile strength Tensile strength measurements shows a higher Young's modulus for PCL/PEG/TCP than PCL/PEG, confirming that TCP loading improves the mechanical properties of the PCL/PEG composite, as shown in Fig. 10. Low PEG loading (8%) does not compromise the material's mechanical properties, with respect to PCL/TCP (80/20) composite. Higher PEG loading (16%) diminishes the material's tensile strength, which nevertheless remains higher than for a similar composite that does not contain TCP particles.

At modest PEG loading, the mechanical properties of the PCL/PEG/TCP composite are comparable to those of the PCL/TCP composite, which has been specifically developed to fabricate porous load-bearing scaffolds, for use in bone regeneration applications. The uniform distribution of PEG domains within the PCL matrix, induced by TCP particles, does not compromise the mechanical properties, as compared to composites that only containing PCL and PEG.

Example 7: In vitro activity of released heparin

The in vitro bioactivity of the released heparin was assayed by measuring the levels of alkaline phosphatase (ALP), expressed in C2C12 cells, in the presence and absence of BMP-2 (100 ng/ml). Fig. 11 depicts ALP activity, induced by BMP-2, being enhanced 10-fold in the presence of released heparin. This effect is similar to that observed when the cells are treated with as-received heparin. Hence, this assay provides a preliminary validation of heparin bioactivity, putatively arising from its unmitigated ability to interact with BMP-2, following its release from the composite material.

ALP is a well-established marker for the initial stages of osteoblast differentiation, and C2C12 myoblasts adopt an osteogenic cell fate/differentiation upon BMP-2 administration. Moreover, heparin enhances BMP-2 ability to induce osteoblast differentiation in C2C12 myoblasts in vitro. The significant enhancement in the BMP-2 induced ALP by the released heparin confirms that the bioactivity of the heparin released from the composite was not affected by previously being incorporated into the composite. .

Industrial Applicability

The composite material of the present disclosure may be useful in tissue engineering, such as bone and cartilage regeneration, or surgical patches wound/skin care applications and the implementation of localized drug release, where the release of bioactive molecules may be of therapeutic benefit.

The composite material may be used as part of a stent, where it may release bioactive agents. For example, heparin may prevent blood clotting or thrombosis. Thus, it may be a component of a therapeutic device. The therapeutic device may be an occlusive device for stopping blood flow, which may be subsequently removed to resume blood flow.

The composite material may form part of a pacemaker electrode, or prosthetic joints. Further, the composite material may also have applications in dental surgery, where bone regeneration may be essential to form a rubust base. In addition, the composite material may have applications in osteointegration of prosthetics.

It will be apparent that various other modifications and adaptations of the invention will be apparent to the person skilled in the art after reading the foregoing disclosure without departing from the spirit and scope of the invention and it is intended that all such modifications and adaptations come within the scope of the appended claims.