The subject of the invention is a method of forming a multilayer coating with a controlled drug release function intended for the functionalization of implant materials.

A dynamic growth in demand for implant materials is being observed both in Poland and globally. This is associated with a number of factors: the aging of the population, the desire to maintain quality of life, the increased number of traffic accidents, and civilization progress. A significant increase in demand is particularly seen with respect to metal implants, which effectively restore the role of damaged bones and allow for the proper functioning of patients in everyday life. Available data indicate that in Poland about 25,000 hip joint replacement surgeries, 13,000 spine surgeries involving the use of implants, and 13,000 knee repair surgeries using fixation implants are performed each year.

The most popular alloys used in orthopaedics are stainless steel, TiAlV and NiTi, which contain heavy metals, even toxic ones. Due to the conditions inside the body promoting corrosion (36.6°C, physiological fluids of a composition similar to sea water and therefore characterised by high ionic strength) (Injury 27 (1996) S/C16), the surface of implants undergoes slow destruction, and heavy metal ions migrate to the body at levels up to 10 μL/rrlL/week (Corros. Sci 51 (2009) 1 157). This concentration may be harmful to patients, leading to a number of diseases, including cancer (Mater. Sci. Eng. C, 24 (2004) 745). The negative effect of released metal ions on the human body continues to be the subject of many studies.

Implantation procedures are complicated and associated with the risk of implant rejection by the host, due to a number of complex processes on the implant-tissue interface (J Pharm Sci, 97 (2008) 2892). The most common post-operative complications include prolonged inflammation persisting for about 3 weeks, and infection (Biomol. Eng, 19 (2002) 21 1). In such cases, patients receive a variety of oral and intramuscular anti-inflammatory and anti-infective drugs. For this reason solutions are considered that rely on controlled local drug release from the surface of the inserted implant, bringing a number of benefits to patients. The most important benefits include the application of lower doses of the drug, and their activity limited to the target tissue, which reduces the risk of side effects associated with the oral administration of high doses of medication. Scientific literature and patent applications present many attempts to solve the highlighted problem by using physical and chemical methods, such as coating the surface of implants with natural (Biomacromolecules 1 1 (2010) 1254) or synthetic polymers (J. Biomed. Mater. Res., 93B (2010) 266) to increase cell adherence, electropolishing to reduce the risk of biofilm formation on the surface (Surf. Coat. Technol., 206 (2012) 3165), or covalent immobilisation of drug-loaded liposomes (Colloids Surf ., B, 84 (2011) 214). One of the most promising techniques is coating metal implants with one or more layers of polymers whose role is to protect the surface on which they are deposited. Suitable polymers should be easy to deposit, have strong adherence to the implant's surface, be chemically passive in the environment of physiological fluids, biocompatible, and have suitable mechanical strength.

One of the polymers meeting these criteria is parylene C (poly(chloro-para-xylene)), whose many medical applications are disclosed in patent documents. Coatings made of this material are used, e.g. on stents to limit the formation of blood clots on their surface, as disclosed in US patent no. 6776792. US patent no. 2013001 1456A1 discloses the use of parylene C as a carrier for an antibacterial agent on the surface of a cochlear implant, an ocular implant, and a pacemaker. Moreover, parylene C is used in multilayer polymer coatings intended for drug release. According to the description in US patent no. 20130004560 Al, parylene C is a solid foundation for such a coating, to which in the next step a mixture of drug and copolymer is attached, and another layer of suitably prepared parylene is deposited. A polymer layer prepared in such a way successfully prolongs drug release time inside the body.

Moreover, biodegradable polymers are used for drug encapsulation. As disclosed in US patent no. 8003125, controlled drug release can be successfully achieved using copolymers of PEG (polyethylene glycol) and another polymer. According to the description in US patent no. 201 10070320A1, a copolymer composed of PEG and a lactic acid copolymer can create an efficient barrier preventing rapid drug release to the body. Moreover, by controlling its composition, the degradation time of the copolymer and drug release kinetics can be adjusted. Copolymers of L-lactide and glycolide (PLGA) are biocompatible and biodegradable. Because their biodegradation time can be controlled by the use of various molar ratios of individual comonomers, they can be successfully used as carriers in controlled drug release systems. PGLA loaded with an immunosuppressive drug is used as a biodegradable polymer layer on a non-biodegradable polyamide coating protecting biocorrosive implants, as disclosed in patent EP 2433660 Al . US patent no. 20120303057 Al discloses a method of manufacturing an ibuprofen-loaded PGLA film in order to coat resorbable sutures made of the same resorbable material, i.e. PGLA. Deposition of a drug- loaded PGLA film on both degradable and non-degradable materials ensures stable drug release thanks to the gradual process of hydrolytic degradation.

The major problem related to drug release from the implant surface is the engineering of its surface. The amount and speed of drug release is associated with the composition of the implant surface, its morphology, the adherence power of cells and susceptibility to the formation of bacterial biofilm. Therefore, there is an ongoing need to improve the surface of implants in order to achieve drug release at the appropriate therapeutic level and within a sufficiently long time. For layers made of parylene, chemical passivity determining parylene application as a protective layer is problematic at the same time, because bioactive substances and a biodegradable polymer cannot be bound directly to it. Parylene C is a crystalline polymer and has a non-porous structure, with poor permeability to small molecules.

To solve the above-described problem, for the present invention we developed a treatment method for parylene coating by inserting anchoring sites for drug molecules, or for a drug-loaded biodegradable polymer. It has to be emphasized that the treatment itself has to be a closely controlled process, so that the achieved porosity and inserted polar groups do not destroy the protective function of parylene. This leads to the conclusion that the several- micrometer-thick layer necessary to ensure anticorrosive protection may be modified only within the thickness of a few dozen nanometers. Over-modification of the polymer structure is reflected by the decrease in the impedance of the polymer coating (measured in electrochemical tests) and change in its glass transition temperature (TG DTA).

The method of forming a multilayer protective coating for implant materials with a controlled drug release function according to the invention is characterised in that a 6 to 20 μηι-thick parylene layer is applied onto the surface of an implant by chemical vapour deposition, and then treated with oxygen plasma under 0.2 to 1 mbar pressure for 15 to 60 minutes, using a plasma generator of 10 to 60 W power. Onto such a modified parylene coating a mixture of lactide and glycolide solutions is deposited in the molar ratio of lactide to glycolide from 1 : 1 to 100: 1, together with the drug substance in the amount of 5% to 10% w/w of the polymer material and a polymerisation initiator, and then the polymerisation is performed, and the obtained layer is dried.

Preferably, poly(monochloro-p-xylylene) (parylene C) is used for the formation of the layer.

Preferably, the formation of the polymer layer is carried out in a process wherein the parylene coating is prepared by chemical vapour deposition. Depending on the type of parylene, a starting substance (dimer) is used: [2,2]-paracyclophane (dimer of parylene N), dichloro-[2,2]-paracyclophane (dimer of parylene C). At 150°C the dimer is vaporised from the solid phase to the gaseous phase. Next, the vapour is heated up to 650°C. At this temperature the dimer is degraded to monomer molecules. The monomer molecules are passed into the next chamber, where the substrate is exposed for deposition. In the coating chamber, at room temperature and under low vacuum of 10 ^"3 mbar, monomer molecules deposit spontaneously on the surface of the implant and the polymerisation occurs. The thickness of the deposited coating is controlled throughout the deposition time.

Preferably, the parylene coating before oxygen plasma treatment is cleaned using organic solvents such as isopropanol, ethanol or acetone, preferably ethanol.

Preferably, oxygen plasma treatment is performed in a chamber with a quartz vessel fitted in a rotor, in order to ensure uniform plasma distribution and to optimise the concentration of excited oxygen forms.

Preferably, oxygen plasma treatment is carried out at 0.2 mbar pressure for 60 minutes, using a plasma generator of 50 W power.

The mixture of lactide and glycolide solutions with the drug substance and the initiator is applied onto the treated layer of parylene by immersion or spraying. Copolymers loaded with the drug substance are bound with the substrate and form a biodegradable layer on the surface of parylene.

Preferably, the molar ratio of lactide to glycolide is from 7: 1 to 4: 1, most preferably

5 : 1.

Preferably, Zr(acac) ₄ is used as an initiator.

The drug substances used are anti-inflammatory, anti-infective, analgesic and antithrombotic agents, preferably ibuprofen.

Preferably, an 8 to 10 μπι-thick parylene layer and an 0.12 to 0.19 mm-thick copolymer layer are applied onto the surface of the implant.

Preferably, the coating according to the invention, i.e. a parylene layer with a copolymer layer, is 0.14 to 0.39 mm thick. The thickness of the therapeutic layer is a compromise between the need to store an adjusted amount of the drug substance and to minimize thickness - an excessively thick layer leaves a void after degradation, which may result in loosening of the implant.

The copolymer matrix loaded with the drug substance is obtained by the method wherein copolymers are dissolved in organic solvent, preferably methylene chloride, and the drug substance in a relevant solvent (chloroform for ibuprofen), and then both solutions are brought together under vacuum and adjusted to an appropriate density.

A coating manufactured using the method according to the invention comprises two parylene layers, wherein the first layer (near the surface of an implant) is a solid, non-porous and hydrophobic parylene, and the second is a non-porous hydrophilic parylene modified by the method according to the invention. The layer of solid parylene protects the metallic surface against corrosion and increases its biocompatibility. In turn, oxygen plasma treatment of the parylene surface leads to the formation of micro- and nanopores and insertion of functional groups containing oxygen, which are the adsorption sites facilitating binding with a degradable copolymer responsible for the function of drug release. It should be emphasized that the declared effect can be achieved only for the surface of parylene treated with oxygen plasma, as this is the only type of plasma ensuring the insertion of functional groups containing oxygen.

It is very important that the hydrophilic parylene layer is formed as a result of modification of the first layer of hydrophobic parylene deposited on the implant. The plasma treatment of the first parylene layer, under conditions specified according to the invention, is limited only to the supraficial part of this layer, while the part in direct contact with the implant remains untreated. With such a modification, the oxygen adsorption sites and pores are distributed only on the surface and serve for connecting of the copolymer layer, while the deeper part of the layer maintains its initial hydrophobic properties. Oxygen functional groups generated in the process of oxygen plasma treatment act as binding sites between the parylene C layer and the copolymer layer. Attempts to deposit the drug substance alone onto the treated parylene in the first step and the copolymer layer in the second step did not provide a satisfactory effect. Oxygen functional groups were saturated with the loaded drug substance and the copolymer layer did not have appropriate adherence. Moreover, the deposition of the copolymer layer on the drug layer does not allow the achievement of the effect of controlled drug release.

The specific modification of the parylene coating, resulting in the formation of nanopores and functional groups containing oxygen, not only increases cell adherence to the parylene surface, thus facilitating faster wound healing, but also does not promote significantly bacterial growth. There is a very narrow range of optimum conditions under which the coating can be modified in a desired manner, thus the choice of plasma treatment is the key factor for their generation. It has to be emphasized that the properties of the parylene surface treated with oxygen plasma and exposed to ambient air remain stable only for a specific time (2-3 days), and therefore the second layer should preferably be deposited within this time period.

The therapeutic layer combined with the polymer layer ensures controlled drug release for 3-7 weeks. The coating manufactured by the method according to the invention forms an integrated anticorrosive and therapeutic system.

The therapeutic layer combined with the polymer layer can be deposited onto orthopaedic implants for both short-term (screws, nails) and long-term use (hip joint or knee joint prostheses) made of various alloys, e.g. SS 316L (stainless steel) or titanium alloys (e.g. T.6A1V).

The subject of the invention is presented in more detail in examples of manufacture. Example 1

Samples of parylene C, 8 μπι-thick, were manufactured by chemical vapour deposition, a method known in the art. At 150°C the dimer of parylene C (dichloro-[2,2]-paracyclophane) was vaporised from the gaseous phase, and then the vapour was heated to 650°C. At this temperature the dimer degraded to monomer molecules. The monomer gas was passed into the next chamber, where the surface for deposition was exposed; then at room temperature and under low vacuum a spontaneous deposition of monomer molecules and the polymerisation occurred. The thickness of the deposited coating was controlled throughout the deposition time. We obtained an 8 μιτι-thick parylene layer. Samples prepared in such a manner were cleaned with ethanol, and then treated with oxygen plasma inside a specially adapted chamber with a quartz vessel fitted in a rotor in order to ensure uniform plasma distribution and optimise the concentration of excited oxygen forms. Process conditions were as follows: pressure 0.2 mbar, time: 60 min, plasma generator power: 50 W. Alternatively, a similar treatment effect can be obtained using the following parameters: oxygen pressure: 0.2 mbar, time: 5 min, generator power: 50 W, without the use of a rotor and placing samples in a glass container.

Example 2.

We carried out biological tests on the treated Parylene C. For that purpose, we used untreated Parylene C and Parylene C treated according to the parameters specified in Example 1. We carried out an adherence test on MG-63 cells after 24 h. On the surface of untreated Parylene C the cell count was 80/mm ² , while for the treated surface the count was 120/mm ² . Moreover, we tested the viability of MG-63 cells using a standard MTT assay. The assay demonstrated that the viability of cells on the surface of untreated parylene C was 79% in comparison to the control (cells cultured directly in the plate), and 114% for the treated parylene C. The stronger adherence of cells and their greater viability on the treated surface of parylene C indicates that the material is non-toxic - it does not cause cell death, but instead promotes cell growth. The presented results are significant from the point of application, as they prove that the tested material under in vivo conditions can accelerate the integration of tissue at the bone-implant interface.

Next, we carried out microbial adherence tests for three selected strains. S. aureus, S. epidermidis and P. aeruginosa. Treated and untreated parylene films were incubated for lh and 4h in cultures of S. aureus, S. epidermidis and P. aeruginosa, the most common strains responsible for implant-associated infections. Adherence of & aureus to treated samples of parylene C was lxlO ⁶ , while for S. epidermidis and P. aeruginosa it was lxl0 ⁷ and did not differ significantly from the adherence to untreated parylene C (reference sample). The carried out experiment did not demonstrate any statistically significant differences between the treated and untreated samples. This result is particularly important in view of the application, as it indicates that modification does not promote the growth of strains responsible for infections.

Example 3

We prepared copolymer matrices comprising lactide and glycolide (85: 15) obtained in the process of ring-opening polymerisation using non-toxic Zr(acac) ₄ as an initiator, and loaded with ibuprofen as a model drug, in a dose 5% and 10% w/w of copolymer according to the following procedure.

Copolymers were dissolved in methylene chloride, and the drug in chloroform. Next, the two solutions were brought together and mixed. The mixture without air bubbles was poured onto the surface of parylene C treated according to Example 1. Gas bubbles released during mixing were removed under vacuum conditions.

The prepared matrices on parylene C treated according to Example 1, of 0.12 to 0.19 mm thickness, were air-dried for 7 days, and then dried for another week in a vacuum dryer in order to completely remove the solvent. Next, the samples were exposed to hydrolytic degradation in phosphate buffer, pH=7.4, at 37°C.

Simultaneously, we analysed changes occurring in the microstructure of copolymer chains during hydrolytic degradation using NMR, drug release from copolymer matrices using HPLC, and changes in molecular masses using GPC. As the degradation of copolymer matrices progressed, a significant release of ibuprofen was observed in up to 7 weeks of the experiment. During that period up to 13% of the drug substance was released. Simultaneously, we did not observe any rapid drug release from the copolymer matrix (Fig. 1 - graphs of cumulative ibuprofen release).

The thickness of coatings was measured using Scanning Electron Microscopy, for treated parylene it was about 8 μ ^η ι, and for the copolymer layer it was 0.12-0.19 mm.

Example 4

Untreated parylene C, parylene C treated according to Example 1, and parylene C overexposed to oxygen plasma (pressure: 0.2 mbar, time: 120 min, power: 50 W) was tested by electrochemical impedance spectroscopy (EIS). The tested sample (working electrode) was mounted into an electrolytic cell filled with 30 mL of electrolyte (synthetic physiological fluid). Additionally, a counter electrode (platinum grid) was mounted in the cell, and a saturated calomel electrode Ag/AgCl//KCl (SCE) was used as a reference electrode. All three electrodes were connected to a potentiostat, integrated with the analyser and PC. Measurements were taken at room temperature (25°C). Samples for measurements were rinsed with isopropanol and acetone, and to ensure electrical contact the polymer was removed from one side of the sample with sandpaper. The exposed surface of a tested sample had a size of 1 cm ² . EIS measurements were taken for 30 minutes at the frequency range from 0.01 Hz to 10 kHz, 50 mV amplitude alternating current, with data register every 10 points per decade.

The analysis of results from EIS enabled us to assess the effect of parylene C treatment on the impedance of the tested systems. The impedance for the untreated sample and optimally treated sample was 1 >< 10 ⁹ Ω/cm ² , while the impedance for the overmodified sample reduced to the level of l x lO ³ Ω/cm ² . These results clearly indicate that overmodification has a negative effect on the protective (anticorrosive) properties of the parylene C coating.

Example 5.

Samples prepared using CVD, as in Example 1 and cleaned with ethanol, were treated with oxygen plasma under the following process parameters: pressure 0.2 mbar, time: 10 min, plasma generator power: 50 W. Next, a layer of drug-loaded copolymer was deposited on the samples, according to the description presented in Example 3. Because of insufficiently long treatment time the concentration of surface functional groups was too low, which resulted in unsatisfactory adherence between the layers of parylene C and drug-loaded copolymer. We observed a delamination of the drug-loaded copolymer from the treated parylene C layer, and the formation of numerous gas bubbles at the interface of the layers, which burst, uncovering the layer of parylene C.

Example 6. On the layer of parylene C coated using CVD and untreated with oxygen plasma, we deposited a solution of ibuprofen in ethanol, concentration 40 mg/mL, and after vaporisation the mass of the drug deposited on the surface was 1 mg. Next, we deposited a layer of copolymer comprising L-lactide and glycolide (85.15) obtained in ring-opening polymerisation using nontoxic Zr(acac) ₄ as an initiator. We found that the layer of copolymer separated from the layer of the drug-loaded parylene.

Example 7.

Samples prepared using CVD, as in Example 1 and cleaned with ethanol, were treated with argon plasma under the following process parameters: pressure 0.2 mbar, time: 60 min, plasma generator power: 50 W (conditions identical to those for oxygen plasma treatment). After modification, we did not identify any new functional groups comprising oxygen, or observe increased hydrophilicity of the parylene C surface.