WO2021222974A1 - An implantable device - Google Patents

Abstract

An implantable device and methods for preparing and implanting said device into a subject for use in treating a medical condition when implanted therein is disclosed. The device comprises at least one composite fibre of a hydrophilic polymer and a hydrophobic polymer, wherein at least one of said polymers is loaded with an agent that is active towards treating the medical condition.

Description

AN IMPLANTABLE DEVICE

Technical Field

The present invention relates to composite fibres and an implantable device fabricated from said composite fibres and their use in treating a medical condition, particularly cancerous tumours.

Background of Invention

In spite of remarkable improvements in cancer treatments and survivorship, cancer still remains one of the major causes of death worldwide. Although current standards of care provide encouraging results, they still cause severe systemic toxicity and also fail in preventing recurrence of the disease.

In order to address these issues, implantable drug delivery systems (DDSs) based on biopolymers, bioceramics and nanoparticles, have emerged as promising therapeutic platlorms. The use of these materials allows local administration of drugs directly to the tumour site which significantly reduces the required drug dosage and often the side/off-target effects, and also enhances tumour treatment and could prevent cancer reoccurrence. These DDSs also enhance tumour treatment and could potentially prevent cancer reoccurrence by increasing the drug dosage at the diseased site. DDSs to be utilised in the context of pancreatic cancer in particular are desirable as it has the lowest survival rate among any solid tumour, with less than 5% cumulative five-year survival, and may benefit from localized delivery of drugs.

Hydrogels, particularly biopolymer-based hydrogels have emerged as promising platforms for DDSs due to their inherent biocompatibility, tunable physical properties and controllable degradability. Consequently, the clinical application of implantable hydrogels for drug delivery have been progressed in recent years. In considering the use of hydrogel materials, however, the need for complicated fabrication methods, slow and expensive manufacturing and relatively low mechanical strength, could be obstacles to the realization of practical drug delivery applications. In addition, investigations into the use of hydrogels often report low drug loading capacity, and poor control over drug release.

Methods for incorporation of drugs into hydrogels generally fall into three categories: (I) physical encapsulation, where drugs are simply entrapped within the cross-linked polymer network; (II) covalent conjugation, where drugs are covalently bound to the polymer network; and (III) affinity binding, where hydrophobic, ionic, or hydrogen bonding are utilised to retain drugs within the hydrogel network.

Though physical encapsulation renders a simplistic approach for designing drug loaded hydrogels, these systems often face uncontrolled drug diffusion out of the bulk hydrogel immediately after formation (burst release). On the other hand, covalent conjugation of drugs to the hydrogel network yields a much slower drug release profile, where release is facilitated by cleavage of the covalent linkage or by complete hydrogel degradation. However, this strategy often involves chemical modification of drugs which not only can impact their FDA approval but could also alter their biological activity. Along those lines, affinity-based drug delivery has emerged as an alternative to achieve controlled release of small-molecule drugs, without compromising their activity. Affinity-controlled release is usually achieved by immobilising a drug-binding ligand within a hydrogel matrix. Based on the type of drug molecule, a variety of ligands have been explored for affinity-controlled release, this includes heparin-like moieties, binding peptides, cyclodextrins, and thiols. Recently, researchers have utilised 3,4-dihydroxyphenylalanine (DOPA) to confer affinity- controlled release to hydrogels. The catechol group of dopamine is capable of adhering to drug molecules by using a variety of mechanisms including hydrogen bonds, p-p interactions, cation-p interactions and dynamic covalent bonds.

The present invention seeks to provide multiaxial fibres, particularly coaxial fibres, particularly an implantable device fabricated from such fibres, and a method for implanting said device in a subject for use in treating a medical condition, which will overcome or substantially ameliorate at least some of the deficiencies of the prior art, or to at least provide an alternative.

A reference herein to a patent document or any other matter identified as prior art, is not to be taken as an admission that the document or other matter was known or that the information it contains was part of the common general knowledge as at the priority date of any of the claims.

Where any or all of the terms "comprise", "comprises", "comprised" or "comprising" are used in this specification (including the claims) they are to be interpreted as specifying the presence of the stated leatures, integers, steps or components, but not precluding the presence of one or more other features, integers, steps or components.

Summary of Invention

The invention is based on composite fibres composed of coaxial or multiaxial polymer fibres or mixtures thereof, which have therapeutic effect per se, particularly anti-cancer or anti-tumour effects per se. Furthermore, when loaded with one or more therapeutics, the composite fibres can simultaneously exert both affinity and diffusion control over the release of the therapeutics or drugs, particularly in the case of chemotherapeutic drugs. The composite fibres of the invention are particularly suited as implantable devices for the treatment or prevention of medical conditions. Preferred uses of the composite fibres are biomedical uses including phototherapy. Preferably, the composite fibres may be used as neoadjuvant to shrink a tumour prior to surgery or adjuvant therapies for controlled delivery of chemotherapeutic drugs locally to the tumour sites. The fibres can also be used to prevent reoccurrence of a tumour after surgical removal and/or treatment with therapeutics that reduce or eliminate a tumour, particularly a pancreatic or breast tumour. Preferably, the composite fibre is provided in the form of an implantable device.

In a first aspect the invention provides at least one composite fibre, preferably in the form of a therapeutic implantable device, the at least one composite fibre comprising: at least one core component comprising at least one swellable modified hydrophilic hydrogel polymer, and at least one shell component, wherein at least a first shell component comprises at least one swellable modified hydrophilic hydrogel polymer, wherein at least one of the polymers of the core are modified with at least one crosslinkable alkene functional group and at least one of the polymers of the at least one core are modified with at least one hydroxylatedphenyl functional group to form a crosslinkable alkene-modified- hydroxylatedphenyl-modified polymer, and wherein, at least one of the polymers of the shell are modified with at least one crosslinkable alkene functional group to form a crosslinkable alkene-modified polymer; and wherein, the at least one core component is encapsulated by the at least one shell component.

In a second aspect the invention provides a therapeutic implantable device, the device comprising at least one composite fibre comprising: a core component comprising at least one hydrophilic hydrogel polymer selected from alginate or chitosan, and a shell component, a first shell component comprising at least one hydrophilic hydrogel polymer selected from alginate or chitosan, wherein the hydrogel polymers of the core are modified with at least one crosslinkable alkene functional group and the polymers ol the core are modified with at least one hydroxylatedphenyl functional group to form a core comprising crosslinkable alkene-modified-hydroxylatedphenyl-modified hydrogel polymers, and wherein at least one of the hydrogel polymers of the shell are modified with at least one crosslinkable alkene functional group to form an crosslinkable alkene-modified polymer, wherein the core component is encapsulated by the shell component.

In a third aspect the invention provides therapeutic implantable device comprising: at least one composite fibre comprising: a core component comprising a hydrophilic hydrogel polymer which is alginate, and a shell component comprising a hydrophilic hydrogel polymer which is alginate, wherein the hydrogel polymer of the core are acrylate-catechol-modified alginate polymers, and the hydrogel polymer of the shell are acrylate modified hydrogel alginate polymers, and wherein the core component is encapsulated by the shell component.

In a fourth aspect the invention provides a therapeutic implantable device, the device comprising at least one composite fibre comprising: a core component comprising alginate, and a shell component comprising alginate, wherein the core comprises a mixture of alginate modified with at least one methacrylate functional group and alginate modified with at least one 3,4-dihydroxyphenylalanine (DOPA) group to form a methacrylate-DOPA modified alginate in the form of a polymer mixture comprising methacrylate-modified alginate and DOPA-modified alginate in a ratio of 75:25, and wherein alginate of the shell is modified with at least one methacrylate functional group to form an acrylate modified alginate, and wherein the core component is encapsulated by the shell component.

In a fifth aspect of the invention, there is provided at least one composite fibre, preferably in the form of an implantable device, comprising at least one composite fibre comprising: at least one core component comprising modified alginate, and at least one shell component comprising modified alginate, wherein at least one of the core components comprises a mixture of alginate modified with at least one methacrylate functional group and alginate modified with at least one 3,4- dihydroxyphenylalanine (DOPA) group to form a methacrylate-3, 4-dihydroxyphenylalanine (DOPA) modified alginate, and wherein alginate of the shell is modified with at least one methacrylate functional group to form an acrylate modified alginate, and wherein the at least one core component is encapsulated by the at least one shell component.

In a sixth aspect of the invention, there is provided at least one composite fibre, preferably in the form of an implantable device, comprising at least one composite fibre comprising: at least one core component comprising a physically crosslinked and chemically crosslinked alginate hydrogel and a therapeutic selected from : DOX and GEM, or a combination of DOX and GEM entrained in the hydrogel, and at least one shell component comprising a physically crosslinked and chemically crosslinked alginate hydrogel, wherein at least one core comprises a mixture of alginate modified with at least one methacrylate functional group and alginate modified with at least one 3, 4-dihydroxyphenylalanine (DOPA) group to form a methacrylate-3, 4-dihydroxyphenylalanine (DOPA)-modified alginate, and wherein alginate of the shell is modified with at least one methacrylate functional group to form a methacrylate-modified alginate, and wherein the at least one core component is encapsulated by the at least one shell component; and wherein the methacrylate is activatable to chemically crosslink strands of the alginate hydrogel.

The fibres of the invention are composite fibres composed of one or more core components surrounded by at least one shell component. Preferably, the multiaxial fibre is a coaxial fibre or a triaxial fibre. Preferably, the fibres are one or more of: swelled, dried and sterilised, particularly when packaged for use. Preferably, the fibres comprise two, three or four independent core components, which may be the same or different. Suitably, the fibres comprise a single (one) core component.

Suitably, the invention includes at least one composite fibre comprising: at least one core component comprising at least one swellable hydrophilic hydrogel polymer, and at least one shell component, wherein a first shell comprises at least one swellable hydrophilic hydrogel polymer, wherein at least one of the polymers of at least one of the cores comprise polymer strands which are physically crosslinked and chemically crosslinked and at least one of the polymers of at least one of the cores are modified with at least one hydroxylatedphenyl group to form a crosslinked hydroxylatedphenyl modified polymer, and wherein at least one of the polymers of the at least one shells are physically crosslinked and chemically crosslinked, and wherein the at least one core component is encapsulated by the at least one shell component.

As described herein the hydroxylatedphenyl group may comprise two or more hydroxyl substituents on a phenyl or phenol ring. For example, the ring may have 2, 3, 4 or 5 hydroxyl substituents on the ring. Desirably, the hydroxylatedphenyl group is a dihydroxyphenyl containing functional group, trihydroxyphenyl containing functional group, or a polyhydroxyphenyl groups with more than two hydroxyl substituents. Further, the hydroxylatedphenyl group may be in the form of a polyphenol moiety having a plurality of phenol groups. Examples include polyphenols containing repeating phenolic moieties of catechol, pyrocatechol, resorcinol, pyrogallol, and phloroglucinol. Desirably, the hydroxylatedphenyl is a disubstituted hydroxyphenyl group, such as catechol.

Suitably, the invention includes at least one composite fibre comprising: at least one core component comprising at least one swellable hydrophilic hydrogel polymer, and at least one shell component, wherein at least a first shell comprises at least one swellable hydrophilic hydrogel polymer, wherein at least one of the polymers of at least one of the cores comprise polymer strands which are physically crosslinked and chemically crosslinked and at least one ol the polymers of at least one of the cores are modified with at least one catechol group to form a crosslinked catechol modified polymer, wherein at least one of the core components are encapsulated by the at least one shell component.

Desirably, the fibres include a single core, a pair of cores or three core components encased in at least one shell. Single cored fibres are particularly preferred. Composite fibres having a single shell are particularly preferred. Preferably one or more strands of the polymers may intertwine forming physical crosslinks in the polymer network. Suitably, the polymer strands are chemically crosslinked through activation of one or more of the crosslinking functional groups. It will be understood that these crosslinkable groups when activated to crosslink, form crosslinking bridges between at least two polymer strands. Examples of crosslinkable groups include acrylate, bisacrylamide or bisimide. Preferably, the acrylate functional groups are crosslinkable to form bridges between different strands of the polymer which are derived from the acrylate crosslinkable groups. Suitably, strands of the polymer may also be ionically crosslinked by ionic interactions between metal ions, such as cations including alkali cations such as Na⁺ or divalent cations, preferably divalent metal cations, including Ca²⁺ or Mg²⁺, for example provided during fibre manufacture.

Suitably, one or more of the polymers of the core polymers and the at least one shell polymers comprise crosslinks. Desirably, the crosslinks in the core component include one or more of physical crosslinks, chemical crosslink or a combination of physical crosslinks and chemical crosslinks. Desirably, the crosslinks in the shell component include physical crosslinks, chemical crosslink or a combination of physical crosslinks and chemical crosslinks. Advantageously, this results in a dual cross-linked network in the polymers of the fibres. Suitably, the shell component has more crosslinks than the core component. Alternatively, the core component has more crosslinks than the shell component. The majority of crosslinks in one or more of core and shell components can be physical crosslinks or chemical crosslinks. In some embodiment, the number of crosslinks, particularly physical or chemical crosslinks in the core and shell components are substantially the same.

Desirably, the acrylate is activatable to chemically crosslink polymer strands. Desirably, the crosslinkable alkene functional group, preferably acrylate, is activatable to chemically crosslink polymer strands. For example, the acrylate groups can be activated, for example, by light, particularly UV light, to from covalent bonds bridging polymer strands. The degree of crosslinking can be controlled by the crosslinking reaction time and the degree of incorporation of the acrylate into the polymer and where light is used, by the wavelength of the light. Particularly preferred wavelengths include UV light, for example at 365 nm.

The composite fibres may have a 3D structure or 3D network which can be build up from assembly of the polymer strands into a construct, for example, by fibre weaving or other assembly.

Preferably, one or more of the core components are encapsulated by at least one shell component. However, a second or even a third shell component may be included where further modified release profiles are desirable. Other than for the first shell, subsequent shells can be fabricated from hydrophobic polymers and if desired can be loaded with one or more hydrophobic therapeutics.

Surprisingly, the composite fibres of the invention have been found to provide anti-cancer properties, particularly against pancreatic or breast cancer even without any loaded therapeutics. This is thought to at least partially result from the hydroxylatedphenyl groups, preferably dihydroxyphenyl or catechol groups, provided in the modified polymer of the fibres. Furthermore, the composite fibres of the invention have been found to readily attach themselves to tumour tissue and/or to grow around or associate tumour tissue. These properties prevent undesirable relocation of the fibres from an implant site, for example, at a targeted tumour

In another embodiment, the fibres can be used in photo thermal therapy, for example, where there are combined with nanoparticles, which are activatable to generate heat. Such nanofibers may be a near infra-red or light sensitive nanoparticle, for instance, copper-selenium particles or gold nanoparticles. Additionally, nanoparticles can optionally themselves be loaded with one or more therapeutics.

If desired, one or more of the core and shell components can be loaded with one or more therapeutics.

Preferably, one or more of the cores of the composite fibres functions as an affinity release core. Preferably, at least one of the shells of the composite fibres functions as a diffusion barrier to form a controlled diffusion barrier for the controlled release of one or more therapeutics from the composite fibres.

Preferably, one or more of the core components and at least one of the shell components are loaded with at least one therapeutic agent. Therapeutics agents include pharmaceutical drugs such as small molecule drugs, protein, peptide or polypeptides drugs, and biopharmaceutical drugs such as biologies or nucleic acid entities such as siRNA, mRNA or RNAi, for example. Therapeutic nanoparticles such as copper selenium or gold nanoparticles are also included within the meaning of therapeutic as used herein. Desirably, one or more of the therapeutic agents are active towards treating or preventing a disease and/or a medical condition. Suitably, at least one therapeutic agent is a chemotherapeutic agent. Desirably, the medical condition is cancer, particularly pancreatic cancer or breast cancer.

Preferably, the core only is loaded with at least one therapeutic. Preferably, only at least one shell is loaded with at least one therapeutic agent. Preferably, the core and the shell are loaded with at least one therapeutic agent.

The hydroxylatedphenyl group, preferably catechol group, should have an affinity for a variety of therapeutics. The hydroxylatedphenyl group, preferably a dihydroxylphenyl or a catechol functional group, when covalently bonded to the polymer, supports slow release of one or more of the therapeutics from the relevant component, due to affinity interactions between the molecules of the therapeutic agent and the polyhydroxyphenyl moieties. Indeed, the hydroxylatedphenyl group makes the composite fibres of the invention compatible with a variety of therapeutic agents, regardless of charge on the therapeutic, because the affinity arises from specific interactions between the therapeutic molecules and the hydroxylatedphenyl groups including hydrogen bonds, aromatic interactions and tt-p interactions. Thus, specific interactions between the therapeutic and the modified polymer matrix of the fibres including hydroxylatedphenyl groups results an affinity between the therapeutic molecules to the hydroxylatedphenyl group modified polymer of the core component. Release of the therapeutic from the relevant component is affected by the degree of affinity between the therapeutic and the polymer of the composite fibres. The controlled release profile is additionally supported by the degree of physical and/or chemical crosslinking in the core and/or shell.

Preferably, the core comprises at least one acrylate catechol-modified polymer. Preferably, the shell comprises at least one acrylate-catechol-modified polymer. Preferably, the core comprises at least one acrylate-modified polymer. Preferably, the shell comprises at least one acrylate-modified polymer. Preferably, the core and the shell comprise at least one acrylate-catechol-modified polymer. Preferably, the core and the shell comprise at least one acrylate-catechol-modified polymer. Preferably, the core and the shell comprise at least one acrylate-modified polymer. Preferably, the core and the shell comprise at least one acrylate-modified polymer. Preferably, the core comprises at least one acrylate-modified polymer and at least one acrylate-catechol-modified polymer. Preferably, the shell comprises at least one acrylate-modified polymer and at least one acrylate-catechol-modified polymer. Most preferably, the core comprises at least one acrylate-modified polymer and at least one acrylate-catechol-modified polymer and the shell comprises at least one acrylate-modified polymer.

Suitably, the acrylate-catechol-modified polymer when loaded with at least one therapeutic/drug confers affinity-controlled release, while an acrylate-modified polymer provides a controlled diffusion barrier. Suitably, the acrylate-catechol-modified polymer when loaded with at least one therapeutic/drug is the core component which confers affinity-controlled release, while an acrylate- modified polymer is the shell component which provides a controlled diffusion barrier.

Preferably, in a dry state, the coaxial fibres have a modulus ol from 2 to 4 GPa, preferably 2.8 GPa. Preferably, in a dry state, the coaxial fibres have a tensile strength of from 80 to 200 MPa, preferably 85 MPa. Preferably, in a wet state, the coaxial fibres have a modulus of from 0.1 to 0.6 GPa, preferably 0.3 GPa. Preferably, in a wet state, the coaxial fibres have a tensile strength of from 0.2 to 2 MPa, preferably 0.6 MPa. Preferably, in a dry state, the coaxial fibres have a storage modulus of from 1000 to 10000 MPa, preferably 4000 Pa.

Suitably, the at least one swellable hydrophilic or swellable hydrophobic polymer is a swellable modified hydrophilic or swellable modified hydrophobic polymer wherein the polymer are modified with one or more of: crosslinkable functional groups and functional groups which have an affinity for one or more therapeutic or drugs.

Preferably, the core component comprises the acrylate-catechol modified polymer, and the shell component comprises the acrylate-modified polymer. Alternatively, the shell component comprises the acrylate-modified polymer, and the core component comprises the acrylate-catechol modified polymer. In one embodiment, both the shell component polymers and the core component polymer comprise the acrylate-catechol modified polymer. In one embodiment, both the shell component polymers and the core component polymer comprise the acrylate-modified polymer.

Preferably, the crosslinkable alkene-mcdified, hydroxylatedphenyl-modified polymer is in the form of a mixture of crosslinkable alkene-modified polymers and crosslinkable alkene-modified, hydroxylatedphenyl-modified polymers. Preferably, the hydroxylatedphenyl-modified polymers and the crosslinkable alkene-modified polymers in the mixture are present in a ratio of 50:50 to 85:15, most preferably 75:25. Desirably, the crosslinkable alkene-modified, hydroxylatedphenyl-modified polymer is in the form of a single polymer which is simultaneously crosslinkable alkene -modified and hydroxylatedphenyl-modified. In another embodiment, a mixture of the hydroxylatedphenyl modified and hydroxylatedphenyl-modified and the single polymer which is simultaneously crosslinkable alkene- modified and hydroxylatedphenyl-modified can be used. More preferably, the acrylate-catechol- modified polymer is in the form of a mixture of acrylate-modified polymers and catechol-modified polymers. Preferably, the catechol-modified polymers and the acrylate-modified polymers in the mixture are present in a ratio of 50:50 to 85:15, most preferably 75:25. Desirably, the acrylate-catechol- modified polymer is in the form of a single polymer which is simultaneously acrylate-modified and catechol-modified. In another embodiment, a mixture of the acrylate-modified polymers and catechol- modified polymers and the single polymer which is simultaneously acrylate-modified and catechol- modified can be used.

Preferably, the amount of crosslinkable functional groups e.g. acrylate derived groups in the shell is higher than in the core component. Alternatively, the amount of crosslinkable functional groups in the core is higher than that in the shell component. This means that more chemical crosslinking can occur in the component having the highest amount of or crosslinkable functional groups.

Suitably, the polymers of one or more of the shell and core adopt a porous structure when the polymers are in a partially or fully swelled state. This can readily be determined by SEM analysis. The more swelling that occurs, the bigger the pore size becomes. Suitably, core pore size distribution ranges from about 20 microns to about 150 microns on full swelling. Desirably, the shell pore size distribution ranges from about 5 microns to about 20 microns on full swelling. The porosity can be readily adjusted as required by changing the concentration of cross- linkable groups in the polymer and/or by incorporating additional cross-link agents, for example, bis- acryl amide, into the fibres. Acrylate crosslinking agent are preferred due to less toxicity. Suitably, on swelling to a fully swelled state, the polymers of the shell have a less porous structure in the shell component when compared to the core section. Preferably, on swelling to a fully swelled state, the polymers of the shell component have smaller pores than the pores of the core component. Suitably, on swelling to a fully swelled state, the shell polymers have a more compact structure with smaller pores than the core polymer.

Suitably, the catechol-modified polymer has a theoretical substitution degree of from about 5% to 32%, preferably 16% in terms of polymer repeating units. Suitably, the acrylate-modified polymer has a theoretical substitution degree of from about 20% to 50%, preferably 40% in terms of polymer repeating units. The acrylate and catechol groups can be provided on a single polymer at the same substitution degree.

In preferred embodiments, at least one of the polymers of the core component is a hydrophilic polymer and wherein at least one of the polymers of at least one shell component is a hydrophilic polymer

Suitably, the hydrophilic hydrogel polymer in the core and the first shell can be the same or different. Preferably, the core comprises at least one hydrophilic polymer and the shell comprises at least one hydrophilic polymer. Most preferably, the core and the at least one shell comprise at least one hydrophilic hydrogel polymer. Where a second or subsequent shell is included, hydrophobic polymers can be included therein, together with hydrophobic therapeutics if desired.

In embodiments, wherein the composite fibre is provided with at least one therapeutic agent, preferably, the at least one therapeutic agent is: a hydrophilic drug loaded into a hydrophilic polymer of the device or a hydrophobic drug loaded into a hydrophobic polymer of the device, where a hydrophobic polymer is present, for example in a second or subsequent shell.

While hydrophobic drugs can be incorporated/loaded into the fibres by first providing an amphipathic polymeric coating (e.g. PLGA) around fibres of the hydrophilic polymer, hydrophilic drugs are preferred particularly in the case of the cores and first shell components surrounding the cores.

Preferably, the at least one hydrophilic polymer is selected from the group consisting of: alginate or chitosan. A particularly preferred hydrophilic polymer is alginate.

Preferably, the at least one hydrophobic polymer is selected from the group consisting of: polycapro lactone, poly(lactic acid), poly(lactic-co-glycolic acid), poly(2-oxazoline), polyglycerol sebacate, polypropylene glycol and a poly(l-amino acid).

Preferably, the polymer in one or more of the core and the shell are both biocompatible and/or biodegradable such that the implant does not need to be surgically removed. Desirably, the hydrophilic polymer in the core and the shell are both alginate or chitosan, particularly when the composite fibres are to be implanted. An alginate based composite fibre has a faster degradation rate and consequently can be completely cleared from the body, while avoiding surgical removal steps. Alginate is a biocompatible and biodegradable polymer that can be located safely in tissue until it degraded, typical after a period of up to 12 weeks. In particular, biocompatible fibres comprising hydrogel fibres of alginate demonstrate favourable properties including controlled swelling, enhanced mechanical properties, and high drug loading capacity, particularly when compared against single fibres made from unmodified alginate.

Desirably, the core component and/or at least one shell component is loaded with at least one therapeutic agent. Desirably, the core component is loaded with at least one therapeutic agent and the at least one shell component is not loaded with therapeutic. Suitably, the core component and the shell component are each loaded with at least one therapeutic agent.

Desirably, the at least one polymer of the core component is alginate or chitosan, and the at least one polymer of at least one shell component is alginate or chitosan. In an embodiment wherein the at least one polymer of the core component is alginate, and wherein the at least one polymer of the shell component is alginate, preferably one or more of the alginates of the core component and the alginate of the shell component is loaded with at least one therapeutic agent.

Suitably, the alginate of the core component is loaded with a first therapeutic and the alginate of the shell component is not loaded with any therapeutic(s). Suitably, the alginate of the core component is loaded with a first therapeutic and a second therapeutic and the alginate of the shell component is not loaded with any therapeutics Suitably, the alginate of the core component is loaded with a first therapeutic and the alginate of the shell component is loaded with a second therapeutic agent.

Suitably, the first and second therapeutics are the same or different. Most preferably, the first and second therapeutics are different. Preferably, at least one of therapeutic is selected the group of hydrophilic drugs consisting of: doxorubicin (DOX), gemcitabine, paclitaxel, camptothecan, everolimus, epothilone, curcumin, docetaxel, rituximab, cetuximab, trastuzumab, pertuzumab, sunitinib, bevacizumab, an anti-EGFR molecule, an anti-CTLA4 antibody, an anti-PDI or anti-PDL1 antibody or inhibitor, an immune modulating agent, and any combination thereof. Preferably, at least one of the therapeutic is selected from the group of hydrophobic drugs consisting of taxane, epothilone, and curcumin. Preferably, at least one of the therapeutic is a therapeutic nanoparticle, such as copper- selenium nanoparticles or gold nanoparticles. Such nanoparticles themselves can be loaded with one or more therapeutics. Preferably, the first therapeutic is doxorubicin. Suitably, the second therapeutic is gemcitabine. Suitably, the core component comprises acrylate-catechol-modified alginate loaded with doxorubicin or gemcitabine. Suitably, the core component comprise acrylate-catechol-modified alginate loaded with a combination of doxorubicin and gemcitabine. Desirably, the shell component comprises acrylate-catechol-modified alginate loaded with doxorubicin or gemcitabine, or a combination of doxorubicin or gemcitabine. In some embodiments, only the core component comprises doxorubicin or gemcitabine or a combination of doxorubicin and gemcitabine. In some embodiments, the shell does not comprise any therapeutic. In a preferred embodiment, the core component comprises acrylate-catechol-modified alginate loaded with doxorubicin and shell component comprises acrylate-modified alginate loaded with gemcitabine. In a preferred embodiment, the core component comprises acrylate-catechol-modified alginate loaded with gemcitabine and shell component comprises acrylate-modified alginate loaded with doxorubicin. In another preferred embodiment, the core component comprises acrylate-catechol- modified alginate loaded with doxorubicin and gemcitabine, and the shell component is not loaded with any therapeutic.

In one embodiment, the amount of doxorubicin is in the range from 0.1 mM to 10 mM, preferably 2 mM. In one embodiment, the rate of release of doxorubicin from the implantable device is in the range from about 0.03 mg to about 0.05 mg in a first hour of measurement, the rate of release of doxorubicin from the implantable device is in the range from about 0.075 mg to about 1 mg in the first day of measurement, most preferably the rate of release of doxorubicin from the implantable device is in the range from about 0.3 mg to about 0.4 mg in a week measurement.

In one embodiment, the amount of gemcitabine is in the range from 5 to 100 mM, preferably 50 mM. In one embodiment, the rate of release of gemcitabine from the implantable device is in the range from about 1 mg to about 2 mg in a first hour of measurement, most preferably the rate of release of gemcitabine from the implantable device is in the range from about 8 mg to about 10 mg after 10 hours of measurement.

In one embodiment, the coaxial fibre has a diameter in the wet state that falls within a range from about 1500 pm to about 2000 pm. In one embodiment, the core component of the coaxial fibre has a diameter in a range from about 1200 pm to about 1600 p and the shell component that surrounds the polymer core has a wall thickness that falls within a range from about 300 pm to about 400 pm. In one embodiment, the core component of the coaxial fibre has a pore size of about 20 pm to about 150 pm. In one embodiment, the shell component of the coaxial fibre has a pore size of about 5 pm to about 20 pm.

Preferably, the at least one acrylate functional group is derived from methacrylic acid, ethylacrylic acid, propylacrylic acid or acryl amide.

Preferably, the at least one catechol group is derived from a functional group comprising 1 ,2- dihydroxybenzene, for example, 3,4-dihydroxyphenylalanine (DOPA), gallic acid, and caffeic acid or other polyphenol moieties.

Preferably, the chemical crosslinks include covalently bonded bridges between polymer strands which are derived from acrylate groups. Suitably, the acrylate groups are converted to crosslinks between strands of polymer, prelerably by application of electromagnetic radiation, for example, UV light, particularly at 365 nm. Preferably, one or more type of metal cation facilitates additional chemical crosslinks between strands of polymer.

In a preferred embodiment, the implantable devices takes the form of a plurality of seeds or particles, a patch or a web, preferably a woven patch or a woven web. Typical preferred dimension/lengths range from about 0.5 cm to about 1 cm for each seed. Preferably, the implantable device of the invention further comprises at least one sheath encapsulating a least one composite fibre substantially therein. Suitably sheath materials include PLA or PCL, for example. Suitably, at least one sheath is modified to comprise a plurality of apertures disposed along the length of the sheath in spaced apart arrangement. Preferably, each of the plurality of apertures has a diameter of from about 74 pm to about 124 pm, preferably about 100 pm.

In a preferred embodiment, the implantable device further comprising locating means for use in locating the implantable device when implanted into a subject. Suitably, the locating means is configured for use in ultrasound detection and comprises a plurality of metal nanoparticles embedded substantially within at least one of the hydrophilic polymer and the hydrophobic polymer of the composite fibre. Desirably, the metal nanoparticles are selected from the group consisting of: platinum, gold, silver, and any combination thereof. Preferably, the locating means is configured for use in ultrasound detection and comprises a metal coating formed on an external surface of at least one of the hydrophilic polymer and the hydrophobic polymer of the composite fibre. Where the composite fibres further comprises a sheath encapsulating the composite fibre substantially therein, the sheath may comprises locating means in the form of a metal coating deposited on an external surface of the sheath. Suitably, the metal coating comprises a metal selected from the group consisting of: platinum, gold, and any combination thereof.

Suitably, the composite fibre is formed using a method selected from the group consisting of 3D printing, wet spinning, electrospinning, coaxial melt extrusion printing, coaxial melt electro-writing, hot melt extrusion and pulsatile fibre spinning. Preferably, the composite fibre is formed using a wet spinning technique. Among these methods, wet-spinning is the simplest with fast and inexpensive manufacturing while providing excellent control over the size, morphology, and composition of the fibres. Advantageously, wet spun fibres are amenable to conventional textile fabrication technologies, such as weaving, braiding and knitting, to create finely tuned 2D or 3D composite fibre constructs which can be used as implantable DDSs. The 2D constructs have a single fibre thickness in one direction, whereas the 3D construct has a thickness in one direction of two or more composite fibres.

Preferably, the first solution is a mixture of at least one swellable acrylate modified polymer and at least one swellable catechol modified polymer. Alternatively, a single hydrophilic hydrogel polymer may be simultaneously modified with acrylate and catechol may be used.

Desirably, the extrusion step involves extruding both solutions through the extrusion means into a coagulation bath comprising a coagulation agent, preferably a metal salt coagulation agent, most preferably a CaC coagulation agent, in the case of alginate. In one embodiment, the coagulation bath comprises a solution of ethanol and water. In one embodiment, the coagulation agent is CaC^. Where chitosan is used the coagulation involves a sodium salt, for example sodium hydroxide.

In a seventh aspect of the invention, there is provided a method of preparing a composite fibre, preferably in the form of a therapeutic implantable device, the method comprising the steps of: preparing a spinnable first solution of at least one hydroxylatedphenyl-modified hydrophilic polymer and at least one further solution of at least one swellable crosslinkable alkene-modified hydrophilic polymer; and extruding the first and the at least one further spinnable solutions through a coaxial spinneret into a coagulation bath comprising a coagulation agent; and collecting from the coagulation bath, via a rotating mandrel, at least one composite fibre spun from the modified polymers.

Preferably, the method comprising the steps of: preparing a spinnable first solution of at least one swellable acrylate-catechol modified hydrophilic hydrogel polymer and at least one further solution of at least one swellable acrylate modified hydrophilic hydrogel polymer; and extruding the first and the at least one further spinnable solutions through a coaxial or multiaxial spinneret into a coagulation bath comprising a coagulation agent; and collecting from the coagulation bath, via a rotating mandrel, at least one coaxial or multiaxial fibre spun from the modified polymers.

Desirably, the method further comprises the step of activating the acrylate groups or other crosslinkable functional groups to crosslink polymer strands in the composite fibre.

Suitably, the acrylate groups are activated by light, suitably UV light.

Preferably, one or more of the first solution and the at least one further solution comprise one or more therapeutics. Where additional shells are required, additional solutions corresponding to each additional shell are prepared combined with a desired therapeutic, if required, and included in the extrusion step using a suitable extrusion means for multiaxial extrusion.

Desirably, wherein on extrusion, the one or more therapeutics are entrained in the polymer.

Preferably, the polymer of the swellable acrylate-catechol modified polymer is a hydrophilic polymer and the at least one swellable acrylate modified polymer is a hydrophilic polymer.

Preferably, the extrusion step is such that the swellable acrylate-catechol modified polymer forms a core component of the composite fibre and the swellable acrylate modified polymer forms a shell component of the composite fibre.

Desirably, the hydrophilic polymers are alginate. Preferably, the polymers are hydrophilic polymers which are alginate, the acrylate group is derived from methacrylate, and the catechol group is derived from DOPA.

Suitably, where doxorubicin and/or gemcitabine are used, they are included in solution at a concentration of 2 mM and 50 mM respectively. Desirably, the solution for the core comprises 3% (w/v) of alginate-dopamine/alginate-methacrylate mixture (75/25% w/w). Desirably, the solution for the shell comprises 3% w/v alginate-methacrylate. Suitably, where doxorubicin and gemcitabine are used, they are included in solution at a concentration of 2 mM of DOX in the core solution, and in solution at a concentration of 50 mM of GEM for the shell. Desirably, the method further comprises the step of: coating the at least one composite fibre with at least one polymer to form at least one sheath encapsulating the composite fibre substantially therein. Suitable sheath materials include PCL or PLA. Preferably, the sheath is modified to comprise a plurality of apertures disposed along the length of the sheath in spaced apart arrangement.

In an eight aspect of the invention, there is provided a method of delivering at least one therapeutic to a subject, the method comprising the step of: implanting a therapeutic implantable device according to the lirst to sixth aspects into a subject presenting with a medical condition that is treatable or preventable with the at least one therapeutic.

Preferably, the implanting step is carried out using endoscopic ultrasound-guided implantation.

In a ninth aspect of the invention, there is provided a method of treatment or prevention of a disease or condition in a subject in need thereof, comprising the steps of: providing a therapeutic implantable device to the subject in need thereof, wherein the therapeutic implantable device is according to the first to sixth aspect.

In a tenth aspect ol the invention, there is provided a use of therapeutic implantable device is according to the first to sixth in the treatment or prevention of a disease or condition.

In an eleventh aspect the invention provides a use of a therapeutic implantable device for controlled release of a therapeutically effective amount of least one therapeutic in the treatment of a disease or condition, wherein the device is according to the first to sixth aspects.

Desirably, the disease or condition is cancer, particularly pancreatic cancer, particularly pancreatic cancer or breast cancer.

Notably, the composite fibres release drugs in a slower manner, when compared to single fibres made from the corresponding unmodified polymer. The may be due to stronger interactions of drugs with hydroxylatedphenyl-modified polymer as observed from zeta-potential measurements. It was further shown that the drug-loaded composite fibres have anticancer activity both in vitro and in vivo cancer tumours, e.g., using various pancreatic cancer cell lines. Most remarkably, drug loaded composite fibres, particularly doxorubicin-containing fibres, had higher anticancer effect in vivo compared to systemic injection of equivalent dosage of the drugs. Altogether, these biocompatible and robust hydrogel fibres may be further used as neoadjuvant or adjuvant therapies for controlled delivery of chemotherapeutic drugs locally to the tumour sites.

Brief Description of Drawings

Figure 1 illustrates (a) schematic illustration of coaxial hydrogel fibres of the invention, showing the alginate-dopamine (75%, v/v) + alginate-acrylate (25%, v/v) mixture of the core and the acrylate modified alginate of the shell, (b) schematic illustration of coaxial hydrogel fibres of the invention, showing the alginate-dopamine-acrylate of the core and the acrylate modified alginate of the shell. Figure 2 illustrates (a) viscometry of 3% (w/v) solution of i) alginate-methacrylate, ii) alginate- dopamine/ alginate methacrylate (75/25 w/w), iii) alginate-dopamine, and iv) pure alginate; and illustrates (b) oscilatory rheology of 3% (w/v) solutions (containing 0.05% (w/v) irgacure D-2959) of i) alginate-methacrylate, and ii) alginate-dopamine/ alginate methacrylate (75/25 w/w). Microscopy images showing (c & d) top view of freshly made core-shell fibres (CS) loaded with DOX in the core (scale bars 2 mm and 1 mm, respectively).

Figure 3. SEM images of core-shell mussel-inspired fibers (CS), (a & b) SEM image of CS fibers at dry state (c & d) SEM images of CS fibers fully swollen in simulated body fluid (SBF).

Figure 4. illustrates (a) Final swelling ratio of single fibers made from pure alginate (Alg), and CS fibers in SBF. (b) In vitro degradation profiles of Alg and CS fibers in SBF as a function of time, (n = 4, mean ± SD) (^****P = 0.0001, ^**P < 0.01)

Figure 5. Mechanical properties of fibers following a static tensile test, pure alginate fibers (Alg) or core-shell mussel-inspired fibers (CS), (a) Stress-strain curve for fibers in dry state (b & c) Tensile strength and modulus of fibers in dry state (n = 3, mean ± SD). (d) Stress-strain curve for fibers in wet state (e & f) Tensile strength and modulus of fibers in wet state (n = 3, mean ± SD) (^**P < 0.01 , "^***P = 0.0001 ).

Figure 6. illustrates dynamic mechanical analysis of hydrogel fibres, single fibres made from pure alginate (Alg) or core-shell fibres (CS), by using a tension clamp. The samples were exposed to a constant strain (0.1 %) of increasing frequencies in the range of 0.1 to 10 Hz for 30 min. (a) Storage modulus and (b) Loss modulus of fibres at different frequencies (c) Storage and loss modulus at the frequency of 10 Hz. (d) Tan delta (damping factor) at the frequency of 10 Hz (n = 3, mean ± SD) (^*P = 0.05, ^**P < 0.01).

Figure 7. illustrates In vitro drug release studies from hydrogel fibers including single fibers made from pure alginate (Alg), core-shell mussel-inspired fibers (CS), and core shell fibers in which both core and shell were entirely made of UV-crosslinkable alginate-methacrylate (CS·). (a) Release profile of gemcitabine (GEM) (b) The first 2 hr of GEM release (c) Release profile of doxorubicin (DOX). (d) The first 10 hr of DOX release (n = 3, mean ± SD)

Figure 8 illustrates in vitro biocompatibility of fibres including empty CS fibres (Control), Gemcitabine loaded CS fibres (GEM treated), Doxorubicin loaded CS fibres (DOX treated), and dual-loaded CS fibres (GEM-, DOX-, Dual-treated) (a-c) MTS cell viability assay of MIA PaCa-2 human pancreatic adenocarcinoma cells when treated with the fibres (d-f) MTS cell viability assay of PANC-1 human pancreatic adenocarcinoma cells when treated with the fibres, (g) & (h) Corresponding live/dead cell staining of treated MIA PaCa-2 and PANC-1 human pancreatic adenocarcinoma cells, respectively. Values are the mean (±SEM) of quadruplicate. ^*P = 0.05, ^**P < 001 , ^***P = 0.001 , ^****P = 0.0001 . Figure 9 illustrates the therapeutic effect of CS fibres on inhibition of cancer cell growth. The viability of (a) MIA PaCa-2, (b) PANC-1 , and (c) BxPC3-Luc human pancreatic adenocarcinoma cells at the different time points treated with Empty CS fibres (Control), Gemcitabine-loaded CS fibres (GEM treated), Doxorubicin-loaded CS fibres (DOX treated), and dual-loaded CS fibres (GEM-, DOX-, Dual- treated). Values are the mean (+SEM) of quadruplicate.

Figure 10 illustrates the in vivo performance of drug loaded CS fibres in MIA PaCa-2 mice xenografts compared to intravenous injection of same drugs at equal dose (n= 5 in each group, ^*P = 0.05, ^**P < 0.01 , ^***P = 0.001 , ^****P = 0.0001). (a) Schematic representation of localized therapeutic approach with drug loaded CS fibres in mice xenograft model (MIA-PaCa-2 tumours), (b) Tumour volume at day 0 and day 14 after the treatment, (c) Tumour weight at day 14 after the treatment, (d) Body weights of mice in different groups after treatment, (e) Images of hematoxylin and eosin (H&E) staining along with KI-67 staining of pancreatic-cancer tissue (MIA-PaCa-2 tumours) sections from treated mice. The scale bar in images with 4x, 10x, and 20x magnification is 1000 pm, 400 pm and 200 pm, respectively. Figure 11 illustrates the in-vivo performance of drug loaded CS fibres in BxPC3-Luc mice xenografts (n= 6 in each group, ^*P = 0 05, ^**P < 0.01 , ^***P = 0.001 , ^****P = 0.0001). (a) Schematic representation of localized therapeutic approach with drug loaded CS fibres in mice xenograft model (BxPC3-Luc tumours) (b) Tumour volume at day 0 and day 14 after the treatment (c) Tumour weight at day 14 after the treatment, (d) Body weights of mice in different groups after treatment, (e) Images of hematoxylin and eosin (H&E) staining along with KI-67 staining of pancreatic-cancer tissue (BxPC3- Luc tumours) sections from treated mice. The scale bar in images with 4x, 10x, and 20x magnification is 1000 pm, 400 pm and 200 pm, respectively.

Figure 11 illustrates (a & b) release profile and cumulative drug release of GEM loaded fibres, respectively (n = 3, mean ± SD). (c & d) Release profile and cumulative drug release of DOX loaded fibres, respectively (n = 3, mean ± SD). (e & f) Release profile and cumulative drug release comparisons of GEM being encapsulated in the core or in the shell section of CS fibres (n = 3, mean ± SD) (^*P = 0.05, ^**P < 0.01).

Figure 12 illustrates the anti-cancer properties on the fibers, the fibroblast cells are healthy skin cells and the empty fibers have no negative impact on them. Whereas MIA PaCa 2 and Panc-1 are pancreatic cancer cells and empty fibers showed their anti-cancer properties against these cell lines. Figure 13 illustrates (a & b) schematic of the coaxial hydrogel fibers of CS fibres loaded with nanoparticles for photo-thermal therapy of tumours.

Figure 14 In vivo photo-thermal conversion effect of 4 cm of the fibers (a) Infrared thermal images and (b) corresponding temperature curves upon exposure to 808 nm NIR laser for 10 min, under power densities of 1 .0 W-cnr².

Figure 15 illustrates various fabrication methods for fabrication of coaxial hydrogel structure, including wet-spinning and 3Dprinting.

Detailed Description of the Invention

It is to be understood that the following description is for the purpose of describing particular embodiments only and is not intended to be limiting with respect to the above description.

Herein is provided a new generation of coaxial hydrogel fibres which preferably have anti cancer effects per se and that when loaded with therapeutics can simultaneously exert both affinity and diffusion control over the release of drugs, such as chemotherapeutic drugs. Specifically, acrylate- dopamine-modified alginate hydrogel along with chemotherapeutic drugs (doxorubicin or gemcitabine) was used as the main core component to confer affinity-controlled release, while an acrylated-alginate hydrogel was used as the shell composition to provide the controlled diffusion barrier. It was shown that the coaxial biofibres yielded biocompatible hydrogel fibres (as indicated by comprehensive in vitro and in vivo experiments) with favourable properties including controlled swelling, enhanced mechanical properties, and high drug loading capacity, when compared against single fibres made from unmodified alginate. Notably, it was observed that these composite fibres were capable of releasing the two drugs in a slower manner, when compared to single fibres made from pure alginate, which was partly attributed to stronger interactions of drugs with dopamine-modified alginate (the core element of composite fibres) as observed from zeta-potential measurements. It was further shown that these drug-loaded composite fibres had optimal anticancer activity both in vitro and in vivo using various pancreatic cancer cell lines. Most remarkably, drug loaded composite fibres, particularly doxorubicin-containing fibres, had higher anticancer effect in vivo compared to systemic injection of equivalent dosage of the drugs. Altogether, these biocompatible and robust hydrogel fibres may be further used as neoadjuvant or adjuvant therapies for controlled delivery of chemotherapeutic drugs locally to the tumour sites.

In this study, with the aim of developing a suitable hydrogel platform for local delivery of therapeutics, particularly anti-cancer drugs, a wet-spinning method was used to engineer novel coaxial biofibres. Specifically, methacrylate-dopamine-modified alginate hydrogel, along with the chemotherapeutic drugs, was used as the main core component to confer affinity-controlled release, while a methacrylated-alginate hydrogel was used as the shell composition to provide the controlled diffusion barrier (Figure 1). The use of two anti-cancer drugs, gemcitabine (GEM) and doxorubicin (DOX) were investigated, as these drugs are in routine clinical use solely or in combination with other drugs, to control the progression of cancer. It was demonstrated that the proposed hierarchical structure enabled the development of biocompatible hydrogel fibres (as indicated by both in vitro and in vivo experiments) with controlled swelling, enhanced mechanical properties (as indicated by both static and dynamic tension tests), and high drug loading capacity, when compared to single fibres made from pure alginate. Particularly, it was observed that these composite fibres were capable of releasing the two drugs in a slower manner, when compared to single fibres made from pure alginate, which was partly attributed to stronger interactions of drugs with the components of core element of the composite fibres as observed from zeta-potential measurements. Remarkably, it was shown that drug-loaded fibres, particularly doxorubicin-containing fibres, had higher anticancer effect in vivo compared to systemic injection of equivalent dosage of the drugs. Also, simultaneous loading of the two drugs in the composite fibres (DOX and GEM) also inhibited the growth of the tumours in animal models. Overall, these biocompatible and robust coaxial hydrogel fibres may be further used as neoadjuvant or adjuvant therapies for controlled delivery of chemotherapeutic drugs locally to the tumour sites with the aim of suppressing cancer growth, with reduced side effects.

Experimental section Materials

Alginic acid sodium salt from brown algea (medium viscosity), N-(3-Dimethylaminopropyl)-N’- ethylcarbodiimide hydrochloride (³99.0%), N-Hydroxysuccinimide, Dopamine hydrochloride (98%), MES hydrate (> 99.5%), Methacrylic anhydride, and Irgacure D-2959 were purchased from Sigma Aldrich. Calcium chloride (CaCh) and Sodium hydroxide (NaOH) were purchased from Chem-Supply. Doxorubicin hydrochloride and Gemcitabine hydrochloride were purchased from Focus Bioscience. Simulated body fluid (SBF) was prepared as explained before. D₂0 (100% - CIL) was supplied by Novachem. The MIA-PaCa-2 human pancreatic adenocarcinoma cell was kindly supplied by Gillian Lehrbach from the Garvan Institute of Medical Research (Aust). The PANC-1 human pancreatic adenocarcinoma cell line was purchased from ATCC. The CellTiter 96® Aqueous One Solution Cell Proliferation Assay (MTS) was from purchased from Promega, Australia. DMEM-Fligh glucose media and foetal calf serum (FCS) were purchased from Invitrogen, USA. Trypsin/EDTA was purchased from Life Technologies, Australia.

Synthesis and characterization of alginate-dopamine

Alginate-dopamine with different substitution degrees was synthesized via carbodiimide chemistry. Briefly, 1 g of alginic acid (5 mmol in terms of repeating unit) was dissolved in 100 ml of 0.1 M MES buffer with pH of 5.6. EDC and NHS were separately dissolved in 2 ml of MES buffer (1 .25 and 25 mmol for 25% and 50%, respectively) and subsequently added to the alginic acid solution while the reaction were allowed to continue for 30 min . Dopamine hydrochloride (at equal concentration to EDC and NHS) was separately dissolved in 2 ml of MES buffer and added to the mixture subsequently. The reaction was allowed to happen for 1 hr under nitrogen flushing. Afterwards the solution was dialyzed (spectra/ por membrane tubing; MWCO 12-14 kD) for 7 days against acidic water (pH of 6) and subsequently precipitated in ethanol. The precipitated polymers were next lyophilized using a freeze dryer. Incorporation of dopamine into alginate was further confirmed using FTIR (Shimadzu IRPrestige-21 infrared spectrometer) and HNMR(Bruker 400 MHz) spectroscopy. Synthesis and characterization of alainate-methacrylate

Alginate-methacrylate was synthesized by using methacrylic anhydride. Briefly, 3 g of alginic acid was dissolved in 300 ml of distilled water (1% w/v) to which 24 ml of methacrylic anhydride (8% v/v) was added and the pH was maintained at 8 for 6 hrs using 5.0 M NaOH solution. Alterwards, the solution was dialyzed (spectra/ por membrane tubing; MWCO 12-14 kD) for 7 days and precipitated in ethanol followed by freeze-drying. Grafting of methacrylate groups onto the alginate backbone was further confirmed using FTIR (Shimadzu IRPrestige-21 infrared spectrometer) and HNMR (Bruker 400 MHz) spectroscopy.

Rheology

All rheology experiments were conducted on a Physica MCR 301 Rheometer (Anton Paar) in parallel plate geometry (50 mm disk, 0.097 mm measuring distance) and at room temperature (23 °C). Flow experiment was performed to evaluate the viscosity of polymer solution (shear rate varying from 1 to 100 s^~1). Oscillatory experiments as a function of time (at constant shear strain of 1 % and constant frequency of 1 Hz) were performed to measure the storage and loss modulus of the hydrogels upon UV irradiation (365 nm, DYMAX BlueWave 75).

Fabrication of dve loaded fibres After synthesis of alginate-dopamine and alginate-methacrylate, we have fabricated the dye- loaded (fluorescein sodium salt) fibres to evaluate the effect of dopamine grafting on subsequent release and encapsulation of the dye. Accordingly, 3% (w/v) of pure alginate, alginate-dopamine 25% and alginate-dopamine 50%, were separately dissolved in distilled water. Next 2 mM of fluorescein sodium salt were added to each solution to prepare the spinning mixtures. The single and core-shell fibres were fabricated using a coagulation bath of H₂0/Ethanol (4:1 ) containing 2% CaCl2 (w/v). The single fibres were fabricated by simply extruding the dye-containing solutions (50 ml/hr) into the coagulation bath using a blunt needle (gauge 19). On the other hand, the core-shell fibres were fabricated using a novel coaxial spinneret with two input ports. Alginate-dopamine/alginate- methacrylate solution (75/25 w/w, 3% w/v, containing 0.05% irgacure D-2959) containing the dye (2 mM) was extruded (50 ml/hr) through the centre outlet nozzle (gauge 19) into the coagulation bath, while alginate-methacrylate solution (3% w/v, containing 0.05% irgacure D-2959) was simultaneously being extruded (50 ml/hr) as the shell of the fibre (gauge 15). Right after exiting the coagulation bath, the core-shell fibre were irradiated with UV light (365 nm, DYMAX BlueWave 75) as they were being collected on the rotating drum. To measure the dye release from the fibres, 30 cm of fibres were immersed in 1 ml of SBF solution and the supernatant was collected at certain time points and replaced with fresh SBF. Eventually these supernatant were tested using UV spectrometer (Shimadzu, UV- 1800) and the unknown concentrations were revealed using a calibration curve achieved previously. Fabrication of drug loaded fibres

Drug loaded fibres were fabricated in the same manner as dye loaded ones with minor differences. Single alginate drug-loaded fibres were fabricated by initially dissolving the drug in distilled water (2 mM for DOX or 50 mM for GEM), next alginate was dissolved in the mixture at a 3% (w/v) concentration. Subsequently, the prepared mixture was extruded into the coagulation bath using a blunt needle. For core-shell fibres, the drugs (2 mM for DOX or 50 mM for GEM) were first dissolved in distilled water, to which 3% (w/v) of alginate-dopamine/alginate-methacrylate mixture (75/25% w/w) and 0.05% (w/v) irgacure D-2959 were added (this solution was used as the core input). Separately, alginate-methacrylate was dissolved in distilled water (3% w/v) and 0.05% (w/v) irgacure D-2959 was added to this mixture (this solution was used as the shell input). The two solutions were simultaneously extruded through the coaxial nozzle into the coagulation bath and the fibres were subsequently irradiated with UV light (365 nm, DYMAX BlueWave 75) as they were collecting on the rotating drum. For dual-loaded fibres, the core constituent was loaded with 2 mM of DOX while the shell constituent was loaded with 50 mM of GEM. Of note, for the sole purpose of drug release studies, we have fabricated another type of core-shell fibres which is indicated as CS·. In these fibres, the core only contained (3% w/v) alginate-methacrylate + 0.05% (w/v) irgacure D-2959 + the drugs (2 mM for DOX or 50 mM for GEM), while the shell only contained (3% w/v) alginate-methacrylate + 0.05% (w/v) irgacure D-2959.

Characterisation of fibres The morphology of fibres were examined using a JSM-6490LV SEM and Leica M205A microscope. For SEM imaging (in dry or wet state), the fibres were cut into small pieces and inserted into special sample holders. The sample holder containing the mounted fibres was then immersed into liquid nitrogen for about 45 s Next, the sample holder was then quickly transferred to the LVSEM for examination. SEM images were taken in high vacuum mode at 15 kV operating voltage and a spot size setting of 60. The swelling properties of the hydrogel fibres were determined by examining their water uptake capacity. The hydrogel fibres were incubated in simulated body fluid (SBF) at 37°C and allowed to fully swell. The swelling ratio was calculated using the following equation: (Ws - Wd)/Wd, where Ws represents the weight of the swollen hydrogel fibres and Wd represents the weight of the dried hydrogel fibres at the beginning. The degradation rate of the fibres was measured by initially freeze-drying the fibres and subsequently weighing them (Wi). Next, the fibres were immersed in 5 ml of SBF at 37°C for in vitro degradation testing. The SBF was replaced every three days. At predetermined time points, the samples were removed, rinsed with distilled water, and lyophilized and weighed (Wd). The percent mass loss was calculated using the equation (Wi-Wd)/Wi x 100. The static mechanical properties of fibres were assessed using a mechanical tester (EZ-L tester from Shimadzu) at 10 mm.min-1 via 50 N and 10 N load cells for dry and wet fibres, respectively. The dynamic mechanical properties of fibres was evaluated using a dynamic mechanical analysis (DMA 242 E Artemis, NETZSCH). Accordingly, by using a tension clamp the samples were exposed to a constant strain (0.1 %) of increasing frequencies in the range of 0.1 to 10 Hz for 30 min, subsequently storage and loss modulus were recorded and Tan delta values were measured as a ratio of loss modulus to storage modulus. Furthermore, hydrodynamic diameter, poly-dispersity index (PDI), and zeta potential of various formulations were measured using a Zetasizer Nano ZS (Malvern Panalytical, UK). Accordingly, prepolymer solutions (with similar drug/polymer ratio as in fibres) were diluted to 100 ng ml.^-1 in ultrapure water for analysis of the zeta potential and hydrodynamic diameter, and poly- dispersity index (PDI).

Release studies from drug loaded fibres

To measure the drug release from the fibres (either DOX or GEM), 30 cm of drug loaded fibres were immersed in 1 ml of SBF solution and the supernatant was collected at certain time points and replaced with fresh SBF. For measuring the DOX release, fluorescence properties of DOX was measured (excitation 485 nm, emission 520 nm) using FLUOstar OPTIMA microplate reader. The absorbance values were eventually converted to concentrations using a previously achieved calibration curve for DOX. For measuring the GEM release, a high-performance liquid chromatography (HPLC, Agilent 1260 infinity) was used. Briefly, Samples were filtered through a 02 pm syringe membrane filter unit before being injected (10 pL) onto a ZORBAX Eclipse Plus column (4.6 c 100 mm, 5 pm particle size). Using an isocratic elution with a water/acetonitrile (95/5), draw and eject speed of 200 pL/min, pressure 300 bar, gemcitabine was detected with the UV-Visible detector at 272 nm. The absorption values were converted to concentrations using a previously observed calibration curve. The release results were plotted as mean value of three repeated tests. In-vitro cell studies

All cell lines were authenticated using short tandem repeat (STR) profiling at the Garvan Institute of Medical Research. Cells (MIA PaCa-2 or PANC-1) were cultured in DMEM-High glucose media containing 10% fetal calf serum (FCS) at 37 °C, 95% humidity and 5% CO2 in a Heracell incubator (Kendro Laboratory Products, Germany). When 80% confluence was reached, cells were detached by incubation with 5 mM trypsin/EDTA and harvested after centrifugation in a Heraeus Megafuge 1.0 (Thermo Scientific, USA) at 1200 rpm for 5 min at RT. Cells were resuspended in media, and viable cells counted using a hemocytometer and trypan blue staining. Cells were confirmed free of mycoplasma contamination. Mia Paca-2 and Panc-1 cells were seeded (8 x 104, 6 x 104, and 4 x 104 cells/well for 24, 48, and 72 hr time points, respectively) in 24-well flat bottomed plates in complete media (1 mL) containing 1% penicillin/streptomycin and kept in the incubator for 24 h prior to addition of empty or drug loaded fibres (1 cm lengths). After each time point the fibres were removed from the wells and subsequently 40 pL of MTS agent was added and allowed to react with the cells for 3 hrs. The absorption of each solution was measured at 490 nm on a microplate reader (SPECTRA max, PLUS). For live cell staining, at each time point the media was removed from the wells and subsequently the cells were washed with PBS. Next 500 pL of PBS solution (Containing 2.5 pL of Calcein AM and 1 pL of Propidium Iodide) was added to the wells and the plates were incubated for 15 min, after which the staining mixture was removed from the wells and replaced with fresh PBS. Immediately after the IncuCyte ZOOM system (Essen BioScience, USA) at 10 x magnification, with green (live cells) and red (dead cells) filters, was used to image the cells.

To confirm the long-term therapeutic elfect on inhibition of cancer cell growth, the effect of the medium supernatant of the drug loaded or empty fibre on the viability of the human pancreatic cancer cell lines, MIA PaCa-2, PANC-1 , and BxPC3 cell lines, was assessed using a cell counting kit-8 method (Sigma, St. Louis, MO, USA) based on manufacturer’s guideline. The BxPC3 cells were cultured in DMEM (Gibco), while PANC-1 and MIA PaCa-2 cells were cultured in RPMI-1640 medium (Gibco) supplemented with 10% v/v fetal bovine serum (Gibco), 100 U/mL penicillin, and 100 pg/mL streptomycin (Invitrogen) at 37 °C in 5% CO2 in a humidified incubator. Cells were plated at a density of 3 x 103 cells per well in a 96-well culture plate and incubated for 24 h before drug treatment. To assess the therapeutic effect of the drug released from the fibres, the fibres (1 cm) incubated in complete culture media at 37 °C in 5% CO2 in a humidified incubator and then the supernatant was collected at determined intervals (1 -, 5-, 8, 11 -, 14-, 17-, and 20-days). At day 1 after cell seeding, the media was changed to corresponding supernatants. Cell viability at each time point was determined using a cell counting kit-8 assay (CCK-8).

In-vivo Studies

To confirm the in vivo effects of the CS-fibres, a subcutaneous tumour model was established using MIA PaCa-2 and BxPC3-Luc cells. MIA PaCa-2 and BxPC-3-Luc cells were harvested (2x106) and resuspended in PBS mixed with Matrigel (1 :1 ratio). To prepare the xenograft tumours were developed in 6-week-old male nude mice by injecting MIA PaCa-2 and BxPC-3-Luc cells subcutaneously into the right posterior flank of mice, respectively. All mouse experiments within the guidelines of the protocol were reviewed by the Institutional Animal Care and Use Committee of Asan Institute for Life Science. Tumour growth was recorded twice a week in three dimensions using a digital caliper. Tumour volume was calculated as [(length x width x height) / 2] and reported in mm³. Tumours were grown in for 11-12 days until average tumour volume 100mm³. Mice were randomly divided into 4 groups at day 11 or 12 after subcutaneous cancer-cell inoculation; Group I: implanted with a drug- free CS fibres; Group II: implanted with gemcitabine loaded CS fibres; Group III: implanted with doxorubicin loaded CS fibres; and Group IV: implanted with gemcitabine and doxorubine loaded CS fibres. For MIA PaCa-2 tumour bearing mice three extra groups were added where equivalent dose of I) doxorubicin or II) gemcitabine or III) dual drugs (doxorubicin+gemcitabine) in saline solution were once injected into tail vein of animals. All experimental groups started with n=6 mice at the time of treatment initiation. To insert the CS fibres in the mouse, CS fibres (12mg each) were cut into 5mm length. The CS fibres were sterilized in 70% alcohol solution and then washed in the PBS solution twice. After washing, the fibres have adhesiveness. Wetted fibres were directly attached under the solid tumour. Tumour growth was observed over a period of 2 weeks. The therapeutic effect and anticancer activity of local drug delivery were determined by tumour volume and tumour weight change. Also, tumour growth was estimated using an in vivo imaging system (IVIS) of luciferase transduced BxPC-3 cell lines. The BxPC-3-Luc cells used in this experiment expressed a bioluminescent signal, and the correlation between cell numbers and bioluminescent signals was confirmed before the experiment.

Optical imaging of the subcutaneous pancreatic tumour

Mice were intraperitoneally injected with D-luciferin (0.3 mg; Perkin Elmer Inc.) Whole body luminescence imaging with an IVIS Spectrum (Caliper Inc., Alameda, CA) was performed every 2 minutes until radiance values reached the maximum. The region of the interested (ROI) level was measured with the radiance (photons/s/cm²/sr) using an analysis program, Living Image 4.4 (Caliper Life Sciences, PerkinElmer Inc.). Immunohistochemical staining: After sacrificing the mice at day 14, the tumours were removed and fixed in 4% neutral buffered paraformaldehyde and embedded in paraffin. Paraffin blocks were cut into 4 pm sections and were reviewed histologically after hematoxylin and eosin staining. Paraffin sections were deparaffinized and then rehydrated. After microwave antigen retrieval, non-specific binding sites were blocked with PBS containing 10% normal goat serum. The sections were further incubated with the primary antibodies against Ki-67 (Dako, Glostrup, Denmark). The samples were mounted using Prolong Gold antifade mountant with DAPI (Thermo Fisher Scientific, Massachusetts, USA). Images were obtained using an EVOS-fluorescence microscope (Thermo Fisher Scientific, Massachusetts, USA).

In vivo toxicity

Retro-orbital blood collection was performed for hematology determinations in tubes with anticoagulants (EDTA-2 K) at day 3 and 14. Hematology determinations included white-blood-cell (WBC) count and differential leucocyte count (neutrophils, lymphocytes, monocytes) using an Advia 120 Hematology Analyzer (Bayer Healthcare, Myerstown, PA, USA).

Statistical analysis

Statistical significance was determined using a two-way ANOVA with a Bonferroni post-test (GraphPad Prism V 6.0; San Diego, CA, USA). P values < 0.05 were considered statistically significant. Values are reported as the average ± standard deviation.

Results and Discussion

Chemical characterization of polymers

For composite fibre fabrication, alginate-dopamine (theoretical substitution degree of 50 % - in terms of alginate repeating units) was first synthesis using standard carbodiimide chemistry. Alginate- methacrylate was then synthesized using methacrylic anhydride. Fourier-transform infrared spectroscopy (FTIR) and proton nuclear magnetic resonance (1 H NMR) were employed to investigate successful synthesis of alginate-dopamine and alginate-methacrylate. Accordingly, FTIR spectrum of both alginate-dopamine and alginate-methacrylate showed the characteristic peaks of alginate associated with C-0 stretching vibration (1030 cm^-1), COO- symmetric and asymmetric stretching vibration (1415 and 1600 cm ¹, respectively), and -OH stretching vibrations (3325 cm ¹). More specifically, alginate-dopamine spectrum showed appearance of peaks at 1544 cm ¹, and 1220 cnr ¹assigned to the N-H deformation and C-N stretching vibrations. In addition, significant reduction in intensity of COO- bands (1415 and 1600 cm ¹) further indicated the successful attachment of dopamine to the alginate back bone via carbodiimide chemistry. Successful synthesis of alginate-dopamine was further confirmed using H-NMR via appearance of catechol protons at around 7 ppm Further, the relative integration of catechol protons (6.6-7 ppm) to anomeric protons of the glucose ring of alginate (4.9 ppm) were used to determine the degree of substitution which was measured to be 16 % (in terms of alginate repeating units). It should be noted that alginate-dopamine was synthesized with the designated degree of substitution (theoretical substitution degree of 50 %) based on preliminary release experiments using a model dye; fluorescein sodium salt, which revealed that degree of substitution of alginate-dopamine have a significant effect on the dye release profile and total amount of dye released.

Along similar lines, alginate-methacrylate was characterized via FTIR spectroscopy. It exhibited -CH stretching vibration band (2980-2850 cm ¹) and the appearance of two shoulders one around 1712 cnr¹and the other at 1161 cm ¹, associated with C=0 and C-0 groups of the esters resulting from the grafting of the methacrylate units. Additionally, H-NMR also proved the grafting of methacrylate groups by appearance of double peaks (vinyl) in the double bond region (5.5-6.5 ppm). Furthermore, the relative integration of methyl protons (of methacrylate groups) to anomeric protons of the glucose ring of alginate (4.9 ppm) were used to determine the degree of substitution which was 40% (in terms of alginate repeating units).

Rheology of polymers To evaluate spinability of the synthesized hydrogels, their rheological properties were investigated using a rheometer (Figure 2a & 2b). Accordingly, modification of alginate with dopamine increased the viscosity ol the solution (Figure 2a-iii) at lower shear rates when compared to pure unmodified alginate (Figure 2a-iv), likely due to noncovalent interactions including p-tt stacking of dopamine moieties. Also, a more pronounced shear thinning effect was observed for alginate- dopamine samples over the entire shear rate range, when compared to that of pure unmodified alginate. This could possibly be associated to potential effect of attached dopamine in further hindering the entanglement of neighbouring alginate polymeric chains. Furthermore, modification of alginate with methacrylate groups led to a significant decrease in viscosity of the resulting solution in comparison to pure alginate (Figure 2a-i), a phenomenon attributed to disruption of the alginate backbone by the methacrylation chemistry. For composite fibre fabrication, effective cross-linking of the core component assures stability of the fibres. Consequently, alginate-dopamine was physically mixed with alginate- methacrylate (75/25% w/w) to allow the formation of a dual cross-linked core constituent by combining chemical and physical crosslinking methods. As a result, the viscosity of the mentioned alginate- dopamine/alginate-methacrylate solution was also assessed (Figure 2a-ii) and the result showed a similar trend to that of alginate-dopamine solution (Figure 2a-iii), yet with a slightly lower viscosity over the entire shear rate range, associated with small portion of alginate-methacrylate in this solution (25% w/w). Furthermore, oscillatory experiments (as a function of time) revealed that upon UV irradiation (365 nm for 5 sec) the storage modulus of the alginate-methacrylate hydrogel increased significantly from lower values (~ 1 Pa) up to 13000 Pa, indicating the formation of methacrylate bonds in the cross-linked hydrogel (Figure 2b-i). Subsequently, the alginate-dopamine/alginate-methacrylate mixture showed a much lower increase in the storage modulus (levelling off at around 3500 Pa) in correspondence with the low volume of UV cross-linkable element in the mixture (Figure 2b-ii). Morphology of the core-shell fibres

Next, the core-shell fibres (CS) were fabricated through a custom-made coaxial nozzle based on a previously described wet-spinning method (see Fabrication of Coaxial Wet-Spun Graphene-Chitosan Biofibers, Col 18, 2, 2016, pg 284-293). Subsequently, the core-shell morphology of the CS fibres was further assessed using optical microscopy and scanning electron microscopy (SEM) (Figure 2c-d & Figure 3). The core-shell boundaries of the doxorubicin-loaded fibres could be clearly observed in these images (Figure 2c-d), as indicated by higher density of red colour (associated with DOX) in the middle region as compared to the adjacent area. The CS fibres were then allowed to dry off and their microscale morphology upon swelling was further investigated using SEM imaging. The core-shell borders of the fibres were only detectable once the fibres started to swell (Figure 3c-h). Accordingly, the fibres demonstrated a time-dependent swelling behaviour as it took them almost 60 hr to completely swell. Initially, after 12 hr of swelling in simulated body fluid (SBF), fibres possessed an almost pentagonal structure with extra folding at the corners (Figure 3c-d). Remarkably, the shell showed a more compact structure with smaller pores, as compared to the core, which was a consequence of dual crosslinking mechanism (ionic and covalent bonds) that was imposed on alginate-methacrylate in the shell region. This pore size difference between the core and the shell could also be attributed to limited diffusion of CaCh crosslinking agent into the core, as well as, the higher concentration of alginate-methacrylate in the shell compared to the core. Nevertheless, after 36 hr of swelling the fibres morphology transformed from a pentagonal structure to a semi-circular structure, but still the shell held a less porous morphology when compared to the core region (Figure 3e-f). Finally, after 60 hr of swelling the fibres had one last transformation in their morphology to become totally circular, with core region showing larger pore sizes as compared to previous time points (Figure 3g-f). Most remarkably, even after the final stage of swelling the fibres still had a less porous structure in the shell when compared to the core section, which further justified the role of shell layer as a diffusion barrier. Of note, the multistep evolution of CS fibres upon swelling was not observed for single fibres made from unmodified alginate (Alg), as single fibres simply reached their maximum swelling after only 2 hr of immersion in SBF.

Swelling and degradation of the core-shell fibres

Considering the direct correlation between the drug release profile, the swelling behaviour, and degradation of the hydrogels, we have measured the swelling ratio and degradation rate of the fibres in SBF (Figure 4). As a result, CS fibres showed to have significantly less swelling (18.05) when compared to alginate fibres (33.07) as shown in Figure 4a, which could be due to dual cross-linked network in the composite fibres. The gradual swelling behaviour of CS fibres (loaded with DOX in the core) in SBF was also recorded over a span of 60 hr. The obtained swelling data for CS fibres in this study were almost 20 times less than the value reported for similar fibrous hydrogel structures. In addition, the degradation rate results revealed that CS fibres possessed a slower mass loss rate over a period of 14 days compared to that of alginate fibres (Figure 4b). More specifically, after 14 days alginate fibres experienced a 63.6 ± 9.1 % of mass loss, while CS fibres showed a 43.9 ± 3.6% of mass loss. This is possibly due to the dual cross-linked hydrogel network in the CS fibres.

Mechanical properties of the core-shell fibres

Given the effect of mechanical properties on release of drugs from hydrogels, the tensile properties of the fibres in both static and dynamic conditions was considered (Figure S4-supporting information and Figure 5, respectively). Initially, mechanical properties of the fibres were assessed using a static tension test (both in dry and wet states) and compared them to that of alginate fibres. In dry state, CS fibres showed higher modulus (2.78 ± 0.03 GPa) and tensile strength (85.44 ± 9.6 MPa) compared to those for single alginate fibres (1 26 ± 0.08 GPa and 68.5 ± 2.6 MPa, respectively). These observations were more pronounced for the fibres in the wet state, as CS fibres had a modulus of 0.21 ± 0.07 MPa and tensile strength of 0.64 ± 0.02 MPa, while single Alginate fibres possessed a modulus and tensile strength of 0.06 ± 0.01 MPa and 0.07 ± 0.02 MPa, respectively. This implied that the proposed core-shell structure provides greater stability in aqueous environments, making them more practical for biomedical applications. It is worth noting that in this study the elastic modulus of wet CS fibres (0.21 MPa) were almost 10 times higher than that observed for fibres produced by microfluidic systems (0.02 MPa). In order to gain a better insight into viscoelastic properties of the fibres, an oscillating tension test (at a constant strain of 0.1%) in a frequency range of 0.1-10 Hz was conducted (Figure 5). Remarkably, CS fibres showed a significantly higher storage modulus (2996 ± 78.1 MPa), over the entire frequency range, when compared to that of single alginate fibres (261 .45 ± 5.4 MPa) (Figure 5a). For both fibres (Alg and CS) the storage modulus values were higher than the loss modulus values, which indicated the domination of elastic response over a viscous behaviour for the two types of samples. However, after measuring the damping factor (Tan d - calculated from the ratio of E7E^') for the fibres, CS fibres demonstrated a significantly lower value (0.031 ) than that of single alginate fibres (0.053), which was associated with higher elasticity of CS libres over single alginate fibres (Figure 5d). Consequently, the observed mechanical properties data suggested that CS fibres possessed a more robust structure (as compared to single alginate fibres), probably due to the dual cross-linked network (chemical and physical) arrangement. Also, the CS fibres were robust enough to be fabricated into a woven structure (using a traditional weaving technique) as a potential 3D implantable patch.

Drug release from the core-shell fibres

Next, in vitro drug release studies were carried out using two specific chemotherapeutic drugs (GEM and DOX) and the results are shown in Figure 6. Specifically, the release of drugs from core shell fibres (CS) was compared to release obtained from single fibres made from unmodified alginate (Alg), as well as core-shell fibres in which both core and shell were made entirely out of UV- crosslinkable alginate-methacrylate (CS·) without dopamine. Of note, the comparison between the normal CS fibres and the specially made CS· fibres, was for the purpose of understanding the specific role of alginate-dopamine (in the core of CS fibres) in the obtained drug release profiles. Consequently, Figure 6a shows the GEM release data from the above-mentioned fibres. Accordingly, CS fibres was able to slow down the release of GEM down to almost 40% when compared to that of alginate fibres (~ 90%). In addition, CS fibres released a much higher amount of GEM (7.6 mg) when compared to that of single alginate fibres (2 mg). This difference could be a consequence of loss of drug in the fabrication process (specifically in the spinning bath), which was further prevented in the CS fibres owing to their specific composition and their coaxial morphology. Similarly, the CS fibres had a slower GEM release profile when compared to CS· fibres (~ 80 %). However, the CS fibres released less amount of GEM (7.6 mg) when compared to CS · fibres (11 g). This could be due to physical entrapment of GEM in the CS fibres which might have been facilitated by intermolecular interactions between GEM and the dopamine moiety in the CS fibres. Nevertheless, release of GEM from CS fibres reached a plateau level after 10 hr. Next, release of DOX from the fibres was evaluated (Figure 6c). Accordingly, DOX release profile from CS fibres exhibited three distinct stages: starting with a burst release of the drug over the first 10 h (-35%), followed by a slower release of the drug up to day 8 (-53%), and then an even slower release rate that lasted up to day 19 (60 %), which was followed by a plateau region up to day 21 (60.5 %). This was in total contrast to single alginate fibres releasing almost the entire DOX in the first 10 h of the experiment. Of note, the DOX release profile obtained from these CS fibres is comparable to other studies in the field. Moreover, cumulative DOX release values revealed that CS libres could encapsulate twice as much DOX (-0.27 mg) in comparison to single alginate fibres (- 0.15 mg). In a similar manner, the CS fibres demonstrated a slower release of DOX when compared to that from CS· fibres, however, in comparison CS· fibres released a slightly higher amount of DOX (~ 0.34 mg). This phenomenon was also observed with GEM release studies which we speculated to be a result of physical entrapment of the drugs in CS fibres due to intermolecular drug-dopamine interactions.

Consequently, to investigate this hypothesis and with the aim of scrutinizing the obtained drug release profiles, a dynamic light scattering (DLS) analysis was conducted (Table 1). Correspondingly, the measurements were carried out for pure unmodified alginate, as well as, dopamine-modified alginate (main component of the core section in CS fibres), and methacrylate-modified alginate (main component of the core section in CS· fibres), to better understand their interactions with the drug molecules. The results revealed that modification of alginate with dopamine led to an increase in hydrodynamic diameter of the polymer from 110.9 ± 0.24 nm to 192.8 ± 1 .83 nm, which was attributed to addition of catechol moieties and possible aggregation of polymeric chains due to tt-p stacking of aromatic rings. Modification of alginate with methacrylate groups also led to an increase in hydrodynamic diameter of the polymer to 205.7 ± 0.40. Compared to pure alginate and alginate- methacrylate, alginate-dopamine had a higher negative zeta potential value (-33.2 ± 0.86 as compared to -28.8 ± 0.36 and -29.9 ± 0.78, respectively), that was ascribed to acidic nature of catechol groups. Of note, addition of drugs (DOX or GEM) to pure unmodified alginate or alginate-methacrylate did not cause any major change in their corresponding hydrodynamic dimeters or zeta potentials, implying the weak interactions of these drugs with pure alginate and alginate-methacrylate, which translated into quick release of these two drugs from those fibres. It should also be noted that slightly slower release of drugs from CS· compared to alginate fibres, could have been due to smaller porosity of these fibres indicated from their cross-sectional SEM images. On the other hand, incorporation of drugs into alginate-dopamine was associated with remarkable changes in hydrodynamic dimeter and zeta potential of this polymer. Specifically, hydrodynamic dimeter of alginate-dopamine increased significantly from 192.8 ± 1 .83 nm to 422.0 ± 0.24 and 388.0 ± 0.35 nm, after addition of DOX and GEM, respectively. The zeta potential value of alginate-dopamine increased from -33.2 ± 0.86 mV to -37.2 ± 0.69 and -34.0 ± 0.69 mv, after addition of DOX and GEM, respectively. The observed data suggested that both drugs could possibly had non-covalent hydrophobic interactions (tt-p interactions) with alginate-dopamine which led to significant increase in hydrodynamic diameter values. However, DOX seemed to also have better ionic interactions with alginate-dopamine, as the value of zeta potential increased significantly (-33.2 ± 0.86 mV to -37.2 ± 0.69 mV) for this drug, whereas, addition of GEM did not generate such changes. Altogether, the obtained data showed that alginate-dopamine had better interactions (physical or chemical) with the drugs, when compared to pure unmodified alginate or alginate-methacrylate, a phenomenon which was also observed in drug release studies. Also, the results confirmed that DOX had stronger interactions with alginate- dopamine, as compared to GEM, which further justified the difference between release profiles of these two drugs from CS fibres.

In vitro biocompatibilitv of fibres

With the aim of evaluating anti-cancer efficacy of the drug loaded CS fibres, in vitro cellular experiments were carried out using two different human pancreatic cancer cell lines (MIA PaCa-2 & PANC-1 ) and the results are shown in Figure 7. The empty CS fibres (indicated as control) were shown to be biocompatible and maintained high cell viability for both cell lines. GEM loaded fibres were able to reduce the viability of the MIA PaCa-2 cells down to 26.6% ± 05, 72 h post treatment (Figure 7a). Remarkably, DOX loaded CS fibres were able to reduce the viability of cancer cell line as early as 24 h post treatment and they showed a higher anti-cancer effect (4.2 ± 3.7) after 72 hr of treatment (Figure 7b). Following the observed results the inventors loaded the two drugs (GEM+DOX) simultaneously in core section of the CS fibres, indicated as dual treated, to investigate any possible synergistic anti-cancer effect of this drug combination. Accordingly, the obtained results suggested that the dual loaded CS fibres were also effective in reducing the viability of cancer cells 72 h post treatment down to 13.8 + 1 .8 (Figure 7c). Lastly, a live/dead staining of the corresponding cells further proved that empty CS fibres (indicated as control) had a good level of biocompatibility after 72 hr of treatment as implied from high density of the live cells (shown in green), whereas all drug-containing CS fibres were capable of reducing the viable cells in correspondence with MTS cell viability results (Figure 7g).

Furthermore, in order to assess the anti-cancer effect of CS fibres on a more resilient pancreatic cancer cell line we have tested these fibres on PANC-1 human pancreatic cancer cell lines. The results showed a similar trend to MIA PaCa-2 cells lines, however herein the cells had a higher level of viability after 72 hr post treatment specially in cells treated with GEM loaded fibres (56.6 ± 4.7) as shown in Figure 7d. Otherwise, cells treated with DOX- and Dual-loaded CS fibres had 13.3 ± 2.6 and 12.6 ± 2.8, respectively (Figure 7e & f). At last, live/dead staining of these cells further verified the obtained MTS cell viability results with DOX- and Dual-loaded fibres having the least number of viable cells 72 hr post treatment (Figure 7h). Of note, the difference in sensitivity to chemotherapy is well documented for MIA-PaCa-2 and PANC-1 cells, and has been attributed to differences in ERK activity. Moreover, the significant difference between anti-cancer performances of GEM loaded fibres compared to DOX-containing fibres could be a result of controlled release of DOX from the fibres and also different mechanism of action of this drug compared to GEM. Overall, the in vitro results suggested that CS fibres possessed a high level of cell biocompatibility, while DOX and dual-loaded CS fibres had superior anti-cancer efficacy when compared to GEM loaded fibres. Moreover, in these short-term in vitro assessments, no statistical difference was observed between the DOX- and dual- loaded fibres. This could be associated with mechanism of action of gemcitabine, as it requires cellular uptake and intracellular phosphorylation. A process that is time-dependent, and also requires prolonged and high intercellular concentration of gemcitabine. This further validates the importance of long-term in vitro studies to evaluate the therapeutic effect of the drug loaded CS fibres. In vitro therapeutic effect of fibres

Next, with the aim of evaluating the long-term therapeutic eflect of CS fibres on inhibition of cancer cell growth, a 20 days cell counting assay (CCK-8) was conducted using three different human pancreatic cancer cell lines including MIA PaCa-2, PANC-1 , and BxPC3-Luc cells (Figure 8). Consequently, in MIA PaCa-2 cell lines the empty CS fibres (indicated as control) showed a high level of biocompatibility throughout the test period and did not hinder the cancer cell growth, whilst, drug- loaded fibres were capable of slowing down the cell growth to different degrees (Figure 8a). Accordingly, Dual-loaded fibres showed a higher degree of cancer cell growth prevention which lasted for almost 14 days before the cells became highly viable again. This was in contrast to DOX- and GEM-loaded fibres which showed a short term inhibition of cell growth lasting only for 5 days. Regardless, in correspondence with previous MTS cell viability results the DOX-loaded fibres had better inhibition of cancer cell growth when compared to the GEM-loaded counterparts. Next, PANC- 1 cells were tested and similar results were obtained where Dual-loaded fibres showed the most inhibition of cell growth that was followed DOX- and GEM-loaded fibres, respectively (Figure 8b). At last, BxPC3-Luc cells further verified the previous observations where Dual-loaded fibres prevented the cancer cell growth the longest which was followed by Dox-loaded and GEM-loaded fibres (Figure 8c). Worth noting that based on these results Dual-loaded fibres showed a notable synergistic anti cancer effect (as compared to GEM- or DOX-loaded fibres), which was not observed in the previous short-term MTS cell viability experiments. This further validates the importance of long-term in-vitro studies in the context of cancer drug delivery as these chemotherapeutic drugs tends to show their therapeutic effects only after weeks of treatment. Overall, the obtained results suggested that these drug loaded fibres can inhibit the cancer cell growth effectively for 14 days.

In vivo therapeutic effect of fibres

Next, in vivo experiments were carried out for a period of 2 weeks using a murine subcutaneous xenograft tumour model using two pancreatic cancer cell lines including MIA PaCa-2, and BxPC3-Luc cancer cells (Figure 9 & 10, respectively). Following tumour formation the drug loaded CS fibres were implanted under the tumour region, through a small incision in the middle of the back of the mice, and after two weeks the animals were sacrificed and their tumour size were measured. The obtained data showed that tumours treated with DOX- and dual-loaded (DOX+GEM) CS fibres showed significant inhibition of tumour growth compared to GEM-loaded or empty CS fibres. Accordingly, MIA PaCa-2 tumour volume increased 79 ± 83% and 28 ± 14% in mice treated with DOX-loaded and dual-loaded CS fibres, respectively, while mice treated with empty and GEM-loaded CS fibres experienced 323 ± 157% and 275 ± 21% increase in the tumour volume, respectively (Figure 9b). Moreover, slightly higher anti-cancer efficacy of dual-loaded fibres, as compared to DOX-loaded fibres, could have possibly been associated to synergistic effect of the two drugs. Most significantly, DOX- and dual- loaded CS fibres were able to inhibit the tumour growth to a much higher degree when compared to injection of equal dosages of DOX and DOX+GEM (277 ± 55% and 194 ± 49% increase in the tumour volume, respectively), which was a testament to importance of local delivery of drugs in a controlled manner. Of note, weak therapeutic effect from GEM loaded fibres was attributed to faster release of this drug from the fibres that prevented steady phosphorylation of GEM by cancer cells leading to its reduced therapeutic effect. Also, all treated mice did not experience any significant weight loss throughout the study, which indicated that all treatments were well tolerated and did not cause any major toxicity (Figure 9d). To further evaluate the anti-cancer effect of drug loaded CS fibres, H&E and Ki67 (proliferating cells marker) staining of tumour tissues post-treatment were evaluated (Figure 9e). In correspondence with the observed results, the DOX- and dual-loaded fibres yielded significant levels of apoptosis as confirmed by H&E staining of tumour tissues post-treatment, which suggests that the level of drugs to which tumours were exposed were sufficient to shrink the tumours and prevent their growth. Similar results were observed from Ki67 staining of the tumour tissues, as tumours treated with empty or GEM loaded fibres showed the highest density of red pigments (as opposed to DOX- or Dual-loaded fibres), indicating the higher population of proliferating cells in tumours treated with these two samples. Furthermore, with the aim of evaluating the cytotoxicity of empty or drug loaded CS fibres, various parameters in the blood of the mice were analysed (Table 2). The obtained results for mice treated with drug loaded or empty fibres showed no significant difference from the haematological values (i.e. white blood cell, red blood cell, hemoglobin, platelet, neurophil and lymphocyte count) of a normal mouse, suggesting that CS fibres did not produce any significant changes in the haematology of mice, implying that CS fibres may be a safe delivery platform for in vivo use.

In order to further assess the anti-cancer performance of drug-loaded CS fibres, a second series of animal studies were conducted using murine subcutaneous xenograft tumour model using BxCP3-Luc light-producing pancreatic cancer cell lines (Figure 10). Similarly, BxCP3-Luc tumour volume increased 77 ± 53% and 47 ± 46% in mice treated with DOX-loaded and dual-loaded CS fibres, respectively, whilst mice treated with empty and GEM-loaded CS fibres experienced 622 ± 589% and 376 ± 138% increase in the tumour volume, respectively (Figure 10b). In these tumour models after 14 days of treatment, GEM-loaded fibres were able to slightly decrease the tumour volume when compared to empty fibres, which was also observed in luciferase images of the growing tumours. The slight difference between BxCP3-Luc and Mia-PaCa-2 tumours responses to GEM-loaded fibres was associated to higher sensitivity of BxCP3-Luc cell lines to gemcitabine. Also, similar to MIA PaCa-2 tumour models, in the BxCP3-Luc tumours dual-loaded fibres failed to show any significantly higher anti-cancer efficacy, when compared to DOX-loaded fibres, which could have been a result of quick release of gemcitabine from these fibres. Additionally, animals treated with empty or drug-loaded fibres maintained their body weight for the whole duration of the study, which showed that these fibres did not cause a systemic toxicity and could potentially serve as a safe in vivo treatment modality (Figure 10d). Lastly, H&E and Ki67 staining of excised tumour tissue, revealed that empty fibres did not cause apoptosis in tumour tissue and they have not interfered with proliferation of cancerous cells, whilst DOX- and Dual-loaded fibres caused the most apoptosis of cancer cells and left the least amount of proliferating cells (Figure 10e). Overall, the in vivo findings showed that CS fibres are well tolerated in the animals and they did not cause any major toxicity. Also, in MIA PaCa-2 tumour models it was shown that drug loaded fibres (especially DOX- and Dual-loaded fibres) had better anti-cancer performance when compared to intravenous injection of the equal dosages of the drugs. In addition, in both tumour models, it was observed that continuous and slow release of DOX from the CS fibres can inhibit the tumour growth more effectively (DOX- and Dual-loaded fibres), when compared to quick release of GEM from CS fibres. The results also showed that dual-loaded fibres could also be effective in inhibiting the tumour growth.

Conclusions

In this study, drug loaded composite fibres were fabricated through a wet-spinning process for the first time, where the core consisted of alginate-dopamine and the shell contained UV cross-linkable alginate-methacrylate with a view to improving both the drug affinity and diffusion, and facilitate high loading capacity and controlled release to function as novel implantable drugs carriers. In this study, coaxial biofibres loaded with chemotherapeutic drugs were fabricated for the first time, where the core consisted of a mixture of UV cross-linkable cross-linkable alginate-methacrylate and alginate- dopamine and the shell contained UV cross-linkable alginate-methacrylate. This was done with a view to improve both the drug affinity and diffusion, and facilitate high loading capacity and controlled release to function as novel implantable drug delivery system. It is proposed that the efficient intermolecular interactions between the chemotherapeutic drugs (DOX and GEM) and catechol moieties of the DOPA in the core conferred high drug loading capability and specifically imparted excellent control on the release of DOX over a period of 21 days. These CS fibres elicited no significant undesired toxicity as evidenced from both in vitro and in vivo experiments. DOX- and dual-loaded (DOX+GEM) fibres achieved significant anti-cancer effect on human pancreatic cancer cells- xenografted tumour models, while GEM loaded fibres showed minimal anti-cancer effect at doses much higher than DOX. Most remarkably, in MIA PaCa-2 tumour bearing mice, Dox- and dual-loaded CS fibres were able to significantly prevent the tumour growth (79% and 28% increase in tumour volume, respectively), when compared to intravenous injection of equivalent dosages of the same drugs (277% and 194% increase in the tumour volume, respectively). Enhanced inhibition of tumour growth in DOX-containing fibres was attributed to prolonged release of the drug from these fibres, and considering the systemic side effects of DOX, the CS fibres could be beneficial for local delivery of this drug and subsequently reducing its cardiotoxicity. Lastly, in this study we did not observe any significantly higher anti-cancer efficacy from dual-loaded fibres, as compared to DOX-loaded fibres, which could be associated with quick release of GEM from the fibres. In conclusion, this study showed that CS coaxial biofibres, as a novel implantable platform, provided an effective and safe strategy towards local delivery of drug cocktails for cancer therapy. Table 1 data supports improved interactions of alginate-dopamine with the drugs in compared to alginate alone or alginate-methacrylate. Table 1. Hydrodynamic dimeter, poly-dispersity index (PDI), and zeta potential of various formulations (n = 5, mean ± SD). Showing the specific interactions of alginate-dopamine with the drugs as compared to weak interactions in pure alginate.

Throughout this application, the term “about’ is used to indicate that a value includes the inherent variation of error for the device, the method being employed to determine the value, or the variation that exists among the study subjects. The indefinite articles “a” and "an,” as used herein in the specification, unless clearly indicated to the contrary, should be understood to mean “at least one.” Where the terms "comprise", "comprises", "comprised" or "comprising" are used in this specification (including the claims) they are to be interpreted as specifying the presence of the stated features, integers, steps or components, but not precluding the presence of one or more other features, integers, steps or components, or group thereof.

Substitute Sheet (Rule 26) RO/AU

Claims

Claims

1 . A therapeutic implantable device, the device comprising : at least one composite fibre comprising: at least one core component comprising at least one swellable modified hydrophilic hydrogel polymer, and at least one shell component, wherein at least a first shell component comprises at least one swellable modified hydrophilic hydrogel polymer, wherein at least one of the polymers of the core are modified with at least one crosslinkable alkene functional group and at least one of the polymers of the at least one core are modified with at least one hydroxylatedphenyl functional group to form a crosslinkable alkene-modified- hydroxylatedphenyl-modified polymer, and wherein, at least one of the polymers of the shell are modified with at least one crosslinkable alkene functional group to lorm a crosslinkable alkene-modified polymer; and wherein, the at least one core component is encapsulated by the at least one shell component.

2. The device according to claim 1 , wherein one or more of the cores and at least a first shell component comprises alginate or chitosan.

3. The device according to claim 1 or claim 2, wherein composite fibre is in the form of a coaxial fibre or a multiaxial fibre.

4. The device according to any one of the preceding claims, wherein the crosslinkable alkene- modified-hydroxylatedphenyl-modified polymer is in the form of a mixture of crosslinkable alkene- modified polymer and hydroxylatedphenyl-modified polymer.

5. The device according to claim 4, wherein the crosslinkable alkene-modified- hydroxylatedphenyl-modified polymer and crosslinkable alkene-modified polymer in the mixture are present in a ratio of from 50:50 to 85:15.

6. The device according to any one of claims 4 or 5, wherein the crosslinkable alkene-modified- hydroxylatedphenyl-modified polymer and crosslinkable alkene-modified polymer in the mixture are present in a ratio of 75:25.

7. The device according to any one of the preceding claims, wherein the crosslinkable alkene- modified-hydroxylatedphenyl-modified polymer is in the form of a single polymer simultaneously modified with the crosslinkable alkene functional groups and the hydroxylatedphenyl functional groups.

8. The device according to any one of the preceding claims, comprising one core component and a first shell component.

9. The device according to any one of the preceding claims, wherein the crosslinkable alkene functional groups are acrylate groups derived from methacrylic acid, ethylacrylic acid, propylacrylic acid or combinations thereof.

10. The device according to any one of the preceding claims, wherein the hydroxylatedphenyl groups comprises a dihydroxylated phenyl group is derived from a polyphenol or a catechol group, for example, derived from a functional group comprising 1 ,2-dihydroxybenzene, for example, 3,4- dihydroxyphenylalanine (DOPA).

11. The device according to any one of the preceding claims, wherein the core component polymers comprises an acrylate-catechol modified hydrophilic hydrogel polymer, and the shell component comprises an acrylate-modified hydrophilic hydrogel polymer.

12. The device according to any one of the preceding claims, wherein one or more of the core component and the shell component is alginate or chitosan.

13. A therapeutic implantable device, the device comprising: at least one composite fibre comprising: a core component comprising at least one hydrophilic hydrogel polymer selected from alginate or chitosan, and a shell component, a first shell component comprising at least one hydrophilic hydrogel polymer selected from alginate or chitosan, wherein the hydrogel polymers of the core are modified with at least one crosslinkable alkene functional group and the polymers ol the core are modified with at least one hydroxylatedphenyl functional group to form a core comprising crosslinkable alkene-modified-hydroxylatedphenyl-modified hydrogel polymers, and wherein at least one of the hydrogel polymers of the shell are modified with at least one crosslinkable alkene functional group to form an crosslinkable alkene-modified polymer, wherein the core component is encapsulated by the shell component.

14. A therapeutic implantable device comprising: at least one composite fibre comprising: a core component comprising a hydrophilic hydrogel polymer which is alginate, and a shell component comprising a hydrophilic hydrogel polymer which is alginate, wherein the hydrogel polymer of the core are acrylate-catechol-modified alginate polymers, and the hydrogel polymer of the shell are acrylate modified hydrogel alginate polymers, and wherein the core component is encapsulated by the shell component.

15. The device according to claim 14, wherein the acrylate-catechol-modified alginate polymers and acrylate-modified alginate polymers and in the mixture are present in a ratio of 50:50 to 85:15, most preferably 75:25.

16. A therapeutic implantable device, the device comprising: at least one composite fibre comprising: a core component comprising alginate, and a shell component comprising alginate, wherein the core comprises a mixture of alginate modified with at least one methacrylate functional group and alginate modified with at least one 3,4-dihydroxyphenylalanine (DOPA) group to form a methacrylate-DOPA modified alginate in the form of a polymer mixture comprising methacrylate-modified alginate and DOPA-modified alginate in a ratio of 75:25, and wherein alginate of the shell is modified with at least one methacrylate functional group to form an acrylate modified alginate, and wherein the core component is encapsulated by the shell component.

17. The device according to claim 16, wherein the DOPA-modified polymer has a theoretical substitution degree of from about 5% to 32% in terms of polymer repeating units.

18. A device according to any one of claims 16 or 17, wherein the acrylate-modified polymer has a theoretical substitution degree of from about 20% to 50% in terms of polymer repeating units.

19. A device according to any one of the preceding claims, wherein one or more of the core component and at least one of shell component is loaded with at least one therapeutic agent.

20. A device according to any one of the preceding claims, further comprising at least one therapeutic selected the group of hydrophilic drugs consisting of: doxorubicin, paclitaxel, camptothecan, everolimus, docetaxel, and any combination thereof, gemcitabine, nivolumab, a platinum agent, 5-fluorouracil, irinotecan, cyclophosphamide, rituximab, cetuximab, trastuzumab, pertuzumab, sunitinib, bevacizumab, an anti-EGFR molecule, an anti-CTLA4 antibody, an anti-PD1 or anti-PDL1 antibody or inhibitor, an immune modulating agent such as siRNA, microRNA, RNAi, or antigens for immunotherapy, and any combination thereof.

21 . A device according to any one of the preceding claims, wherein the core component comprises acrylate-catechol-modified alginate loaded with doxorubicin or gemcitabine.

22. A device according to any one of the preceding claims, wherein the core component comprises acrylate-catechol-modified alginate loaded with doxorubicin and gemcitabine

23. A therapeutic implantable device of any one of the preceding claims, wherein the acrylate groups are converted to crosslinks between strands of polymer, preferably by light, most preferably UV light.

24. A therapeutic implantable device of any one of the preceding claims, wherein a metal cation forms ionic crosslinks between strands of polymer.

25. A device according to any one of the preceding claims, in the form of a patch or a web.

26. A device according to any one of the preceding claims, further comprising at least one sheath encapsulating a least one composite fibre substantially therein, preferably comprising a plurality of apertures disposed along the length of the sheath in spaced apart arrangement.

27. A device according to any one of the preceding claims, wherein the fibre is formed using a method selected from the group consisting of 3D printing, wet spinning, electrospinning, coaxial melt extrusion printing, coaxial melt electro-writing, hot melt extrusion and pulsatile fibre spinning.

28. A method of preparing a therapeutic implantable device, the method comprising the steps of:

(i) preparing a first solution of a mixture of at least one swellable acrylate modified hydrophilic hydrogel polymer and at least one swellable catechol modified hydrophilic hydrogel polymer;

(ii) preparing at least one further solution of at least one swellable acrylate modified hydrophilic hydrogel polymer;

(iii) extruding the first and further solutions through a coaxial or multiaxial extrusion means; and collecting from the extrusion means, at least one coaxial or multiaxial fibre comprising the modified polymers.

29. A method of claim 28, wherein the extrusion step involves extruding both solutions through the extrusion means into a coagulation bath comprising a metal salt coagulation agent.

30. A method of preparing therapeutic implantable device, the method comprising the steps of: preparing a spinnable first solution of at least one swellable acrylate-catechol modified hydrophilic hydrogel polymer and at least one further solution of at least one swellable acrylate modified hydrophilic hydrogel polymer; and extruding the first and the at least one further spinnable solutions through a coaxial or multiaxial spinneret into a coagulation bath comprising a coagulation agent; and collecting from the coagulation bath, via a rotating mandrel, at least one coaxial or multiaxial fibre spun from the modified polymers.

31 . A method according to any one of claims 28 to 30, further comprising the step of activating the acrylate groups to crosslink polymer strands in the coaxial or multiaxial fibre.

32. A method according to claim 31 , wherein the acrylate groups are activated by light.

33. A method according to any one of claims 28 to 32, wherein one or more of the first solution and the at least one further solution comprise one or more therapeutics.

34. A method of any one of claims 28 to 33, wherein the swellable acrylate-catechol modified polymer forms a core component of the composite fibre and the swellable acrylate modified polymer forms a shell component of the composite fibre.

35. A method of any one of claims 28 to 34, wherein the hydrophilic polymers are alginate or chitosan.

36. A method of any one of claims 28 to 35, wherein the acrylate group is derived from methacrylate, and the catechol group is derived from DOPA.

37. A method of delivering at least one therapeutic to a subject, the method comprising the step of: implanting a therapeutic implantable device according to any one of claims 1 to 27 into a subject presenting with a medical condition that is treatable or preventable with the at least one therapeutic.

38. A method of treatment or prevention ol a disease or condition in a subject in need thereof, comprising the steps of: providing a therapeutic implantable device to the subject in need thereof, wherein the therapeutic implantable device is according to any one of claims 1 to 27.

39. A method of claim 38, wherein the disease or condition is cancer, particularly pancreatic cancer or breast cancer.

40. Use of therapeutic implantable device is according to any one of claims 1 to 27 in the treatment or prevention of a disease or condition.

41 . Use according to claim 40, wherein the disease or condition is cancer, particularly pancreatic cancer or breast cancer.

42. Use of a therapeutic implantable device for controlled release of a therapeutically effective amount of least one therapeutic in the treatment of a disease or condition, wherein the device is according to any one of claims 1 to 27.