WO2000050497A1

WO2000050497A1 - Novel polymer materials and their preparation

Info

Publication number: WO2000050497A1
Application number: PCT/FI2000/000135
Authority: WO
Inventors: Eija SÄILYNOJA; Anders SÖDERGÅRD; Jukka Salonen; Peter Holmlund; Pekka HÄYRY; Mika Koskinen; Ilkka Kangasniemi; Marju VÄKIPARTA
Original assignee: Bioxid Oy
Priority date: 1999-02-25
Filing date: 2000-02-22
Publication date: 2000-08-31
Also published as: EP1165672A1; AU2808000A; FI990404A0

Abstract

The invention comprises a polymer material bearing a biologically active agent, a device made of or coated with the polymer material and a method for the preparation of the polymer material. The material comprises a core polymer, which is in bulk form or formed into a body or a coating of said body, which body optionally has the shape of a finished device. The biologically active agent is bound to a functional group on the core polymer. The functional group has been generated by oxidation of the core polymer or it is attached to a vinyl terminated monomer which has been grafted onto the core polymer.

Description

NOVEL POLYMER MATERIALS AND THEIR PREPARATION

This invention relates to novel biodegradable polymer materials, in bulk form or in the form of a shaped body or a coating of said body, said polymer material bearing biologically active agents. Furthermore, the invention concerns polymer materials shaped into a body, into which such biologically active agents can be introduced. Still further, the invention concerns a method for the introduction of biologically active agents into solid polymers.

BACKGROUND

Generation of macroradicals in various core polymers by irradiation or by peroxide treatment, followed by grafting a monomer onto the core polymer, has been described before, e.g. Ullmann's Encyclopedia of Industrial Chemistry, Vol. 22, pages 478 — 497, Finnish patent application FI 922878 and DE 2027178.

Mixtures of polymers and various osteogenic materials in the form of particles, e.g. particles of bioactive ceramics, have been described in many patents, e.g. US 4 595 713.

US 5 610 241 discloses a graft polymer having side chain groups derived from amino acids. The graft polymer obtained is useful for drug dispensing, especially for delivery of 5-fluorouracil. WO 97/44013 describes particles of polymers, in which the core polymer contains an amino acid group to which an amino acid is grafted. The polymer particles are useful for the delivery of a pulmonary drug. In these patents, the grafting process is carried out by chemical reactions not involving radical mechanisms and not by introduction of vinyl terminated monomers subsequent to irradiation or peroxide treatment of the core polymer.

Ceramic alloplasts such as hydroxyl apatite (HA) and bioactive glass (BAG) are materials that are commonly used for orthopedic applications. These inorganic biomaterials exhibit properties that promote bone growth and regeneration. In physical terms, ceramics exhibit good compressive strength but are quite brittle (similar to bone) when exposed to tension. In order to overcome this problem, it would be helpful to be able to design composite, organic-inorganic devices (Verheyen et al, 1992) in which the bone promoting capacity of a ceramic material and the tensile strength of an appropriate organic polymer would be combined.

Immobilization of biologically active molecules onto different surfaces of various medical devices can favorably influence the biocompatibility of implanted devices. Heparin, for instance, has since the late sixties been bound to surfaces in order to utilize the property of the molecule as an anticoagulant. Until now, however, the control of the release of heparin has been largely dependent on the chemical character of the polymer-core-material it has been immobilized on, and complicated reactions in several steps have regularly been used in attempts to prepare devices with favorable characteristics (Kim 1996). The heparin was not attached to functional groups on monomers grafted onto the core polymer. Instead, the heparin was immobilized on special large linker molecules, which had a tendency to cause allergic reactions.

OBJECTS OF THE INVENTION

One object of this invention is to provide devices made of or coated with an biodegradable organic polymer material, especially tailored devices to be used in the mammal body, wherein desired regions of the device are provided with agents having a desired biological activity.

Another object of this invention is to provide a device or coating of said device that is biodegradable and to which the biologically active agents are strongly bound.

Still another object is to achieve a device or coating of said device that is biodegradable and having a strongly bound outer layer into which a therapeutically active agent is loaded, and from which said therapeutically active agent is released at a controlled rate.

SUMMARY OF THE INVENTION

Thus, according to one aspect, this invention concerns a biodegradable polymer material bearing a biologically active agent, said material comprising a core polymer, wherein said core polymer is in bulk form or formed into a body or a coating of said body, which body optionally has the shape of a finished device, wherein the biologically active agent is bound to a functional group to the core polymer. The material is characterized in that,

a) said functional group has been generated by oxidation of the core polymer, or

b) said functional group is attached to a vinyl terminated monomer which has been grafted onto the core polymer.

According to another aspect, this invention concerns a device made of or coated with a biodegradable polymer material, useful for finishing into a device of a polymer material bearing a biologically active agent, wherein said polymer material comprises a core polymer, which has been formed into a body or a coating of said body being the shape of a finished device. The device is characterized in that the core polymer bears a functional group on a desired region of the body, said functional group being capable of binding the biologically active agent, wherein

According to still another aspect, the invention concerns a method for the preparation of a biodegradable polymer material bearing a biologically active agent, said material comprising a solid core polymer, wherein the biologically active agent is bound to the core polymer. The method is characterized by the steps of

a) generating macroradicals on the core polymer, and

i) allowing said macroradicals to react with oxygen to create functional groups, or ii) grafting vinyl terminated monomers bearing functional groups onto the core polymer, and

b) attaching said biologically active agent to said functional group.

BRIEF DESCRIPTION OF THE DRAWINGS

Figure 1 shows the release of heparin from heparinized silica gel in SBF (Simulated Body Fluid) solution as a function of time.

Figure 2 shows the distribution of silver as a function of the film thickness for poly(ε-caprolactone) grafted with acrylic acid and further reacted with silver ions. DETAILED DESCRIPTION OF THE INVENTION

Definitions and preferred embodiments

The term " biodegradable" in this context means that it is degradable upon prolonged implantation when inserted in the mammal body.

The term "biologically active agent" shall be understood as an agent causing a valuable effect in vivo, such as a bioactive effect (i.e. promoting the binding of bone to an artificial implant inserted into the mammal body or other systems containing living organism), a therapeutic effect, or the like. The term covers also agents useful for attaching bioaffinity groups such as antibodies, antigens, nucleotides etc. to a surface of a device for use in a bioaffinity assay. Such "attaching groups" are for example members of an affinity pair such as biotin-streptavidin or the like.

It shall be noted that certain biologically active agents also possess the ability to bind other biologically active agents. As an example can be mentioned silica gel, which as such may or may not, depending on e.g. its dissolution rate, nanoporosity and specific area, be biologically active (i.e. bioactive) in that it promotes the bone binding to an implant coated with silica gel. Furthermore, silica gel can be used as a matrix which can be loaded with a biologically active agent (e.g. a drug), which then is released in certain conditions, e.g. in body fluids. Silica gel can therefore also be used for the controlled release of therapeutically active agents.

The term "functional group" means a group which is able to bind a biologically active agent.

The term "functionalization" means the introduction of a "functional group" into the core polymer or the generation of reactive macroradicals in said core polymer. The "core polymer" is the same as backbone polymer. The core polymer can be a thermoplastic polymer or a thermosetting polymer. The polymer can also be a protein, e.g. collagen, or parts thereof. The thermoplastic polymers can be condensation polymers or addition polymers. The polymer can also be a co-polymer. According to a preferred embodiment, the polymer is a synthetic thermoplastic polymer. Particularly preferable degradable polymers for use in the mammal body are PLLA (poly-L-lactide) and PCL (polycaprolactone) or copolymers of L-lactide and caprolactone units.

The "solid core polymer" means that the core polymer is in bulk for such as powder or granules, or that it has been shaped into a body.

The word "body" shall be understood to be any defined piece or continuous article such as a spherulite, sheet, film, plate, stick, pin, screw, tube, fiber, hollow fiber, woven fabric or non-woven fabric, or the like also when built to resemble human or animal body parts such as ear, nose, joints etc. or parts thereof. The coating of a body shall be understood to be any coat of material on a body as defined above.

It shall be noted that the term "coating" is used in two different contexts in this application. Firstly, it refers to a coating of a body or device which coating is of a biodegradable polymer material. Secondly, it refers, e.g. in examples 1, 2 and 4 to follow, to the outer layer, i.e. coating, of said body, device or biodegradable polymer coating (first meaning) of such body or device. This coating (second meaning), i.e. outer layer of the biodegradable polymer material, can be, as in the examples 1 — 4, e.g. silica gel, heparin or silver. It should, however, be clear from the context, which meaning is valid.

The term "monomer" means a vinyl terminated monomer, wherein the vinyl group is capable of binding to reactive sites, such as macroradicals, generated into the core polymer. According to a preferred embodiment, the polymer material bearing the biologically active agent has been shaped to a device or is a coating of said shaped device. Said device comprises a core polymer formed into a body or a coating of said body, which body optionally has the shape of a finished device, wherein the biologically active agent is bound to the core polymer. The biologically active agent is bound to functional groups on the core polymer on a desired region of said body.

The oxidation of the macroradical means a formation of functional groups containing oxygen (e.g. carboxy or hydroxy group) due to excess of oxygen in the atmosphere., The oxidation can be carried out directly or in a separate step.

As suitable functional groups can be mentioned carboxylic, hydroxy, ester, ether, epoxy, amino, amido, imino, imido, cyano, nitro, sulfono, mercapto, siloxy and phosphoro groups. According to a preferable embodiment, the functional group is bound to the vinyl terminated monomer before the vinyl group is bound to the core polymer. As example of a very useful functional group can be mentioned an alkoxysilyl group, which can be bound to the vinyl terminated monomer before the grafting is carried out. After cleavage of the alkyl group from this alkoxy group, silica gel will via its Si atoms become covalently bound to the - O - Si groups attached to the grafted polymer. Another possibility is that the oxygen atoms of the silica gel bind to the Si atoms of the grafted groups.

According to a preferred embodiment of this invention, the biodegradable polymer material is formed to a device or a coating of said device to be implanted into the mammal body to serve different clinical applications, [e.g. stents; implants (dental or orthopedic); implants for controlled drug delivery; bone fixation pins; fixation plates; regeneration matrixes etc.]. The core polymer of the device or its coating is degradable in the biological environment. If the biodegradable polymer material is formed to a coating of said device the core of the device can be of any material inert or degradable. The biologically active agents must of course be pharmaceutically acceptable.

Generally, the biologically active agent can be a protein, a polypeptide, a polysaccaride, an oligosaccaride, a mono- or disaccaride, an organic compound, an organometallic compound or an inorganic compound containing any element with an atomic number ranging from 3 to 84. As particularly useful biologically active agents in implantable devices can be mentioned non-toxic inorganic ions or polymers thereof; a silica gel as such or a silica gel loaded with a therapeutically agent; heparin; a growth factor; a growth factor inhibitor; an integrin blocker (e.g. a Ila/IIIb inhibitor); an ohgonucleotide; a complete functional gene in sense or antisense orientation in a suitable expression vector or any other expression vector construct for local delivery of therapeutically active agents.

According to a particularly valuable embodiment, the polymer material is formed to a stent or a coating of said stent. A stent is a splint or tube to be placed temporarily or permanently inside a duct, canal or blood vessel to aid healing or relive an obstruction. Stents are inserted in blood vessels in balloon angioplasty, where the balloon is placed inside an expandable stent (tube) which is expanded as the balloon is pressurized. Known stents are typically tubes of metallic networks.

The stent according to this invention can be made, for example, of a biodegradable or inert polymer. The inner wall of the tube can be coated with a silica gel loaded with the anticoagulant drug heparin. According to a preferred embodiment to be described more in detail in the Experimental Section, the silica gel is strongly bound to the inner wall of the tube via cleaved alkoxysilyl substituents on the grafted monomers. A silica gel coating made in this manner is very stable and also very thin, which is an important feature in this field of use. The outer wall of the tube can, if desired, be provided with other biologically active agents. According to another preferable embodiment, heparin is attached to a functional group on a graft attached to the inner wall of the stent.

In this invention the functionalization of the core polymers is performed in solid state. This method has the following advantages over the known melt modification methods:

- performed articles can be functionalized

- the functionalization can be limited to certain areas of the article

- the functionalization process is easily controlled

- several different functional units can be added to the article - further reactions can easily be performed

The concept of grafting to solid polymers includes various methods and modifications, in which the basic principle is the generation of reactive sites on a solid core polymer which then initiates radical polymerization of monomers brought into contact with the activated core polymer. The graft-co-polymers so obtained can be designed, by manipulating the length of selected grafts, to contain a variable number and different types of chemically reactive groups, (i.e. functional groups). The functional groups in the grafts can be further used for attachment of biologically active agents.

The active sites on the core polymer are described as macroradicals, which can be generated by ionizing radiation of different energy and wavelength (e.g. UV-radiation, γ-radiation and accelerated particles). Alternatively, the macroradicals can be generated by exposing the core polymer to a peroxide.

Although the solid core polymer material can be functionalized in bulk form (powder or granulate), the core polymer material has preferable been formed into a body or a coating of said body, which body optionally has the shape of a finished device, before generating of macroradicals optionally followed by grafting.

The biologically active agents can either be attached directly to the oxidized macroradicals in the core polymer (one step method), or alternatively, attached to the functional group on the monomer grafted on the core polymer (multistep method).

The final activity of the material can be controlled by the chemical nature of the monomer selected, the number of grafts attached to the core polymer and the length and distribution of the polymer grafts.

The main field of is invention can briefly be summarized as utilization of novel grafting methods for equipping a biodegradable polymer material with tailored properties, which give raise to a desired respond when brought into contact with living tissues.

The invention is disclosed in more detail by the following experiments.

EXPERIMENTAL SECTION

Example 1 describes a method in which the biologically active agent is attached directly onto the irradiated and oxidized polymer. Examples 2 to 4 describe the binding of the biologically active agents to functional groups on monomers grafted onto the core polymer. The functionalization of the different core polymers in the Examples 2 to 4 was performed according to the same method. Films were prepared from polymer granules by compression molding the polymer in question at about 20 °C above its melting point. The granules were allowed to melt for two minutes before pressing (200 bar) the material to form a film and then cooled rapidly to room temperature. The films were cut into square pieces (thickness 120 — 160 μm), carefully washed with ethanol and dried. Macroradicals in the polymer materials were achieved by irradiating the pre-weighed films by making use of an Electrocurtain^® electron accelerator in air at an acceleration voltage of 175 kV (Sδdergard 1998a and b). The irradiated films were removed from the accelerator and immersed into defined vinyl monomer solutions (vinyl-R / concentration) at 25 °C for variable reaction times without using homopolymerization inhibitors or chain-transferring agents. Before grafting, the monomer solutions were purged with nitrogen for at least 30 minutes in order to minimize the presence of oxygen during the grafting process. The grafted films were washed with deionized water for several hours in order to remove eventual homopolymers and dried to constant weight in vacuum at room temperature. The grafted films were further functionalized in various ways, depending on the desired properties. The methods are described in detail in the Examples 2 — 4.

Example 1

Silica gel coating attached directly to irradiated biodegradable polylactide

By using the grafting technique explained above, the surface properties of hydrophobic polymers such as poly-L-lactide (PLLA) and polycaprolactone (PCL) can be altered to be more hydrophilic for silica gel to attach there to. Silica gel is used as a controllable drug delivery coating as well as to achieve better biocompatibility for implants. In this example, non-pregrafted, only irradiated PLLA fibers were coated with nonheparinized silica gel. Preparation of silica-sol

The silica-sol was prepared by a two step sol-gel process using acid as a catalyst (Brinker 1990). The following reagents were used: tetraethoxysilane (TEOS) (Aldrich), deionized water, ethanol and acetic acid. The r-value (water/TEOS molar ratio) was 3.55. Ethanol was used as a solvent to obtain better viscosity (water/ethanol molar ratio = 1). TEOS, deionized water, ethanol and catalyst (acetic acid) were added to a glass container and stirred as long as the inorganic water phase and the organic TEOS phase had become homogenized to obtain the hydrolysis solution.

Coating of PLLA fibers

The silica gel coating applied to PLLA polymer fibers (0 0.2 mm) immediately after electron beam (EB) radiation (Electrocurtain^® electron accelerator) treatment by dipping the fibers into the sol in the hydrolysis phase. The EB radiation dose used was between 25 — 300 kGy. The uniformity of coating was examined with SEM (Scanning Electron Microscopy) measurements. The PLLA fibers used for the experiments were supplied by Bionx Implants Ltd.

Treatment of hydrophobic PLLA fibers by methods that cause macroradical formation on the polymer surface improved the attachment of the silica gel coating on the polymer compared to direct dipping without foregoing irradiation. The pre-weighed PLLA samples were irradiated in air and dipped in the hydrolysis solution. During the irradiation the hydrophobic surface of PLLA becomes more hydrophilic.

According to information obtained from SEM studies, an EB radiation dose between 25 — 50 kGy was enough to make the surface of the PLLA fiber suitably reactive for coating purposes. The thickness of the coating produced was 0.5 μm and its cracking after bending was minimal.

Biocompatibility test

The biocompatibility of the high grade PLLA fibers and the silica gel coated PLLA fibers were tested on animals. The tests were carried out at the Transplantation laboratory, Hartman Institute, University of Helsinki.

Male Wistar rats were used for the experiments. Their weights were between 300 — 400 g at the time of implantation. Three different time points (7, 14 and 28 days) were used. One or two implants were implanted per one rat and five rats were used per one time point. The samples were implanted either in subcutane or in omentum.

The test showed that when silica gel coated PLLA and plain PLLA sticks were implanted subcutaneously in experimental animals for seven days a difference could be seen in the tissue reactions. In preliminary experiments the histopathological reaction of the tissues involved revealed that the uncoated material caused a clearly stronger inflammatory reaction (mononuclear phagocyte-reaction) as compared to the silica gel coated PLLA and less fibrosis and capillary proliferation. The result of the experiment indicates that the biocompatibility of a biodegradable polymer can be altered by a strongly attached inorganic polymer surface.

Example 2

Application of heparin loaded silica gel onto a grafted PLLA-co-CL copolvmer.

In this example, heparin is immobilized onto silica gel and then the grafted PLLA sheets are coated with the heparinized silica gel. Bulk heparinized silica gel samples were obtained for the drug delivery ability tests. It is known that silica gel can be used as a drug delivery system (Kortesuo et al. 1999).

Preparation of silica-sol

The heparin containing silica sol was prepared by the same two step sol-gel process using acid as a catalyst as mentioned in example 1 (Brinker 1990, Ellerby et al. 1992). Nitric acid (HN0₃) (Merck) was used as a catalyst instead of acetic acid and ammonium hydroxide (NH₄OH) to raise pH up to 4.5. The heparin used was a sodium salt obtained from Orion Corporation (biological activity 139 I.U./mg). To obtain 100 ml hydrolysis solution, 48 g of TEOS, 45 g of deionized water, 4 g of glycerol and 10J g of catalyst (0.04 M HN0₃) were added to glass container and stirred as long as the inorganic water phase and the organic TEOS phase have become homogenized.

To obtain small bulk pieces for the heparin delivery test, the hydrolysis solution was divided to smaller samples (V = 0.5 ml) that were aged in a sealed polystyrene container at 40 °C and 40 % of relative humidity for 3 days. After this gelatination process samples were dried at room temperature for additional 3 days. The heparin content in one gel piece varied from 1 to 15 % (calculated from the dry weight). H₂0/TEOS -ratio (r-value) was between 14 — 16 and pH 4,5. The same sol was used for coating of the graft polymers.

The coating was applied to Poly(L-lactide)-co-ε-caprolactone (PLLA-co-CL) sheets immediately after hydrolysis of TEOS and water by dipping technique (acrylamide grafted polymers) or polymer sheets were added in the reagents mixture already before hydrolysis was started (trimetoxysilane grafted polymers). A strong bond was obtained. The uniformity of coating was examined with SEM measurements. Functionalization of the polymer

The functionalization of the different polymers in the examples 2 — 4 were performed according to the same method. PLLA and PCL were polymerized according to the method described elsewhere (Holmlund 1999) The pre-weighed polymer film samples were irradiated by using an Electrocurtain^® electron accelerator in air at an acceleration voltage of 175 kV. The irradiated films were removed from the accelerator and immersed into the monomer solutions at ambient temperature for various reaction times without using any homopolymerization inhibitor.

Polymer: Poly(L-lactide)-co-ε-caprolactone (PLLA-co-CL)

Monomers: 3-Trimetoxysilyl propylmetacrylate (WITCO)/THF

(Aldrich) (90/10) and acrylamide (Promega)

Extent of grafting: About 20 %

Monomer distribution: 3-Trimetoxysilyl propylmetacrylate: through the matrix

Acrylamide: as a uniform coating

The extent of grafting was gravimetrically determined from the following equation:

, - m_n

E (%) = 100 ^■ — ₀

where m₀ and _; are the weight of the ungrafted and the grafted sample, respectively.

The monomer solutions were purged with nitrogen for at least 30 minutes before the grafting in order to minimize the presence of oxygen during the grafting process. The grafted films were washed with ion-exchanged water for several hours in order to remove homopolymer, and dried to constant weight in vacuum at room temperature. Tests

The dissolution rate of the silica gel was studied by soaking in an SBF (Simulated Body Fluid) solution. The release rate of heparin from the silica gel was also studied. The biological activity of the released heparin was further studied. The coated polymers were also subjected to cell growth and cytotoxicity tests. The coated polymers were additionally tested in vivo in rats.

Dissolution test

SBF was prepared by dissolving NaCl, NaHC0₃, KC1, K₂HP0_{4 *} 3H₂0, MgCl₂ x 6H₂0, CaCl₂, Na₂S0₄, TRIZMA^® HC1 and TRIZMA^® base as shown in Table 1. The fluid was adjusted at physiological pH 7.40 and temperature 37 °C. The composition of inorganic ions emulated that of human blood plasma.

Table 1. Reagents used for the SBF solution.

Reagent Amount 1 dm³ (H₂0) Manufacture/purity

NaCl 7.9951 Riedel-deHaen, pro analysis/ 99.8%

NaHC0₃ 0.3534 Merck, pro analysis/99.5 %

KC1 0.2243 Merck, pro analysis/99.5 %

K₂HP0₄ • 3H₂0 0.2281 Merck, pro analysis/99 %

MgCl₂ • 6H₂0 0.3053 Merck, pro analysis/99 %

CaCl₂ 0.2776 Merck, pro analysis/99 %

Na₂S0₄ 0.0709 Merck, pro analysis/99 %

TRIZMA HC1 * 6.2414 Sigma Ultra/ 99.9 %

TRIZMA base ** 1.2591 Sigma Ultra/99.9 %

* TRIZMA^® HC1: tris[hydroxymethyl]aminomethane hydrochloride (HOCH₂)₃CNH₂ • HC1 ** TRIZMA^® base: tris[hydroxymethyl]aminomethane (HOCH₂)₃CNH₂

20 respectively 25 mg of heparinized silica gel was immersed in 50 ml SBF in a polyethylene bottle covered with a tight lid. Two parallel samples and six different time points were used. All bottles were placed in a shaking water bath at 37 °C. All samples were filtered before the ion concentration (Si) analysis was carried out by spectroscopic method described by Boltz and Mellon 1947.

Toluidine blue method for heparin measurements

The total amount of heparin dissolved was measured by a colorimetric toluidine blue method (Smith et al). In the toluidine blue method toluidine blue reacts with heparin and forms a color complex. The complex molecule has its absorption maxima at a different position than the unreacted toluidine blue. By measuring the visible spectra of the sample the amount of both the complex molecule and unreacted toluidine blue can be measured (Smith et al). The same method can be used for both the immobilized heparin (Park et al.) and heparin in solution (Smith et al.) by making only small modifications. This method can be utilized as qualitative indication as well since it dyes the polymer sheet with specific color. The method is usable for all kind of colorless materials.

Reagents:

1. Heparin standard solution: 20 mg of heparin diluted into 100 ml of 0.2 % NaCl water solution. The standard dilutions should be between 5 and 40 μg of heparin in a sample (100 μl).

2. 0.005 % toluidine blue: 25 mg toluidine blue dissolved in 500 ml 0.01 N HC1 containing 0.2 % NaCl.

3. 0.2 % NaCl

4. hexane

Procedure:

1.25 ml of toluidine blue solution, 100 μl of sample or standard solution and 1 J5 ml of 0.2 % NaCl solution was pipetted into each 10 ml glass tube. The tubes were mixed vigorously by Vortex for 30 s. Next 2.5 ml of hexane was added to the tubes and they were shaken for another 30 s to separate the heparin-dye complex formed. The aqueous layers of the tubes were sampled and if necessary diluted with 0.2 % NaCl water solution or ethanol. The absorbance at 631 nm was measured within 30 min with Shimadzu UV-1601 Spectrophotometer.

Factor Xa Assay

The biological activity of released heparin against thrombin formation was evaluated by a HEPRN^® method using a aca^® discrete clinical analyzer. The HEPRN^® method is a chromogenic substrate assay based on the inhibition of bovine factor Xa (FXa) by heparin activated antithrombin III (ATIII). The uninhibited FXa catalyzes the hydrolysis of the chromogenic substrate, CH₃-0-CO-cyclohexcyl- glycyl-glycyl-arginyl-p-nitroanilide (CHGGA-pNA). The p-nitroaniline has absorption maxima at 405 nm. The reactions can be represented by the following equations:

^X) , ^FXa _^ + ^{ATπI keparin} > {FXa - ATIII) + ^FXa

(excess) ' (active)

FXa

2) CHGGA- pNA ^(ac,ive) > p - nitroaniline + CHGGA

The rate of increase in absorbance at 405 nm due to the appearance of the chromophore (p-nitroaniline), is linear and inversely related to the effective heparin activity (HEPRN^®). The activity measurements were carried out for heparin in solution not for the immobilized heparin. For the measurements some citrated plasma was added to the solution. The measurement was performed at Turku University Central Hospital (TYKS) by using commercial available reagent kits. Cell growth and cytotoxicity testing

The following materials were tested: copolymer as such, PLLA, CL, copolymer grafted in both ways and bulk heparinized silica gel. The materials were in vitro tested by culturing cells on materials and by measuring cytotoxicity of materials. The purpose of this was to evaluate if the grafting process and grafted materials are safe for living tissues, and also to choose the best candidates for in vivo testing. Human gingival fibroblasts (Hakkinen 1995) were routinely cultured in Dulbecco's Modification of Eagle's Medium (DMEM), including 10 % (v/v) Foetal Calf Serum (FCS, kibbutz Beit Haemek, Israel), 4,500 mg/1 glucose, 3.7 g/1 NaHC0₃ and penicillin-streptomycin solution (GibcoBRL, 10,000 U/ml and 10,000 μg/ml in saline) 1 ml/1. Cells were cultured on petridishes (0 10 cm) at +37 °C and 5 % C0₂ atmosphere. The medium was changed every other day and the cells were harvested at confluency. Only cells from nearly confluent dishes were used for experiments.

Materials used for cell culturing were washed with 20 % ethanol and sterile deionized water. Then they were attached to culture dishes with silicone grease. Amount of cells per sample material was about 1/6 of the confluent petridish. Medium was changed every other day and cell growth was followed and investigated with a microscope.

Cytotoxicity of materials was evaluated using a modification of the lactate dehydrogenase (LDH) method (Korzeniewski and Callewaert 1983). Materials were tested as extracts. Standard 6- well plates (Nunc) were used. One confluent petridish (0 10 cm) was used per test. Cells were first washed with 4 ml of EDTA solution [in phosphate buffer solution (PBS), pH 7.4] and then incubated 5 minutes in 4 ml of trypsin EDTA solution (40/0.4) at +37 °C. [Trypsin stock = 2.5 % (w/v) in normal saline, GibcoBRL]. Detached cells were transferred into a 15 ml centrifuge tube and centrifuged 5 minutes at 800 rpm. Finally cells were suspended in 5 ml of medium. From this suspension 200 μl per well was added. Cells were cultured in these plates using 2 ml medium per well. Cells were cultured as mentioned before, changing medium every other day until wells had reached confluency.

When all the wells had almost reached confluency, the test materials were extracted. Materials (ca. 0.5 cm ) were dipped into 20 % ethanol and rinsed with sterile deionized water. Then they were dipped into sterile eppendorf tubes and 1 ml of medium was added. These tubes were incubated 24 hours at +37 °C temperature. After that the wells of the test plates were emptied, dead cells removed with EDTA solution. Then 1 ml of fresh medium was added into each well. The extracts were added, and into wells for spontaneous release 1 ml of medium was added. For wells of maximal LDH release 200 μl of 10 % Triton X-100 and 1 ml of medium was added. This test system was incubated in cell culture conditions (as before mentioned) for 24 hours.

After 24 hours a sample of 500 μl from each well was taken. LDH measurement was carried out at room temperature (20 — 25 °C) with chemicals at the same temperature. Solution A was prepared by pipetting 432 μl of 30 mM sodium- pyruvate (Sigma) in Tris-Cl buffer (TRIZMA^®, Sigma), pH 7.4 and 432 μl of 6.6 mM NADH (FlukaChem) in the same buffer. The reagent was then diluted with 2550 μl of the same buffer. The ready made solution A was kept in dark between single measurements and vortexed carefully before use.

The samples were diluted by mixing 200 μl of the sample with 700 μl of Tris-buffer (same as above). They were vortexed and transferred into cuvettes. Measurements were carried out using a Shimadzu UV/Vis spectrophotometer at 340 nm wavelength. The reaction was started by adding 100 μl of solution A into the cuvette containing a sample, and the reaction was followed for 7 minutes. The slope was measured for the first 120 second period and used for indication of the amount of LDH released into the medium. Cytotoxicity (%D) was calculated by using the equation:

A = sample slope B = spontaneous release well slope and

C = Triton X-100 well slope.

Results

Both the dissolution rate of the silica gel and release rate of heparin from the silica gel was examined by using the bulk gel prepared by sol-gel technique. Up to 15 % (calculated from the theoretical dry weight) of heparin was successfully immobilized to the silica gel produced by an acid catalyzed hydrolysis reaction.

During 24 days 45 % of the heparin loaded was released (Figure 1). Each heparin concentration, i.e. 1, 2, 5, 10 and 15 % (calculated from the theoretical dry weight) used had similar releasing profile. Heparin concentrations were studied by the toluidine blue method. Heparin released retained its biological activity as an anticoagulant when examined by the HEPRN^® method.

According to information obtained from SEM studies, a uniform silica gel coating was obtained on a surface of the acrylamide grafted PLLA-co-CL sheet. The thickness of the coating produced was 0.3 μm and its cracking after bending was minimal. The gel layer on the surface of 3-trimetoxysilyl propylmetacrylate grafted polymer was thinner than the one observed on acrylamide grafted PLLA-co-CL sheet.

When human gingival fibroblasts were grown under cell culture conditions on coverslips without, and together with, small silica gel particles it was found that cell growth was not influenced by the presence of the gel. Cells divided and spread normally and finally covered the silica gel particles even though they were topographically elevated from the substratum surface. From the results of the in vitro test it seems obvious that silica gels do not have any toxic effects or other harmful influences on fibroblasts growing in contact with the material. The present results agree with previous results indicating that silica gels are biocompatible materials and promising dissolvable vehicles for delivery of biologically active molecules. The gels have no harmful effects on living tissues as can also be seen from the results of in vivo testing (Ahola et al, 1997 and Kortesuo et al, 1999).

The heparin coated polymers implanted in rats did not during the 2 weeks of in vivo testing induce any blood clotting demonstrating biological activity of the heparin in vivo.

Example 3

Attachment of heparin to polyethylene (PE) grafted with vinylpyridine (VP) and polyethylene (PE) grafted with vinylacetate (VAc).

This example demonstrates a novel method for tailoring the rate of heparin release by modifying the functionality of polyethylene-core-material by grafting.

X-ray photoelectron (XPS) spectra were obtained by using a Perkin Elmer 5400 electron spectroscopy for chemical analysis (ESCA) spectrometer employing Mg K X-rays. A survey scan spectrum and narrow scans in the Cls, Ols, Nls and S2p were recorded for the different materials. Atomic concentrations were calculated from the survey spectra.

The distribution of the inorganic elements in the functionalized polymer films was determined by using energy dispersive X-ray analysis (EDXA).

The immobilized heparin was qualitatively visualized by using the Laser Scanning Confocal Microscopy (LSCM). For the LSCM measurements the grafted and heparinized polymer sheets were dyed by acridine orange solution. The Ar⁺-laser line 488 nm was used as the excitation line. The resolution of the equipment setup was about 200 nm. The quantitative analysis of the heparin was carried out either by using ESCA or by the toluidine blue method which can be used both as quantitative and qualitative method.

Functionalization of the polymer

The functionalization of the different polymers in the examples was all performed according to the same method. Low density polyethylene (NCPE-7518) film samples (thickness 500 μm), received from Neste Oy Ltd. Finland, were cut into quadratic pieces, washed with ethanol and dried in vacuum. The pre-weighed samples were irradiated by using an Electrocurtain^® electron accelerator in air at an acceleration voltage of 175 kV. The irradiated films were removed from the accelerator and immersed into the different vinyl monomer solutions at ambient temperature for various reaction times without using any homopolymerization inhibitor. The monomers used, grafting times and extent of grafting in the different graft-co-polymerizations are listed in Table 2. All the monomers and solvents used were supplied by Sigma-Aldrich. Table 2. Monomers, solvents and grafting times used in the functionalization of the LDPE films.

Monomer / solvent Parts Grafting time Grafting extent

Vinylacetate 100 / 0 300 min 120 % Vinylpyridine 100 / 0 180 min 200 %

The monomer solutions were purged with nitrogen for at least 30 minutes before the grafting in order to mininiize the presence of oxygen during the grafting process. The grafted films were washed with ion-exchanged water for several hours in order to remove homopolymer, and dried to constant weight in vacuum at room temperature.

Hydrolysis of the acetate groups the poly[ethylene-grø t-(vinyl acetate)] was performed according to a previous used method where the acetate groups were hydrolyzed in a toluene/isopropanol solution (70/30) at pH 8 (with KOH) for two hours (Ekholm 1993).

Immobilization of heparin

The grafted polymer was allowed to react with heparin solution (0.025 mg heparin/ 5 ml pH 5) for 2 to 96 hours at 37 °C. After incubation the films were washed thoroughly by deionized water. The attachment of heparin was confirmed by XPS measurements, where the atomic concentration of sulfur and nitrogen was measured (Table 3), and visualized by using LSCM and toluidine blue method. Table 3. Atomic concentration of the elements for different modified materials.

Material C O N S

PE 97.8 2.1

PE-g-VAc (hydr.) 95.2 4.1

PE-g-VAc (hydr.)+ hep. 89.8 8.5 1.2 0.1

PE-g-VP 95.7 2.6 1.7

PE-g-VP + hep. 76.1 14.4 5.9 1.6

The materials listed in table 3 were subjected to cell growth and cytotoxicity testing described in example 2.

Cell growth and cytotoxicity testing

Cells grown in the presence of the extract from vinylacetate grafted PE in cell culture medium at 37° C for 24h showed no signs of cytotoxicity. The lactate dehydrogenase (LDH) activity remained on a level corresponding to that of spontaneous release of LDH from the fibroblast cells used for the study. Spontaneous release gave an activity between 5.0 and 6.5 while the release in the presence of the extract was 5.2 — 5.3.

Cells grown on vinylacetate grafted PE grew almost in an equal manner to those grown on plain PE. The mitotic activity of the cells attached was identical on each of the materials tested. However, the cells grown on vinylacetate grafted PE seemed to have stronger tendency to form clones than the cells grown on PE.

Cells grown in the presence of the extract from vinylpyridine grafted PE in cell culture medium at 37° C for 24 h showed no signs of cytotoxicity. The LDH-activity remained on a level corresponding to that of spontaneous release of LDH from the fibroblasts used for the study. The release in the presence of the extract was 5.0 — 5.7. Cells grown on vinylpyridine grafted PE grew in an equal manner to those grown on plain PE.

Heparin release

Heparin molecules attach to the functional group on the graft by hydrogen bonds. Heparin was cleaved therefrom during a dissolution test carried out at 37 °C by using SBF as solvent. According to the biological activity test, HEPRN method using discrete clinical analyzer carried out for samples after dissolution tests, the cleaved heparin remains its biological activity against thrombin formation.

The heparin is more strongly bound on the surface of the vinylpyridine grafted polymer than in the case of the vinylacetate grafted polymer. Neither the biological activity test nor the toluidine blue method, indicated heparin present on the solution taken out from the dissolution test after first week. ESCA, LSCM, and the toluidine blue dye showed that heparin was immobilized on the surface. While staining the polymer sheets directly with toluidine blue, clear difference was observed between the heparinized and non-heparinized polymer. The visible spectra of the toluidine blue solution after reaction with heparinized polymer sheets show difference between the samples as well.

Example 4

Attachment of silver onto grafted poly(ε-caprolactone)

Certain metal ions attached to surfaces of polymer materials may have beneficial effects also when judged from biological or therapeutic viewpoints. For example silver coatings can be used in order to reduce the risk that a material becomes infected (Coen et /J996). In this example a biodegradable polyester, poly(ε-caprolactone) was coated with silver ions. In order to demonstrate the extreme possibilities to regulate the extent of functionalization, the polymer was in one example functionalized only at the surface, leaving the bulk unmodified. In another example the polymer was functionalized throughout the whole film.

This example demonstrates how the amount and distribution of an ion (silver) can be controlled by varying the material's (poly-acrylate grafted poly(ε-caprolactone) preparation strategy. By using a short grafting time, the functionalization of the material by silver can be limited to the surface of the article while the bulk of the material remains unchanged (middle curve in figure 2). If a higher degree of functionalization through the whole film-matrix is desired it can be achieved by prolonging the reaction time of the grafting procedure (upper curve in the figure 2). The lowest curve in figure 2 is an non-grafted film.

Grafting of the polymer

Polymer: Poly(ε-caprolactone) (Union Carbide)

Monomer: Acrylic acid/water (30/70)

Method of grafting:

Macroradicals in the poly(ε-caprolactone) bulk were created by irradiating the pre-weighed film by making use of an Electrocurtain^® electron accelerator in air at an acceleration voltage of 175 kV. The irradiated films were removed from the accelerator and immersed into the acrylic acid solution at 25 °C for defined reaction times without using homopolymerization inhibitors.

Extent of grafting: 250 % (upper curve in Figure 2) and

50 % (middle curve in Figure 2) Attachment of silver on the grafted polymer

Silver ions (Sigma Aldrich) were attached to the grafted films in an ion-changed water solution. Silver nitrate was dissolved in water (5 % w/v) and films exhibiting a variable extent of poly-acrylate grafting were immersed into the solution and allowed to react under stirring for 24 hours. The silver distribution in the film was measured by using energy dispersive X-ray analysis (EDXA).

REFERENCES

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Claims

1. A biodegradable polymer material bearing a biologically active agent, said material comprising a core polymer, wherein said core polymer is in bulk form or formed into a body or a coating of said body, which body optionally has the shape of a finished device, wherein the biologically active agent is bound to a functional group to the core polymer, characterized in that,

a) said functional group has been generated by oxidation of the core polymer, or b) said functional group is attached to a vinyl terminated monomer which has been grafted onto the core polymer.

2. The polymer material according to claim 1, wherein the core polymer has been formed into a body or a coating of said body, which body optionally has the shape of a finished device, characterized in that the core polymer on a desired region of the device bears a functional group.

3. The material according to claim 1 or 2 characterized in that the core polymer is a thermoplastic polymer.

4. The material according to claim 3 , characterized in that the thermoplastic polymer is a synthetic biodegradable polymer.

5. The material according to claim 1 or 2 characterized in that the core polymer is a protein or parts thereof.

6. The material according to any of the claims 1 to 5, characterized in that the functional group is a carboxylic, hydroxy, ester, ether, epoxy, amino, amido, cyano, nitro, sulfono, mercapto, siloxy or phosphoro group.

7. The material according to any of the claims 2 to 6, characterized in that said body is in the shape of a spherulite, sheet, film, plate, stick, pin, screw, tube, fiber, hollow fiber, woven fabric or non-woven fabric also when built to resemble human or animal body parts such as ear, nose, joints etc. or parts thereof.

8. The material according to any of the claims 2 to 6, characterized in that said body is a device implantable into the mammalian body and that the biologically active agent is pharmaceutically acceptable.

9. The material according to claim 8, characterized in that said body is shaped to a stent, dental or orthopedic implant; implant for controlled drug delivery; bone fixation pin, fixation plate, regeneration matrix, or the like also when built to resemble human or animal body parts such as ear, nose, joints etc. or parts thereof.

10. The material according to any of the claims 1 to 9, characterized in that the biologically active agent is a polypeptide, a protein, a polysaccaride, an oligosaccaride, a mono- or disaccaride, an organic compound, an organometallic compound or an inorganic compound containing any element with an atomic number ranging from 3 to 84.

11. The material according to claim 10, characterized in that the biologically active agent is an inorganic ion or a polymer thereof, - silica gel as such or silica gel loaded with a therapeutical agent, heparin, a growth factor, a growth factor inhibitor, an integrin blocker (e.g. a Ila/IIIb inhibitor), - an oligonucleotide or a complete functional gene in sense or antisense orientation in a suitable expression vector or any other expression vector construct for local delivery of therapeutically active agents.

12. The material according to claim 9, characterized in that said body is shaped into a stent, the inner wall of which is provided with a biologically active agent, which is an inorganic ion or a polymer thereof, silica gel as such or silica gel loaded with a therapeutical agent, heparin, - a growth factor, a growth factor inhibitor, an integrin blocker (e.g. a Ila/IIIb inhibitor), an oligonucleotide or a complete functional gene in sense or antisense orientation in a suitable expression vector or any other expression vector construct which biologically active agent is released at a controlled rate in in vivo conditions.

13. A device made of or coated with a biodegradable polymer material, useful for finishing into a device of or coated with a polymer material bearing a biologically active agent, wherein said polymer material comprises a core polymer, which has been formed into a body or a coating of said body being the shape of a finished device, characterized in that the core polymer bears a functional group on a desired region of the body, said functional group being capable of binding the biologically active agent, wherein a) said functional group has been generated by oxidation of the core polymer, or b) said functional group is attached to a vinyl terminated monomer which has been grafted onto the core polymer.

14. A method for the preparation of a biodegradable polymer material bearing a biologically active agent, said material comprising a solid core polymer, wherein the biologically active agent is bound to the core polymer, characterized by the steps of

a) generating macroradicals on the core polymer, and

b) attaching said biologically active agent to said functional group.

15. The method according to claim 14, characterized in that the core polymer material is in the bulk form such as a powder or a granulate.

16. The method according to claim 14 wherein the material has been formed into a body or coating of said body, which body optionally has the shape of a finished device, characterized in that the macroradicals are generated on the core polymer on a desired region of the body.

17. The method according to any of the claims 14 to 16 characterized in that the macroradicals are generated by exposing the core polymer to a peroxide or by ionizing radiation thereof, e.g. by accelerated electrons.