Patent Publication Number: US-2009220612-A1

Title: Hollow-fibre-based biocompatible drug delivery device with one or more layers

Description:
The present invention is directed towards the preparation of biocompatible drug delivery devices. Specifically it is directed towards the production of hollow fibres, in particular fibres of nanoporosity without additional coating and a generic synthesis route for the production of a range of hollow fibres with specific properties directed towards the delivery of drugs. Further, the invention is directed towards the use of such fibres to release compositions in a controlled fashion as the fibres degrade. The invention is also directed towards the preparation of microbeads as drug delivery devices and in particular to a synthesis route for the production of microbeads with specific properties. 
     Although chemotherapy can be given orally, most drugs used to treat cancer are given intravenously. The chemotherapy is either injected directly into a vein or through a thin tube called a venous catheter, a tube temporarily put into a large vein of the heart to make injections easier. However, there are many problems associated with this. Firstly, because therapeutics are commonly delivered intravenously, a relatively high dose needs to be administered. Consequently it can cause clotting and so adjustments of dose may be needed to prevent thrombosis. Secondly, many people also experience substantial side effects from the treatment because the drugs used affect normal cells as well as cancer cells. Such side effects include mucositis, immunosuppression, nausea/vomiting and hair loss. Thirdly, several courses of chemotherapy may be required to identify an active drug. There is therefore an enormous medical need for a means to enable sustained local delivery of chemotherapeutics to reduce side effects and improve quality of life 
     It is therefore an object of the present invention to prepare biocompatible drug delivery devices (which may be hollow fibres or microbeads) which can carry drugs for controlled release. The concept is based on the use of biodegradable, implantable devices which are capable of releasing one or a number of drugs over a period of time as the devices (which act as capsules) degrade. This device can either be connected to an inlet port on the surface of the body, thus allowing flushing of the devices and investigation into appropriate treatments/doses, or can be used as a stand alone “stent” implanted within a patient. The device also can be injected into a cavity of the patient as a microemulsion. 
     As well as treatment of cancers, this device could revolutionise the treatment of a number of diseases/disorders which typically require long-term attention. It is a further object of the present invention to enable the provision of effective, convenient, low cost treatment of chronic diseases. It will benefit patients by giving them peace of mind that the treatment is targeted at the problem area, will reduce side effects, and will be more effective and rapid than other treatments. Hospitals will benefit by the reduction in total cost of treatment and drain on resources. 
     According to a first aspect of the present invention, there is provided a biocompatible drug delivery device in which the mean pore size in one or more layers is less than 100 μm and in which a drug is carried. Such devices can be produced economically and reliably i.e. without defects, and are useful to a range of situations where devices with a small pore size, range of wall thickness and a controlled degradation rate are required to prevent dose dumping. 
     The device may be a hollow fibre of one or more layers. The mean pore size may be controlled to be in the claimed range in the outer wall and this may optionally be less than 50 μm, 10 μm, 1 μm, 100 nm or even less than 10 nm. The porosity may be even throughout the cross section of the wall of the fibre or may vary across the fibre. In some cases there will be different porosities across the cross section of the fibre. In particular, there may be a higher porosity towards the centre of the fibre, and a lower porosity in the outer layers to provide the strength of fibre aligned with the drug delivery properties desired. 
     According to a further aspect of the present invention, the device may be a microsphere. The porosity may vary across the radius of the sphere or may be substantially even. There may be a higher porosity towards the centre of the sphere and a lower porosity towards the outer surface. This variable porosity and the thickness of the respective layers play a significant role in the speed of release of the drug from the device. 
     The fibre or microsphere may be biodegradable such that it will decompose substantially completely over a period of 30 days, 60 days, 90 days, 120 days or anything from 1-24 months. The fibre or sphere may consist of two or more layers, in which different drug compositions may be carried within each layer or within the lumen of the fibre or macrovoids of the sphere, such that the medication applied varies over time as the fibre or sphere decomposes. The different compositions may be different chemical formulations or may alternatively be different concentrations of the same drug. 
     The fibre or microsphere may be an organic or polymeric fibre or sphere comprising a polymer, an additive and one or more drug compositions. The fibre may include additional drugs in different layers or contained within the lumen (hollow) or macrovoids to be released in a controlled fashion over a period of time. Details of preferred drug compositions and concentrations may be found in the following examples and in the claims. 
     An advantage to the production of double or triple layer hollow fibres (in addition to an increased mechanical strength which may be beneficial in avoiding breaking or dose dumping the drug delivery device when implanting it in to the patient) is that the fibres are largely defect free. With two or three layers of the same polymeric composition forming the scaffold for the drugs, any defects in one layer are extremely unlikely to be mirrored by a similar defect in the next layer. The net effect is that there are no pinholes which pass through the fibre and it can therefore be used as a scaffold for a drug delivery device without the risk of dose dumping through a defect. 
     Further, as discussed above, it is possible to have different compositions or different concentrations in the two or more layers as well as a drug contained within the lumen. It is therefore possible to produce a fibre where each layer is tailored towards a particular property or treatment which may vary over time. 
     A further aspect of the present invention is the high surface to volume ratio of the biocompatible porous hollow fibre or microsphere which may be greater than 1,000 m 2 /m 3  for each delivery device. The area to volume ratio may be in the range 1,000-30,000 m 2 /m 3 , preferably 1,000-6,000 m 2 /m 3 , and most preferably 4,000-6,000 m 2 /m 3 . 
     The hollow fibres may be arranged in an appropriate configuration for the specific medical application such that they have the desired drug release properties. In particular, the fibres may be arranged such that when they degrade and start to release the drug or drugs, these are targeted at the intended area (e.g. tumour or damaged tissue). In a first arrangement, one or more, optionally 5, 10, 20 or more fibres are bundled together. These may be held tightly in a cylindrical configuration or constrained to another shape which maximises the effectiveness of the delivery device for the release of drugs. The fibres may be held together by a casing which may or may not be biodegradable. 
     As discussed above, the hollow fibres may contain the same or different drugs in the same or different concentrations within the layer or layers of the fibre and in the hollow or lumen of each fibre. Further, the drugs in some or all of the layers and lumen may be microbead encased to enable a double time-release mechanism to be created. Such microbeads may have an average dimension in the range of 10-50 μm, and are preferably of a uniform diameter. For example, in one embodiment at least 90% of the microbeads have a diameter of n±1 μm where n=10-50 μm. These microbeads in the hollow fibre may be beads produced by known techniques, for example by emulsion techniques, or may be beads or microspheres produced according to the present invention. 
     Each hollow fibre may, in addition, have an extra outer layer provided to avoid premature drug release. The individual fibres may also be selected to allow the configuration to act as a stand-alone implant or stent. The fibres are sealed at each end and the drugs are released by a combination of diffusion of the drug through the walls of the fibre and degradation of the wall. In stents, some layers of the fibre are developed to include a stronger polymer (for example, polylactide (PLA)) to maintain the strength, integrity and flexibility of the coil. This may have application in the treatments of colon or oesophagus type cancers. 
     Alternatively, one or both ends of each fibre may be connected to the surface of the body of the patient to allow the fibres to be flushed and refilled with drugs. This connection may be direct, or indirect via an outlet port, lumen or hickman line. The device may be flushed with any suitable material, for example saline or heparin, and the replacement drugs may be the same or different to those previously present. This may have particular application when trying to establish a treatment regime in the early stages of chemotherapy, and different drugs or concentrations of drugs need to be tried to determine what is most effective. 
     The fibres may be woven together in a patch-like configuration which may then be placed or attached to the site of interest. The patch may be attached, for example, with surgical staples. The patch may be configured to be in an appropriate shape for the site of interest, for example, a ring or sphere to encase part or all of the site of interest. Alternatively, the fibres may be prepared such that they are helical and can be used as a stent within narrow orifices such as the throat (oesophagus), arteries, colon, bowel, ovaries etc. The helical arrangement allows the orifice to be kept open while allowing the drug to be delivered directly to the site of interest. 
     The biocompatible porous hollow fibre or microsphere may include a high percentage of drugs to be delivered. According to one embodiment, there is at least 5% by weight of drugs in the polymer solution used to prepare the delivery device, although values of up to 80% or higher may also be present. The polymer solution used to prepare the delivery device may therefore have 5-80% by weight of drugs, or 10-75%, 20-70%, 30-65%, 40-80% or 50-55% as appropriate for the situation. This will depend on the nature of the drugs being released and the desired speed of delivery. 
     The level of drug entrapped within a device (hollow fibre or microsphere) or loaded in the drug delivery device will depend on the encapsulation efficiency which may be from 10-80%, typically 10-40% or 15-30%. This may produce fibres or microspheres having a drug loading of 1-50% by weight or 1-45%, 2-25%, 3-20% or 4-15% by weight. 
     The outer diameter of the biocompatible fibres produced can be 10-200 μm depending on the diameter of the spinneret and the number of layers used. Therefore, lightweight and compact drug delivery devices can be made using a single hollow fibre or a cluster of narrower fibres as appropriate. The hollow fibres are nanoporous or microporous and can be tailored to exhibit significant drug pharmacokinetics, bending strength (flexibility) and liquid bursting pressure (1-7 bar). The properties of the fibre can be tailored to individual situations. 
     As indicated above, the microspheres according to the present invention may have a diameter of 10-70 μm, preferably 20-40 μm or 25-35 μm or 35-50 μm or 50-70 μm depending on the particular circumstances. The microspheres of the present invention are generally produced to be of a uniform size and a very narrow size distribution. 
     According to another aspect of the present invention, there is provided a method for preparing biocompatible porous hollow fibres, in which a spinning dope is prepared in a viscous liquid or gel form, filtered using a mesh, the dope is degassed in a piston delivery vessel attached to a spinneret, the vessel is pressurised using an inert gas using jets, the dope is extruded through the spinneret to form a fibre precursor, the precursor is thoroughly washed to remove any residual solvent and dried. All the equipment is sterilised before use, preferably by steam sterilisation. 
     The hollow fibres formed by the method of the present invention are the result of the controlled solidification process. First a spinning mixture or dope is prepared from a polymer, a solvent, optionally a binder/additive and any drug composition(s) to be included within the layers of fibre. Subsequently, the produced mixture is extruded through a spinneret into a bath of non-solvent. This non-solvent, selected from a number of internal coagulants including distilled water, is also introduced through the bore of the spinneret. Exchange of solvent and non-solvent leads to thermodynamic instability of the spinning mixture and induces liquid-liquid demixing. Further exchange leads to solidification of the polymer-rich phase. The precursor is washed and dried to remove any residual solvent. Subsequently, a further drug composition may be added to the lumen (hollow) of the produced fibre through a syringe pump. 
     The produced compact fibres show very good quality and may have different porosities across the cross section of the fibre with a preferred total porosity in the range 30-55%, in particular 35-45%. Average pore size and effective surface porosity of the hollow fibres can be determined by the Poiseuille flow method. In one embodiment, the fibres produced have a pore size in the range 10 m≦pore diameter≦0.1 mm. In another embodiment, the fibres produced have a pore size in the range 100 nm&lt;pore diameter&lt;1 μm. In a still further embodiment, the fibres produced have a pore size in the range pore diameter≦1 nm. 
     According to another aspect of the present invention, there is provided a method for preparing biocompatible microspheres in which: a polymer is dissolved in a suitable solvent; a drug is dissolved in or dispersed in an organic solvent containing the polymer; the mixture is fed to a high pressure airbrush or spray device with a small nozzle; and the mixture is sprayed under water and microspheres are formed by solvent de-mixing. Similarly, the mixture may be fed to a ceramic hollow fibre bundle with 2-20 nm pores to create micro droplets under water. The feed pressures could be from 2-10 bar. 
     According to a further aspect of the present invention, there is provided a method for preparing biocompatible microspheres in which: a polymer is dissolved in a suitable solvent; a drug is dissolved in or dispersed in an organic solvent containing the polymer; the mixture is fed to a high pressure airbrush with a small nozzle; and the mixture is sprayed into an antistatic chamber with water saturated air and the solvent is extracted from the produced droplets to form microspheres. 
     In each case, the microspheres are removed and dried to remove substantially all of the solvent such that the microsphere may contain 0-2% by weight of solvents. 
     The size and density of the microspheres is controlled and affected by the method used, the spraying pressure and the air gap (gap where beads are exposed to air before polymer solidification or phase inversion) or pore size of the fibres used. In each case, the spheres produced have macrovoids to accommodate more drugs to be delivered. 
     Although the hollow fibres or microspheres do not have to be biodegradable to operate as drug delivery devices, to be most effective they would degrade in the body over time in order to enable more practical control of the dose of drugs over time (and to prevent any implant from causing inflammation or requiring further surgery for removal). 
     The fibres or microspheres may be made primarily from any biodegradable material, i.e., where over time the material will decompose mainly to either CO 2  or water or harmless acid. For example: Polyglycolide (PGA) where the degradation product is glycolic acid (a natural metabolite which may be eliminated from the body through the Krebs cycle) or Poly(L-lactic) acid (PLA) where the degradation product is lactic acid (which is eventually converted to carbon dioxide and water by the Krebs cycle and released through respiration). 
     Examples of suitable materials for the production of the biocompatible fibres or microspheres include: Poly(ethylene terephathalate) (which resists fungal and enzymatic degradation) with the addition of 20-25% PLA (which introduces the biodegradation properties); PLA; PGA; a copolymer of PLA and PGA to form PLGA (poly lactic co glycolic acid); Lactide-glycolide copolymers (PLG); Poly-ε-caprolactone (PCL); Lactide-caprolactone copolymers; and Cellulose-based polymers. 
     Further biocompatible materials that may be suitable include: Acrylate polymers and copolymers: (for example methyl methacrylate, methacrylic acid; hydroxyalkyl acrylates and methacrylates; methylene glycol dimethacrylate; acrylamide, bisacrylamide); Ethylene glycol polymers and copolymers; Oxyethylene and oxypropylene polymers; Poly(vinyl alcohol) and Polyvinylacetate; Polyvinylpyrrolidone and polyvinylpyridine. 
     Other possible materials include Biodegradable polyphosphazenes, Pseudo-poly(amino acids), polyethylene glycol containing poly-carbonates, phosphorous containing biodegradable polymers, polyphosphazenes and poly(phosphate esters) or PCL and/or other polymers/copolymer combinations, polyhydroxyalkanonate (PHA) class of polymers, poly(imino carbonates) or Poly(cyanoacrylates). 
     PCL is a synthetic a-polyester exhibiting a low Tg of around 60° C. which imparts a rubbery characteristic to the material. PCL like other members of this family of polymers such as PLA and PLG, undergoes auto-catalysed bulk hydrolysis). The rubbery characteristics of PCL results in high permeability which may be exploited for delivery of low molecular weight drugs such as steroids and vaccines. PCL is a degradable biopolymer that typically takes more than 1 year to degrade in vivo. However, the semi-crystalline nature of the PCL polymer extends its resorption time to over 2 years. 
     It is thought to degrade in vivo by hydrolysis to caproic acid and its oligimers and by enzymatic action. In vivo studies have shown that PGA &amp; PLGA fibre scaffolds collapsed after 3 or 4 weeks; hence, in order to improve the integrity of the construct, the addition of PCL to PLGA could stabilise the fibres for long term drug release. PCL may be copolymerized and blended with PLA and PLGA in order to accelerate the degradation. 
     The delivery device of the present invention can deliver small molecules, or recombinant proteins (for example, chemotherapy drugs) or anaesthetics (for example lidocaine). In particular, the drugs which may be delivered by means of the delivery device of the present invention include, but are not limited to, fluorouracil (5-FU), cisplatin, oxaliplatin, carboplatin and warfarin. The present invention may therefore be applied to a range of diseases, in particular, chronic diseases including (but not limited to) cancers, inflammatory or autoimmune disease, transplant rejection, spinal damage, bone disease, Parkinsons, Alzheimers, etc and to long term anaesthetic agents. 
     The desired drug may be added to the solution of the matrix material (polymer) by either co-dissolution in a common solvent, dispersion of finely pulverised solid material or emulsification of an aqueous solution of the drug immiscible with the matrix material solution. Dispersion of the solid or dissolved bioactive material in the matrix-containing solution may be achieved by ultrasonication, impeller or static mixing, etc. 
     The present invention also extends to the use of a drug delivery device (for example a hollow fibre or a microsphere) to deliver drugs to treat chronic diseases in mammals. 
    
    
     
       The invention may be put into practice in a number of ways and a number of embodiments are shown here by way of example with reference to the following figures, in which: 
         FIG. 1  shows in schematic form the apparatus for the generic spinning procedure for producing a hollow fibre according to the present invention; 
         FIG. 2  shows in schematic form the apparatus for the generic spinning procedure for a double layer fibre according to another aspect of the present invention; 
         FIGS. 3 to 5  are line drawings of photographs of embodiments of double, triple and quadruple orifice spinnerets according to an aspect of the present invention; 
         FIGS. 6 to 9  show the component parts for a triple orifice spinneret for use in the production of a double layer fibre; 
         FIGS. 10 to 16  show a quadruple orifice spinneret for use in the production of a triple layer fibre; 
         FIG. 17  shows a quadruple orifice spinneret including an additional access point for the introduction of a thin layer of adsorbent as an outer coating; 
         FIG. 18  shows an alternative design for the quadruple orifice spinneret in which each of the chambers has independent feeds; 
         FIG. 19  shows schematically three different designs for producing triple layer fibres using a quadruple orifice spinneret; 
         FIG. 20  shows a pressure vessel suitable as a delivery vessel for any spinneret of the present invention; 
         FIG. 21  shows in schematic form, apparatus for the production of microspheres according to the present invention; 
         FIG. 22  shows a photograph of a high pressure airbrush which may be used in the production of microspheres according to the present invention. 
         FIG. 23  shows the percentage release of Lidocaine from PLA for samples 1 and 2; 
         FIG. 24  shows the percentage release of Lidocaine from sample 3; 
         FIGS. 25   a  and  b  are SEMs of 25:75 PLGA with Lidocaine in the wall; 
         FIG. 26  shows the percentage release of Lidocaine and Lidocaine HCL from sample 3; 
         FIG. 27  shows the percentage release of Lidocaine from sample 4; 
         FIGS. 28   a - d  show SEM micrographs of different samples of hollow fibres before 5-fluorocil injection and before degradation; 
         FIG. 29   a - c  shows SEM micrograph of samples after 6 weeks of degradation; 
         FIG. 30  shows the release of 5-FU from different compositions hollow fibres; 
         FIG. 31  shows the release of Cisplatin from a range of compositions of hollow fibres; 
         FIG. 32  shows the release kinetics for cisplatin from a PLGA hollow fibre; and 
         FIG. 33  shows the cumulative percentage mass of carboplatin released from PLGA fabricated fibres over a three week period. 
     
    
    
     An aspect of the present invention is directed towards a method of production of hollow fibres. This method may generically be described as follows. 
     Generic Method for Production of Hollow Fibres 
     One or more spinning dopes are prepared depending on whether the fibre is to be a single, double, triple, etc layer fibre. For each spinning dope, a suitable solvent is poured into a 500 ml wide-neck bottle, and the desired quantity of polymer is slowly added. The mixture is stirred on a roller to form a polymer solution and once the polymer solution becomes clear, the desired amount of the finely divided powdered drugs to be delivered are slowly added or the drug is dissolved in a suitable solvent first and the drug concentrate is added to the mixture. The mixture is then stirred with an IKA® WERKE stirrer at a speed of 500-1000 rpm for 1-2 days until the drugs are dispersed uniformly in the polymer solution and the mixture is sonicated for 0.5 h in order to obtain an homogeneous mixture, and from the vigorous stirring the mixture is turned into a viscous solution or a gel. The mixture is slightly heated, then filtered through a 100 μm Nylon filter-bag in order to remove any agglomerated or large particles and the mixture is then placed on a rotary pump for 1-2 days to degas and to form a uniform spinning dope. 
     The fibres are then produced by spinning using an appropriate spinneret which may be followed by slight heat treatment (the temperature is generally kept below 30° C.). Referring to  FIG. 1 , the mixture  5  is transferred to a stainless steel piston delivery vessel and degassed using a vacuum pump (optionally at a slightly raised temperature) for two hours at room temperature—this ensures that gas bubbles are removed from the viscous polymer dope. The spinning process is then carried out with the following parameters:
     1. The heated tank  10  (heating wire around the tank) is pressurised to 2-4 bar using a nitrogen jet  12  and this is monitored by means of a pressure gauge  14 . Release of the dope mixture  5  to the spinneret  20  is controlled by means of a piston  16  and valve  18 . The delivery vessel is long and small in diameter to maintain uniform pressure for longer periods with in the vessel. The higher the pressure in the tank, and therefore the pressure of the precursor dope passing through the spinneret, the smaller the fibre produced   2. A tube-in-orifice spinneret  20  is used with an orifice diameter of, for example, 2 mm and an inner tube diameter of 0.72 mm, in order to obtain hollow fibre precursors. This double orifice spinneret is for a single layer fibre. For two or more layers, triple or quadruple spinnerets are used and feeds are arranged appropriately. Bore liquid (or the internal coagulant)  25  is also fed to the spinneret  20  and is controlled by means of a gear pump  22 . If less bore liquid is pumped through the spinneret the hollow core of the fibre will be smaller and the walls will be thicker. By changing the delivery pressure for each feed the properties of the fibres may be changed   3. The air gap  24  between the bottom of the spinneret  20  and the top surface of the coagulation bath  26  is typically varied in the range 0-3 cm. Increasing the air gap will cause the outer “skin” of the fibre produced to be more dense whereas a smaller air gap will produce a product with more open layers and the fibre will be more porous.   4. The fibre  30  (once extruded from the spinneret) is passed over a series of rollers  28  through a washing bath  27  to a fibre storage tank  29 .   5. Water is used as the internal and external coagulator as both bore liquid  25  and as bath liquid in water baths  26  and  27 . A low concentration of other solvents also could be added to improve precipitation rate e.g. ethanol, ethyl acetate, acetone.   

     The precursor is run through the water bath  26  to complete the solidification process and then the hollow fibre  30  is washed thoroughly in the second water bath  27 . Care must be taken to ensure that the hollow fibre is not subject to mechanical dragging during the spinning process. Continuity in the pressure is important to deliver polymer dope gel as well as uniform delivery of the internal coagulant in order to avoid entrapment of air and separation of the fibre which would otherwise result in unsuccessful spinning. A guide motor  31  helps to control the movement of the fibre through the water baths. The hollow fibre precursors are then left to soak for 3-4 days in fresh water in the fibre storage tank  29  in order to remove any residual solvent. The precursors are then dried in ambient conditions for seven days. 
     Apparatus as set out in  FIG. 2  may be used to form double layer fibres. The apparatus includes a triple orifice spinneret and two solution feeds. Typical dimensions of the triple orifice spinneret are external layer (d out  4.0 mm, d in  3.0 mm), internal layer (d out  2.0 mm, d in  1.2 mm), and bore (d out  0.8 mm). Triple and quadruple orifice spinnerets for use in the production of double or triple layer fibres are described in further detail below. For a triple layer fibre, apparatus similar to that shown in  FIG. 2  is used but there will be an additional third solution feed for the third layer. 
     For spinning, two delivery vessels  10   a ,  10   b  (or more as may be required) are prepared, one may be pressurised to 2 bar using nitrogen  12 , and the other delivery vessel may be further pressurised to 2.5-4 bar using a nitrogen jet. In order to maintain uniform pressures two piston delivery vessel pressure controllers were used. These provide gel feeds  5   a  and  5   b  to the triple orifice spinneret with the feed  5   a  providing the inner layer of the fibre and feed  5   b  providing the outer layer. 
     The fibres produced by this method may have two or more layers. This method has the advantage of reduced production costs when compared to prior art methods and also enables the introduction of layers with different functional properties and mixed matrix compositions. 
     Triple Orifice and Quadruple Orifice Spinneret 
       FIG. 3  is a photograph of embodiments of a triple orifice spinneret (left) and a quadruple orifice spinneret. These will be described in further detail below.  FIG. 4  is a photograph of the components of one embodiment of a triple orifice spinneret and  FIG. 5  is a photograph of one embodiment of a quadruple orifice spinneret. 
     Triple Orifice Spinneret 
       FIGS. 6 to 9  show the components for one embodiment of a triple orifice (double layer) spinneret.  FIG. 6  shows a base module  110  to which the precursor feeds are fed and to which the delivery chambers are attached. The precursor feeds may be the same or different and may therefore be fed from the same reservoir (not shown). Alternatively, they may be of different composition and accordingly supplied from different reservoirs under controlled pressure conditions. Feed  112  is for the bore liquid which passes through the centre of the fibre to form the hollow core. Precursor feeds  114 ,  116  are for the two layers of the fibre. At the outlet  118  of the base module  110  is a screw thread (not shown) to which the delivery chambers are secured. 
       FIG. 7  shows the outer delivery chamber  120  which controls the precursor feed for the outer layer of the fibre. At the inlet end  122  of the chamber there is provided an external thread  123  to secure the chamber to the base module  110 , and an internal thread  124  to which the second chamber  130  is secured. At the outlet, there is a circular orifice  126  at the end of a neck region  127 . This orifice  126  will, when the spinneret is assembled, have further outlets passing though it leaving an annular passage through which the material for the outer layer will pass. The outer diameter of this orifice may, for example, be 4 mm. The angle θ of the slope directing the material to the orifice is preferably 60° but may be from 45-65°. Ideally the angles throughout the spinneret should remain constant for all chambers to maintain uniform delivery of the precursor material. 
       FIGS. 8 and 8   a  show the second delivery chamber  130  which together with the outer delivery chamber controls the precursor feed for the outer layer of the fibre. At the inlet end  132  of the chamber there is provided a securing ring  134  which has an external thread dimensioned to cooperate with the internal thread  224  of the first chamber  120 . The ring  134  has channels  135  cut in the ring at regularly spaced intervals. In a preferred embodiment there are eight channels spaced evenly around the circumference of the ring. These channels permit the flow of the precursor feed for the outer layer to pass from the reservoir, through the spinneret to the outlet  136  of the second delivery chamber. 
     The outlet takes the form of a circular orifice  136  and the orifice extends in a neck  137  dimensioned to fit inside the neck  127  of the first delivery chamber thereby forming the channel for the intermediate material. This orifice will, when the spinneret is assembled, have further outlets passing though it thereby leaving an annular passage through which the material for the inner layer will pass. The outer diameter of this orifice  137  may, for example, be 3.9 mm and the internal diameter may be 3.5 mm. The angle θ of the external slope of delivery chamber  130  must be the same as θ in the first delivery chamber to maintain the width of the passage through which the outer layer flows. This will also minimise pressure losses in the spinneret. The angle φ of the internal slope which will direct the inner layer of material to the outlet is preferably the same as θ, namely preferably 60°, but may be from 45-65°. As mentioned above, the angles preferably remain constant throughout the spinneret to ensure uniform flow. 
       FIGS. 9 ,  9   a  and  9   b  show the third delivery chamber  140  which controls the precursor feed for the inner layer of the fibre. At the inlet end  142  of the chamber there is provided a ring  144  which rests against the ring  134  of the second delivery chamber  130 . The ring  144  has channels  145  cut in the ring at regularly spaced intervals. In a preferred embodiment there are four channels spaced evenly around the circumference of the ring. These channels permit the flow of the precursor feed for the inner layer to pass from the source, through the spinneret to the outlet  146  of the third delivery chamber. Corresponding channels  145   a  are also found on the cap at the front end of the chamber which includes the outlet  146 . 
     Again, the outlet takes the form of a circular orifice  146  and the orifice extends in a neck  147  dimensioned to fit inside the neck  137  of the second delivery chamber thereby forming the channel for the intermediate material. The outer diameter of this orifice  147  may, for example, be 2.5 mm and the internal diameter (i.e. the diameter of the hollow core of the produced fibre) may be 2.1 mm. The angle φ of the external slope of delivery chamber  140  must be the same as φ in the second delivery chamber to maintain the width of the passage through which the inner layer flows. This will also minimise pressure losses in the spinneret. The angle α of the internal slope which will direct the bore fluid to the outlet of the spinneret is preferably the same as θ and φ, namely preferably 60°, but may be from 45-65°. Constant angles throughout the spinneret enable uniform delivery of precursor. 
     The precursor for the inner layer of the fibre passes on the outside of the third delivery chamber, bounded on the other side by the second delivery chamber. The bore liquid passes through the centre of the third delivery chamber to the needle outlet  146 . 
     Quadruple Orifice Spinneret 
       FIGS. 10 to 16  show the components for one embodiment of a quadruple orifice, triple layer fibre spinneret.  FIG. 10  shows the spinneret  200  assembled. It comprises six members each of which is shown in greater detail in the following figures. Typical dimensions of the spinneret are 140 mm length by 70 mm diameter. 
       FIG. 11  shows the base module  210  to which the precursor feeds are fed and to which the delivery chambers are attached. The precursor is fed through three feed inlets spaced around the perimeter of the module  210 . Two of these inlets are shown as  214 ,  216 . The third (not shown) may be arranged such that it extends out in an orthogonal direction from feeds  214 ,  216 . The feeds may all be the same composition thereby producing a fibre of one composition, but greater strength and with fewer defects, and in this case the inlets are fed from the same reservoir (not shown). Alternatively, the feeds may be of two or three different compositions and accordingly supplied from different reservoirs (not shown) under controlled pressure conditions. Feed  212  is for the bore liquid feed which passes through the precursor material and forms the hollow core in the finished product. At the outlet  218  of the base module  210  is a screw thread (not shown) to which the delivery chambers are secured. 
       FIG. 12  shows the outer delivery chamber  220  which controls the precursor feed for the outer layer of the fibre. At the inlet end  222  of the chamber there is provided an external thread  223  to secure the chamber to the base module  210 , and an internal ridge  224  to support the second chamber  230 . At the outlet, there is a circular orifice  226  at the end of a neck region  227 . This orifice  226  will, when the spinneret is assembled, have further outlets passing though it leaving an annular passage through which the material for the outer layer will pass. The outer diameter of this orifice may, for example, be 4 mm. The angle θ of the slope directing the material to the orifice is preferably 60° but may be from 45-65°. Ideally the angles throughout the spinneret should remain constant for all chambers to maintain uniform delivery of the precursor material. 
       FIGS. 13 and 13   a  show the second delivery chamber  230  which controls the precursor feed for the intermediate layer of the fibre. At the inlet end  232  of the chamber there is provided a securing ring  234  which has an external thread dimensioned to cooperate with the internal thread  224  of the first chamber  220 . The ring  234  has channels  235  cut in the ring at regularly spaced intervals. In a preferred embodiment there are 8 channels spaced around the circumference of the ring. These channels permit the flow of the precursor feed for the intermediate layer to pass from the source, through the spinneret to the outlet  236  of the second delivery chamber. 
     Again, the outlet takes the form of a circular orifice  236  and the orifice extends in a neck  237  dimensioned to fit inside the neck  227  of the first delivery chamber thereby forming the channel for the intermediate material. Again, this orifice will, when the spinneret is assembled, have further outlets passing through it thereby leaving an annular passage through which the material for the inner layer will pass. The outer diameter of this orifice  237  may, for example, be 3.9 mm and the internal diameter may be 3.5 mm. The angle θ of the external slope of delivery chamber  230  must be the same as θ in the first delivery chamber to maintain the width of the passage through which the intermediate layer flows. This will also minimise pressure losses in the spinneret. The angle φ of the internal slope which will direct the inner layer of material to the outlet  205  is preferably the same as θ, namely preferably 60°, but may be from 45-65°. As mentioned above, the angles preferably remain constant throughout the spinneret to ensure uniform flow. 
       FIGS. 14 ,  14   a  and  14   b  show the third delivery chamber  240  which controls the precursor feed for the inner layer of the fibre. At the inlet end  242  of the chamber there is provided a ring  244  which rests against the ring  234  of the second delivery chamber  230 . The ring  244  has channels  245  cut in the ring at regularly spaced intervals. In a preferred embodiment there are four channels spaced around the circumference of the ring. These channels permit the flow of the precursor feed for the intermediate layer to pass from the source, through the spinneret to the outlet  246  of the third delivery chamber. Corresponding channels  245   a  are also found on the cap at the front end of the chamber which includes the outlet  246 . 
     Again, the outlet takes the form of a circular orifice  246  and the orifice extends in a neck  247  dimensioned to fit inside the neck  237  of the second delivery chamber thereby forming the channel for the intermediate material. Again, this orifice will, when the spinneret is assembled, have further outlets passing though it thereby leaving an annular passage through which the material which will form hollow core of the fibre will pass. The outer diameter of this orifice  247  may, for example, be 2.5 mm and the internal diameter may be 2.1 mm. The angle φ of the external slope of delivery chamber  240  must be the same as φ in the second delivery chamber to maintain the width of the passage through which the inner layer flows. This will also minimise pressure losses in the spinneret. The angle α of the internal slope which will direct the bore fluid to the outlet  205  is preferably the same as θ and φ, namely preferably 60°, but may be from 45-65°. Constant angles throughout the spinneret enable uniform delivery of precursor. 
       FIG. 15  shows a further chamber  250  through which the bore fluid flows. Attached to this chamber at the front end is a bore needle  260  as shown in enlarged form in  FIG. 16 . The bore needle  260  will define the dimension of the inner hollow core of the fibre and may therefore be varied from embodiment to embodiment as appropriate. The inlet  252  of the chamber  250  is arranged to cooperate with the bore liquid inlet feed  212  of base module  210 . The shoulders  253  abut the inner surface  213  of the base module  210 . The shoulders  254  abut the ring  244  at the inlet end of third chamber  240 . The shoulder portion  254  has matching channels which line up with the channels  245  in ring  244 . There is also a small gap below the shoulder  254  to allow further passage of the precursor fluid. The core  255  of the chamber  250  has an external diameter d 1  and an internal diameter d 2 . Preferred values for d 1  and d 2  may be 8 mm and 4 mm respectively but any values in the range 1-20 mm may be appropriate for a specific embodiment. 
     At the front end of chamber  250  there is an outlet  256 . Towards this end the core may increase in internal diameter to accommodate the bore needle  260  (see  FIG. 16 ). For example the internal diameter may increase from 4 mm to 5.2 mm. The bore needle  260  is arranged to fit inside the outlet  256  of chamber  250  as a snug push fitting. The dimension d 3  of the bore needle may, for example, be 5 mm to fit inside the outlet end  256  of the chamber  250  having an internal diameter of 5.2 mm. The diameter of the needle d out  may be in the range 0.1-5 mm, more preferably 0.5-3 mm, for example 1 mm. This defines the size of the hollow core of the fibre. The angle α should be the same as in the third delivery chamber  240  to maintain the width of the passage through which the precursor fluid flows. As indicated above, α is preferably 60°, but may be in the range from 45-65°. 
       FIG. 17  shows a similar view to  FIG. 10 , but the spinneret has an additional access point  270  for the introduction of a thin layer of adsorbent or other functional material. This will form an outer coating in addition to the three layers of the fibre. This thin layer may be present to help the selectivity of the fibre for a particular adsorbate. 
       FIG. 18  shows an alternative design for the quadruple orifice spinneret in which each of the chambers has independent feeds. The arrangement of the chambers is similar to that described with respect to  FIGS. 10 to 16  above but each chamber has clear and separate precursor feeds which do not all pass through the base module. 
       FIG. 19  shows schematically three different designs for producing triple layer fibres using the quadruple orifice spinneret.  FIG. 19   a  shows a spinneret  280  for the delivery of gel precursor of one composition from a single piston pressure vessel  282 . The fibre produced is stronger and defect free. The bore liquid passes from reservoir  281  through the centre of the spinneret  280  to form the hollow core of the fibre. Each of the channels leads to the outlet  290  which may take the form shown in  FIG. 19   d .  FIG. 19   b  shows a spinneret  280  for the delivery of three different compositions from three different delivery vessels  283 ,  284 ,  285 , the pressure of which is controlled independently. Each composition may have different types of adsorbent with different functional properties attached to them. Using this system it is possible to produce compact fibres with very small particles to achieve small pores of the order of 1-2 nm. 
       FIG. 19   c  shows a design for the delivery of two different compositions. The two inner layers are of the same composition fed from delivery vessel  286  and the outer layer is of a different composition from delivery vessel  287 . This fibre has the advantage of a stronger fibre with fewer defects of a first composition, with an outer layer which is specifically chosen to have the functional properties required, for example in the choice of adsorbent. 
       FIG. 19   d  shows a typical arrangement of the outlet  290  of the spinneret  280 . The three concentric rings of precursor each have a thickness of 0.5 mm. For example, the inner core formed by the bore liquid may have a diameter a of 1.1 mm. The outer diameter of the first layer of precursor then has a diameter b of 2.1 mm. The intermediate layer has an inner diameter c of 2.5 mm and an outer diameter d of 3.5 mm. The outer layer has an inner diameter of 3.9 mm and an outer diameter of 4.9 mm. 
       FIG. 20  shows a pressure vessel  300  suitable as a delivery vessel to the spinneret. The precursor gel  301  is maintained under pressure by means of jets of nitrogen  302  being applied to a plate type piston  303 . The pressure is measured by means of a pressure gauge  304 . The precursor is fed out of the vessel  300  through the outlet means  305  to the appropriate feed of the spinneret. As shown in  FIG. 20 , the outlet is conically shaped and may be, for example, 5-15 mm in diameter, preferably 10 mm. The vessel  300  is made of stainless steel and may have dimensions of 150-200 mm height by 60-80 mm diameter. The delivery vessel can also be heated by a heating tape. 
     Referring to  FIG. 21 , an apparatus suitable for producing microspheres according to the present invention is shown schematically. A biocompatible polymer is dissolved in a suitable solvent  510 , and a finely divided drug  512  is dissolved or dispersed in the solvent containing the polymer material and the materials are mixed by a mixer  514  in a beaker  516 . The mixture is placed in an airbrush jar and is then sprayed via line  502  into an antistatic chamber  500 . Simultaneously, water saturated air  504  is introduced into the chamber to achieve solidification of the beads. The two sprays are introduced into the antistatic chamber at an angle to achieve a cyclonic action to ensure good contact between the sprays. 
     The mixture has a substantial residence time in the chamber (for example 2-10 minutes) to allow the droplets formed to undergo phase inversion. The two sprays are fed into the chamber in the same direction so that they are co-current. The beads are formed into microspheres by extraction of the solvent. This may be by any suitable means, for example heating or vacuum treatment. The collected microspheres are then dried  520  to reduce the level of solvent still further to 0-2% by weight. 
       FIG. 22  shows an example of an airbrush jar which may be used in this method or the bath method or producing beads (not shown). The airbrush jar has a very small nozzle to enable the production of small droplets (10-50 μm). The mixture is generally sprayed at a high pressure (3-3.5 bar) using compressed air. The size of the beads is controlled by both the pressure of the spray and the viscosity of the polymer dope. 
     In the bath method, the mixture of solvent, polymer and drug is sprayed directly into a water bath and not into an antistatic chamber. Water-solvent demixing occurs and drug encapsulated beads are produced. The microspheres produced by both methods were found to be of a uniform size, the volume of which is determined by the spraying pressure. Generally the spheres produced by the bath method have a smaller pore size and have a more dense outer skin. 
     EXAMPLES 
     In the following examples, the following abbreviations may be used for the different chemicals used.
         DCM—Dichloro-methane   EA—Ethyl acetate   PBS—Phosphate buffer saline (adjusted to pH 7.4 with 0.1M NaOH)   PVA—Poly-vinyl-alcohol   PLA—Poly(L-lactic) acid   PLGA—Poly(lactic-co-glycolic) acid (75:25 and 65:35)       

     Experimental Method 
     Spinning of Hollow Fibre 
     Materials 
     PLA and PLGA (16-30 wt %) were dissolved in DCM, acetone or ethyl-acetate solvent. Finely divided Lidocaine powder or Cisplatin was sieved and a known amount was added to the polymer solution. The internal and external coagulants are distilled water and tap water respectively. 
     Preparation of Polymer Dope 
     Spinning solution was prepared by dissolving a specific amount of polymer to DCM, acetone or ethyl-acetate solvent. The polymer was then allowed to dissolve completely while being mixed on a roller mixer. 
     Hollow Fibre Spinning 
     The polymer solution was degassed in a vacuum chamber to completely remove gas bubbles and then transferred to the piston delivery vessel ( FIG. 1 ). The delivery vessel was then sealed and pressurised to 2-4 bar using nitrogen gas. The solution was extruded through a tube-in-orifice spinneret with 1 mm orifice diameter and 0.3 mm inner tube diameter. Water was used as an internal coagulant to form a hollow fibre and the flowrate pressure and other parameters were altered to produce fibres with ranges of properties (eg diameters, porosity and wall thickness). 
     The hollow fibre was then passed through a series of water baths to aid the phase inversion process (as shown in  FIG. 1 ). The air gap between the spinneret and water bath was usually at 0-10 cm. A suitable roller rotation rate needs to be selected to prevent mechanical dragging of the fibre. After spinning, the fibres were left to soak in fresh water for 2-4 days to remove residual solvent. The fibres were then dried at ambient conditions. 
     Drug Loading and Encapsulation Efficiency for Lidocaine 
     Hollow fibres, either newly spun or from post-release experiments, were dissolved in 10-20 mL of DCM. After complete dissolution, a known amount (usually 10 mL) of 0.1M H 2 SO 4  solution was added to the solvent to extract the lidocaine (as lidocaine is more soluble in acidic solutions). The UV absorbance (at 262 nm) of the acidic solution was measured and this lidocaine concentration was multiplied with the amount of H 2 SO 4  solution used (i.e. dilution factor) in order to give the total amount of lidocaine in fibre. 
     In order to determine the amount of lidocaine encapsulated in microspheres, approximately 0.1 g of microspheres was weighed and dissolved in 5 mL of DCM followed by 0.5 h sonication. To this solution, 5 mL of 0.1M H 2 SO 4  was added. 3 mL of sample solution from the H 2 SO 4  solution was withdrawn for lidocaine concentration determination taking into account of the dilution factor. H 2 SO 4  solution was used as the blank sample while measuring absorbance of the sample at 262 nm and as a result, the drug loading and encapsulation efficiency could be estimated. 
     
       
         
           
             
               Drug 
                
               
                   
               
                
               Loading 
                
               
                   
               
                
               
                 ( 
                 % 
                 ) 
               
             
             = 
             
               
                 
                   Total 
                    
                   
                       
                   
                    
                   Drug 
                    
                   
                       
                   
                    
                   Encapsulated 
                    
                   
                       
                   
                    
                   
                     ( 
                     mg 
                     ) 
                   
                 
                 
                   Total 
                    
                   
                       
                   
                    
                   Microsphere 
                    
                   
                       
                   
                    
                   Produced 
                    
                   
                       
                   
                    
                   
                     ( 
                     mg 
                     ) 
                   
                 
               
               × 
               100 
                
               % 
             
           
         
       
       
         
           
             or 
              
             
               
 
             
             = 
             
               
                 
                   Total 
                    
                   
                       
                   
                    
                   Drug 
                    
                   
                       
                   
                    
                   Encapsulated 
                    
                   
                       
                   
                    
                   
                     ( 
                     mg 
                     ) 
                   
                 
                 
                   Total 
                    
                   
                       
                   
                    
                   Weight 
                    
                   
                       
                   
                    
                   of 
                    
                   
                       
                   
                    
                   Fibre 
                    
                   
                       
                   
                    
                   used 
                    
                   
                       
                   
                    
                   
                     ( 
                     mg 
                     ) 
                   
                 
               
               × 
               100 
                
               % 
             
           
         
       
       
         
           
             
               Encapsulation 
                
               
                   
               
                
               Efficiency 
                
               
                   
               
                
               
                 ( 
                 % 
                 ) 
               
             
             = 
             
               
                 
                   Total 
                    
                   
                       
                   
                    
                   Drug 
                    
                   
                       
                   
                    
                   Encapsulated 
                    
                   
                       
                   
                    
                   
                     ( 
                     mg 
                     ) 
                   
                 
                 
                   
                     
                       
                         
                           
                             Theoretical 
                              
                             
                                 
                             
                              
                             Maximum 
                           
                            
                           
                               
                           
                         
                       
                     
                     
                       
                         
                           Drug 
                            
                           
                               
                           
                            
                           Used 
                            
                           
                               
                           
                            
                           
                             ( 
                             mg 
                             ) 
                           
                         
                       
                     
                   
                    
                   
                       
                   
                 
               
               × 
               100 
                
               % 
             
           
         
       
     
     The effectiveness of this method was tested by dissolving 9 mg of lidocaine powder in DCM and extracting it into 10 ml of 0.1M H 2 SO 4  solution. The absorption reading obtained was 1.6 which in turn gave a lidocaine concentration of 0.93 mg/mL. The corresponding ±2% error may be due to the variation of the instrument and this method was found to be reliable. 
     In Vitro Drug Release Analysis 
     Microspheres weighing between 0.1 and 0.3 g were suspended in a suspension of 5 mL of PBS (pH 7.4) incubated in a shaking water bath (100 rpm) at 37° C. for up to three weeks. Absorbance measurement was taken at different time intervals. I.e. Initially 0.5 h, then hourly and daily etc. In the case of hollow fibres, 20 mL of PBS (pH 7.4) was used. 
     Selected fibres were cut into equal lengths, weighed and injected with lidocaine HCL solution, before sealing the ends with Araldite epoxy resin. The fibres were then rinsed with distilled water to wash away any drug residue on the surface before immersing into 20 mL of PBS buffer solution (pH 7.4). For release experiments involving only incorporated lidocaine, the fibres were first weighed before being placed into buffer solution. The fibre samples were placed in a shaking water bath (100 rpm). At regular intervals, the concentration of lidocaine in the buffer was measured using an UV-visible spectrophotometer at 262 nm. The buffer was refreshed when the UV absorbance indicated plateau. 
     A similar procedure was adapted for cisplatin. At regular intervals, the concentration of Cisplatin in the buffer was measured using UV-visible spectrophotometer at 298 nm. The results are presented in  FIG. 31 . 
     Particle Size Analysis (for Beads) 
     Microparticle size was measured using a laser diffracting analyser, Malvern Mastersizer X. Each sample (suspended in cyclohexane with 1 mg/mL lecithin) was analysed a total of 2000 times to give an average value for the particle diameter or undersize. Briefly, in a laser diffraction particle analyser, a representative ‘ensemble’ or particle passes through a broadened beam of laser light which scatters the incident light onto a lens. This lens focuses the scattered light onto a detector array and using Mie theory (which solves Maxwell&#39;s equations exactly for the boundary conditions of a spherical particle) the particle size distribution is inferred from the collected diffracted light data. 
     Examples 1-4 
     Lidocaine Hollow Fibres 
     Spinning of Hollow Fibre Membranes 
     Hollow fibre membranes were spun with lidocaine powder incorporated into the polymer solution in order to produce fibres with a high drug loading (in the examples, lidocaine loadings were carried out to fulfil the long term anaesthetic requirements). Range of polymers/compositions with varying degradation rates were used. 
     
       
         
           
               
               
               
               
               
             
               
                 TABLE 1 
               
               
                   
               
               
                 Ex 
                   
                 Polymer 
                 Lidocaine 
                   
               
               
                 No. 
                 Polymer 
                 wt % 
                 wt % 
                 Notes 
               
               
                   
               
             
            
               
                 1. 
                 100DL-PLA 
                 20 
                 20 
                 Lidocaine powder added to 
               
               
                   
                   
                   
                   
                 solution 
               
               
                 2. 
                 100DL-PLA 
                 16 
                 20 
                 Lidocaine powder added to 
               
               
                   
                   
                   
                   
                 solution 
               
               
                 3. 
                 75:25 PLGA 
                 20 
                 40 
                 Lidocaine powder added to 
               
               
                   
                   
                   
                   
                 solution 
               
               
                 4. 
                 65:35 PLGA 
                 25 
                 40 
                 Drug pre-dissolved in ethanol 
               
               
                   
               
            
           
         
       
     
     By altering the concentration, composition, molecular weight of the polymer and the spinning parameters, hollow fibre membranes of varying thicknesses and controlled porosity are produced; the less viscous the solution (lower polymer concentration), the finer the fibres spun. 100DL-PLA fibres were found to be very stiff due to the high modulus of PLA and seem to be suitable for stents. 
     Drug Encapsulation Efficiency 
     The amount of drug entrapped successfully in the walls of the hollow fibre membrane was determined. The effectiveness of this method was tested by dissolving 9 mg of lidocaine powder in DCM and extracting it into 10 mL of 0.1M H 2 SO 4  solution. The test was repeated 5 times and the resultant concentration was 0.93 mg a standard error of 0.10 mg/mL. 
     
       
         
           
               
             
               
                 TABLE 2 
               
             
            
               
                   
               
               
                 Drug loading and encapsulation efficiency 
               
               
                 of hollow fibre membranes 
               
            
           
           
               
               
               
               
            
               
                 Ex 
                   
                 Drug Loading 
                 Encapsulation Efficiency 
               
               
                 No. 
                 Polymer 
                 (wt %) 
                 (%) 
               
               
                   
               
            
           
           
               
               
               
               
            
               
                 1. 
                 PLA 20% 
                 4.43 
                 22.15 
               
               
                 2. 
                 PLA 16% 
                 5.12 
                 25.60 
               
               
                 3. 
                 75:25 PLGA 
                 7.72 
                 19.23 
               
               
                 4. 
                 65:35 PLGA 
                 6.22 
                 15.55 
               
               
                   
               
            
           
         
       
     
     Sample No 1 
     0.012 g of Ex. 1 was dissolved in DCM and 10 mL of 0.1M H 2 SO 4  solution added to the solvent to extract the lidocaine. After vigorous shaking, the UV absorbance (at 262 nm) of the acidic solution was found to be 0.114, which correlated to a concentration of 0.0532 mg/mL. Therefore the total amount of lidocaine present was 0.532 mg. The drug loading of Sample No 1=(0.532 mg Lid/12 mg fibre)×100%=4.43%. The encapsulation efficiency=4.43/20×100%=22.15% 
     Sample No 2 
     0.0326 g of Ex. 2 was dissolved in DCM and 20 mL  0 . 1 M H 2 SO 4  solution was added. UV absorbance of the acidic solution was 0.165 indicating a concentration of 0.0835 mg/mL. Therefore there was 1.67 mg of lidocaine present which resulted in a drug loading of 5.12%. Encapsulation efficiency 5.12/20×100%=25.6% 
     Sample No 3 
     0.02 g of Ex. 3 was dissolved and lidocaine extracted into 10 mL of 0.1M H 2 SO 4  solution. Absorbance at 262 nm was 0.284 therefore there was 1.543 mg of drug in the solution. Drug loading was calculated as 7.72% and the encapsulation efficiency was 19.23%. 
     Sample No 4 
     0.062 g of Ex. 4 was dissolved and lidocaine extracted into 10 mL of 0.1M H 2 SO 4  solution. Absorbance at 262 nm was 0.673 therefore there was 3.856 mg of drug in the solution. Drug loading was calculated as 6.22% and the encapsulation efficiency was 15.55%. 
     In Vitro Drug Release Studies 
     The release of lidocaine into PBS (pH 7.4) was measured by UV-Visible spectrophotometry to determine its in vitro release profile. Lidocaine was incorporated into the walls of hollow fibres Sample 1, Sample 2, Sample 3 and Sample 4 and release was facilitated by diffusion initially, followed by degradation of the polymer. Additionally, lidocaine HCl solution (20 mg/mL) was injected into the lumen of sample 3 with the ends sealed in order to determine the diffusion release of the drug. 
     At the end of each experiment, the amount of drug remaining in the polymer was calculated by dissolving the hollow fibres in DCM, followed by extraction of the drug into 0.1M H 2 SO 4  solution. The release profile results were then represented as a percentage release over time, where: 
     
       
         
           
             
               % 
                
               
                   
               
                
               Release 
             
             = 
             
               
                 
                   Cumulative 
                    
                   
                       
                   
                    
                   amount 
                    
                   
                       
                   
                    
                   of 
                    
                   
                       
                   
                    
                   drug 
                    
                   
                       
                   
                    
                   released 
                 
                 
                   ( 
                   
                     
                       Final 
                        
                       
                           
                       
                        
                       amount 
                        
                       
                           
                       
                        
                       released 
                     
                     + 
                     
                       Amount 
                        
                       
                           
                       
                        
                       remaining 
                     
                   
                   ) 
                 
               
               × 
               100 
                
               % 
             
           
         
       
     
     Release of Lidocaine from Sample 1 and 2 
     Lidocaine powder was incorporated into the walls of the fibre (44.3 mg of lidocaine per 1 g of fibre). Two experiments were run concurrently and both their drug release profiles were obtained. 
     After incubation into PBS (pH 7.4), there was an immediate release of lidocaine into the surrounding medium, with 15% of the total contained drug released within the first 30 minutes (see  FIG. 23 ). By 4 hours, 30% had been released. This was due to the diffusion of non-trapped drug or drug near the surface of the polymer wall, as water diffuses into the polymer matrix to form a homogenous distribution. 24 hours after immersion, a further 10% of drug had diffused out of the hollow fibre. Subsequent release was slow with barely any lidocaine released after the third day. 
     However, when the residual drug content of the fibres was determined after 17 days into the experiment, it was found that only 45% of the lidocaine had been released with a further 55% still remaining in the hollow fibre wall. Due to the slow degradation of 100-DL PLA, the drug was entrapped within the polymer matrix and there was insufficient degradation during the period of experiment to facilitate the release of the remaining drug confirming that PLA is suitable for long term drug release. 
     Release of Lidocaine from Sample 3 
     Lidocaine powder was incorporated into the polymer matrix of sample 3 which was made of 75:25 PLGA. Initial drug release was similar to that of sample-2 as 18% of lidocaine was released in the solution. After 24 hours the total amount released had increased to 21% (see  FIG. 24 ). 
     This was followed by a slow gradual release (zero order kinetics), reaching 100% release after 20 days post-immersion. This was accompanied with degradation of the fibres into smaller particles therefore the release of lidocaine was diffusion as well as erosion-controlled. Further absorption readings could not be taken accurately due to the presence of polymer particles suspended in the solution.  FIGS. 25   a  and  25   b  are SEMs of 25:75 PLGA with Lidocaine in the wall. 
     Release of Lidocaine and Lidocaine HCl from Sample 3 
     To increase the drug loading of the hollow fibres, lidocaine HCl solution was injected into the hollow fibres in addition to having drug particles in the fibre wall. The resulting release profile showed a very quick release effect with 54% of the total drug loaded released in the first 2 days. The results shows that initially the liquid drug diffused quickly into bulk solution (see  FIG. 26 ). 
     After 24 days, 72% of lidocaine was released and this was followed by a very gradual release for the next 22 days. As the ends of the fibres were sealed initially, water uptake into the hollow fibre did not proceed as quickly as plainly immersed fibres and therefore degradation was also slower. At the end of 23 days, 80% of the drug had been released. 
     The injection of solution into the hollow fibre lumen was able to increase the drug loading considerably, as 0.1 g of fibres was able to contain 20 mg of lidocaine. 
     Release of Lidocaine from Sample No 4 
     Lidocaine was incorporated into the walls of Sample 4 which was made of 65:35 PLGA, the fastest degrading polymer used in the experiments conducted. The drug powder was first dissolved in ethanol before it was added into the polymer solution to ensure that there was good dispersion of drug. The resulting release profile showed a more gradual release of drug initially; only 7% was released in the first 60 minutes. After one day around 20% of the total drug had diffused out into the PBS solution (pH 7.4) (see  FIG. 27 ). 
     Subsequent release continued to be gradual while degradation of the polymer took place more rapidly than the previous experiments. Degradation of the polymer was quite extensive leading to a particle suspension in the buffer. 
     Development of Fibres for 5-Fluorouracil Tests and Cisplatin Tests 
     Materials 
     The polymers used in the spinning processes were poly(D,L-lactic acid) (PLA) and polycaprolactone (PCL) and PLGA. Dichloromethane, ethyl acetate, 1-methyl-2-pyrrolidone (NMP) (99.5%) was used as solvents to dissolve the PLA, PLGA polymer and prepare the polymer solution. Acetone on the other hand, was used to prepare the PCL polymer solution. Distilled water purified was used as an internal coagulant, as well as the external coagulant. Commercial liquid form 2.5 g/100 ml 5-florouracil (5FU) and cisplatin 10 mg/ml was used for pharmacokinetic studies. Also cisplatin, carboplatin and 5FU powdered form used for fibre and microsphere preparations. For the medium in drug release experiment, phosphate buffer saline (PBS) pH 7.4 was used. 
     Spinning Solution Preparation 
     The required quantity of organic solvent (ethyl acetate, dichloromethane, acetone, or NMP), was poured into a one-litre wide-neck bottle and then the desired quantity of polymer (PLA, PLGA PCL and the mixture of co-polymers) was slowly added. The mixture was stirred on a rotary roller to form the polymer solution. The finely divided powdered drug (eg. cisplatin, carboplatin, 5FU or oxaliplatin and other) was added to the polymer mixture and vigorously stirred to achieve a homogeneous mixture. This ensured that drugs were only exposed to solvents limited time. During the triple layer spinning the drug particles were suspended in a saline solution or in purified water which was injected into the middle layer of the fibre through a syringe pump. This process aids the solidification process entrapping the drugs efficiently within the polymer matrix. For PCL, the polymer solution was needed to be heated in a water bath to 55° C. to improve dissolution of poly s-caprolactone. 
     
       
         
           
               
             
               
                 TABLE 3 
               
               
                   
               
               
                 Spinning Parameters 
               
               
                   
               
             
            
               
                   
               
            
           
           
               
               
               
            
               
                   
                 Coagulation bath temperature (° C.) 
                 20 
               
               
                   
                 Injection rate of internal coagulant (ml/min) 
                 from 2 to 14 
               
               
                   
                 Nitrogen Pressure (bar) 
                 2-4 
               
               
                   
                 Air gap (cm) 
                 0 and 3 
               
               
                   
                 Linear extrusion speed (rpm) 
                 varies 
               
               
                   
                 Bore liquid 
                 purified water 
               
               
                   
                 External coagulant 
                 purified water 
               
               
                   
                   
               
            
           
         
       
     
     Bio Polymeric Hollow Fibre Spinning 
     The polymer solution was degassed for 24 hours at room temperature before the spinning process in order to completely get rid of any gas bubbles of viscous polymer solution in the spinning process. The polymer solution was then transferred to a piston delivery vessel. The tank was pressurised to 2-4 bar using nitrogen during the spinning process (see  FIG. 1 ). 
     A tube-in-orifice spinneret with orifice diameter and inner tube diameter of 2.0/0.72 (mm) was used to produce the hollow fibres. The air gap was kept at 0 cm and 3 cm, and water was used as the internal and external coagulant for all spinning runs. Finally, in forming the hollow fibre, it was passed through a series of water baths to aid the solidification process. The hollow fibre was then washed thoroughly in a second water bath. During spinning care was taken to ensure continuity of the pressure and internal water support in order to avoid entrapment of air and separation of the fibre, which would eventually result in an unsuccessful spinning. The hollow fibres were left to soak in fresh water; this was essential for thorough removal of residual solvent and then vacuum dried to remove any residual solvents. Prepared drug encapsulated fibres were freeze dried/vacuumed packed for characterisation 
     Characterisation 
     Scanning Electron Microscope (SEM) 
     The scanning electron microscope (SEM) was used for the characterisation of hollow fibres before drug (5-fluorouracil) injection and after drug injection into lumen of the fibres. The surface structure, particle size and porosity distribution in the matrixes were observed using the JEOL JSM6310 model. All samples were dry before use. Firstly, the sample was frozen in liquid nitrogen for 20-30 seconds and then sectioned using a sharp blade. Then, a specimen plate was coated with a thin layer of gold under 3 mbar pressure for 3-5 minutes with the Edwards Sputter Coater (S150B). The SEM was operated in the range 10-20 kV and micrographs were taken of a number of areas on each sample. 
       FIGS. 28   a - d  show SEM micrographs of different samples before any degradation.  FIG. 28   a  shows a Cross section of 85:15 PLGA at 250× magnification.  FIG. 28   b  shows a Cross-section of 75 L: 25 PLGA at 250×magnification.  FIG. 28   c  shows a cross section of a Polymer blend of 65:35 PLA/PCL.  FIG. 28   d  shows a SEM micrograph of 85:15 PLGA at 1000× magnification showing microporous structure. 
       FIG. 29   a - c  show SEM micrographs of different samples after 6 weeks of degradation.  FIG. 29   a  shows crystallised 65:35 PLGA,  FIG. 29   b  shows 75 L:25 PLGA and  FIG. 29   c  shows PLGA 85:15. 
     Viscosity Test 
     The viscosity values of the spinning dopes were obtained by using a Bohlin CS 50 Rheometer (Stress Viscometry Model). In order to spin fibres, the viscosity of the polymer solution should be generally between 1-5 Pa·s for non-biodegradable polymers. For biodegradable polymers, the viscosity of the spinning dope might be lower. It is important when applying polymer dope, that the correct amount is used. 
     Polymer Degradation Study 
     0.07 g of each hollow fibre was weighed and immersed in 20 ml of de-ionised water in a screw-topped bottle. These were then placed on an orbital shaker maintained at a temperature of 37° C. The pH of the degradation medium was measured every two days to monitor the rate of generation of acidic byproducts. A Jenway (Model 3051) pH meter was used for this purpose. 
     Drug Release Study 
     In vitro drug release studies were carried out in phosphate buffer saline (PBS) at both room temperature (˜25° C.) and 37° C. PBS solution was applied to keep the degrading system at a constant pH 7.4 value. Every type of polymer fibre was cut into 10 pieces, with 5 cm each. At room temperature, each fibre piece was first sealed at one end with bio-adhesive, and was injected with liquid 5-FU with a syringe pump from the other open end before it was sealed. Every 10 pieces of the same type of fibre were placed into a bottle which was filled with 20 ml of PBS solution. The bottles were placed at room temperature and at a constant temperature water bath, maintained at 37±0.1° C. for 6 months. The amount of drug release was measured every 24 hours within the 3 months. The quantitative measurement of the amount of drug released was measured with a UV Spectrometer. A calibration curve was performed as the concentration versus absorbance for the drug in buffer (wavelength of absorption, □=266 nm (5-FU) and λ=298 nm (cisplatin and carboplatin) or atomic absorption spectroscopy) which was plotted in a logarithmic scale gave a straight line. Samples were drawn from the release medium every 24 hours within the predetermined release period of 3 months to full-fill the cuvette, which would be put into the UV Spectrometer. The absorbance of the drug was noted. The sample in the cuvette was then put back into the release medium. 
     Hollow fibre membranes of different polymer compositions were investigated to determine their morphology and porosity, response to hydrolytic degradation as well as release rates of the anti-coagulant drug, 5FU and cisplatin and carboplatin. 
     The following table shows the compositions of the hollow fibres studied and their corresponding spinning conditions. 
     
       
         
           
               
             
               
                 TABLE 4 
               
             
            
               
                   
               
               
                 Hollow fibres investigated, their compositions and spinning 
               
               
                 conditions with range of chemotherapeutics. 
               
            
           
           
               
               
               
               
               
               
            
               
                   
                   
                   
                   
                   
                 Internal 
               
               
                   
                   
                   
                 Polymer/ 
                 Air 
                 Water 
               
               
                   
                   
                   
                 solvent 
                 Gap 
                 Flowrate 
               
               
                 No. 
                 Fibre 
                 Composition 
                 wt % 
                 (cm) 
                 (ml/min) 
               
               
                   
               
               
                 1 
                 50:50 PLGA 
                 50% DL-PLA, 50% PGA 
                 25 
                 0 
                 3 
               
               
                 2 
                 65:35 PLGA 
                 65% DL-PLA, 35% PGA 
                 20 
                 0 
                 3 
               
               
                 3 
                 75:25 PLGA 
                 75% DL-PLA, 25% PGA 
                 20 
                 0 
                 3 
               
               
                 4 
                 85:15 PLGA 
                 85% DL-PLA, 15% PGA 
                 20 
                 0 
                 3 
               
               
                 5 
                 75L:25 PLGA 
                 75% L-PLA, 25% PGA 
                 10 
                 0 
                 3 
               
               
                 6 
                 75:25 PLA/PCL 
                 75% DL-PLA, 25% PCL 
                 20 
                 0 
                 3 
               
               
                 7 
                 100 PLA 
                 100% DL-PLA 
                 30 
                 3 
                 4 
               
               
                 8 
                 75L:25 PLGA 
                 75% L-PLA, 25% PGA 
                 15 
                 5 
                 4 
               
               
                   
                   
                 35% Cisplatin loading 
               
               
                 9 
                 75L:25 PLGA 
                 75% L-PLA, 25% PGA 
                   
                 5 
                 4 
               
               
                   
                   
                 30%, 50% and 70% Carboplatin 
               
               
                   
               
            
           
         
       
     
     Fluorouracil (5-FU) (see FIG. 30) 
     Hydrophilic 5-FU accelerates the degradation and drug release rate from polymer hollow fibres as shown in the figure. 
     Amorphous fibres (PLGA) degrades and releases 5-FU faster than crystalline PLA fibres, 75dL:25PCL has the second slowest degradation and drug release due to its hydrophobicity, but releases twice the 5-FU amount released from PLA fibre. 
       FIG. 31  shows the percentage of Cisplatin release from hollow fibres of different concentrations where the drug was in liquid form within the lumen of the fibre.  FIG. 32  shows the released kinetics of 35% powder carboplatin loaded PLGA 75:25 fibre.  FIG. 33  shows the Carboplatin three weeks dose; Percentage mass released (cumulative) of carboplatin against time from carboplatin loaded-fibres fabricated using PLGA (75:25) and 30%, 50% and 70% carboplatin, incubated at 37° C. in PBS (pH 7.4) respectively.