WO2001070293A1

WO2001070293A1 - Polymeric composite materials and their manufacture

Info

Publication number: WO2001070293A1
Application number: PCT/GB2001/001177
Authority: WO
Inventors: Allan Gerald Arthur Coombes; Sandra Downes; Martin Griffin
Original assignee: The University Of Nottingham; The Nottingham Trent University
Priority date: 2000-03-18
Filing date: 2001-03-19
Publication date: 2001-09-27
Also published as: AU4087801A; GB0006439D0

Abstract

A method for the preparation of a polymeric composite material comprises the steps of a) forming a porous body of a first polymer; b) impregnating said porous body with a solution of a second polymer; and c) causing or allowing solvent to evaporate from said body. The first polymer is preferably a natural polymer, eg collagen, and the second polymer is preferably a synthetic polymer, eg a polymer selected from the group consisting of poly(α-hydroxy acid) such as polylactide, poly(DL lactide coglycolide), poly(ε-caprolactone), polyorthoesters, polyphosphazines, hyaluronic acid esters, polyanhydrides, copolymers of such polymers and blends thereof.

Description

Title: Polymeric Composite Materials and their Manufacture

This invention relates to composite polymeric materials of particular utility in medical and biological applications including tissue engineering.

Tissue engineering continues to attract considerable and growing interest from researchers across a wide range of disciplines because of the potential for providing improved biomaterials for hard and soft tissue repair and implantable devices for mimicing the function of organs such as the liver and pancreas. Production of scaffolds to support and encourage cell development and correct function depends critically on the design, physico-chemical nature and the material architecture (micro/macroporosity). Autograft and banked bone are the best materials for repairing hard tissue but there are problems of supply and spread of infection respectively. Collagen is currently the most popular material for scaffold production in connective tissue repair but is difficult to formulate reproducibly. Concerns also exist over the immune response to implanted collagen.

Three dimensional collagen matrices incorporating various bioactive materials such as glycosaminoglycans (GAGs, eg chondroitin-6-sulphate), fibronectin and hyaluronate have been investigated extensively as templates for regeneration of dermal tissue. Such materials have also been formulated to incorporate RGD- containing peptides (the integrin-mediated cell attachment domain found in many extracellular matrix proteins, RGD- being arginine-glycine-aspartine-), the aim being to promote cell attachment and thus wound healing (Grzesiak et al, 1997).

Hyaluronic acid (HA) is a polysaccharide made up of molecules of N-acetylglucosamine and D-glucuronic acid. HA is involved in cell migration, adhesion, aggregation, proliferation and cell function and has been widely investigated for modulating cell-biomaterial interactions. HA dissolves on contact with biological fluids. As a result, thermal crosslinking has been applied to blends of HA and carboxyl-containing polymers such as polyacrylamide (PAA) to create stabilising, intermolecular bridges. Natural polymers offer advantages of good biocompatibility but their use as biomaterials is often limited by their poor mechanical properties. The need to preserve biological properties also complicates their formulation into biomaterials. Native collagen, for example, possesses good tensile properties but these are reduced by the chemical processes used to isolate the material which results in non-fibrous materials. Biodegradation, post-implantation, also results in rapid deterioration of structural and biological properties.

Blends, interpenetrating polymer networks (IPNs) and composites of collagen with synthetic materials have been widely investigated in attempts to overcome the deficiencies of collagen as a biomaterial.

Collagen matrices (used for repair of connective tissue such as skin) are generally chemically crosslinked (eg using glutaraldehyde) to improve biological stability, ease of handling and mechanical properties (Grzesiak et al, 1997). This process can cause problems of toxicity due to residual unreacted or partially reacted crosslinking agent.

The relative advantages and disadvantages of natural and synthetic polymers as biomaterials are exemplified by the case of artificial skin grafts. Composite skin replacements have been described comprising human keratinocytes and fibroblasts seeded on a collagen-glycosaminoglycan (GAG) matrix. The collagen matrix was expected to function as a dermal template allowing vascular ingrowth from the wound bed. The collagen-GAG matrix is however highly susceptible to bacterial and enzymatic attack and is generally unsuccessful when applied to chronic wounds heavily colonised with bacteria (Hansbrough, 1994).

Improved dermal matrix has been developed by culturing human fibroblasts on biodegradable synthetic polymer mesh produced from polyglactin 9-10 (Vicryl). The material has been termed Dermagraft by the manufacturers (Advanced Tissue Sciences, La Jolla, California). The fibroblasts secrete proteins and glycoproteins as they grow in the Vicryl mesh, forming an extracellular matrix which fills the mesh interstices. Composite structures consisting of Dermagraft and overlying meshed skin grafts are expected to be resistant to wound proteases and thus allow more efficient wound closure.

Biocomposites comprising synthetic polymers such as poly(vinyl alcohol) (PVA) or poly(2- hydroxyethylmethacrylate) [poly(HEMA)] and natural polymers such as collagen and gelatin have been widely investigated with the aim of exploiting the advantageous properties of each component and compensating for less desirable characteristics (Giusti et al, 1993(a)). For example, the poor cell adhesion normally associated with poly(HEMA) has been mitigated by blending the synthetic polymer with collagen (Santin et al, 1996). This approach subsequently enables application of the biocomposite for tissue engineering or biological scaffolding where growing cells are supported during wound repair to encourage integration of host tissue and implant.

Biocomposites of synthetic polymers and natural polymers (or synthetic analogues thereof) benefit potentially from the wide range of mechanical properties and processing techniques applicable to synthetic polymers. These advantages may be augmented by the biocompatibility of natural polymers and the wide diversity of biological reactions which may be induced by contact between the natural polymer and host cells and tissue.

A decided disadvantage of many existing biocomposites of synthetic and natural polymers, however, concerns the requirement for chemical crosslinking of the components during manufacture which can cause problems of toxicity due to residual catalysts, initiators and unreacted or partially reacted crosslinking agent. The toxic effect of glutaraldehyde residues following crosslinking of proteins is well documented.

Hydrogels are three dimensional polymeric networks generally stabilised by covalent crosslinking and weak cohesive forces, particularly hydrogen bonds. These networks imbibe large quantities of water or organic liquids without dissolution. In turn, the large water content imparts a very low interfacial tension with biological fluids. This feature, along with high permeability to small molecules (such as tissue metabolites) and their viscoelastic nature makes hydrogels behave similarly to biological tissues. As a result, they have been used as biomaterials in a wide variety of biomedical applications such as opthalmology, drug delivery and orthopaedics.

Freeze drying of hydrogels such as poly(HEMA) has been used to form macroporous sponges for subsequent impregnation by natural polymers (Santin et al, 1996).

IPNs may be described as multi-phase materials wherein each phase is independent, continuous and contacts all portions of the sample space. IPNs are often manufactured by a sequential method involving swelling of a pre-synthesised network in a solution of the second polymer to be crosslinked. A high porosity of the first network is desirable to achieve satisfactory loading or impregnation of the second polymer.

IPNs based on poly(HEMA) and gelatin have been described (Santin et al, 1996). They were prepared by impregnating freeze-dried, crosslinked, poly(HEMA) hydrogel films (0.5-3.0 mm thick) with gelatin solution. The gelatin phase was subsequently crosslinked using glutaraldehyde.

IPNs of fibrin and polyurethane have been prepared by spraying a suspension of the natural and synthetic polymers simultaneously with water onto a substrate. The fibrinogen was subsequently covalently crosslinked using thrombin, factor XIII and calcium ions (Giusti et al 1993(a)).

Blends of crosslinked fibrinogen and PVA, obtained by mixing aqueous solutions of the respective polymers, have been converted to hydrogels by repeated freeze- thawing (Giusti et al 1993(a)).

PVA-collagen blends may also be produced by repeated freeze-thawing of a solution containing both polymers (Giusti et al 1993(b)). The freeze-thaw process results in the formation of crystallites of PVA which act as crosslinking sites between polymer chains and gives rise to hydrogel formation. The hydrogel is capable of entrapping the natural polymer (eg collagen) within the PVA network.

However, PVA is not biodegradable and its use for implant manufacture has been associated with adverse reactions in vivo.

There has now been devised a method for the preparation of polymeric composite materials, and composite materials so produced, which overcomes or substantially mitigates the above-mentioned or other disadvantages of the prior art.

According to a first aspect of the invention, there is provided a method for the preparation of a polymeric composite material, which method comprises the steps of a) forming a porous body of a first polymer; b) impregnating said porous body with a solution of a second polymer; and c) causing or allowing solvent to evaporate from said body.

Preferably, one of the first and second polymers is a natural polymer or a synthetic analogue thereof and the other is a synthetic polymer. Most preferably, it is the first polymer which is the natural polymer or synthetic analogue and the second polymer is synthetic.

The porous body of first polymer is preferably prepared by forming a solution, eg an aqueous solution, of the first polymer and lyophilising that solution. The body prepared in this manner will commonly have the form of a porous mat.

The method of biocomposite manufacture avoids chemical, thermal or irradiation- induced stabilisation of the natural polymer, thereby eliminating structural damage or modification due to crosslinking. The natural polymer phase is expected to be stabilised due to localised coating by the synthetic polymer phase.

The synthetic polymer component may be produced from biodegradable polymers such as polylactide which potentially allows complete replacement of the implant by repair tissue. In addition, the synthetic polymer is not crosslinked during biocomposite manufacture to confer structural stability, which eliminates problems of toxicity due to residual crosslinking agent.

The solvent used to form the second polymer solution for impregnation of the porous body may be any suitable solvent in which the second polymer is sufficiently soluble and which is sufficiently volatile subsequently to be removed by evaporation or sublimation.

The structural properties of the biocomposite enables applications for bone repair where some load bearing role can be expected and is indeed desirable to expose the repair tissue to stress fields for optimal development (the bone remodelling phase of bone repair).

Consolidation and compaction of the biocomposite is achieved during solvent evaporation removing the need for a separate compaction stage.

Variation in exposure/presentation of the natural polymer can be achieved by controlling process conditions such as concentration and volume of the polymer solutions. This facility should also be useful for controlling the delivery of bioactive factors (eg growth factors) incorporated in the biocomposite.

The open porous structure of the biocomposites, having pore sizes ranging from 50-1 OOμm, is expected to facilitate cell ingrowth and good integration of the biocomposite with the host tissue.

Large cell adhesion proteins such as fibronectin are considered to adsorb at a low packing density at biomaterial surfaces which consequently limits the number of attached cells. The use of suitably immobilised RGD sequences may allow higher packing at the material surface and hence increase the number of attached cells. The incorporation of peptide fragments containing cell adhesion (RGD) sequences in biocomposites is facilitated by blending with the natural polymer component or by post-adsorption. The invention is anticipated to be generally useful for producing biomaterials and biodegradable scaffolds or templates for supporting cell attachment and spreading, thereby encouraging growth and repair of hard and soft tissue. Applications are also foreseen in drug delivery.

Specific areas of use could include implants for bone and cartilage repair, bone graft substitutes, cardiovascular devices, nerve guides, connective tissue repair and artificial skin grafts, wound and burn dressings and controlled release systems for delivery of bioactive materials such as steroids, oligonucleotides and DNA, and growth factors.

The biocomposites produced by the method of the invention could find application as bioactive constructs for tissue engineering where controlled release of growth factors encourages and guides tissue repair.

Development of synthetic alternatives to bone graft would eliminate the risk of spreading Creutzfeldt-Jacob disease and AIDS and reduce the need for bone banking. The biocomposites of the invention may be useful as a replacement for allogeneic bone obtained from bone banks.

The method for preparing the biocomposites is expected to be useful for coating glass and plasticware for cell culture. Applications in drug delivery such as transdermal administration of hormones via patch-type devices are also envisaged.

Natural polymers envisaged as being useful in the method include gelatin and extra cellular matrix proteins (collagen, elastin, laminin), cell adhesion proteins (such as fibronectin, vitronectin, vinculin, fibrinogen), polysaccharides (eg hyaluronic acid, heparin), glycosaminoglycans (such as chondroitin-4-sulphate) and combinations of natural polymers or natural polymer-synthetic polymer conjugates. The presently preferred natural polymer is collagen.

Synthetic analogues of natural polymers such as silk-like and elastin-like protein polymers (Capello, 1997) may be substituted in total or in part for the natural polymer.

Heparin-like synthetic polymers (Miggoney et al, 1988) are further examples of synthetic analogues of natural polymers.

Synthetic polymers include poly(α-hydroxy acid) such as polylactide, poly(DL lactide co-glycolide), poly(ε-caprolactone), polyorthoesters, polyphosphazines, hyaluronic acid esters, polyanhydrides, copolymers of the above polymers and blends.

Suitable polymers may thus be members of the class of polyesters formed by ring- opening polymerisation. Precursors to such polymers may thus have the generic formula

in which R represents an optionally substituted alkylene chain, eg a chain (CH₂)_n in which n is an integer of from about 4 to 10. In the particularly preferred case of ε- polycaprolactone, n is 5.

Another specific group of suitable polymer precursors are those represented by the generic formula

in which R¹ and R², which may be the same or different, represent optionally substituted lower alkyl groups, ie alkyl groups of 1 to 6 carbon atoms. In one preferred case, at least one, and preferably both, of R¹ and R² represents methyl.

The presently preferred synthetic polymer is poly(ε-caprolactone), a biodegradable polymer from the same family of poly (α-hydroxy acids) as polylactide (PLA) and polyglycolide (PGA).

Polycaprolactone (PCL) is characterised by a slower degradation rate and increased permeability relative to PLA and PGA. However, PCL has not been investigated for biomedical implants and drug delivery systems to the same degree as PLA and PGA, possibly because of its lower mechanical properties and longer degradation times.

According to a second aspect of the invention, there is provided a composite polymeric material prepared by the method defined above.

The characteristics of the composite material produced in accordance with the invention will generally depend on factors such as the first polymer : second polymer weight ratio. That ratio may vary widely but for many applications a first polymer : second polymer (eg collagen : PCL) ratio of 1 :40 or less (eg 1 :8 or 1 :4) is preferred. It is found that biocomposites having such a composition may have a highly porous morphology which would be expected to facilitate ingress of, for instance, enzyme solutions, and a high degree of interaction with proteins or cells contacting the surface.

The invention will now be described in greater detail, by way of illustration only, with reference to the following Examples.

Example 1

Production of collagen.polycaprolactone biocomposites

Collagen:PCL biocomposites were produced by impregnation of lyophilised collagen mats using a solution of PCL in dichloromethane (DCM), followed by solvent evaporation. The process stages are described in detail below.

Preparation of lyophilised collagen mats

Collagen solutions (0.25, 0.5 and 1% w/v) were prepared by dissolving Type 1 , acid-soluble collagen from calf skin (Sigma C-3511) in 1 % acetic acid. The pH was adjusted to 2.9 using 0.1M NaOH and dissolution was facilitated by stirring with a magnetic stirrer overnight at room temperature.

Aliquots (1 , 2 or 4 ml) of the collagen solution were added to 7ml, glass squat vials and freeze dried for 24 hours.

1.1.2 Solution impregnation of collagen mats

2ml of a solution of PCL in DCM (0.5, 1 , 2, 5 and 10% w/v) were added to freeze dried collagen mats (prepared using a 0.25% collagen solution) which were retained in 7ml squat vials. The vials were kept stoppered for 30-60 minutes before allowing solvent evaporation overnight. The characteristics of the resulting biocomposite may be controlled to a large extent simply by variation of the collagen:PCL ratio and collagen solution volume. Further investigations concentrated on use of a 0.25% collagen solution since this resulted in an open mat structure on freeze drying which was expected to facilitate impregnation by polymer solutions. However there is wide scope for fine adjustment of the properties of collagen:PCL blends by variation of the basic formulation approach described here. In addition, substitution of PCL by other biodegradable polymers such as PLA and PLG should extend the range of useful properties and applications for the biocomposites.

1.2 SEM examination of collagen:PCL biocomposites

Samples of collagen:PCL biocomposites were attached to aluminium SEM stubs using carbon tabs (Agar Scientific). Silver electrodag (Acheson Electrodag 1415M, Agar Scientific) was applied to improve conductivity. Specimens were sputter coated with gold prior to examination using a JEOL 6400 SEM. The principal morphological features of the collagen:PCL biocomposites are summarised in Table 1. Scanning electron micrographs of collagen:PCL biocomposites are shown in Figure 1.

SEM examination of collagen:PCL biocomposites revealed the significant influence of formulation variables on material morphology which in turn, and importantly, affects the presentation of collagen. In particular, the increased porosity and lower collagen content of the freeze dried mats prepared using 1ml of collagen solution resulted in more efficient impregnation by low concentration PCL solutions and extensive coverage of the collagen phase. Table 1 The morphological features of collagen:PCL biocomposites revealed by SEM.

^* 2ml PCL solution used to impregnate collagen mats

1.3 Estimation of collagen exposure in collagen: PCL biocomposites using a collagenase-based assay.

The extent of exposure of collagen at the surface of collagen:PCL biocomposites can be expected to exert a major influence on the interaction of cells with the biomaterial through, for example, binding of fibronectin or related cell adhesion proteins. In addition, controlled changes in coating efficiency may also be used to influence the pattern of release or presentation of co-factors such as peptide fragments or growth factors. SEM analysis clearly demonstrated the changes in morphology which could be achieved by variation of processing parameters such as collagen: PCL ratio. A collagenase digestion assay was applied to the collagen:PCL biocomposites to provide further insights regarding collagen presentation/exposure.

1.3.1 Method

Collagen:PCL biocomposites were prepared in 7ml squat vials by freeze drying 2ml, 0.25%, collagen solution and impregnating the dried mat with 2ml PCL solution. The materials were washed in PBS for 4 hours and left immersed in fresh PBS for 48 hours prior to testing for exposed collagen.

Collagen:PCL biocomposites were also prepared in 4ml glass shell vials by freeze drying 0.5ml, 0.25% collagen solution and impregnating the dried mat with 0.5 ml of PCL solution. The materials were washed in PBS and left immersed in fresh PBS overnight.

Samples were cut from the mats and analysed for exposed/presented collagen using a collagenase digestion technique.

In brief, the BCA total protein assay (Sigma) was used to measure the amount of collagen digested from collagen:PCL biocomposites after incubation in enzyme solution at 37°C.

Collagen calibration samples (0.5-2.0mg) were also added to 2ml of digestion medium (0.1 mg/ml collagenase solution in HBSS) and retained at 37°C in 20ml glass vials until dissolved (2.5-3.0 hours). Test and calibration samples were allowed to cool to room temperature and tested immediately using the BCA total protein assay.

BCA reagent (2ml bicinchoninic acid solution, 40μl copper II sulphate) was added to 100μl aliquots of the digestion solution and retained in a water bath at 37°C for approximately 30 minutes. The absorbance at 562nm was recorded using a UV spectrophotometer (Unicam UV/VIS Spectrometer, UV4) and used to construct a calibration curve.

Table 2 Estimation of collagen exposure in collagen:PCL biocomposites by collagenase digestion

1.3.2 Results

The results presented in Table 2 indicate that a major fraction (approximately 70- 100%) of the collagen content of 1 :4 and 1 :8 collagen:PCL biocomposites is accessible for digestion by collagenase indicating a high degree of collagen exposure/presentation for interaction with other ECM proteins or cells contacting the biomaterial surface.

These findings are in line with the SEM examination of collagen:PCL biocomposites (Figure 1) which revealed a highly porous morphology for the 1 :4 and 1 :8 blends which would be expected to facilitate ingress of enzyme solutions and collagen digestion. A much reduced collagen presentation was measured for the 1 :20 collagen:PCL materials (35-50%), again consistent with the SEM analysis which revealed virtually complete coverage of the collagen component by PCL (Figure 1).

Example 2

Incorporation of Fibronectin (FN) in Collagen/PCL Biocomposites

2.1 50ug and IQO ig FN content a. Collagen solution (0.25% w/v) was prepared by dissolution in 1 % acetic acid adjusted to pH 2.7 using 0.1 M NaOH.

b. The collagen solution was adjusted to a pH of 7.4 using 1 M NaOH followed by O.IM NaOH.

c. 100 μl and 50μl of 1 mg/ml fibronectin solution (from bovine plasma, Sigma F-4759) respectively were added to 0.5 ml aliquots of collagen solution in 4ml glass vials.

d. The samples were held at room temperature for 2 hours to allow gel formation.

e. The gels were frozen by holding at -80°C for 1-2 hours prior to freeze- drying, solution impregnation and solvent evaporation as described in Example 1.

2.2 25, 12,5 and 6.25 ιιg FN content

a. Collagen solution (0.25% w/v) was prepared by dissolution in 1 % acetic acid adjusted to pH 2.7 using 0.1 M NaOH.

b. The collagen solution was adjusted to a pH of 6.2 using 1M NaOH followed by O.IM NaOH. c. 100μl aliquots of solution containing 25 μg, 12.5 μg and 6.25 μg FN respectively, were added to 0.4 ml aliquots of collagen solution in 4ml glass vials.

d. The samples were frozen immediately prior to freeze-drying, solution impregnation and solvent evaporation as described in Example 1.

Example 3

Production of biocomposites of collagen, chondroitin sulphate (ChSO₄) and PCL

3. 1 Background

Hyaluronic acid (HA) is a natural glycosaminoglycan (GAG) widely distributed in animal tissues and is also found in the synovial fluid and the vitreous and aqueous humors of the eye. HA has an isoelectric point of 8.6. Its sulphation provides the polysaccharide with good anticoagulant activity and improves its haemocompatibility. Dermatan and chondroitin sulphate are natural sulphated glycosaminoglycans. GAGs are found in connective tissues at concentrations of less than 10% by weight of the fibrous proteins. They form porous hydrated gels and GAG chains fill most of the extracellular matrix space, providing mechanical support to tissues, while still allowing diffusion of molecules, and cell migration.

3.2 Materials and Formulation

Collagen/GAG/PCL biocomposites were prepared with collagen/CHS0₄ ratios of 1/1(50% GAG), 4/1(20% GAG), 8/1 (11%) and 16/1(6%). Ch₆SO4 does not bind to reconstituted collagen under physiological conditions (Hanthamrongwit et al, 1986). Therefore, collagen/ChSO₄ matrices were prepared at 3 different pH values in an attempt to promote collagen/GAG interaction and complexation and to investigate the effect on cell-biocomposite interaction.

Chondroitin sulphate A (from bovine trachea, approx, 70%, balance is chondroitin sulphate C) was obtained from Sigma (C-8529). Hyaluronic acid from bovine vitreous humor was also obtained from Sigma (H-7630). 3.2.1 20% ChSQ₄, system, pH 2.7

a. Collagen solution (0.5% w/v) in 1% acetic acid (pH 2.7) was added to an equal volume of ChSO₄ solution in water (0.125%) to produce a 4:1 blend of collagen and GAG.

b. A turbid suspension of gel/precipitates was obtained at pH 2.7. The suspension was homogenised for 2 minutes using a Silverson homogeniser fitted with a mini-micro mixing head to reduce the size of the precipitates and improve their dispersion.

A second suspension was stirred with a magnetic stirrer for 1 hour for comparison with and as an alternative to homogenisation.

c. Aliquots of the suspension (0.5m) were frozen prior to freeze drying and solution impregnation as described in Example 1.

A scanning electron micrograph of a collagen/20% ChSO /PCL biocomposite (magnetic stirrer conditions) is shown in Figure 2.

3.2.2 20% ChSO₄ system, pH 4.8 and pH 7.1 a. Collagen solution (0.5% w/v) in 1% acetic acid (pH 2.7) was added to an equal volume of ChSO solution in water (0.125%) to produce a 4:1 blend of collagen and GAG.

b. The pH of the suspension of gel/precipitates obtained was adjusted to 4.8 and 7.1 respectively using 1M NaOH followed by 0.1 M NaOH. This had the effect of gradually dissolving the particles of gel/precipitate to yield a turbid solution containing few precipitates.

c. Aliquots of the suspension (0.5ml) were added to 4ml glass vials and frozen prior to freeze drying, solution impregnation and solvent evaporation as described in Example 1. 3.2.3 6. 11 , 20% CHSO₄ system. pH 2.7 a. Collagen solution (0.5% w/v) in 1% acetic acid (pH 2.7) was added to an equal volume of ChSO₄ solution in water (0.125%, 0.0625% and 0.0313) to produce 20%, 11% and 6% GAG content systems respectively.

b. The dispersion of gel/precipitates obtained was homogenised for 2 minutes using a Silverson homogeniser fitted with a mini-micro mixing head to reduce the size of the precipitates and improve their dispersion.

c. Aliquots of the suspension (0.5ml) were frozen prior to freeze drying and solution impregnation as described in Example 1.

3.2.4 50% ChSO₄ system, pH 3, 4.8 and 7.5

a. Collagen solution (0.5% w/v) in 1% acetic acid (pH 2.7) was added to an equal volume of ChSO₄ solution in water (0.5%) to produce a 1 :1 blend of collagen and GAG.

b. The suspension of gel/precipitates obtained was stirred with a magnetic stirrer for 1 hour to improve the dispersion.

c. The pH of separate suspensions of gel/precipitates was adjusted to 4.8 and 7.5 respectively using 1 M NaOH followed by 0.1 M NaOH.

d. Aliquots of suspension (0.5ml) were added to 4ml glass vials and frozen prior to freeze-drying, solution impregnation and solvent evaporation as described in Example 1.

3.2.5 50% ChSQ₄ system, production of a gel phase

b. The pH of the suspension of gel/precipitates obtained was adjusted to 7.3 using 1M NaOH followed by 0.1 M NaOH.

c. Aliquots of the suspension (0.5ml) were added to 4ml glass vials and allowed to stand at room temperature for 2 hours resulting in gel formation.

d. The gels were frozen prior to freeze-drying, solution impregnation and solvent evaporation as described in Example 1.

Example 4

Production of biocomposites of collagen, hyaluronic acid (HA) and PCL

4.1 6, 11 , 20% HA system-homogenised

a. Collagen solution (0.5% w/v) in 1% acetic acid (pH 2.7) was added to an equal volume of HA solution in water (0.125%, 0.0625% and 0.0313%) to produce 20%, 11% and 6% HA content systems respectively.

The precipitate obtained was reduced to smaller scale precipitates by adjusting the pH of the medium to 3.9 using 1M NaOH. The suspension was homogenised for 2 minutes using a Silverson homogeniser fitted with a mini-micro mixing head to improve dispersion.

Aliquots of the suspension (0.5ml) were frozen prior to freeze-drying and solution impregnation as described in Example 1.

4.2 6, 11 , 20% HA system-magnetic stirrer a. Collagen solution (0.5% w/v) in 1% acetic acid (pH 2.7) was added to an equal volume of HA solution in water (0.125%, 0.0625% and 0.0313) to produce 20%, 11% and 6% HA content systems respectively. b. The precipitate obtained was reduced to smaller scale precipitates by adjusting the pH of the medium to 3.7-3.9 (6% HA system), 4.1 (1 1 % HA system) and 4.9 (20% HA .system). The suspension was stirred with a magnetic stirrer for 30 minutes to improve the dispersion.

c. Aliquots of the suspension (0.5ml) were frozen prior to freeze-drying and solution impregnation as described in Example 1.

References Grzesiak, J.J. et al, Biomaterials, 18 (1997) 1625-1632.

J. Hansbrough, Keratinocyte Methods. Eds. I. Leigh, F. Watt. Cambridge University Press, (1994) 67-70.

Giusti, P. et al, (a), Bioartificial polymeric materials. TRIP, Vol 1 (1993) 261-267.

Santin, M. et al, Biomaterials, 17 (1996) 1459-1467.

Giusti, P. et al, (b) J. Mater. Sci. Mater, in Med. 4 (1993) 538-542.

Hanthamrongwit, M. et al, Biomaterials, 17 (1996), 775-780

Cappello, J., "Synthetically Designed Proein-Polymer Biomaterials in Controlled Drug Delivery. Challenges & Strategies", K. Park (Ed), Amer. Chem. Soc. (1997), 439-453.

Miggoney et al, Biomaterials 9 (1988) p.145, p.230 & p.413

Claims

1. A method for the preparation of a polymeric composite material, which method comprises the steps of a) forming a porous body of a first polymer; b) impregnating said porous body with a solution of a second polymer; and c) causing or allowing solvent to evaporate from said body.

2. A method as claimed in Claim 1 , wherein one of the first and second polymers is a natural polymer or a synthetic analogue thereof and the other is a synthetic polymer.

3. A method as claimed in Claim 2, wherein the first polymer is the natural polymer or synthetic analogue and the second polymer is synthetic.

4. A method as claimed in any preceding claim, wherein the porous body of first polymer is prepared by forming a solution of the first polymer and lyophilising that solution.

5. A method as claimed in any preceding claim, wherein the first polymer is selected from the group consisting of gelatin and extra cellular matrix proteins (collagen, elastin, laminin), cell adhesion proteins (such as fibronectin, vitronectin, vinculin, fibrinogen), polysaccharides (eg hyaluronic acid, heparin), glycosaminoglycans (such as chondroitin-4-sulphate), synthetic analogues of natural polymers such as silk-like and elastin-like protein polymers and heparin- like synthetic polymers.

6. A method as claimed in any preceding claim, wherein the first polymer is collagen.

7. A method as claimed in any preceding claim, wherein the second polymer is selected from the group consisting of poly(α-hydroxy acid) such as polylactide, poly(DL lactide co-glycolide), poly(ε-caprolactone), polyorthoesters, polyphosphazines, hyaluronic acid esters, polyanhydrides, copolymers of the above polymers and blends.

8. A method as claimed in Claim 7, wherein the second polymer is a polyester formed by ring-opening polymerisation of a compound of the generic formula

in which R represents an optionally substituted alkylene chain (CH₂)_n in which n is an integer of from about 4 to 10.

9. A method as claimed in Claim 7, wherein the second polymer is a polymer of a compound of the generic formula

in which R¹ and R², which may be the same or different, represent optionally substituted lower alkyl groups, ie alkyl groups of 1 to 6 carbon atoms.

10. A method as claimed in Claim 8, wherein the second polymer is poly(ε-caprolactone).

11. A composite polymeric material prepared by the method of any preceding claim.

12. A material according to Claim 11 , wherein the first polymer : second polymer weight ratio is 1 :40 or less.

13. A material as claimed in Claim 11 or Claim 12, which has a porous morphology.

14. A material as claimed in Claim 13, wherein the material has pores with sizes ranging from 50-1 OOμm.

15. A material as claimed in any one of Claims 11 to 14, which comprises collagen and poly(ε-caprolactone).