WO2019166087A1 - Implantable nerve guidance conduit for nerve repair - Google Patents

Abstract

The invention relates to an implantable nerve guidance conduit comprising: a hollow tube, and optionally, an inner microstructure, characterized in that the hollow tube is made of a blend of biopolymers comprising a copolymer of poly (3-hydroxyoctanoate- co-3-hydroxydecanoate) and poly(ε-caprolactone). The inner microstructure can be present in the form of a microchanneled inner tube made of methacrylated polycaprolactone, or a bundle of threaded microfibres made of a blend of poly (3- hydroxyoctanoate) and poly (3-hydroxybutyrate). Said implantable nerve guidance conduit is useful for treating and reconstructing nerve injuries, in particular peripheral nerve injuries of up to 50mm.

Description

DESCRIPCION

IMPLANTABLE NERVE GUIDANCE CONDUIT FOR NERVE REPAIR FIELD OF THE INVENTION

The present invention concerns an implantable nerve guidance conduit (INGC) for nerve repair. In particular, the present invention is directed to a new and improved nerve guidance conduit useful for nerve regeneration in peripheral nerve injuries. The present invention is also directed to the method of obtaining such implantable nerve guidance conduit.

BACKGROUND PRIOR ART Post-traumatic peripheral nerve repair is one of the major challenges in restorative medicine and microsurgery. Primary causes of damage are traumatic accidents, tumor resection, iatrogenic side effects of surgery or repetitive compression (tunnel syndromes). At present, peripheral nerve injuries are a cause of medical consultation in more than 1 ,000,000 patients per year in the United States and Europe, with more than 100,000 cases undergoing surgery. Severe nerve injury has a devastating impact on patients’ quality of life. Typical symptoms are sensory and motor function defects that could result in complete paralysis of an affected limb or development of intractable neuropathic pain. Despite the progress in understanding the pathophysiology of peripheral nervous system injury and regeneration, as well as advancements in microsurgical techniques, peripheral nerve injuries are still a major challenge for reconstructive surgeons.

The surgical treatment for the complete severing of a nerve with small gap length (<5mm) and no loss of tissue is direct suturing of opposite nerve stumps. In this particular case, nerve co-aptation with fascicle alignment and tension-free suturing is feasible because peripheral nerves are phenotypically driven to regenerate spontaneously following injury. However, when direct suturing is not possible because it would cause tissue tension affecting nerve regeneration or when there is a discontinuity (> 1 cm) between both distal and proximal stumps, tissue engineering approaches are required. During nerve regeneration, axons grow randomly forming a nerve fibre mass called bands of Bungner. Unless surgically intervened, these regenerative sprouts will result in complete axonal degeneration affecting motor control and sensory perception. Nerve autografting is still the gold standard technique for nerve gap repair. Autografts are primarily taken from purely sensory nerves, since this allows the obtaining of longer grafts with lower donor-site morbidity than from motor or mix nerves, as the primary complication is often temporary localized numbness rather than a motor deficit. The most common donor source is the sural nerve, which allows for the harvest of up to 50 mm of nerve graft (up to 30 mm nerve gaps), with quite well- tolerated adverse effects ranging from sensory deficit around the lateral foot (9,1 - 41% of patients), to neuroma formation and unbearable pain (6,1 - 8,1 % of cases). Autograft has several disadvantages such as limited sources of donor nerve, the need for a second surgery to obtain the donor nerve, loss of nerve function in transplantation, and a lack of correspondence between the repaired nerve and the graft for the cross sectional area. According to these drawbacks, the success rate in patients treated with sural nerves is limited to 50%. Thanks to the progress in the field of tissue engineering, it appears increasingly possible to use artificial conduits for reconstruction of nerve gaps. Implantable nerve guidance conduits offer a promising alternative to conventional treatments, supporting and guiding the axons during their growth, while avoiding scar tissue infiltration in the gap. Fundamental requirements for effective nerve tissue regeneration demands a tubular scaffold that should be biocompatible, have sufficient mechanical stability during nerve regeneration, be flexible (with mechanical properties close to that of nerve tissues to prevent compression of the regenerating nerve), be porous to ensure supply of nutrients, and degrade at an appropriate time (after nerve regeneration) into nontoxic products to prevent long-term adverse reactions. Biodegradable materials, of either synthetic or natural origin, have been studied for INGCs production because they offer several advantages. The most interesting one is that biodegradation acts on biomaterial as a time-controlled elimination system as well as avoiding nerve compression and fibrosis formation. Between them, synthetic polymers show highly tunable possibilities, as variations in their chemical or materials properties may change biocompatibility, degradation behaviour, flexibility, porosity, and mechanical strength. On the other hand, natural-derived materials show good biocompatibility, although most lack adequate mechanical strength and water stability and thus need cross linking. To date, several single tubes developed using synthetic or natural polymers have shown to be capable of physically guiding the linear growth of regenerated axons to some extent. A major problem of these hollow INGC comes from axon misdirection: Oriented structures still differ substantially from the guiding basal lamina microchannels in nerve autografts, and hence lead to very limited positive outcomes in the patient’s recovery, too short degradation time and inappropriately high Young^'s modulus (linked to the limited range of biomaterials used for hollow tube production). Thus, for these commercial tubes, there is a mismatch in mechanical properties between the device and native nerve tissue (Ultimate tensile stress of =11.7 MPa for rabbit tibial nerve). Another drawback of tubulization with empty hollow tubes is the maximum nerve gap allowable for successful recovery: 3cm nerve defects. Due to these limitations, commercially available hollow conduits fail to match the regenerative levels of autograft. For pronounced discontinuities (> 3cm), 3D scaffolds based on processed nerve allografts are used for bridging the nerve gap. Disadvantages of this nerve allografting include the risk of rejection and complications related to immunosuppression. An improved alternative to hollow tubes, similar to nerve allografts that are currently in the market, are decellularized mammalian implants that contain the structure of a mature nerve. However, these implants cost thousands of Euros for each unit and are exposed to anatomical variability. Since the basal lamina microchannels in nerves are known to play a significant guiding role in the linear growth of regenerating axons, INGCs with architecture and dimensions resembling the basal lamina microchannels sutured to nerve stumps are described as alternative for bridging nerve gaps. However, the designs proposed to date lack the mechanical properties required to mimic those of the native nerve tissue. Additionally, proper scale - up of the manufacturing technologies used for obtaining biomimetic tubular devices at lab scale is required to ensure the final industrialization of the product/device, covering a range of gaps in a cost - effective manner.

US20100291 180 describes a cell guidance tube that comprises an inner layer, wherein the inner layer comprises at least one biodegradable polymer, wherein the tube comprises a lumen that comprises at least one immobilized peptide mimic of a carbohydrate. The materials used in this invention include two different nerve graft components: bioresorbable “PolymerDrugs” that release salicylic acid and/or other nonsteroidal anti-inflammatory drugs (NSAIDs) as they degrade, enabling them to serve not only as a structural scaffolding but also as a drug delivery device, and 2) immobilized peptide that mimics the structure of carbohydrates.

US20170296193 describes nerve guides that comprise three components: 1 ) a biodegradable membrane 2) a hydrogel attached to the inner surface of the biodegradable membrane, wherein the hydrogel comprises one or more neurotrophic factors; and 3) nanofibres lining the lumen of the nerve guide. The biodegradable membrane is selected from a group consisting of biodegradable polyesters, poly(amino acid)s or derivatives thereof, and natural biodegradable polymers. The hydrogel layer is comprised of a material selected from the group consisting of PVA hydrogel or a mixture or PVA and one or more of pluronic polymers, heparin, heparan sulfate, chitosan, alginate, and dextran sulfate. The nanofibre layer comprises one or more polymers selected from biodegradable polyesters, poly(amino acid)s and derivatives thereof.

US5019087 describes a hollow conduit comprised of a matrix of Type I collagen and laminin-containing material.

EP2465472 relates to a silk nanofibre nerve conduit characterized in that fibroin nanofibres having a diameter of 200 to 400 nm, originated from silk fibre, are stacked layer upon layer to form a porous conduit-shape.

WO2011004971 describes a nerve conduit for peripheral nerve regeneration composed of a polymeric material, which is coated on its inner surface with collagen and nerve growth factor, wherein the polymeric material is selected from the group consisting of paralene, SU-8, polynorbornene and polyimide.

US8926886 describes collagen multichannel nerve conduits which are suitable for use in repair of peripheral nerves. US9327054 refers to a conduit including an insulating material in the form of a tube having an inner surface and an outer surface, the inner surface having carbon nanotubes.

And, finally, EP1472301 refers to DL-lactide-e-caprolactone copolymers for use in medical application and the application of these polymers in the production of biodegradable medical applications, such as the commercially available artificial nerve guide, Neurolac™.

In this scenario, the present invention provides a novel biomimetic nerve prostheses paying special attention to device structure, biomaterials and their combination with high throughput manufacturing methods, which play a vital role in the industrialization process. The implantable nerve guidance conduit of the present invention overcomes the limitations of state of the art INGCs in terms of regenerative capacity, biodegradability, physical properties and manufacturability. More specifically, the present invention provides a 3D scaffold mimicking the nerve tissue leading to positive outcomes close to 100% for short (<20 mm) and large gaps (>30 mm and up to 50 mm). The aim of the present invention is also to reduce the cost for peripheral nerve implant as well as to avoid dependence on potential donors or complicated and expensive removal of cellular components as needed for allograft implants. The implantable nerve guidance conduit of the present invention promotes an earlier and more effective regeneration in large and short nerve gaps, overcoming the limitations of single tube conduits linked to axon misdirection.

SUMMARY OF THE INVENTION

Accordingly in a first aspect, the invention relates to an implantable nerve guidance conduit comprising: a hollow tube, and optionally, an inner microstructure, characterized in that the hollow tube is made of a blend of biopolymers comprising a copolymer of poly(3-hydroxyoctanoate-co-3-hydroxydecanoate) and poly(e- caprolactone) (P(3HO-co-3HD)/PCL).

In a further aspect, the blend of biopolymers of the hollow tube comprises from 60 to 95% by weight, more preferably, 75% by weight of a copolymer of poly(3- hydroxyoctanoate-co-3-hydroxydecanoate) and from 5 to 40% by weight, more preferably, 25% by weight of poly(e-caprolactone).

In a further aspect, the hollow tube presents a Young’s Modulus of from 0.5 - 20 MPa, preferably between 2 and 3 MPa. In a further aspect the wall of the hollow tube is porous, being the volume of pores in the wall of the hollow tube from 40 to 80%, preferably 65% and the size of the pores is from 1 to 50pm, preferably 25 - 50pm.

In a further aspect, the inner diameter of the hollow tube is from 0.5 to 20mm, preferably 1.6mm, and the thickness of the wall of the hollow tube is from 0.1 to 0.8mm, preferably 0.5mm.

Further, the length of the hollow tube can be up to 50 mm.

And further, the implantable nerve guidance tube presents a biodegradation rate of 18 to 24 months after implantation at the injured site of the human body. A further embodiment of the present invention refers to an implantable nerve guidance conduit which incorporates an inner microstructure which is in the form of a microchanneled inner tube or a bundle of threaded microfibres.

Where the inner microstructure is in the form of a microchanneled inner tube, this inner structure presents from 2 to 50 aligned parallel microchannels with a diameter of from 50 to 500pm, and separated from each other by a distance of from 50 to 500pm embodied in a sponge like tubular structure.

In a further aspect, the inner microstructure in the form of a microchanneled inner tube is made of a biopolymer comprising methacrylated polycaprolactone (mPCL).

In a further aspect, the microchanneled inner tube is a porous structure, where the pore size is from 10 to 70pm and the volume of pores in the microchanneled inner tube is from 40 to 80%.

In still a further aspect of the invention, the microchanneled inner tube presents a Young’s Modulus of from 0.05 to 5MPa, preferably from 0.1 to 0.35MPa.

In a further aspect, the microchanneled inner tube has a diameter of 0.5 to 20mm and a length of from 5 to 50mm.

A further embodiment of the present invention refers to an implantable nerve guidance conduit with an inner microstructure is in the form of a bundle of microfibres. In a further aspect of this embodiment the inner microstructure is in the form of a bundle of microfibres and is made of a blend of biopolymers which comprises poly(3- hydroxyoctanoate) and poly(3-hydroxybutyrate) (P(3HO):P(3HB)), preferably, from 25 to 75% by weight of poly (3-hydroxyoctanoate) and from 25 to 75% by weight of poly(3- hydroxybutyrate), more preferably, 50% by weight of poly(3-hydroxyoctanoate) and 50% by weight of poly-(3-hydroxybutyrate).

In a further aspect of the invention the microfibres have a diameter of from 5 to 10pm, preferably from 5 to 8pm.

The present invention also refers to a method of obtaining the implantable nerve guidance conduit of the present invention, comprising obtaining a hollow tube by microextrusion, and optionally, obtaining an inner microstructure in the form of a microchanneled inner tube or a mat of microfibres, and introducing the microchanneled inner tube or mat of microfibres into the hollow tube.

In a further aspect of the invention, obtaining a hollow tube by microextrusion comprises the following steps:

- preparing a blend (mixture) of biopolymers comprising a copolymer of poly(3- hydroxyoctanoate-co-3-hydroxydecanoate) and poly(e-caprolactone) and plasticizing the blend at a temperature of 90 - 120°C by means of mechanical shaker in an extruder system, - pressing the blend out of the extrusion die head with the desired hollow tube configuration

- cooling and solidifying the hollow tube extrudate at ambient temperature for about 5 to 10 minutes.

- pulling, cutting the extrudate and collecting the finished hollow tubes. In a further aspect, the blend of biopolymers of poly(3-hydroxyoctanoate-co-3- hydroxydecanoate) and poly(e-caprolactone) prepared for obtaining the hollow tube comprises additionally a porogen selected from the group consisting of glucose and NaCI, preferably the blend of biopolymers comprises 30% by weight of poly(3- hydroxyoctanoate-co-3-hydroxydecanoate) and poly(e-caprolactone) and 70% by weight of porogen selected from the group consisting of glucose and NaCI.

In a further aspect, the method of the invention comprises obtaining an inner microstructure in the form of a microchanneled tube, by means of UV-3D micromoulding and comprises the following steps:

- injection of a mixture comprising methacrylated polycaprolactone in a negative mould, comprising stretched stainless steel wires having each from 50 to 500pm diameter, preferably 200pm, aligned in a silicone container with a diameter of from 5 to 20mm, preferably 1.6mm - curing of the biopolymer inside it and mould disassembling.

In a further aspect, the mixture comprising methacrylated polycaprolactone used for obtaining the inner microstructure in the form of a microchanneled tube further contains glucose as porogen to create porosity and a photoinitiator to cure the polymer by UV- VIS exposure. In still a further embodiment of the present invention, the mixture comprising methacrylated polycaprolactone used for obtaining the inner microstructure in the form of a microchanneled tube further contains a surfactant, chloroform, toluene and a photoinitiator and water droplets at 35°C. . These componenets are subjected to vigorous stirring until an emulsion is obtained to create porosity through High Internal Phase Emulsion Templating (polyHIPE)..

In a further embodiment of the invention the method comprises obtaining an inner microstructure in the form of a mat of microfibres by means of electrospinning and comprises the following steps:

-preparing a solution comprising poly(3-hydroxyoctanoate) and poly(3-hydroxybutyrate),

- electrospinning the solution at a flow rate of from 1 to 10ml/hr, preferably 3ml/hr, speed of from 1000 to 5000rpm, preferably 2000rpm and voltage of 1 to 20kV, preferably 12.5kV to obtain electrospun fibres, - laser cutting the electrospun fibres to obtain mat pieces of 100mm width and 6mm length.

Still another object of the present invention comprises introducing the mat of microfibres into the hollow tube by threading the fibres into the hollow tube by tightly wrapping a 100 x 6mm fibre mat around a syringe needlde with the fibres parallel to the needle, placing the fibres into the tube and removing the needle by a sharp pair of forceps. matmatFinally, a further embodiment of the invention refers to the implantable nerve guidance conduit as defined in the previous paragraphs for use in a medical treatment or diagnosis. It is a further aspect of the present invention the implantable nerve guidance conduit as defined in preceding paragraphs for treating a nerve injury in a mammal, such as animals and humans. In particular, the nerve injury includes peripheral nerve injury.

In a further aspect of the present invention the implantable nerve guidance conduit is for use in reconstructing nerve gaps of up to 50mm.

A still additional aspect of the present invention is a method for regenerating a damage nerve in a patient in need thereof, comprising placing the implantable nerve guidance conduit as defined in previous paragraphs, at the site of neuronal injury so as to regenerate the nerve. FIGURES

Figure 1 : Diagram of the extrusion line configuration

Figure 2: Set up for the fabrication of channeled microstructure

Figure 3: Optical images of PCLm polyHIPE channeled inner microstructure showing A) length and B) cross-section. SEM images of C) PCLm polyHIPE channeled inner microstructure inserted in a hollow conduit and D) a detail of the porosity of the PCLm polyHIPE channeled microstructure.

Figure 4: MicroCT images of 5pm fibres, spun for 20 minutes, and 8pm spun for 15 minutes Figure 5: Histological analysis results P(3HO-3HD)/PCL: comparative data of animals regenerated vs. autograft and Neurolac™

Figure 6: Immunohistochemistry of Fibronectin (green) and Iba1 (red), for labelling extracellular matrix supporting nerve growth and infiltrated macrophages. Cell nuclei are stained with DAPI (blue). Scale bar = 250 urn.

Figure 7: Immunohistochemistry of s1003 (green) and Neurofilament (red), for labelling Schwann cells and axons, respectively. Cell nuclei are stained with DAPI (blue). Scale bar = 250 urn for low magnification images (top panels) and 50 urn for higher magnification images (lower panels). Figure 8: Change in the wall thickness at 4 months after injury and repair.

Figure 9: Live/Dead analysis confocal micrographs of rat primary Schwann cells cultured on (A) PCL spin coated film, (B) P(3HO):P(3HB) (50:50)film, (C) P(3HO- co3HD):PLLA (50:50) flat film and (D) Tissue culture plastic. Neuronal cells were labelled with Syto-9 (green) and propidium iodide (red) for live and dead cells respectively. Scale bar=100pm.

Figure 10: Live/dead analysis of NG108-15 neuronal cells and rat primary Schwann cells on candidate material fibres expressing the average total number of live versus dead cells.

Figure 1 1 : Average neurite length of NG108-15 neuronal cells cultured on different fibre materials and diameter sizes, compared to the average neurite length when cultured with primary Schwann cells. (Mean ± SD, n=3 independent experiments)

Figure 12: Representative images of implanted PCLm 6 mm long tube held with fibrin, for a 6 mm gap (LEFT). It is showed the presence of fibronectin protein, myelinated axons and Schwann cell inside the microchannels.

Figure13. Histogram of the pore size distribution of PCLm PolyHIPE

Figure 14: Development of a nerve regeneration platform for three dimensional (3D) testing of different candidate nerve guide conduit materials and geometries showing testing in a P200 pipette tip box (A) via filling of the bottom chamber with media, removal of sterile pipette tips and insertion of candidate nerve guides onto which DRG are placed (B) and reinsertion into a pipette tip box (C) which is then placed in a humidified C02 incubator.

Figure 15: Computed Tommography of Guidance conduit consisting of hollow tube incorporating a microchanneled inner tube.

Figure 16: Immunostaining analysis of DRGs regenerating on electrospun fibres inserted inside P(3HO-3HD)/PCL 75/25 hollow tubes (A) and subsequently removed for microscopy analysis (B) DRG’s were placed on candidate internal materials using our 3D ex-vivo nerve regeneration system. A series of different candidate fibres were tested with blue representing DAPI staining, green representing Schwann cells (S100) and Red representing neurons (Beta III tubulin). Cell migration distances was measured via visualizing multiple fields in the z direction using DAPI and analyzed using a 1 way ANOVA with a Kruskal-Wallis post-test ^*p<0.05.

DETAILED DESCRIPTION OF THE INVENTION

The following disclosure is presented to provide an illustration of the general principles of the present invention and is not meant to limit, in any way, the inventive concepts contained herein.

All terms defined herein should be afforded their broadest possible interpretation, including any implied meanings.

The experimental data provided in the present specification has been performed in rat sciatic nerve models. The results thereof are an indisputable means to prove same activity in humans. Where the experimental results relate to small nerve gaps or gaps of 6mm, these results serve to proof same activity in human nerve gaps of 1 - 2 cm. Where the experimental results relate to medium or large nerve gaps or gaps of up to 10mm, these results serve to proof same results in human nerve gaps of 3 - 5 cm. Further, it should be stated that, as recited herein, the singular forms“a”,“an”, and “the” include the plural referents unless otherwise stated. Additionally, the terms “comprises” and “comprising” when used herein specify that certain features are present in that embodiment, however, this phrase should not be interpreted to preclude the presence or addition of additional steps, operations, features and/or components. In the present specification, weight % of biopolymers are calculated with respect to the total weight of biopolymer composition. Success of regeneration in a neural guide is mainly linked to the material properties and the physical characteristics of the conduit, and the length of the gap to be bridged. The improvement of some parameters such as the internal diameter of the conduit, permeability of the outer wall of the conduit, re-absorbable materials, etc. will help to improve the final outcome.

The present invention refers to a new and advanced implantable nerve guidance conduit made of specific biomaterials and/or blends thereof combined with a specific device structure or configuration with improved biocompatibility and biodegradability, regenerative capacity, mechanical and physical properties required to mimic endoneural tubes for an efficient regeneration of both sensory and motor axons of small and even large-gap transected nerves of up to 50mm.

In particular, the present invention provides an implantable nerve guidance conduit which presents improved properties in terms of biocompatibility, mechanically stable at time zero, similar mechanical and physical properties of native nerve, flexibility to allow bending of joints without nerve compression, limit scar infiltration, semi-permeable to allow nutrients to enter and the wastes to exit but keep the inflammatory cells out, prevent fibrous ingrowth but maintain neurotrophic factors inside, exhibit low immune response, be surgeon friendly in terms of handling, accommodate nerve swelling without excessive compression, correct breakdown rate to match the rate of neural regeneration, resorbable to remove the need for secondary surgery and to prevent chronic inflammation and pain caused by nerve compression due to the eventual collapse of the conduit and have non-toxic degradative products.

Structure, design and dimensions of the implantable nerve guidance conduit

It has been found by the inventors of the present invention that an implantable nerve guidance conduit comprising a hollow tube made of a blend of biopolymers comprising a copolymer of poly(3-hydroxyoctanoate-co-3-hydroxydecanoate) and poly(e- caprolactone) is sufficient for providing the benefits of the present invention.

In a more particular embodiment, the blend of biopolymers of the hollow tube comprises from 60 to 95% by weight, more preferably, 75% by weight of a copolymer of poly (3-hydroxyoctanoate-co-3-hydroxydecanoate) and from 5 to 40% by weight, more preferably, 25% by weight of poly(e-caprolactone). In order to achieve the mechanical and physical properties of native nerves, while maintaining biocompatibility, the hollow tube presents a Young’s Modulus of from 0.5 - 20 MPa, preferably between 2 and 3 MPa.

Further, selective porosity for nutrient exchange and a slow degrading rate to reduce fibrosis is achieved by a porous wall of the hollow tube, being the volume of pores in the wall of the hollow tube from 40 to 80%, preferably 65% and the size of the pores is from 1 to 50pm, preferably 25 - 50pm.

Additionally, as the implantable nerve guidance conduit must serve as a physical guidance along the length of the nerve injury site, so as to better organize axon alignment for successful reinnervation, the inner diameter of the hollow tube is from 0.5 to 20mm, preferably 1 6mm, and the thickness of the wall of the hollow tube is from 0.1 to 0.8mm, preferably 0.5mm. Further, the length of the hollow tube can be up to 50 mm for effective healing and reconstruction of large peripheral nerve injuries. In each case there are around 2mm in each side for suturing the tube to the nerve. And further, the implantable nerve guidance tube presents a biodegradation rate of 18 to 24 months after implantation at the injured site of human body, which corresponds to the correct rate of conduit breakdown to match the rate of neural regeneration.

In order to enhance the regenerative ability of hollow INGCs, the applicants of the present invention have increased the surface area of the device for allowing efficient incorporation of specific cells and growth factors. To this respect, the implantable nerve guidance conduits of the present invention may, optionally, include an inner microstructure used as a structural component incorporated into the hollow tube, which provides an improved effect when reconstructing nerve injuries, in particular to avoid misdirection during nerve regeneration. The said inner microstructure may be represented by the following structures:

- microfibres of poly(3-hydroxyoctanoate-co-3-hydroxybutanoate), (P(3HO):P(3HB)),

- microchanneled PCLm tubes, preferably a porous structure made by addition of a porogen during its manufacturing process, and

- microchanneled PCLm polyHIPE tubes, also a porous structure but made by the polyHipes technology Where the inner microstructure is in the form of microfibres, these fibres are made of a blend of biopolymers which comprises poly(3-hydroxyoctanoate) and poly(3- hydroxybutyrate), preferably, from 25 to 75% by weight of poly-(3-hydroxyoctanoate) and from 25 to 75% by weight of poly-(3-hydroxybutyrate), more preferably, 50% by weight of poly (3-hydroxyoctanoate) and 50% by weight of poly(3-hydroxybutyrate).

In a further aspect of the invention the microfibres have a diameter of from 5 to 10pm, preferably from 5 to 8pm.

Where the inner microstructure is in the form of a microchanneled inner tube, this inner structure presents from 2 to 50 aligned parallel microchannels with each a diameter of from 50 to 500pm, and separated from each other by a distance of from 50 to 500pm. The resulting structure consists of a tubular sponge like structure with aligned microchannels embodied in said structure. The inner microstructure in the form of a microchanneled inner tube is made of a biopolymer comprising polycaprolactone methacrylated and is a porous structure, where the pore size is from 10 to 70pm and the volume of pores is from 40 to 80%.

Again, in order to achieve the desired mechanical and physical properties of native nerves, while maintaining biocompatibility, the microchanneled inner tube presents a Young’s Modulus of from 0.05 to 5MPa, preferably from 0.1 to 0.35MPa.

Finally, the microchanneled inner tube has a diameter of 0.5 to 20mm and a length of from 5 to 50mm.

A preferred embodiment of the present invention is an implantable nerve guidance conduit for treating peripheral nerve injuries comprising a hollow tube made of a blend of biopolymers comprising 75% by weight of a copolymer of poly (3-hydroxyoctanoate- co-3-hydroxydecanoate) and 25% by weight of poly(e-caprolactone), and an inner microstructure in the form of a microchanneled inner tube, which presents from 2 to 50 aligned parallel microchannels with each a diameter of from 50 to 500pm, and separated from each other by a distance of from 50 to 500pm.

Another preferred embodiment of the present invention is an implantable nerve guidance conduit for treating peripheral nerve injuries comprising a hollow tube made of a blend of biopolymers comprising 75% by weight of a copolymer of poly (3- hydroxyoctanoate-co-3-hydroxydecanoate) and 25% by weight of poly(e-caprolactone), and an inner microstructure in the form of microfibres made of a blend of biopolymers which comprises 50% by weight of poly (3-hydroxyoctanoate) and 50% by weight poly (3-hydroxybutyrate), having each a diameter of from 5 to 10pm.

Fabrication process of the different biopolvmers and corresponding blends

Before proceeding with the fabrication of the different structures composing the implantable nerve guidance conduit of the present invention, the different biopolymers and blends used are to be obtained.

In general, polyhydroxyalkanoates (PHAs) are synthesized by bacteria as storage compounds for energy and carbon, normally in the presence of excess carbon and with at least one nutrient essential for growth, such as nitrogen, phosphorus, sulphur or oxygen, present in limiting concentration.

In particular, the following processes are used for obtaining the specific biopolymers and blends used in the present invention: Production of Polv(3-hvdroxyoctanoate-co-3-hvdroxydecanoate or P(3HO-co-3HD):

P(3HO-co-3HD) is produced by Pseudomonas mendocina using glucose as the sole carbon source via batch fermentation. P(3HO-3HD) production occurs in three stages. First stage involves the preparation of the seed culture or inoculum. This inoculum is used to inoculate second stage seed culture. During the third stage, P(3HO-3HD) production media (mineral salt media) is inoculated using second stage seed culture as the inoculum Fermentation is carried out in the bioreactors for over a period of 48 hours after which the cells are harvested by centrifugation. P(3HO-co-3HD) is extracted from the lyophilized cells using a soxhlet apparatus. It is precipitated from the solvent using ice-cold methanol. Production of Polv(3-hvdroxyoctanoate) or P(3HO): P(3HO) is produced by Pseudomonas mendocina CH50 using sodium octanoate as the sole carbon source via batch fermentation. P(3HO) production also occurs in three stages. The first stage involves the preparation of the seed culture or inoculum. This inoculum is used to inoculate the second stage seed culture. During the third stage, P(3HO-3HD) production media (mineral salt media) is inoculated using the second stage seed culture as the inoculum Fermentation is carried out in the bioreactors for over a period of 48 hours after which the cells are harvested by centrifugation. P(3HO) is extracted from the lyophilized bcells using the soxhlet apparatus. It is precipitated from the solvent using ice-cold methanol. Production of PolvO-hvdroxybutyrate^') or ROHB^'): P(3HB) is produced by Bacillus subtilis OK2 using glucose as the sole carbon source via batch fermentation. P(3HB) production occurs in two stages. First stage involves the preparation of the seed culture or inoculum. This inoculum is used to inoculate P(3HB) production media (modified Kannan and Rehacek (KR) media). Fermentation is carried out in the bioreactors for over a period of 48 hours. P(3HB) is extracted from the lyophilized cells using a soxhlet apparatus. It is precipitated from the solvent using ice-cold methanol.

PHAs produced are characterized for their chemical and structural properties using GCMS and NMR respectively. Their thermal properties are assessed using DSC whereas their mechanical properties are assessed using tensile testing. Molecular weight analysis is done using GPC.

Production of P(3HO:3HD)/PCL 75/25 blend with glucose poroqen:

For the production of the blends, the first component is weighed and dissolved in the chloroform solution while stirring on a magnetic stirrer. Once the first component is completely dissolved, the second component is weighed and added to the polymer solution. Once the components are completely dissolved, the glucose porogen is added to the polymer solution and stirred for 24 hours in the fumehood. The polymer blend solution containing glucose porogen is poured into a clean, sterile glass tray. The tray is covered with an aluminium foil to prevent contamination. Perforations are made in the foil to allow solvent evaporation. Once the blend is dry, it is removed from the tray using a clean, sterile scalpel and cut into specific dimension.

Production of methacrylated PCL with poroqen:

PCLm and a photoinitiator is mixed with glucose in a certain proportion to obtain a solution. Production of methacrylated PCL with porosity defined bv High Internal Phase

Emulsion.

• PCLm is mixed with a surfactant, chloroform, toluene and photoinitiator in a certain proportion. Water droplets at 35°C are added under vigorous stirring resulting in an emulsion ready for the injection into the mould. The stirring is defined by the revolutions per minute of the mixer and it is correlated with the degree of porosity obtained finally on the polymer. A pore size distribution between 10 and 90 pm is obtained according to the process definition. It is showed on figure 13. Production of P(3HO):P(3HB) blend solution

On a small scale, 0.4g of P(3HO) and 0.4g of P(3HB) were cut up and dissolved in chloroform. Chloroform was added 1 g at a time, to make a solution of 10g (equal to a 8wt% solution). The solution was dissolved overnight on a magnetic stirrer platform at room temperature and 1000rpm. Fabrication of the parts composing the implantable nerve guidance conduits

The fabrication of the hollow tubes is performed by microextrusion of the biopolymer mixture blended with glucose as the porogen.

In this type of processing, the polymer is plasticized inside the machine by means of adding heat and mechanical stirring. For mechanical stirring, a spindle is used that can have different cross sections and geometries, but generally acts as a worm screw, mixing, compressing and moving the polymer towards the mouthpiece as it is melted or plasticized.

As shown in Fig. 1 , the material, once it has reached a malleable state, is forced through the extruder outlet. The mouthpiece, head or matrix that has the shape of the extruded profile is at the outlet of the mouthpiece.

Next to the mouthpiece is a cooling system for the extruded profile, such that the polymer is cooled in order to set the final shape thereof. Depending on the quality or tolerance required by the extruded profile, different systems for modifying the dimensions and controlling the dimensions and quality of the extrudate can be used. Once cooled, the profile passes through the drive system that serves to stretch the profile and help remove the already extruded material from both the extruder itself and from the cooling system. Next to the cooling system is the storage system, which can have different forms, from a winding to a sectioning part for cutting profiles to the length desired.

To extrude the external tube of the neural regeneration device, a Thermo Fisher Scientific Haake minilab II conical double-spindle mini-extruder model was used.

A mouthpiece was designed that makes it possible to produce tubular sections with an adjustable inner diameter and wall thickness in order to be able to extrude the tube section required. The designed mouthpiece has three temperature-controlled areas to regulate the plasticized state of the polymer.

The material used to extrude the external tube is a blend of P(3HO-3HD)/PCL 75:25 with a sifted glucose load with a particle size in the range of 25-50 pm and a ratio of 70% by weight. The extrusion conditions of P(3HO-3HD)/PCL 75:25 to obtain the required tube are:

Machine temperature 1 10°C; Area 3 temperature: 1 15°C; Area 2 temperature: 115°C; Area 1 temperature: 105°C,

- Spindle rotation 25rpm, and

- Drive speed 8-10 The tube collected after the drive system is sectioned into short lengths of several tenths of a centimetre in order to have pieces of a manageable size.

After leaving the segments of the extruded tube to air cool for several minutes, they are then cut into an appropriate size for the packaging thereof in airtight bags that are subsequently frozen in a no-frost freezer at temperature of 20°C below 0. Since the tubes have glucose as a porogen, it is not necessary to wash and extract the glucose before they are implanted. The tubes may be implanted with the presence of glucose. The leaching of the glucose on the gave a pore size between 25 and 50 pm and the porosity generated was around 70% in weight, which means a porous volume around 65%.

As regards fabrication of the inner microstructure formed by a microchanneled inner tube this part is made by UV-3D micromoulding technology.

Micromoulding is a widely used technology for generating micro structures in three dimensional (3D) tissue engineering constructs. Thus, it was selected for the fabrication of porous and biodegradable microchanneled microstructure. It consists of the fabrication of a negative mould, injection and cured of a biomaterial inside it and the mould disassembled.

As to the negative mould to be used for the said purpose a specific set up is needed.

As shown in Figure 2, stainless steel wires having a 200 pm diameter (1 ) are interwoven and aligned by means of two microperforated steel surfaces with the desired configuration. One of the microperforated surfaces is placed on the steel base (2) and is covered with the part (3). The part (4) is placed and the other microperforated surface is fitted on the part (5).

The threaded bars (6) are introduced through the lateral holes of the plates (2), (3), (4) and (5). The lower ends of the bars (6) are threaded on (7) and the microperforated plates are fastened to (3) and (5) with nuts (8). The edges of the cables are tied with aluminium bushings (9) and (9^'), (9^') inside of the part (7), and the nuts (8) are adjusted until the cables are perfectly stretched.

The entire system is fastened by introducing the alignment bars (10) into (2), (4) and (5). Finally, a flexible outer Polydimethylsiloxane (PDMS) mould that is transparent to ultraviolet (UV) light (1 1 ) and with an inner channel with a 1.5 mm diameter and 4 cm height is placed surrounding the wires (Note: the diameter of the PDMS inner channel and the height thereof may be changed as needed). The part (4) is freely adjusted in height and rests on the PDMS by keeping it fastened to and aligned with the wires.

With respect to the method used, it is made up of the following steps: The fabrication of channeled nerve guide conduits by means of the wires set up was evaluated with several biomaterials (synthetic, natural and blends). The final decision of the biomaterial was based on the fluidity of the biomaterial, the ease of the injection, the properties of the biomaterial used. Methacrylated polycaprolactone (PCLm) was chosen as the biomaterial for the fabrication of the porous inner microstructure in the form of a channeled inner tube. PCLm is structured by means of two different approaches:

Leaching of embedded particles: PCLm and photoinitiator was mixed with glucose in a certain proportion (50:50;; 70:30 Glucose: PCLm) to make the final device porous, being the final proportion of polymer and glucose in weight of

60:40. The mixture was injected into the PDMS container embracing tense and aligned wires. The injection of the biomaterial was manual with a syringe. Upon UV radiation and disassemble of the set up, the final microchanneled inner tube was fabricated. To finally obtain large, flexible and porous conduits, the microchanneled inner tube was immersed in distilled-water.

In this particular case, the leaching of the porogen gave a pore size between 25 and 50 pm (sieved glucose size) and the porosity generated was around 60% in weight, which means a porous volume around 50%.

- Alternatively, the channeled inner microstructure can be carried out by polymerization of high internal phase emulsion technique (polyHIPE). In that case, PCLm is mixed with a surfactant, chloroform, toluene and photoinitiator in a certain proportion. Water droplets at 35°C are added under vigorous stirring resulting in an emulsion ready for the injection using a plastic syringe. Upon UV radiation and disassembling of the setup, the final microchanneled inner structure is fabricated obtaining an hydrophilic, flexible and compressible sponge conduit (Figure 3).

In this particular case, the pore size distribution generated using this technique is around 10 to 90-100 microns and a porous volume nearly 70%.

The liquid light-curing material (PCLm + porogen or PCLm polyHIPES) is injected with a syringe in the PDMS mould (12) through an injection point and is cured under UV light (350-450 nm) for 15 minutes. Under this radiation, the material criss-crosses and solidifies.

According to the method explained above and depending on the type of process to be followed, the disassembly of the mould is carried out according to one of the following two options:

PCLm+porogen: The wires are cut on both sides of the PDMS mould and are submerged in ethanol for 20 minutes. After this time, the wires are extracted and the tube structured with the inner channels is removed from the PDMS.

PCLm+polyHIPES: The wires are cut on both sides of the PDMS mould, the wires are extracted and the tube with inner channels removed is removed from the PDMS.

In order to thread the microchanneled inner tube into the hollow tube fibre mat of different dimensions, were tightly wrapped around a Hamilton syringe with the fibres parallel to the needle. The syringe was a Hamilton 80300 Standard Microliter Syringes, 10 uL, Cemented-Needle, 1/ea, with a needle gauge of 26S. The fibres were placed into the hollow tube, and removed off the needle by a sharp pair of forceps. From 1 to 4mm was marked onto a pair of sharp forceps, to push the fibre mat into place. In order to determine that alignment of fibres was achieved, as well as the 1 to 4mm overhang, conduits containing fibres were then analyzed using microCT. Once alignment, and design was fabricated, conduits were packaged up and sent for in vivo analysis.

Optionally, in order to secure fibres inside the hollow tube, it may be possible to add a little fibrin glue, for example, at the middle of the fibre piece before threading. Alternatively, it is also possible to pinch the middle of the tubes, to ensure fibres fill the entire conduit if moved during transport. As regards, fabrication of the inner microstructure in the form of a bundle of microfibers was carried out by electrospinning techniques.

Basically, there are two main electrodes connected to an electrostatic field of very high intensity in the common electrospinning method: the first one (so-called spinneret), is usually connected to a positive high voltage potential. The second electrode (so-called collector), opposite the first one, most frequently has a plate shape (4SPIN) connected to a lower electric potential (in most cases it is grounded, otherwise on a negative potential).

Fabricating P(3HO):P(3HB) (50:501 Fibres On a small scale, 0.4g of P(3HO) and 0.4g of P(3HB) were cut up and dissolved in chloroform. Chloroform was added 1 g at a time, to make a solution of 10g (equal to a 8wt% solution). The solution was dissolved overnight on a magnetic stirrer platform at room temperature and 1000rpm.

The solution was then electrospun using the corresponding equipment. Using a 1 ml terumo syringe, the solution was spun at a flow rate of 3ml/hr, speed of 2000rpm and voltage of 12.5 kV. The cylinder drum collector had a length of 16cm and diameter of 6cm. The electrospinning conditions are below (Table 1 ).

Determining correct density for threading P(3HO^'):P(3HB^') Fibres

In order to determine the correct density of fibres, of 5 and 8pm, to fill P(3HO:3HD):PCL external tubes, fibres were electrospun for different times, to create a mat, threaded into the tubes, and then analyzed using microCT to determine the correct filling times.

Table 2. Table of spinning times and Mat thickness

Table 2 shows the different diameters, and different times chosen for threading. Figure 4 shows two microCT images of 5pm (3A) and 8pm fibres (4B) threaded into tubes and imaged to determine how full the tube should be. Figure 4C is an image of how the fibres appear after electrospinning. Due to the close distance of the needle tip to the collector, solutions do not need to electrospun for long to produce a thick mat of fibres. Also, due to closeness of the needle tip to the collector, fibres are concentrated into the middle of the collecting material, fabricating a mat of fibres with a width of 1-1.5cm, and length of 15cm. In order to thread fibres into tubes, the correct length of the fibres was determined due to conduit design (short, medium or large gap). For example, a short gap conduit, tube length of 26mm, required 16mm of electropsun fibres in order to fabricate conduits with a 5mm gap either side for suturing. Electrospun fibres were laser cut, to fabricate mat pieces of 100mm width and 16mm length. In order to thread fibres into the hollow tubes, the 100x6mm fibre mat, was tightly wrapped around a Hamilton syringe with the fibres parallel to the needle. The syringe was a Hamilton 80300 Standard Microliter Syringes, 10 uL, Cemented-Needle, 1/ea, with a needle gauge of 26S.

The fibres were placed into the hollow tube, and removed off the needle by a sharp pair of forceps. 2mm was marked onto a pair of sharp forceps, to push the fibre mat into place. In order to determine that alignment of fibres was achieved, as well as the 2mm overhang, conduits containing fibres were then analyzed using microCT. Once alignment, and design was fabricated, conduits were packaged up and sent for in vivo analysis. Optionally, in order to secure fibres inside the hollow tube, it may be possible to add a little fibrin glue, for example, at the middle of the fibre piece before threading. Alternatively, it is also possible to pinch the middle of the tubes, to ensure fibres fill the entire conduit if moved during transport. It is to be understood that variants of the below described examples of the invention in its various aspects, such as would be readily apparent to the skilled person, may be made without departing from the scope of the invention in any of its aspects.

EXAMPLES

Example 1 : Regeneration of peripheral nerves through the hollow tube

Pre-clinical model in the rat sciatic nerve is used to assess regeneration through designed devices. P(3HO-3HD)/PCL prototypes either Leached or Non-Leached, either of 500 or 250 pm of wall thickness and PLCL Leached of 250 and 500 pm of wall thickness were used to bridge 10 mm gap defects in the rat sciatic nerve. Regeneration capacity was assessed during 4 months follow-up following electrophysiological analysis once a month. Regenerated cables and tubes were collected to analyze the area of the regenerated nerves, number of myelinated fibres and density of them at mid point of the graft/tube or distally (out of the tube).

When comparing P(3HO-3HD)/PCL prototypes, differences in electrophysiological data and histological data were obtained when comparing 500 and 250 pm of wall thickness tubes, being higher in the 500 pm ones (either Leached or Non-Leached).

When comparing Neurolac (a commercial and FDA approved device, biodegradable and biocompatible) with the implantable nerve guidance conduit of the present invention (hollow tube) made of P(3HO-3HD)/PCL 500 pm of wall thickness prototypes revealed similar degrees of nerve regeneration (electrophysiological and histological data).

Main results obtained include the demonstration of a good capacity to promote regeneration of hollow P(3HO-3HD)/PCL 500 pm of wall thickness tubes.

Material and methods

Animals: Female Wistar rats (200-220 g) were used in this experiment. Animals were housed in plastic cages, maintained at 22°C in a 12 h light/dark cycle and allowed to free access to water and food. The experimental procedures were approved by the Ethical Committee of our institution and followed the rules of the European Communities Council Directive.

Surgical procedure:

All surgical procedures were performed with aseptic operating conditions and under pentobarbital/xylacine anesthesia (40/10 mg/kg i.p.; respectively). The right sciatic nerve was exposed and cut 6 mm distal to the exit of the gluteal nerve. Then, in the case of the hollow guides 14 mm long tubes were implanted (for a 10 mm gap) with a 10-0 epineural suture. The tubes were filled with sterile physiologic saline solution. In the autograft group, 10 mm of the right sciatic nerve was excised, flipped and sutured back to the remaining stumps with 10-0 epineural sutures. The muscle plane was sutured with resorbable 5-0 sutures, and the skin with 2-0 silk sutures. The wound was disinfected and the animals were treated with amitriptyline for preventing autotomy. The animals were randomly distributed in the different experimental groups described below (Table 3).

Electrophysioloqical analysis:

Functional reinnervation of the target muscles (gastrocnemius muscle, tibialis anterior muscle and plantar muscle) was assessed at 7, 60, 90 and 120 days post injury (dpi). Animals were anesthetized with pentobarbital/xylacine anesthesia (40/10 mg/kg i.p.; respectively). The sciatic nerve was stimulated by transcutaneous electrodes placed in the sciatic notch, and the Compound Muscle Action Potential (CMAP) of the target muscles was recorded using monopolar needle electrodes, placing the active one in the muscle belly and the reference in the 4^th toe. During the tests, the rat body temperature was maintained by a thermostated warming flat coil. Latencies of the responses and the amplitudes of the M-waves were measured. The contralateral limb of each animal was used as values pre-injury.

Histological analysis:

Four months after injury animals were euthanized and transcardially perfused with 4% PFA in PBS. After perfusion, the regenerated nerves were collected and postfixed in 3% glutaraldehyde. Nerves were postfixed in osmium tetroxide (2%) and glucose (10% in PB 0.2M) (1 :1 ), 90 minutes at room temperature. Nerves were washed three times with distilled water and kept overnight at room temperature. Nerves were dehydrated in ascending series of Acetone and embedded in Araldite resin. Nerves were sectioned using a ultramicrotome in 1 pm sections. Collected sections were stained with toluidine blue and examined under light microscopy. Images of the whole nerve were acquired at 4x with a digital camera, while sets of images chosen randomly to represent at least the 30% of nerve cross-sectional area were acquired at 60x from mid and distal parts of the regenerated nerves. Measurements of the cross-sectional area of the whole nerve and the estimation of myelinated fibres were performed by using Image J Software.

Immunohistochemistry:

Regenerated nerves were processed for IHC analysis. Cross sectional sections (30 pm) of the regenerated nerves were used to label rmacrophage (lba-1 ) and the presence of extracellular matrix (Fibronectin).

Statistical analysis:

GraphPad (GraphPad Software, USA) was used to perform the statistical analysis on the collected data. Two-way anova analysis of variance with Bonferroni post hoc test was conducted to analyse differences in the electrophysiological data. One-way analysis of variance with Bonferroni post hoc test was conducted to analyse differences on histological analysis. Data was considered significantly different when P < 0.05.

Electrophvsioloav. P(3HO-3HDVPCL experiment.

Electrophysiological measurements were performed in anesthetized animals once a month during the 4 months follow-up. An electrical impulse was applied at the level of the sciatic notch and the electrical activity of targeted muscles due to their activation by the electrical impulse was recorded.

In the case of no regenerated animals, no electrical impulse is recorded. In the case of regenerated animals, the amplitude of the M wave was measured, because of the direct activation of the axons innervating the targeted muscles. Results revealed that autograft group has the best performance after injury. Onset of responses were sooner (60 dpi) were all animals positively responded to the electrical impulse and M-waves were recorded. At 4 months, these animals reached the plateau phase, recovering part of the lost function. In the case of tubular devices, animals responded to the electrical impulse later on time (90 dpi) with lower responses due to the slight poorer reinnervation of the targeted muscles. Differences were observed between P(3HO-3HD)/PCL 500 (leached and no-leached) and 250 urn no leached tubes, with higher values in the 500 urn groups.

Histological analysis. P(3HO-3HD)/PCL experiment: Animals were euthanized at 4 months. Peripheral regenerated nerves were collected and processed for histological analysis. Nerves were sectioned and separated into 3 different pieces. Proximal pieces were used to measure the area, estimation of myelinated fibres and density of axons and the central level of the graft or tube. Distal portions were used to measure the area, estimation of myelinated fibres and density of axons and the distal level, 6mm distal to the final suture of the graft or tube. Mid pieces were used for immunohistochemistry (IHC) analysis.

Cross sectional area of regenerated nerves was higher in the autograft group when compared to tubular device groups. This is due to the flip of the graft during the surgery. The trifurcation of the sciatic nerve (tibial, peroneal and sural branches) which originally (healthy nerve) is located in the distal part, in the autograft (after surgery) is located in the proximal part, increasing the cross sectional effective area for nerve regeneration. The estimation of myelinated fibres either in the mid nerve or in the distal region was higher in the autograft group when compared to the rest of the groups. However, the density of myelinated fibres is higher in the P(3HO— co-3HD)/PCL tubes than in the Neurolac’s for both histological analysis. Similar results were obtained when analysing density of axons, being increased in the autograft group vs tubular devices, while the density in the case of the P(3HO-3HD)/PCL tubes is higher than in the Neurolac’s tube.

As shown in Fig. 5, when analyzing the percentage of animals that regenerated after 4 months (presented a regenerated cable and positive responses after electrical stimulation during electrophysiological tests), 100% of animals regenerated in the autograft group. In the Neurolac group, 7 of 10 animals regenerated, in the P(3HO-co- 3HD)/PCL 500pm L group 7 out of 8 animals regenerated, in the P(3HO-co-3HD)/PCL 500pm NL group 6 out of 8 animals regenerated and in P(3HO-co-3HD)/PCL 250pm NL group, 6 out of 8 animals regenerated. The 500 pm wall thickness was chosen instead of 250 pm wall thickness due to the biodegradation observed after 4 months implanted. The wall of the tube was slightly degraded and even though the fibrin cable was intact, the stability of the wall for the thicker geometry is clearly superior. On top of that, the area of the regenerated nerve is obtained from hstolofical data is higher for the thicker wall thickness.

Example 2: Immunohistochemical tests with the hollow tube P(3HO-co-3HD)/PCL

Pre-clinical model in the rat sciatic nerve is used to assess regeneration through designed devices. In this section, immunohistochemical analysis was performed to analyze the presence of extracellular matrix supporting regeneration (Fibronectin positive matrix) and macrophage infiltration (Iba1 , cell type responsible of the degradation of implanted prototypes), Schwann cell presence and axonal growth throughout the tubular devices. Furthermore, measurements of wall thickness were performed after implantation to assess changes in diameter of implanted prototypes, as an indirect measurement of prototype degradation. Results demonstrated that P(3HO-co-3HD)/PCL of 500 pm resisted the 4 months of implantation. Neither fractures nor changes in the wall thickness were observed in the leached and no-leached materials.

Material and methods

Animals: Female Wistar rats (200-220 g) were used in this experiment. Animals were housed in plastic cages, maintained at 22°C in a 12 h light/dark cycle and given free access to water and food. The experimental procedures were approved by the Ethical Committee of our institution and followed the rules of the European Communities Council Directive.

Surgical procedure: All surgical procedures were performed with aseptic operating conditions and under pentobarbital/xylacine anaesthesia (40/10 mg/kg i.p.; respectively). The right sciatic nerve was exposed and cut 6 mm distal to the exit of the gluteal nerve. Then, in the case of the hollow guides 14 mm long tubes were implanted (for a 10 mm gap) with a 10-0 epineural suture. The tubes were filled with sterile physiological saline solution. In the autograft group, 10 mm of the right sciatic nerve was excised, flipped and sutured back to the remaining stumps with 10-0 epineural sutures. The muscle plane was sutured with resorbable 5-0 sutures, and the skin with 2-0 silk sutures. The wound was disinfected and the animals were treated with amitriptyline for preventing autotomy. The animals were randomly distributed in the different experimental groups as described in the preceding example.

Determination of changes in wall thickness:

Measurements of wall thickness pre-implantation and post-implantation were performed in hollow prototypes. Degradability of implanted devices was measured as a direct correlation of wall thickness lost after 4 months of implantation.

Immunohistochemistry

Regenerated nerves were processed for IHC analysis. Cross sectional sections (30 pm) of the regenerated nerves were used to label macrophage (lba-1 ) and the presence of extracellular matrix (Fibronectin). Statistical analysis:

GraphPad (GraphPad Software, USA) was used to perform the statistical analysis on the collected data. One-way analysis of variance with Bonferroni post hoc test was conducted to analyse differences on wall thickness change and biocompatibility of implanted prototypes. Data was considered significantly different when P < 0.05. Results:

Compatibility of P(3HO-co-3HD)/PCL prototypes in vivo:

Regenerated nerves were collected and sectioned in the cryostat (30 pm thick sections). Immunohistochemistry was performed to label fibronectin, a protein of the extracellular matrix needed for the formation of the fibrin cable which support nerve growth; Iba1 , a marker for macrophage, cell type needed for either the degradation of biodegradable components of the implanted devices; neurofilament 200, a marker for regenerated myelinated axons; and S1003, a marker for Schwann cells, responsible of guiding axon growth and myelination.

Figure 6 shows the positive staining for fibronectin and Iba1 in each implanted prototypes. The presence of the extracellular matrix (stained in green for fibronectin) demonstrate the biocompatibility of this material to support nerve growth, since this protein is essential for the formation of the fibrin cable in the first weeks after injury. This structure acts as a physical structure where the regenerated cable is supported and axons may grow easily.

Iba1 positive staining demonstrate that the tube is biodegradable, since this mechanism is the primary mechanism in which biopolymers are degraded in vivo.

Figure 7 show Schwann cells and axons distribution in the regenerated animals at 4 months after injury and repair. Higher magnification images, show that Schwann cell density was higher in P(3HO-co-3HD)/PCL 500 urn tubes than those with 250 urn.

Degradability of implanted P(3HO-co-3HDVPCL prototypes in vivo

4 months after injury and repair, implanted devices were collected and sectioned in the cryostat. Wall thickness was measured and % of wall thickness maintenance was analysed.

Results illustrated in Figure 8 reveal that after 4 months of implantation, P(3HO-co- 3HD)/PCL devices did not suffer from a relevant degradation. Slightly change in wall thickness was observed in Leached devices (either 500 urn or 250 urn), but no significant differences were observed between the groups. Example 3: In vitro analysis of electrospun fibres made with P(3HO):P(3HB) (50:50) using rat primary Schwann cells

For determining which biopolymer material or blend of the electrospun microfibres is most efficient in terms of nerve regeneration, the following assay was performed. In vitro analysis of candidate electrospun fibres was performed using rat primary Schwann cells. Schwann cells play an important role in peripheral nerve regeneration after injury, as well as maintaining the nerves. Rat primary Schwann cells were isolated and cultured via methods from Kaewkhaw et al (2012), in which the sciatic nerves were dissected and dissociated. Using a specialised DMEM media, Schwann cells could be expanded to higher passage numbers, whilst the media inhibited fibroblast growth. Schwann cells between passages 4-7 were used for these experiments. Live/dead analysis was used to confirm the cell viability of primary Schwann cells cultured on the candidate fibres and diameters, as well as flat material films.

Primary rat Schwann cells were cultured on candidate material flat polymers films for 6 days before staining to label live and dead cells. Candidate materials tested were: P(3HO):P(3HB) (50:50), P(3HO-co-3HD):PLLA (50:50) film and PCL film.

Figure 9 visually shows live and dead Schwann cells on samples, as well as confirming their predicted morphology and samples were confirmed as biocompatible.

Results show that the most elongated Schwann cells were seen on the tissue culture plastic control and the P(3HO):P(3HB) (50:50) polymer flat films.

Images were quantified for live cells against dead cell numbers. Cell viability was determined, as a percentage, in which all polymer samples were confirmed as biocompatible. P(3HO):P(3HB) (50:50) had the highest cell viability of 96.29 ± 2.14%, followed by P(3HO-co-3HD):PLLA (50:50) film and PCL film, 93.20 ± 5.88% and 91.62 ± 5.95% respectively.

Results in this example show the effectiveness of using the blend P(3HO):P(3HB) over other biopolymer blends for nerve regeneration purposes.

Example 4: In vitro analysis of electrospun fibres using co-cultures of NG108-15 Neuronal cells and rat primary Schwann cells.

For determining which biopolymer material or blend of the electrospun microfibres is most efficient in terms of nerve regeneration, the following assay was performed. 40,000 rat primary Schwann cells were cultured onto candidate material flat films and candidate electrospun fibres for 7 days in specialized Schwann cell media. (DMEM D- valine media, Biosera, containing 2 mM glutamine, 10% FCS, 1% N2 supplement, 20 pg/ml bovine pituitary extract, 5 mM forskolin, 1% penicillin/streptomycin and 0.25% amphotericin) On day 7, Schwann cell media was removed and 6000 NG108-15 neuronal cells were seeded onto samples. A co-culture media, (DMEM (Biosera) and Ham’s F12 medium (Biosera) at a 1 :1 ratio (v/v) was supplemented with 1 % L- glutamine, 1 % penicillin/streptomycin, and 1 % N2 supplement) was added to samples and replaced every 2 days. Co-cultures were left for 6 days, before being labelled with 0.001% Syto-9 (Invitrogen) and 0.0015% propidium iodide (Invitrogen) for live/dead analysis, as well as being fixed and immunolabelled for DAPI (cell nuclei), b III tubulin (for neurite outgrowth) and S100 (Schwann cell marker).

As shown in Figure 10, both rat primary Schwann cells and NG108-15 neuronal cells adhere and align themselves to the fibres. The highest number of live cells were cultured on the P(3HO):P(3HB) (50:50) 8pm fibres, 269.55 ± 1.12 cells, compared to the other materials and fibre sizes.

As shown in Figure 11 , in NG108-15 neuronal cell cultured with the addition of Schwann cells, the highest average neurite length was also found on 5pm P(3HO):P(3HB) fibres.

Results in this example show the effectiveness of using the blend P(3HO):P(3HB) over other biopolymer blends for nerve regeneration purposes.

Example 5: In vivo testing of microchanneled inner tube made of methacrylated PCL for nerve regeneration

Nerve regeneration of the microchanneled inner tube obtained according to the present invention was tested in the following assay. For that purpose, female Wistar rats (200- 220 g) were used. The right sciatic nerve was exposed and cut 6 mm distal to the exit of the gluteal nerve. Then, in the case of the hollow guides 8 mm long tubes were implanted (for a 6 mm gap). In the case of the PCLm tube, 6 mm long devices were implanted. Animals were followed-up for 4 weeks. After this time, the right sciatic nerves were exposed and cut 6 mm distal to the distal end of the silicone tube. Then, the proximal cut stump was dipped in 2 pi of the tracers (True Blue at the tibial branch and Dil at the peroneal branch) placed in a vaseline pool for 1 h. One week later, animals were perfused and DRG (L4, L5), lumbar spinal cord section, and the regenerated sciatic nerves were obtained for their analysis. Experimental design and qualitative observations.

Table 4. Summary of experimental groups, number of implanted tubes and observations found during the implantation process.

As shown in Fig. 12, immunohistochemistry was performed to label fibronectin, a protein of the extracellular matrix needed for the formation of the fibrin cable which support nerve growth; Iba1 , a marker for macrophage, cell type needed for either the degradation of biodegradable components of the implanted devices; neurofilament 200, a marker for regenerated myelinated axons; and S1003, a marker for Schwann cells, responsible of guiding axon growth and myelinate them. The fluorescent signal coming from the inner part of the channels shows how fibrin protein, myelinated regenerated axons and Schwann cells clearly exist in the inner part of the microchannels. Then, the PCLm microchannels support clearly nerve regeneration. Example 6: Guidance conduit consisting of a hollow tube incorporating an inner microstructure in the form of a microchanneled inner tube

Pairs of dorsal root ganglion (DRGs, 26 maximum), which lie just outside the spinal column have been shown to regenerate in vitro and comprise of multiple cell types (neurons and Schwann cells) known to co-ordinate the regenerative response seen in the peripheral nervous system. Therefore utilisation of whole tissue explants (referred to as‘ex vivo’ tissue culture) of DRG in combination with development of a 3D culture system is more likely to mirror the in vivo environment.

The testing was carried out by adapting a P200 pipette tip box to perform 3D ex vivo cell culture via filling with media to the level of the tips, inserting nerve guide conduits into tips, placing a DRG on top and adding media above this and placing into a humidified incubator perfused with 5% C02 (Figure 14).

In the ex-vivo protocol performed, 6 mm long P(3HO-3HD)/PCL 75/25 porous hollow tubes and an inner part comprised of PCLm with porosity created by both glucose (60% in weight) and high internal phase emulsion were used as candidates for ex-vivo nerve regeneration.

For validation of the ex-vivo model, Enhanced Antibody Computed Tomography (ENACT) was used. This approach involves using lanthanide conjugated antibodies which bind to FITC (Fluorescein isothiocyanate). These antibodies can then be used in whole mount antibody staining by the addition of a step to the staining procedure, whilst removing the necessity to embed or section any of the constructs.

Using the nerve regeneration platform. (Figure 14), DRGs were placed inside the produced conduits and the ENACT approach which permits the specific detection of beta III tubulin positive neurons throughout the sample in a non-destructive manner was carried out. Thus, primary antibodies raised against beta III tubulin were applied to the sample, followed by FITC conjugated antibodies to which a lanthanide conjugated antibody binds and thereby permits contrast within Computed Tommography. As it can be observed in Figure 15, there is a clear evidence of neural like extensions through the tube length. This behavior is observed for both degree of porosity, created by glucose porogen and by high internal phase emulsion.

Example 7: Guidance conduit consisting of a hollow tube incorporating an inner structure in the form of a bundle of threaded microfibres. Pairs of dorsal root ganglion (DRGs, 26 maximum), which lie just outside the spinal column have been shown to regenerate in vitro and comprise of multiple cell types (neurons and Schwann cells) known to co-ordinate the regenerative response seen in the peripheral nervous system. Therefore utilisation of whole tissue explants (referred to as‘ex vivo’ tissue culture) of DRG in combination with development of a 3D culture system is more likely to mirror the in vivo environment.

The testing was carried out by adapting a P200 pipette tip box to perform 3D ex vivo cell culture via filling with media to the level of the tips, inserting nerve guide conduits into tips, placing a DRG on top and adding media above this and placing into a humidified incubator perfused with 5% C02 (Figure 14).

In the ex-vivo protocol performed, different electrospun fibres were threaded into 6 mm P(3HO-3HD)/PCL 75/25 porous hollow tubes. Later, fibres were removed carefully as accurate as possible from the hollow tubes and performing immunostaining and confocal analysis directly on the fibres (Figure 16). This approach yielded useful information including whether a material can potentially support DRG adhesion, regeneration, maximum cell outgrowth distance and the extent of approximate migration of Schwann and neuronal cells. Qualitatively, neurons and Schwann cells were closely associated with DAPI positive cells, which progressed furthest through the conduit. There were technical challenges trying to accurately measure either neuronal or Schwann cell outgrowth as based on beta III tubulin and S100 respectively, due to autofluorescence of electrospun materials and the complex dimensionality of the samples when compared to conventional flat films. Therefore, the maximum distance of migration as based on DAPI positive nuclei throughout immobilised electrospun fibres removed from the conduit was used as a surrogate measure for neural regeneration.

When comparing different candidate materials 5pm P(3HO):P(3HB)(50:50) was found to perform significantly better than other materials when compared against 5 and 8pm PCL, 8 pm P(3HO):P(3HB)(50:50), 4pm P(3HB):P(3HO-co-3HD) and gyrated PHA. It is due to the larger cell migration distance obtained when these microfibres are threaded into the hollow tube.

Claims

1.- Implantable nerve guidance conduit comprising

- a hollow tube, and optionally,

- an inner microstructure,

characterized in that the hollow tube is made of a blend of biopolymers comprising a copolymer of poly(3-hydroxyoctanoate-co-3-hydroxydecanoate) and poly(e- caprolactone).

2.- Implantable nerve guidance conduit according to claim 1 , characterized in that the blend of biopolymers of the hollow tube comprises from 60 to 95% by weight of a copolymer of poly(3-hydroxyoctanoate-co-3-hydroxydecanoate) and from 5 to 40% by weight of poly(e-caprolactone).

3.- Implantable nerve guidance conduit according to claim 2, characterized in that the blend of biopolymers of the hollow tube comprises 75% by weight of a copolymer of poly(3-hydroxyoctanoate-co-3-hydroxydecanoate) and 25% by weight of poly(e- caprolactone).

4.- Implantable nerve guidance conduit according to any previous claim, characterized in that the hollow tube presents a Young’s Modulus of from 0.5 - 20MPa.

5.- Implantable nerve guidance conduit according to any previous claim, characterized in that the wall of the hollow tube is porous.

6. - Implantable nerve guidance conduit according to claim 5, characterized in that the pore size is froml to 50pm.

7.- Implantable nerve guidance conduit according to any of claims 5 and 6, characterized in that the volume of pores in the wall of the hollow tube is from 40 to 80%.

8.- Implantable nerve guidance conduit according to any previous claim, characterized in that the inner diameter of the hollow tube is from 0.5 to 20mm.

9.- Implantable nerve guidance conduit according to any previous claim, characterized in that the thickness of the wall of the hollow tube is from 0.1 to 0.8mm.

10.- Implantable nerve guidance conduit according to any previous claim, characterized in that the length of the hollow tube is up to 50 mm.

11.- Implantable nerve guidance conduit according to any previous claim, characterized in that the inner microstructure is in the form of a microchanneled inner tube or a bundle of threaded microfibres.

12.- Implantable nerve guidance conduit according to claim 1 1 , characterized in that the inner microstructure is in the form of a microchanneled inner tube having from 2 to 50 aligned parallel microchannels with each a diameter of 50 to 500pm, and separated from each other by 50 to 500pm.

13.- Implantable nerve guidance tube conduit according to any of claims 11 and 12, characterized in that the inner tube is made of a biopolymer comprising methacrylated polycaprolactone.

14.- Implantable nerve guidance tube conduit according to any of claims 1 1 to 13, characterized in that the microchanneled inner tube is porous.

15.- Implantable nerve guidance conduit according to claim 14, characterized in that the pore size is from 10 to 70pm.

16.- Implantable nerve guidance tube conduit according to claim 14, characterized in that the microchanneled inner tube has a volume of pores of 40 to 80%.

17.-lmplantable nerve guidance conduit tube according to any previous claim, characterized in that the microchanneled inner tube presents a Young’s Modulus of from 0,05 - 5 MPa.

18.- Implantable nerve guidance tube conduit according to any of claims 12 to 17, characterized in that the inner tube has a diameter of 0,5 to 20mm.

19.- Implantable nerve guidance tube conduit according to any of claims 12 to 18, characterized in that the inner tube has a length from 5 to 50mm.

20.- Implantable nerve guidance conduit according to claim 1 1 , characterized in that the inner microstructure is in the form of a bundle of microfibres.

21.- Implantable nerve guidance conduit according to claim 20, characterized in that the microfibres are made of a blend of biopolymers comprising poly (3- hydroxyoctanoate) and poly (3-hydroxybutyrate).

22.- Implantable nerve guidance conduit according to claim 21 , characterized in that the blend of biopolymers comprises from 25 to 75% by weight of poly (3- hydroxyoctanoate) and from 25 to 75% by weight of poly (3-hydroxybutyrate).

23. - Implantable nerve guidance conduit according to claim 22, characterized in that the blend of biopolymers consists of 50% by weight of poly (3-hydroxyoctanoate) and 50% by weight of poly (3-hydroxybutyrate).

24.- Implantable nerve guidance conduit according to any of claims 20 to 23, characterized in that the microfibres have a diameter from 5 to 10pm.

25.- A method of obtaining the implantable nerve guidance conduit of claims 1 to 24 comprising:

- obtaining a hollow tube by microextrusion, and optionally,

- obtaining an inner microstructure in the form of a microchanneled inner tube or a mat of microfibres, and

- introducing the microchanneled inner tube or mat of microfibres into the hollow tube.

26.- A method according to claim 25, characterized in that obtaining a hollow tube by microextrusion comprises the following steps:

- preparing a blend of biopolymers comprising a copolymer of poly (3- hydroxyoctanoate-co-3-hydroxydecanoate) and poly( -caprolactone) and plasticizing the blend at a temperature of 90 - 120°C by means of mechanical shaker in the extruder system, - pressing the blend out of the extrusion die head with the desired hollow tube configuration - cooling and solidifying the hollow tube extrudate at ambient temperature for several minutes

- pulling, cutting the extrudate and collecting the finished hollow tubes.

27.- A method according to claim 26, characterized in that the blend of biopolymers of poly(3-hydroxyoctanoate-co-3-hydroxydecanoate) and poly(e-caprolactone) comprises additionally a porogen selected from the group consisting of glucose and NaCI.

28.- A method according to claim 27, characterized in that the blend of biopolymers comprises 30% by weight of poly(3-hydroxyoctanoate-co-3-hydroxydecanoate) and poly(e-caprolactone) and 70% by weight of porogen selected from the group consisting of glucose and NaCI.

29.- A method according to claims 25, characterized in that obtaining an inner microstructure in the form of a microchanneled tube is done by means of UV-3D micromoulding and comprises the following steps:

- injection of a mixture comprising polycaprolactone methacrylated in a negative mould, comprising stretched stainless steel wires having from 50 to 500pm (pref200pm) diameters, aligned in a silicone container with a diameter from 5 to 20mm.

- curing of the biopolymer inside it and mould disassembling.

30.- A method according to claim 29, characterized in that the mixture comprising methacrylated polycaprolactone further contains glucose as porogen to create porosity and a photoinitiator to cure the polymer by UV-VIS exposure.

31.- A method according to claim 29, characterized in that the mixture comprising methacrylated polycaprolactone further contains a surfactant, chloroform, toluene, a photoinitiator and water droplets at 35°C are added under vigorous stirring until an emulsion is obtained to create porosity through High Internal Phase Emulsion Templating (polyHIPE)

32.- A method according to claim 25, characterized in that obtaining an inner microstructure in the form of a mat of microfibres is done by means of electrospinning and comprises the following steps: preparing a solution comprising poly(3-hydroxyoctanoate) and poly(3- hydroxybutyrate),

- electrospinning the solution at a flow rate of from 1 to 10ml/hr, speed of from 1000 to 5000rpm and voltage of 1 to 20kV to obtain electrospun fibres, - laser cutting the electrospun fibres to obtain mat pieces of 100 mm width and 6 mm length.

33.- A method according to claims 25 to 28 and 32, characterized in that introducing the mat of microfibres into the hollow tube comprises:

- threading fibres into the hollow tube, by tightly wrapping around a syringe, a 100 x 6 mm fibre mat with the fibres parallel to the needle, placing the fibres into the tube, and removing off the needle by a sharp pair of forceps.

34.- Implantable nerve guidance conduit according to any of claims 1 to 24 for use in medical treatment or diagnosis.

35.- I mplantable nerve guidance conduit for use according to claim 34 for treating a nerve injury in a human.

36.- I mplantable nerve guidance conduit for use according to claim 35, for treating a peripheral nerve injury.

37.- Implantable nerve guidance conduit for use according to claims 35 and 36, for reconstructing nerve gaps of up to 50mm.