WO2015027274A1 - Coating for an implantable biomaterial, implantable biomaterial and method of making the coating and biomaterial - Google Patents

Abstract

The invention provides a coating for a biomaterial, the coating comprising monomeric units comprising a terpene or terpene analogue or derivative polymerised by plasma polymerisation. The invention also provides a method of coating a biomaterial in which monomeric units comprising a terpene or terpene analogue or derivative are polymerised with plasma polymerisation to form a polymer and the polymer is deposited as a coating on a biomaterial substrate. Also provided is a biomaterial comprising a substrate and the polymeric coating along with a method of making a biomaterial comprising polymerising terpene or terpene analogue or derivative monomeric units and depositing the polymer as a coating on a. biomaterial substrate. The invention also provides a method of reducing the degradation rate of a biomaterial.

Description

TITLE

COATING FOR AN IMPLANTABLE BIOMATERIAL, IMPLANTABLE

BIOMATERIAL AND METHOD OF MAKING THE COATING AND

BIOMATERIAL FIELD OF THE INVENTION

[001 ] The present invention relates to a coating for a biomaterial and a biomaterial suitable for implantation. The coating and implant may he used i tissue regeneration applications, in particular, the inventio is directed to a coating for an implantable biomaterial and an implantable biomateriai comprising a polymeric coating comprising tnonomeric units from a terpene or terpene analogue or derivative polymerised by plasma polymerisation.

BACKGROUND TO THE INVENTION

[002] Metallic biomaterials have found a plethora of applications as medical devices, owing to a favourable combination of mechanical properties, particularly fracture toughness and fatigue strength, biological inertness, and corrosion resistance in vivo.¹ Materials such as, 316L stainless steel titanium and its alloys, and cobalt based alloys are well suited to long-term implantation under load-bearing conditions, especially where regeneration of host tissues is not expected.

[003] Indeed, minimal degradation and mechanical stability over extended period of time are crucial for success of such procedures as total joint arthroplasty and restorative dentistry. These metals are also commonl employed for fabrication of implants designed to provide a temporary platform to support tissue regeneration and function restoration. Metallic screws, plates, and other fixatio devices are utilised t support fractured bone and thus ensure well -aligned tissue re-growth; stents are introduced into tubular vessels to address abnormal narrowing by restoring flow and providing physical support for the vessel through the healing process.

[004] Once this has been accomplished, the implant is no longer required and is either surgically removed or left to reside within the host permanently. Both of these scenarios present certain challenges. An invasive removal procedure is detrimental to the patient's recovery, has been associated with prolonged hospital stays and increased morbidity, and is of additional cost burden to the healthcare system/ Equally, the existence of a permanent metallic prosthesis presents significant disadvantages to the host. Even though most of the metallic devices are fabricated from inert materials, the extra-cell ular environment is a chemically aggressive space. Over time, both titanium- and chrom i um -based biomaterials have been shown to corrode, leaching ions and particl es into the peri-implant space. ^"These are potentially toxi and exhibit chronic inflammations in. some patients. In addition, the surfaces of implants may serve as a colonization ground for bacteria. Moreover, even most advanced antiproliferative drug - eluting stents have been associated with an irregular endothelialisation, requiring prolonged double antiplatelet tlierapy to reduce the risk of late and very late stent thrombosis."

SUMMARY OF THE INVENTION

[005] The present invention is broadly directed to a. coating for a biomaterial and a biomaterial suitable tor implantation. In one embodiment the coatin for a biomaterial and biomaterial comprises a polymeric coating comprising monomeric units comprising a terpene or terpene analogue or derivative polymerised by plasma polymerisation. The biomaterial and coating of the present invention may be resorbable.

[006] The present invention is also broadly directed to a method of making a coating for a biomaterial and a method of making a biomaterial comprising polymerising monomeric units comprising a terpene or terpene analogue or derivative with plasma polymerisation. These methods may be for making a coating for a resorbable biomaterial and a method of making a resorbable biomaterial.

[007] A preferred advantage of the present invention is that. it. provides coatings and biomaterials which are c tQcompatible and/or antibacterial or antimicrobial. Another preferred advantage of the present invention is that it provides coatings and biomaterials that are resorbable and/or degradabie.

[008] In one aspect, although not necessarily the broadest aspect, the present invention provides a coating for a biomaterial, the coating comprising monomeric units comprising a terpene or terpene analogue or derivative polymerised by plasma polymerisation.

[009] In one embodiment the invention provides the coating of the first aspect when used or for use as a coating for a biomaterial substrate.

[010] The biomaterial substrate may comprise the entire biomaterial or a part of the biomaterial.

[Oi l ] In a preferred embodiment of the first aspect, the biomaterial substrate is a resorbable biomaterial substrate. [012] The resorbable biomateriai substrate of the first aspect may comprise or consist of a resorbable metal. The resorbable metal may comprise one or more of a magnesium; a niaganesrara alloy; an iron; an iron alloy; a zinc; and a zinc alloy.

[013] In a suitable embodiment, the resorbable metal is magnesium or a magnesium alloy.

[014] In another suitable embodiment of the first aspect, the monomelic units are polymerised and deposited by plasma-enhanced chemical vapour deposition.

[015] hi a second aspect, the invention provides a method of coating a biomateriai, the method comprising polymerising monomelic units comprising a terpene or terpene analogue or derivative with plasma polymerisation to form a polymer and depositing the polymer as a coating on a biomateriai substrate to thereby coat the biomateriai

[016] In a preferred embodiment of the second aspect, the biomateriai substrate comprises a resorbable biomateriai

[017] The biomateriai substrate of the second aspect may comprise the entire biomateriai or a part of the biomateriai.

[018] According to the second aspect, the biomateriai substrate may comprise or consist of a resorbable metal. The resorbable metal may comprise one or more of a magnesium; a maganesium alloy; an iron; an iron alloy; a zinc; and a zinc alloy.

[019] In a suitable embodiment of the second aspect, the resorbable metal comprises magnesium or a magnesium alloy.

[020] In another suitable embodiment of the second aspect, the polymerising and depositing are accomplished by plasma-enhanced chemical vapour deposition.

[02 Ϊ ] In one embodiment of the second aspect, the method further comprises cleaning the biomateriai substrate to remove surface contaminants.

[022] The cleaning may comprise etching.

[023] The etching may comprise plasma etching,

[024] In a third aspect, the invention provides a biomateriai coated by the method of the second aspect.

[025] In a fourth aspect, the invention provides a biomateriai comprising a substrate and a polymeric coating, the polymeric coating comprising monomeric units comprising a terpene or terpene analogue or derivative polymerised by plasma polymerisation and deposited on the substrate.

[026] In a preferred embodiment of the fourth aspect, the biomateriai substrate comprises a resorbable biomaterial substrate.

[027] According to the fourth aspect, the biomaterial substrate may comprise the entire biomaterial or a part of the biomaterial.

[028] The biomaterial substrate of the fourth aspect, may comprise or consist of a resorbable metal. The resorbable metal may comprise one or more of a magnesium; a maganesium alloy; an iron; an iron alloy; a zinc; and a zinc alloy.

[029] In a suitable embodiment of the fourth aspect, the resorbable metal may comprise magnesium or a magnesium alloy.

[030] In another suitable embodiment of the fourth aspect, the raonomeric units are polymerised and deposited by plasma-enhanced chemical vapour deposition.

[031] in a fifth aspect, the invention provides a method of making a biomaterial comprising polymerising monomelic units comprising a terpene or terpene analogue or derivative with a plasma polymerisation to form a polymer and depositing the polymer as a coating on a biomaterial substrate to coat the biomaterial and thereby make the biomaterial.

[032] In a preferred embodiment of the fifth aspect, the biomaterial substrate comprises a resorbable biomaterial substrate.

[033] The biomaterial substrate of the fifth aspect may comprise the entire biomaterial or a part of the biomaterial

[034] According to the fifth aspect, the biomaterial substrate may comprise or consist of a resorbable metal. The resorbable metal ma comprise one or more of a magnesium; a maganesium alloy; an iron; an iron alloy; a zinc; and a zinc alloy.

[035] In a suitable embodiment of the fifth aspect, the resorbable metal comprises magnesium or a magnesium alloy.

[036] In another suitable embodiment of the fifth aspect, the polymerising and depositing are accomplished by plasma-enhanced chemical vapour deposition.

[037] I one embodiment of the fifth aspect, the method may further comprise cleaning the biomaterial substrate to remove surface contaminants,

[038] The cleaning may comprise etching.

[039] The etching may comprise plasma etching,

[040] In a sixth aspect, the invention provides a biomaterial made by the method of the fifth aspect.

[04.1 ] In a seventh aspect, the invention provides a method of reducing the degradation rate of a biornaterial, the method comprising polymerising monomeric units comprising a terpene or terpene analogue or derivative with plasma polymerisation to form a polymer and depositing the polymer as a coating on a biornaterial substrate to coat the biornaterial and thereby reduce the degradation rate of the biornaterial .

[042] In a preferred embodiment of the seventh aspect, the biornaterial substrate comprises a resorbable biornaterial.

[043] According to the seventh aspect the biornaterial substrate may comprise the entire biornaterial or a part of the biornaterial.

[044] The biornaterial substrate of the seventh aspect may comprise or consist of a resorbable metal. The resorbable metal may comprise one or more of a magnesium; a maganesium alloy; an iron;_, an iron alloy; a zinc; and a zinc alloy.

[045] In a suitable embodiment of the seventh aspect, the resorbable metal comprises magnesium or a magnesium alloy.

[046] In a suitable embodiment of the seventh aspect, the polymerising and depositing are accomplished by plasma-enhanced chemical vap ur deposition.

[047] In an eighth aspect, the invention provides a biornaterial comprising a reduced degradation rate made according to the method of the seventh aspect.

[048] In one embodiment of any above aspect, the biornaterial comprises a permanent biomaierial. The peiinanent biornaterial may comprise one or more of a non-resorbable metal: a non-resorbable polymeric material; and a non-resorbable silk. The non- resorbable metal may comprise one or more of nitinol (nickel titanium) and titanium. The non-resorbable polymeric material may comprise one or more of polypropylene; polyester; polyethylene; a polyamkie (Nylon); polytetrailuoroethylene; polyetherefherketone; polyethetketoneketone; silk; stainless steel; and a platinum iridium alloy.

[049] In another embodiment of any above aspect, the resorbable biornaterial substrate comprises one or more of a resorbable silk; a resorbable polymer; and a resorbable calcium phosphate ceramic.

[050] In another embodiment of any above aspect, the biornaterial may comprise a processed metal. The processed metal may comprise equal channel angular pressing.

[051] In a preferred embodiment of any one of the above aspects a source of the terpene or terpene analogue or derivati ve comprises an essential oil or a constituent of an essential oil. [052] In one embodiment, the essential oil may comprise tea tree oil, lavender oil, eucalyptus oil, d-Iemonene, and/or a-pinene.

[053] According to any of the above aspects, the terpene or terpene analogue o derivative may comprise an acyclic terpene or acyclic terpene analogue or derivative or a cyclic terpene or cyclic terpene analogue or derivative.

[054] The terpene or terpene analogue or derivative may comprise a -monoterpene or monoterpene analogue or derivative.

[055] The 'monoterpene or monoterpene analogue or derivative may comprise an acyclic monoterpene or an acyclic monoterpene analogue or derivative or a cyclic monoterpene or cyclic monoterpene analogue or derivative.

[056] The acyclic monoterpene or acyclic monoterpene analogue or derivative may comprise geraniol, ocimene, a myrcene, citral, citronellal, citronellol and/or halomon.

[057] The cyclic monoterpene or cyclic monoterpene analogue or derivative may comprise a monocyclic monoterpene or monocyclic monoterpene or analogue or derivative or a bicyclic monoterpene or tricyclic monoterpene analogue or derivative.

[058] The monocytic monoterpene or monocytic monoterpene analogue or derivative may comprise a terpinene, phellandrene, terpinolene and/or p-cymene.

[059] The bicyclic monoterpene or bicyclic monoterpene analogue or derivative may comprise one or more of pinene, carene, sabinene, camphene, tiii ene, camphor, borneol and/or eucalyptol. Tire pinene may comprise one or both of a- pinene and β-pinene,

[060] The monoterpene analogue or derivative may comprise a monoterpene alcohol or a monoterpene alcohol derivative.

[061] The monoterpene alcohol may comprise one or more of terpinene-4-ol, Hnatoot and a terpinene.

[062] The terpinene may comprise one or more of ct-terpinene, β-terpinene, γ-terpinene or δ-teqrinene. In preferred embodiment the terpineae^' or terpinene derivative comprises a-terpinene.

[063] The monoterpene alcohol analogue or derivative may comprise or consist of linalyl acetate.

[064] The cyclic terpene or cyclic terpene analogue or derivative may comprise or consist of a timonene or a timonene derivative. In one embodiment the limonene comprises D-limonene.

[065] According to an of the above aspects the coating may comprise an antibacterial or antimicrobial, coating.

[066] According to any of the above aspects the biomaterial substrate may be polished and or cleaned before the coating is deposited.

[067] The polishing may comprise mechanical polishing,

[068] The cleaning may comprise ultrasonic cleaning.

[069] According to any of the above aspects the plasma-enhanced chemical vapour deposition may comprise radio frequency plasma-enhanced chemical vapour deposition, [070] In a preferred embodiment the radio frequency plasma-enhanced chemical vapour deposition comprises a radio frequency between 13 - 14 MHz; between 13,4 - 13,7; or between 13.5 - 13.6. to a preferred embodiment the radio frequency is 13.56 MHz. The radio frequency may be 13.0, 13.1 , 13.2, 13.3, 13.4, 13.5, 13.51 , 13.52, 13.53, 13.54, 13.55, 13.56, 13.57, 1.3.58, 13.59, 1.3.6, 13.7, 13.8., 13.9 or 34.

[071] Whe radio frequency plasma-enhanced chemical vapour deposition is used an input power may be between 5 and 100 RF power, in preferred embodiments, the input power is between 5 and 50 RF or between 7 and 25 RF power. In. a particularly preferred, embodiment the input power is between 7 and 25 RF. The input power may be 5, 6, 7, 8, 9, 10, 11, .12, 13, 14, 15, 16, 17, 1.8, 19, 20, 21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 32, 33, 34, 35, 36, 37, 38, 39, 40, 41 , 42, 43, 44, 45, 46, 47, 48, 49 or 50 RF.

[072] In a suitable embodiment the input power may be controlled or varied to increase or decrease the amount of cross-linking in the polymer.

[073] In another preferred embodiment the RF power is sufficient to provide in situ sterilisation and/or functionalise or increase a mechanical property of a substrate surface.

[074] In another embodiment of any above aspect, the hydropMlicity and/or hydrophobics ty of the coating may be varied. The hydrophobicity may be increased by addition of a carrier gas. The carrier gas may comprise nitrogen. The hydrophilicity may be increased by controlling carrier gas amount and/or controlling surface roughness. The carrier gas may comprise oxygen.

[075] Also when radio frequency plasma-enhanced chemical vapour depositio is used polymerisation may be performed at a pressure between 25 - 400 raTorr; between 30 - 100 mTorr; or between 40 - 60 mTorr. In a preferred embodiment the polymerisation pressure comprises 50 mTorr. The polymerisation pressure ma comprise 25, 30, 31 , 32, 33, 34, 35, 36, 37, 38, 39, 40, 41, 42, 43, 44, 45, 46, 47, 48, 49, 50, 51, 52, 53, 54, 55,

56, 57, 58, 59, 60, 61 , 62, 63, 64, 65, 66, 67, 68, 69, 70, 71 , 72, 73, 74, 75, 76, 77, 78,

79, 80, 81 , 82, 83, 84, 85, 86, 87, 88, 89, 90, 91 , 92, 93, 94, 95, 96, 97, 98, 99, 100, 105, 110, 115, 120, 125, 150, 175, 200, 225, 250, 275, 300, 325, 350, 375 or 400 mTorr.

[076] The radio frequency plasma-enhanced chemical vapour deposition may comprise depositing the coating at a pressure between 50 - 400 mTorr; between 100 - 300 mTorr; or between 125 - 250 mTorr. In a preferred embodiment the deposition pressure comprises 150 mTorr. The deposition pressure may comprise 50, 51 , 52, 53, 54, 55, 56,

57, 58, 59, 60, 61, 62, 63, 64, 65, 66, 67, 68, 69, 70, 7.1 , 72, 73, 74, 75, 76, 77, 78, 79,

80, 81, 82, 83, 84, 85, 86, 87, 88, 89, 90, 91, 92, 93, 94, 95, 96, 97, 98, 99, 100, 101, 102, 103, 104, 105, 106, 107, 108, 109, 110, 111 , 112, 1 13, 114, 115, 116, 1 17, 118, 1 19, 120, 121, 122, 123, 124, 125, 126, Ϊ27, 128, 129, 130, 131 , 132, 133, 134, 135, 136, 137, 138, 139, 140, 141 , 142, 143, 144, 145, 146, 147, 148, 149, 150, 151 , 152, 153, 154, 155, 156, 157, 158, 159, 160, 161, 162, 163, 164, 165, 166, 167, 168, 169, 170, 171 , 172, 173, 174, 175, 176, 177, 178, 179, 180, 181, 182, 183, 184, 185, 186, 387, 188, 189, 190, 191, 192, 193, 194, 195, 196, 197, 198, 199, 200, 201 , 202, 203, 204, 205, 206, 207, 208, 209, 210, 21 1 , 212, 21.3, 214, 215, 2.16, 217, 218, 219, 220, 221, 222, 223, 224, 225, 226, 227, 228, 229, 230, 231 , 232, 233, 234, 235, 236, 237, 238, 239, 240, 241, 242, 243, 244, 245, 246, 247, 248, 249, 250, 275, 300, 325, 350, 375 or 400 mTorr.

[077] The radio frequency plasma-enhanced chemical vapour deposition may be performed at a low temperature. The temperature ma e between 10 and 50°C; between 15 and 40°C or between 20 and 30°C. In a preferred embodiment the temperature comprises room temperature or between 22 - 25°C. The temperature may comprise 10, 15, .20, 21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 35, 40, 45 or 50 °C.

[078] The monomer source may be introduced at a rate of between 0.5 and 5 cm", in; 1 and 4 cm^'Vmin; or 1.5 and 2.5 cm³/min. In a preferred embodiment tire rate may comprise 2 cnrVmin, The rate may comprise 0.5, 0.6, 0.7, 0.8, 0.9, 1.0, 1 . 1 , 1.2, 1.3, 1.4, L5, 1.6, 1.7, 1.8, 1.9, 2.0, 2.1, 2.2, 2.3, 2.4 or 2.5 en /mm,

[079] in a preferred embodiment of any of the above aspects the coating comprises a degradable coating.

[080] The degradable coating may be a biodegradable coating. [081] In one embodiment of any of the above aspects the coating comprises a thin film.

[082] The thin film may comprise a thickness between 5 - 2000 ara.

[083] The thickness of the coating according to an of the above aspects may be control led by varying one or more of lime, monomer flow, monomer concentration, pressure, radio frequency, input power or temperature.

[084] In. another embodiment of any of the above aspects the coating may comprise an amorphous, highly branched and/or cross- linked coating.

[085] hi one embodiment of any of the above aspects the degree of cross-linking may increase with the deposition power.

[086] In yet another embodiment of any of the abo ve aspects the coating may comprise a smooth, dense and/or uniform coatmg.

[087] In yet another embodiment of any of the above aspects the coating does not alter or significantly alter the nanotopography of the biomateriai substrate.

[088] In another embodiment of any of the above aspects the coating degrades slower than the substrate.

[089] In yet another embodiment of any of the abo ve aspects the coating may comprise a cytocompatihie coating. The c tocompatible coating may be more cytocompatible tha the substrate.

[090] In another embodiment of an of the above aspects the coating may comprise a bioactive agent. The bioactive agent may comprise a biologically active pharmaceutical ingredient such as, one or more of an anti -bacterial; an anti -inflammatory, a antiproliferative, an anti-angiogenic, an antirestenotic and a thrombogenic agent

[091 ] I still another embodiment of any above aspect the coating may comprise two or more deposition layers. Each of the two or more layers may comprise differin parameters.

[092] The biomateriai according to the invention may comprise a fixation device or a stent In a preferred embodiment the biomateriai of the invention comprises a stent. The fixation device ma be a soft-tissue suture.

[093] Where the terms "comprise", comprises", "comprising", "include", "includes", "included" or "including" are used in this specification, they are to be interpreted, as specifying the presence of the stated features, integers, steps or components referred to, but not to preclude the presence or addition of one or more other feature, integer, step, component or group thereof. [094] Further, any prior art reference or statement provided in the specification is not to be taken as an admission that such art constitutes, or is to be understood as constituting, part of the common general knowledge.

MM_. DESCRIPTION OF THE^AWiNGS

[095] In order that the present invention ma be readily understood and put into practical effect, reference will now be made to the accompanying illustrations, wherein like reference numerals refer to like features and wherein:

[096] Fig. 1. (A) AFM visualisation of surface topographical features of polymer coating fabricated from tea tree oil at 50W and N₂ earner gas. (B) Phase contrast images of L929 mouse fibroblast cells after 48hr of incubation at 37°C in the presence of polymer-coated poiyediylene terephthalate substrate. Images were taken with cells and the polymer-coated substrates in the incubation wells. Gentle rinsing of the substrate removes >99% of the cells, suggesting minimal surface fouling.

[097] Fig, 2. SEM micrographs of unmodified (left) and polymer-coated (right) Mg after 360 hours of immersion.

[098] Fig. 3. Total weight lost from unmodified (»), plasma-etched (■) and polymer- coated (♦) Mg samples over time as a result of immersion.

[099] Fig. 4. Human aortic smooth muscle ceils incubated in the absence of substrate (A, control) and in the presence of micoated (B) and polymer-coated (C) Mg substrates for 14 days. Top panel: the colour of the media indicated more acidic environment for controls, with media containing the uncoated Mg samples being most alkaline. Bottom panel: optical microscopy of the cells attached to the surfaces of (A) plastic control, (B) uncoated and (C) polymer-coated Mg substrates.

[0100] Fig. 5 SEM images of uncoated and polymer-coated Mg samples incubated i n the presence of THP^~1 cells for 3, 7, and 14 days.

[0101] Fig. 6 SE miages of uncoated and polymer-coated Mg samples incubated in the presence of Human aortic smooth muscle ceils for 3, 7, and 14 days.

[0102] Fig. 7 Serum 1L-6 levels for a sample pool of 8 mice implanted with uncoated

(clear circles and bars) and polymer-coated (grey circles and bars) Mg samples.

[0103] Fig, 8 SEM images of uncoated and polymer-coated Mg samples after 35 days of implantation.

[01.04] Fig. 9 Elution of rhOPG in vitro. Time-dependent release over 168 hours (7 days). P=0.051, ruskal-Wallis test. π

[0105] Fig. 10 Stimulation of IL-6 production in THP-1 cells incubated in rhOPG-eluted conditioned media over 24 hour in vitro. Data presented as median, interquartile range of 6 repeat cultures. P=O.002, Mann- Whitney lJ test.

[0106] Fig. 1.1 MG63 bone cells cultured for 14 days on uncoated and polyterpenol- coated commercial pure Ti, control - tissue culture plastic.

[0107] Fig. 12 MG63 bone cells cultured for 14 days on uncoated and polyterpenol- coated commercial pure Ti, control - tissue culture plastic

[0108] Fig. 13 Polyterpenol film deposited at 25 W for lOmins on ep-Ti (A) and 3Qmms on iTi (B).

[0109] Fig. 14 Polyterpenol-coated Mg, 14A; Uncoated: - Corrosion rates were too fast. After only a day, p^'H reached highest level in both AR and ECAP processed Mg.; .14B: Polyterpenol-coated: Corrosion occurred steadily over the period, which implies polymer coating acted as a protection barrier.

DETAILED DESCRIPTION OF THE INVENTION

[0.1 .10] The following description refers to specific embodiments of the present invention and is in no way intended to limit the scope of the present invention to those specific embodiments.

[O i l 1] Surprisingly, the present inventors have overcome at least one deficiency of the prior art by using implantable materials that may undergo controlled degradation In vivo. Similar to inert implants, the resorbable materials of the present invention provide necessary mechanical support to tissues and organs for the time required for the healing process to occur. Of sign ificant advantage, however, they gradually resorb and are safely metabolised and cleared from the body,

[01 12] The biomaterial according to th present invention may be a resorbable biomaterial. Suitable resorbable biomaterials include resorbable metals; resorbable silk; resorbable polymers; and resorbable calcium phosphate ceramics. The resorbable metal may comprise one or more of a magnesium; a maganesium alloy; an iron; an iron alloy; a zinc; and a zinc alloy. A preferable resorbable metal is magnesium or a magnesium alloy which hold great promise in tissue regeneration applications, particularly where load bearing function is required. The resorbable polymer may comprise one or more of poly (lactic acid) and poly(glycolic acid). The resorbable ceramic may comprise tricalcium phosphate.

[01 13] In another embodiment the biomaterial according to the present invention may be a permanent biomaterial such as, a non -resorbable metal and/or a. non-resorbable polymeric material. Examples of suitable non-resorbable metals include one or more of nitinoi (nickei titanium} and titanium, in one embodiment the titanium comprises commercial pure titanium. Examples of suitable polymeric materials include one or more of polypropylene; polyester; polyethylene; a polyamide (Nylon); polytetrafluoroetJhylene; polyetlieretherketone; polyewerketoneketorte silk; stainless steel; and a platinum iridium alloy.

[0114] Magnesium is particularly suited to some embodiments of the present invention becaus it is a bioresorbable material with a suitable lifetime in the body to allow it to be of therapeutic use. Many other materials either degrade too fast or don't degrade.

[0115] Advantageously, magnesium is highly biocompatible and nontoxic, with Mg ions being essential for metabolic processes. Magnesium is highly suited for fabrication of fully resorbable intravascular stents for the treatment of arterial disease, minimising the risk of chronic inflammation and late thrombosis associated with permanent metallic stents. For osteosynthesis, Mg and its alloys offer high primary stability, high tensile strength, and fracture toughness, it is also lightweight, with a density of l.74 g/cm^'\ which is 1.6 and 4.5 times less dense than, aluminium and steel, respectively. The specific gravity and elastic modulus of Mg is very close to those of human cortical bone, decreasing stress shielding effects in bone tissue, whereas a relatively high Young's modulus of commonly used metallic biomateriais means these are unable to homogeneously transfer stress between themselves and the bone, which may impede adequate bone regeneration. Furthermore, Mg bivalent ions are intimately involved in formation of biological apatites and thus determine bone fragility, bone healing and regeneration.

[0116] Through diligent research the present inventors have surprisingly discovered that application of a biodegr adable coating may overcome the barriers to clinical applications of Mg which result from its rapid corrosion rate in vivo, especially in physiological environment with pH value of 7.4-7.6 and biological fluid with chloride ions at levels of 350 mmol/L.

[0117] Advantageously, biomateriais coated according to the present invention may have a reduced degradation rate. This is of significant advantage because the coating may be used to prolong the viability of an implanted biomaterial.

[01 18] The present invention may be used to coat all or a part of the biomaterial. In certain applications the biomaterial may comprise a non-resorbable par and a resorbable part. The invention may be used to coat the entire biomaterial or only a part of the biomaterial. For example, it may be desired to coat only the resorbable part of the biomaterial. because the non-resorbable part is inert and designed to be implanted permanently.

[01 19] The part of the biomaterial that is to be coated is referred to as the substrate.

[0120] The biomaterial substrate may be polished and or cleaned before the coating is deposited. The polishing may comprise mechanical polishing to remove oxide. The cleaning may comprise ultrasonic cleaning. The ultrasonic cleaning may be done first in isopropyl alcohol and then in distilled water. After cleaning the biomaterial substrate may be air dried at room temperature.

[0121] The biomaterial of the invention may comprise a fixation device such as, a screw or plate. In a suitable embodiment the biomaterial comprises a resorbable stent. In another suitable embodiment the biomaterial comprises soft-tissue suture.

[0122] The present invention makes use of plasma polymerisation. In a suitable embodiment the polymerisation and deposition of the coating is performed using plasma-enhanced chemical vapour depositio (PECVD). A suitable form of PECVD is radio frequency plasma-enhanced, chemical vapour deposition (RF PECVD).

[0.123] The thickness of the coating according to any of the above aspects may be controlled by varying one or more of time, monomer flow rate, monomer concentration, pressure, radio frequency, input power or temperature. From the teaching herein the skilled person will understand that varying the thickness of the coating will vary the time before the coating is completely resorbed or degraded. This is because a. thicker coating will take longer to degrade. As will be explained below, a coating with an increased amount of cross-linking will also take longer to be resorbed or degraded. This highlights yet another advantage of the present invention in that it provides two mechanisms of controlling the time taken for the coating to resorb or degrade, thickness and amount of crosslinking.

[0124] Time may be used to vary the thickness by increasing the deposition time to increase the thickness. The reverse is also true, reducing the deposition time will reduce the thickness of the coating.

[0125] The monomer flow may be varied by increasing the monomer flow rate to increase the thickness or decreasing the monomer flow rate to decrease the thickness. [0126] The RF PECVD may use a radio frequency between 33-14 MHz. between 13.4 - 33.7 or between 13.5 - 13.6. in a suitable embodiment the radio frequency is 13.56 MHz. The radio frequency may be 13.0, 13.1, 13.2, 13.3, 13.4, 13.5, 13.51 , 13.52, 13.53, 13.54, 13.55, 13.56, 13.57, 13.58, 13,59, 13,6, 13.7, 13.8., 13,9 or 14.

[0127] When radio frequency plasma-enhanced chemical vapour deposition is used an input power may be between 5 and 100 RF power. In a suitable embodiment the input power is between 5 and 50 RF or between 5 and 30 RF. In a particularly suitable embodiment the input power is between 7 and 25 RF. The input power may be 5, 6, 7, 8, 9, 10, . 1 1 , 12, .13, 14, 15, 16, 37, 18, 19, 20, 21 , 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 32, 33, 34, 35, 36, 37, 38, 39, 40, 41 , 42, 43, 44, 45, 46, 47, 48, 4 o 50 RF.

[0128] Significantly the inventors have shown that the rate of degradation can be controlled by varying one or both or the thickness of the coating and the power use to produce the coating. The higher the RF power, the more cross-linked and less soluble the polyme becomes.

[0129] This is of great advantage because it means that the input powe may be controlled or varied to increase or decrease the amount of cross-linking in the polymer. The ability to provide a polymer coating with a controllable thickness or a defined amount of cross-linking and thereby a defined lifetime in the body is of great advantage because it allows the biomaterial of the invention to provide its therapeutic or structural benefit for a desired time period before being resorbed or degraded.

[0130] In a suitable embodiment the RF power is sufficient to provide in situ sterilisation and/or fiinctionalise or increase a mechanical properties of a substrate surface.

[013.1 ] Of significant advantage, both th hydrophilicity and hydrophobicity of the coating of the invention may be controlled.

[0132] The hydrophilicity may be controlled by varying feeder gas amount and/or surface roughness. The feeder gas may be introduced and the amount varied by introducing the feeder gas along with, the monomer. The feeder gas may comprise oxygen. The presence of oxygen may result in increased oxygen containing functional groups which increase the hydrophilicity.

[0133] The presence and amount of carrier gas may also be used to control the surface roughness.

[0134] The inventors have also show that coatings fabricated in the presence of a carrier gas results in a nano structured surface which is more hydrophobic and less susceptible to colonisation by cells. This can be advantageous in applications where cell attachment is detrimental. The carrier gas may comprise one or more of nitrogen; helium, argon, carbon dioxide; hydrogen and air. Nitrogen has been shown to be a particularly suttabie carrier gas.

[0135] RF PECVD polymerisation may be performed at a pressure betwee 25 - 400 mTorr, between 30 to 100 mTorr or between 40 to 60 Torr. Preferably the polymerisation pressure is 50 mTorr. The polymerisation pressure may be 25, 30, 31, 32,

33, 34, 35, 36, 37, 38, 39, 40, 41 , 42, 43, 44, 45, 46, 47, 48, 49, 50, 51 , 52, 53, 54, 55, 56, 57, 58, 59, 60, 61 , 62, 63, 64, 65, 66, 67, 68, 69, 70, 71 , 72, 73, 74, 75, 76, 77, 78, 79, 80, 81, 82, 83, 84, 85, 86, 87, 88, 89, 90, 91, 92, 93, 94, 95, 96, 97, 98, 99, 100, 105, 1 10, 1 1 , 120, 125, 1 0, 175, 200, 225, 250, 275, 300, 325, 350, 375 or 400 mTorr.

[0136] The RF PECVD may comprise depositing the coating at a pressure between 50 - 400 mTorr; between 100 - 300 mTorr; or between 125 - 250 mTorr. In a preferred embodiment the deposition pressure comprises 150 mTorr. The deposition pressure may comprise 50, 51, 52, 53, 54, 55, 56, 57, 58, 59, 60, 61, 62, 63, 64, 65, 66, 67, 68, 69, 70, 71, 72, 73, 74, 75, 76, 77, 78, 79, 80, 81 , 82, 83, 84, 85, 86, 87, 88, 89, 90, 91, 92, 93, 94, 95, 96, 97, 98, 99, 100, 101, 102, 103, 104, 105, 106, 107, 108, 109, 1 10, 1 1 1, 1 12, 1 13, 1 14, 1 15, 116, 117, 1 .18, 119, 120, 121 , 122, 123, 124, 125, 126, 127, 128, 1.29, 130, 131, 132, 133, 134, 135, 136, 137, 138, 139, 140, 141, 142, 143, 144, 145, 146, 147, 148, 149, 150, 151, 152, 153, 154, 155, 156, 157, 158, 159, 160, 161, 162, 163, 164, 165, 166, 167, 168, 169, 170, .171 , 1 72, 173, 174, 175, 176, 177, 178, 179, .180, 181 , 182, 183, 184, 185, 186, 187, 188, 189, 190, 191 , 192, 193, 194, 195, 196, 197, 198, 199, 200, 201, 202, 203, 204, 205, 206, 207, 208, 209, 210, 23 1 , 212, 213, 214, 215, 216, 217, 218, 219, 220, 221 , 222, 223, 224, 225, 226, 227, 228, 229, 230, 231 , 232, 233, 234, 235, 236, 237, 238, 239, 240, 241, 242, 243, 244, 245, 246, 247, 248, 249, 250, 275, 300, 325, 350, 375 o 400 mTorr.

[0137] RF PECVD may be performed at a temperature between 10 and 50°C; between 15 - 40°C or between 20 - 30°C. In a preferred embodiment the temperature comprises room temperature or between 22 - 25°C. The temperature may comprise 30, 15, 20, 21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 35, 40, 45 or 50 ^QC,

[0.138] Another .mechanism of controlling the amount of polymer deposited and thereby the lifetime of the coating once implanted is by varying the monomer flow rate. Based on the teachings herein the skilled person will, understand that increasing the flow rate of monomer introduction will increase the amount of polymer deposited and decreasing the flow rate of monomer introduction will decrease the amount of polymer deposited.

[0139] The monomer source may be introduced at a rate of between 0.5 - 5 cm^'Vmin; 1 — 4 cnv /min; or 1.5 - 2.5 cm /ram, In a preferred embodiment the rate may comprise 2 em'Vmin, The rate may comprise 0,5, 0.6, 0.7, 0.8, 0.9, 1.0, U, 1.2, 1.3, 3.4, 1.5, 1.6, 1.7, 1.8, 1.9, 2.0, 2.1, 2.2, 2.3, 2.4 or 2.5 cmVrnin.

[0140] .Another mechanism of varying the amount of monomer introduced is by varying the concentration of the monomer. From the teaching herein the skilled person understands that increasing the monomer concentration will increase the amount of polymer deposited and decreasing the monomer concentration will decrease the amount of polymer deposited.

[0141] The skilled person readily understands that the concentration will depend on the monomer source. Exemplified below is a monomer source comprising a distilled essential oil at a purity of 99.8%. The skilled person is readily able to select suitable concentrations depending on the monomer source from the teachings herein,

[0142] A source of the terpene or terpene analogue or derivative used in the invention may be an essential oil or a constituent of an essential oil. The present inventor's use of essential oils or constituents thereof is of significant commercial advantage due to their cost-effectiveness and read availability. The essential oil may comprise tea tree oil, lavender oil, eucalyptus oil, d-lemonene, and/or a-pinene.

[0143] The terpene or terpene analogue or derivative of the invention may comprise an acyclic terpene or acyclic terpene analogue or derivative or a cyclic terpene or cycl c terpene analogue or derivative.

[0144] The terpene or terpene analogue or derivative may comprise a monoterpene or monoterpene analogue or derivative. The monoterpene or monoterpene analogue or derivative may comprise an. acyclic monoterpene or an. acyclic monoterpene analogue or derivative or a cyclic monoterpene or cyclic monterpene analogue or derivative. The monoterpene analogue or derivative may comprise a monoterpene alcohol or a monoterpene alcohol derivative. The monoterpene alcohol may comprise terpinene-4-ol, Hnalool and/or a terpinene. The monoterpene alcohol analogue or derivative may comprise linalyl acetate. [0145] The acyclic monoterpene or acyclic monoterpene analogue or derivative may comprise geraniol, ocimene, a myrcene, citrai, citronellal, citronellol and/or halomon.

[0146] The cyclic monoterpene or cyclic monoterpene analogue or derivative may comprise a monocyclic monoterpene or monocyclic monoterpene or analogue or derivative or a bicyclic monoterpene or bicyclic monoterpene analogue or derivative. The monocytic monoterpene or monocytic monoterpene analogue or derivative may comprise a terpinene, phellandrene, terpinolene and/or p-cymene. The bicyclic monoterpene or bicyclic monoterpene analogue or derivative may comprise one or more of pineiie, carene, sabinene, e mphene, thujene, camphor, borneol and/or euealyptoL The pinene may comprise one or more of -pinene and β-pinene.

[0147] The terpinene may comprise one or more of a-terpinene, β-ierpinetie, y-terpinene or δ-terpinene. In a preferred embodiment the terpinene or terpinene derivative comprises -terpinene.

[0148] The cyclic terpene or cyclic terpene analogue or derivative may comprise a limonene or a limonene derivative. In one embodiment the limonene comprises D- limonene.

[01 9] An analogue or derivative according to the invention may comprise a functional analogue or a structural analogue.

[0150] Another significant advantage of the present invention is that the coatings are antibacterial or antimicrobial.

[0151 ] The coating according to the invention is degradahle. Because the degradable coating degrades over time when implanted it may be referred to as a biodegradable coating,

[0152] The coating of the invention may comprises a thin film comprising a thickness of between 5 and 2000nm. Being able to control the thickness of the coating is of great advantage because resorbing or degradation of the polymer coating will be faster for a thinner coating and slower for a thicker coating. This provides the ability to deliver a coating mat will provide protection of the substrate for a defined period of time. The coating according to the invention may comprise an amorphous, highly branched, and/or cross-linked coating. I one embodiment the degree of cross-linking increases with the deposition power.

[0153] The coating of the inventio ma comprise a smooth, dense and/or uniform coating. Advantageously, in one embodiment the coating of the invention does not alter or significantly alter the nanotopogfaphy of the biomaterial substrate.

[0154] In one embodiment of the invention the coating degrades slower than the substrate.

[0155] As well as being antibacterial or antimicrobial the coating of the invention may be a cytocompatible. This is of particular advantage when the cytocompatible coating is more cytocompatible than the substrate to which it is applied.

[0156] The coating of the invention may comprise a bioactive agent. The bioactive agent incorporated into the coating of the invention may include, for example, a antimicrobial, an antibiotic, a antimyobacterial an antifungal, an antiviral, an antineoplastic agent, an antitumor agent, an agent affecting the immune response, a blood calcium regulator, an agent useful in glucose regulation, an anticoagulant, an antithrombotic, an antih erlipidernic agent, a cardiac drag, a th omimetic and/or antithyroid drug, an adrenergic, an antihypertensive agent, a cholnergic, an anticholinergics, an antispasmodic, an antiulcer agent, a skeletal and/or smooth muscle relaxant, a prostaglandin, a general inhibitor of the allergic response, an antihistamine, a local anaesthetic, an analgesic, a. narcotic antagonist, an antitussive, a sedative-hypnotic agent, an anticonvulsant, an antipsychotic, an anti -anxiety agent, a antidepressant agent, an anorexigenle, a non-steroidal anti-inflammatory agent, a steroidal antiinflammatory agent, an antioxidant, a vaso-active agent, a bone-active agent, an osteogenic factor, an antiaithritics, and one or more diagnostic agents.

[01 7] The following non-limiting examples illustrate the products and methods of the invention. These examples should not be construed as limiting: the examples are included for the purposes of illustration only. The products and methods discussed in the Examples will be understood to represent an exemplification of the invention.

EXAMPLES

MA TERIALS AND METHODS

[0158] Monomer: A monoterpene alcohol derived from Melaleuca a!ternifolia essentia! oil was chosen as the precursor due to its biocompa ability and broad-spectrum antimicrobial and anti-inflammator activity.

[0159] Sample Preparation; Magnesium sheets (98 wt.% Mg) were obtained from Advent Research Materials, UK, with thickness of 0.5 mm. They were cut into smaller samples weighing between 0.05 to 1 g per sample depending on the type of studies to be conducted. Some samples were i cm x 1 cm, weighing 0.087 g per sample. Both sides of all samples were mechanically polished to remove oxide. Some samples were first ulnasonically cleaned first in isopropyl alcohol and then double distilled water for 30 mm, and air dried at room temperature. Thin polymer coatings of monoierpene alcohol were deposited onto the prepared Mg samples using plasma enhanced chemical vapour deposition technique with a radio frequency of 13.56 MHz and varying input power between 7 and..25 KW. hi particular embodiments a power of 7W, .10 W or 25W was used. For each deposition on Mg, glass slides or Si substrates were also placed into the deposition chamber, and then used for characterisation of just the coating. Masking during the deposition was also used to confirm the presence of the coating on the surfaces of Mg samples, although these samples were used to confirm the procedure and not for the actual in vitro testing. Prior to each deposition, surfaces of Mg samples were etched in situ using Ar plasma for 20 min at 50 W and ambient temperature to completely remove surface contaminants. Monoterpene alcohol (99.8% pure, distilled from Melaleuca alte ifo a essential oil by Australian Botanical Products) vapours were introduced directly into the deposition chamber at the rate of 2 cm'/min. The deposition of the nanostructure was performed at ambient temperature of 22 ^QC and working pressure of 150 mTorr.

[0160] Coating characterisation: The thickness, surface composition and chemical structure of the nanoseopicaliy thin coating were confirmed using spectroscopic eliipsometry (model M-2000D, J. A. Woollam Co. inc.). X-ray photoelectron spectroscopy (AXIS Ultra spectrometer, ratos Analytical Ltd., UK), and ATR-FTIR spectrometer (Perkin Elmer, Waltham, M,\), respectively. The wetting preference of the unmodified, plasma-etched and polymer-coated Mg samples were studied using a contact angle system (KS V 101, CCD camera) employing the sessile drop method. An average of ten measurements per sample was obtained using sterilized nanopure TLO (18.2 M Ω cm⁴), diiodomethane (Sigma Aldrich) and ethylene glycol (Sigma ASdrich). The surface free energy was calculated using Lewis acid base method. The samples were also investigated using an atomic force microscope (AFM, NT-MDT Co, Zelenograd, Russia) to visualise surface micro- and nano-topography and to quantitatively estimate surface roughness parameters. A minimum of three samples of each surface type was scanned initially to assess the overall homogeneity of the surface, and subsequently the micro- and nano-topographical profiles were studied in detail at different randomly chosen locations. Scanning was performed in semi-contact mode, using antimony doped silicon NSG 10 probes with a force constant of 1 1.8 N/m, tip curvature of 10 nm, and a resonance frequency of 240 kHz. The roughness parameters were taken as an average of a minimum^' of ten scans. Surface visualisation was performed using SEM (Jeol JSM54I0LV, Japan).

[0161 Structure and composition of polymer coatings: Morsoterpene alcohols and other volatile components of essential oils were polymerised onto a variety of substrates at low temperatures by plasma enhanced chemical vapour deposition and that microstructuxe of the resultant nanoscale coatings is typically amorphous. The results from FTIR studies confirmed that polymer coatings deposited on the surface of g were amorphous, highly branched and cross-linked, with the degree of cross-linking increasing with the deposition power. The -OH content significantly reduced compared to the original monomer. The XPS data indicated the polymer was hydrocarbon-ridi, with C and O dominating the surface. These findings are consistent with previously reported results for plasma-polymerised raoiioterpene alcohols.

[0162] As shown in the data presented in Table 1, the plasma etching significantly decreased the water contact angle of the surface, from 73.8° for unmodified Mg to 47.4° for Ar etched sample. The water contact angle of 60.3° obtained on the polymer-coated. M was consistent with previously repotted data for the RF plasma-polymerised monoterpene alcohol film. The total surface free energy values for the polymer-coated sample were significantly higher compared to both unmodified and plasma-etched Mg, with all surfaces displaying a strong electron donor component and minor electron acceptor fraction (Table 1). Surface free energy is an important indicator of the type of interaction that occur at the solid-liquid interface, such as surface wettability. It has been shown that surface events thai take place immediately after the insertion of a material into biological fluids predetermine subsequent response. Such events include wetting by physiological liquids and adsorption of proteins and cells to the biomaterials surface. There is a correlative relationship between surface wettability and blood- cell-, or tissue-compatibility, with higher degree of wettability corresponding to a higher level of cell attachment and subsequent spreading rates. The friction behaviour of an implantable rribological system is highl affected by the surface wettability, with higher wettability generally resulting in better tolerance of the biomaterial by the body.

[0163] Surface topographical analysis using SEM and AFM visualisation showed a smooth, dense and uniform polymer coating on the surface of Mg (Table 1 ). The data obtained from 10 μηι · · 10 μηι scanning areas indicated that application of the coating did not significantl alter the naiiQtopogtaphy of Mg, whereas plasma etching resul ted in a smoother surface. The maximum peak height values were very similar between, unmodified and polymer-coaled Mg samples, and significantly higher than those of plasma-etched surfaces. Similarly, the average surface roughness (R_a) and root mean square (R_¾) values of the etched surfaces of 19.6 rim and 26.7 run, respectively, were notably lower than those of unmodified and polymer-coated Mg. All of the examined surfaces displayed a surface skewness i R.u greater than zero, indicating that the surfaces were composed of a disproportionate number of peaks. The coefficient of kurtosis { R ·..·,.,·,·» was above three for the plasma-etched sample, suggesting a considerable number of sharp peaks and shallow valleys; whereas both unmodified Mg and polymer-coated sample exhibited ¾_ΙΓ of less than 3.25 Closer examination (scanning area 1 pm * 1 μτη) demonstrated that all surfaces were nanoscopically smooth. The R_Triax values were the highest for the unmodified samples, at 69.8 nm, compared to polymer-coated and plasma-etched surfaces, at 53.1 nm and 35.7 nm, respectively. Likewise, the » and R_¾ values were higher for the unmodified Mg, at 7.4 nm and 9.7 nm, against respective values for the polymer coating, at 5.4 nm and 6.9 nm, and plasma-etched Mg, at 4.0 nm and 5.3 nm, respectivel .

RESUL TS AND DISCUSSION

[0164] Coating biocompatibility and toxicity in vitro: One of the key attributes of protective coating, along with controlled degradation and excellent attachment to the underlying substrate, is its biocompatibility with cells and tissues that come into direct contact with said material The products of biodegradation should also be cytocompatible with cells in the immediate proximity and or elsewhere in the body. Polymers from essential oils, such as tea tree and lavender oil and their individual constituents, are biocompatible, being fabricated from substances that have been extensively used in aromatherapy and cosmetics industries.

[0165] Cytocompatlbility of polymer coating: Essential oil-based compounds advantageously have anti-inflammatory properties and broad-spectrum antimicrobial activity. Advantageously, the present inventions deposition at low temperature makes these coatings attractive for surface modification of temperature-sensitive substrates, such as polymers, in addition to metals, ceramics and their composites. At low levels of input RF power, the process allows for retention of a significant degree of chemical functionalities pertinent to the original molecule. Reduced energy intensity (just sufficient enoug to initiate dissociation) results in less fragmentation of the molecule, wit recombination process resembling conventional polymerisation, and also favours the entrapment of ^'unfragmented molecules within the polymer matrix,

[0166] Yet another benefit of the present invention is In situ sterilisation. Plasma treatment has been show to interfere with the membranes of man pathogenic bacteria thus killing the ceils that may have attached onto the biomaterial surface during fabrication. The in situ plasma etching not only removes the organic debris and other contaminants left behind from the cleaning process, but can also be used to functionalise the surface of implant or to increase the mechanical properties of the top surface layer.

[0167] The cytocompatibility of the polymer thin films fabricated onto glass under 7 W, 10 W, and 25 W was assessed in vitro. Polymer thin films deposited at 7 W and 10 W have been previously shown to suppress bacterial adhesion and reduce the ability of attached cells to form three dimensional biofilms. Films fabricated at or above 25 W have been previously shown to have an opposite effect, promoting adhesion and proliferation of the bacterial ceils. As shown in Table 2, after 24 hours of incubation, the viabilit of THP-1 cells was similar for wells containing inert glass and the polymer samples. The cell numbers were lower when incubated in the presence of the polymer fabricated at 25 W compared to both the inert glass and the polymer samples deposited at 7 W and 10 W. Similarly, after 48 hours of incubation, there were no significant differences in viability between all four treatment groups, although the cell number was lower lor the welts containing the 25 \V polymer-coated samples.

[0168] Similar response has been observed for mouse macrophage cells, where cells cultured in the wells containing inert glass control or polymers for 24 hours showed similar numbers and adhesive properties. Independent of the sample type, the cells exhibited healthy well-spread morphology. After 24 hours, the cell monolayers were washed to remove non-adherent ceils, their culture medium was refreshed and the attached cells were left for additional 24 hours under the same incubation conditions. The addition, as a positive control, the endotoxin LPS, a potent stimulator of inflammatory responses in macrophages was added to half the number of wells containing all four groups. After 48 hours from initial seeding, the number of attached viable cells was slightly higher for the glass control compared to polymer samples, with 25 W polymer displaying the lowest cell count. There were no significant differences between the adhesive properties of ceils in following LPS stimulation within all four groups.

[0169] Degradation of polymer coated Mg: At lower pH levels, such as that of physiological fluids, magnesium and its alloys are known to degrade following the hydrogen evolution mechanism, for which the following electrochemical equation applies: Mg + 2!¾0→ g(OH)2 + H2. The evolution of gaseous hydrogen and release of OH- ions will lead to localise alkalisatk and gas accumulation, which may be detrimental to the injured and recovering tissues. Lower degradation rate, particularly at the early stages of implantation, may lessen the possibility of further deterioration of the tissues adjacent to the implant.

[0170] AFM examination of the surfaces subsequent to 72 hours of immersion into SBF is presented i Table 3. Compared to the respective rougliness values collected for samples prior to immersion, the R_&, R and R_nmK increased for all sample types. The increase was most profound in the case of the unmodified Mg samples, where R_miiX increased from 406.4 nm to 704.4. The R_a and _q also increased significantly, from 43.5 nm to 72,6 nm, and from 55.7 am to 94.6 nm, respectively. For plasma-etched samples, the a and Ri, more than doubled, although the maximum peak height for these samples increased from 300.4 nm to 367.3 nm. The significant increase in the roughness profile of the unmodified Mg and plasma-etched Mg surfaces suggests that the samples were undergoing surface degradation, with the associated evolution of ¾ bubbles obsen ed in the fluid. In contrast, the same parameter for the polymer-coated sample increased only slightly, suggesting that these samples were significantly more stable when in contact with the simulated body fluid compared to unmodified Mg.

[0171 ] Example: Polymer coatings were subjected to in vitro cytotoxicity study to determine if leachables from the test extract would cause cytotoxicity to several types of eukaryotic cells, namel non-adherent THP-.1 monocyte cells, healthy human aortic smooth muscle ceils, MG63 bone cells, L929 mouse fibroblast cells, and Balb/c mouse macrophage cells. Phase contrast confocal fluorescent and electronic scanning microscopies showed no signs of coatings and/or leachables and/or biodegradat on products causing cell lysis or toxicity. The positive, negative, and reagent controls behaved as anticipated.

[0172] Coating anti-inflammatory behaviour in vitro: Implantation of medical devices and bioniateriais results in injury and initiation of the inflammatory response. The extent of the inj ur is dependent on the invasiveness of the surgical procedure. Pressure exerted by the physical presence of the implant, onto the adjacent tissues can worse the injur and slow down the recovery process. In addition to its shape, dimension and weight, the specific surface physical , chemical and mechanical properties of the implanted materials and the ionic and particulate products of biomaterial degradation can incite inflammatory response. The foreign body reaction involves macrophages and foreign body giant cells at the surface of the implant with subjacent fibroblastic proliferation and collagen deposition, and capillary formation. Here, macrophages are believed to play a pivotal role in the response of tissue to implants. S lowly degrading polymer coating that is made of monomers with anti-inflammatory activity can limit the implant- associated inflammation by limiting macrophage activation.

[0173] Example." Potential anti-inflanimatory activity of polymer coatings was evaluated using murine peritoneal exudates ceils (PEC) that were stimulated with bacterial LPS. To enrich PEC. 2,5 ml of 3 % (w/v) sterile Brewer thioglycoUate medium was injected into the peritoneal cavity of each BALB/c mouse. After three days, the mice were euthanized by means of CG₂ asphyxiation and sprayed with 70 % ethanol. Then, 10 ml of ice cold lavage medium (10 v/ml heparin in PBS) were injected into the peritoneal cavity of each animal, the peritoneum was gently massaged to dislodge attached cells into the lavage solution, and then the solution was extracted. The erythrocyte lysing buffer (Sigma Aldrich, Australia) was then added to remove red blood cells from the cell pallet. After repeated washing, the cells were restispended in PlM-1640 medium as described above. The cells were seeded at density of 5 ^x 10^s cells per ml into 24- well plates in the presence of unmodified and polymer-coated glass samples. After 24 hours of incubation, the samples were replenished with fres h culture media.

[0174] Half of the wells were stimulated with Escherichia coli lipopolysacchartde ( LPS) at a concentration of 50 ng/rni Prior to staining, the samples were washed three times with PBS to remove nonadherent macrophage cells. Ceils were stained using Diff-Quik and their morphology and attachment properties examined using a light microscope (BH2 Olympus, Center Valley, PA) fitted with a Qlmaging camera at 24 and 48 hours of incubation. Data were analysed using SPSS 16.0 statistical package; all values were expressed as the mean ± SD. Differences between groups were analysed by the analysis of variance (ANQVA). The value ofp< 0.05 was defined as statistically significant.

[0175] Without stimulation, cells cultured in the wells containing inert control and polymer-coated for 24 hours showed similar numbers and healthy morphology. After 24 hours, the cell monolayers were washed to remove non-adherent celts, their culture medium was refreshed and the attached cells were left for additional 24 hours under the same incubation conditions. The endotoxin LPS, a potent stimulator of inflammatory responses in macrophages was added to half the number of wells containing all four groups, After 48 hours from initial seeding, the number of attached viable cells was significantly higher for the control samples than that for polymer-coated samples. This suggests the coating may be used to prevent the attachment and subsequent activation of macrophages in vitro.

[0176] Biafaulfng and ceil attachment in vitro: The attachment of the cells could be controlled by changing the monomer chemistry and/or deposition conditions, especially deposition Ri power and the carrier gas. For example, coatings fabricated at 10W are more hydrophilic compared to those fabricated at l OGW, due to higher content of oxygen containing functional groups and slightly higher roughness. As such coatings fabricated at iOW are more readily settled on by certain types of ceils. The oxygen content can be further increased by using oxygen as a feeder gas, along with the monomer.

[0177] Example: Coatings fabricated in the presence of nitrogen as a carrier gas results in a . nanostructured surface, such as that fabricated from tea tree oil at SOW, depicted in FIG. 1. The specific chemistry and micro and nanotopography make these coatings more hydrophobic and less susceptible to colonisation by cells. This can be used advantageously in the applications where ceil attachment is detrimental to the proper functioning of the device, e.g. in devices surfaces of which comes into contact with bodily fluids.

[0178] Degradation of polymer-coated Mg in saline: Corrosion in a humid environment, such as that to which an implanted matenal is exposed, involves first the formation of metal ions, and subsequently the removal of metal atoms from the metal surface. Water and oxygen present in the bodily fluids reacts with the metal surface, causing the latter to lose electrons and tor the positivel charged metal ions to form. The ions then leave the metal to form salts in solution . Metal oxidation is followed by reduction of hydroge ions and oxygen in solution; with the reduction reactions driving the oxidation reactions.

[0179] Limiting the direct contact between the metallic biomatertal and bodil fluid can significantly slow down the corrosion. By using a slowly degrading protective coating. the degradation onset can be delayed rather than hindered completely. By optimising the properties of the coating system, a suitable degradation schedule of events can be attained.

[0180] Example: Mechanically polished magnesium samples -were etched using argon plasma for 20 min at 50 W and ambient temperature to completely remove surface contaminants. Half the samples were then coated on both sides with terpinene-4-ol using RF plasma polymerisation at 25W. Unmodified polished Mg samples were used as a control. The samples were then immersed into simulated body fluid comprising of Na ^: 142.0 mM, fC 5.0 mM, Mg ^r 1.5 mM, Ca^{2 i'} 2.5 mM, HCO^3" 4.2 mM, CI 147.8 mM, HPQ-f 1.0 mM and SO₄" ^" 0.5 mM, and incubated at 37 ± 1 °C [37]. The surface morphology of the samples after 72 hours of immersion was visualised using AFM; the samples were air dried at room temperature prior to visualisation. Measurements of the sample weight were also performed prior to immersion and after 72 and 240 hours. After 360 hrs of immersion, the samples were visualised using SEM.

[0181] Immersion of unmodified Mg samples into the simulated body fluid for extended periods of time resulted in the visible corrosion of the surface. As can be seen from SEM micrographs, FIG. 2, the smooth and compact surface of Mg was significantly transformed as a result of the immersion test, with many large, deep pits forming on the surface of corroding Mg. In view of Mg being considered for a multitude of primarily load-bearing applications, e.g. as coronary stent or orthopaedic fixation material, such corrosion is highly undesirable, as it may result in local stress concentration and rupture [37].

[0182] TH.P-1 cells incubated in the presence of the unmodified, plasma-etched and polymer-coated samples for 24 hours showed good viability and replication in all test wells. After 48 hours of incubation, however, the viability of tire cells was significantly lower in the wells containing unmodified and plasma-etched Mg samples, compared to those containing polymer-coated samples (Table 4). Hydrogen gas bubbles were also observed near the surface of the former two sample types, and the culture medium had changed colour compared to that of the wells containing either polymer-coated samples or positive control. It is believed that this change in pH and not die toxicity of Mg ions being released in the course of incubation was responsible for the reduced viability of the cells.

[0183] The surface corrosion resulted in a weight loss of up to 10 % of the initial weight of the sample (0.087 g ± 0.003 g) over 360 hours of exposure to the simulated physiological environment (FIG. 3). The extent of the corrosion taking place on the surfaces of the plasma-etched samples was similar to that on the surfaces of unmodified Mg, However, the weight loss was slightly more pronounced, with the total weight loss approaching 12 % for the plasma-etched samples. For polymer-coated samples, the total weight loss afte 360 hours of immersion was approximately 2 mg, significantly less compared to both the unmodified and plasma-etched samples. The SEM images collected from the polymer-coated samples also showed less corrosion on the surface, with some sites of polymer deterioration.

[0184] The observed slow increase in the hydrogen evolution at the initial stages of immersion, and almost constant deterioration of the polymer-coated samples after a certain period of time are important considering the importance of mechanical integrity for cardiovascular and orthopaedic applications. The amount of detached corrosion products was also significantly less in the case of polymer-coated samples compared to unmodified and plasma-etched Mg. The corrosion products of Mg in simulated body fluid are known to be composed of high concentrations of O, Ca, P, and Mg [37], These degradation products have been shown to slow down the corrosion process. The increase in solution alkalinity is known to contribute to the formation of hydroxy apatite on the surfaces of Mg(OH)? and the polymer coating, further contributing to the stability of the material in the solution. Whereas high concentration of CI ions in the physiological environment positively contributes to the corrosion of Mg, by reacting with Mg(OJ% corrosion product to form a highly soluble MgC¾ [29].

[0185] Biacompatibillty of polymer-coated Mg substrates: to addition to untimely loss of mechanical strength, fast degradation of magnesium can result in the significant evolution of hydrogen gas and release of hydroxyl ions, leading to undesirable alkalisation of the peri-implant milieu. Such alkalisation can hinder the recovery of injured or diseased tissues. Lower degradation rate, particularly at the early stages of implantation, may lessen the possibility of further deterioration of the tissues adjacent to the implant Polymer coatings can delay the onset of Mg degradation, serving as a protective coating, and/or slow down the rate of Mg degradation, by partially revealing the surface of the substrate overtime and thus controlling the area of Mg that comes into direct contact with physiological media and cells. This should allow the peri-implant space to maintain a reasonable pH by providing sufficient time for the released ions of Mg aid other degradation by-products to be metabolised and/or removed from the peri- implant milieu.

[0186] Example: Human aortic smooth muscle ceils were placed into flasks containing control (no substrate; uncoated Mg substrate) and polymer-coated Mg samples. Each flask received 1 x 10 cells, with complete confluency attained at 2.5¾ 10 ^' cells, and had a growth area of 25 cm⁴. The Dulbecco's modified Eagle's medium with 10% foetal bovine serum was used, with a working volume of 5ml. The flasks were incubated at 37°C and 5% CO2 for 14 days. Medium was refreshed once on day 7 of incubation.

[0187] Cells were found to attach well to all of the tested substrates, as can be seen hi. FIG. 4. In the case of substrate-free control and polymer-coated Mg samples, the attached cells displayed normal elongated morphology. The shape of the cells attached to the uncoated Mg surfaces was more spread-out, suggesting the cells were responding to a stiessor. The colour of the media (Figure 4, top panel) is different in the samples containing unmodified and polymer-coated Mg samples, with the media in the former being more alkaline. At pH levels characteristic of physiological fluids, magnesium and its alloys are known to degrade following the hydrogen evolution mechanism, for which the following electrochemical equation applies: Mg + 2¾0 -* Mg(OH)₂ + ¾. The evolution of gaseous hydrogen and release of OH ions will lead to localise alkalisation and gas- accumulation, which may be detrimental to the cells in the proximity of the material [50].

[0188] Higher alkalinity of the media containing uncoated Mg substrates suggests that the rate of degradation (and hydrogen evolution) in these samples is significantly higher compared to that in polymer-coated samples.

[0189] Biodegmdation of polymer-coated Mg in vitro: Immersion of pure Mg samples into the simulated body fluid for extended periods of time has been shown to result in the visible corrosion of the surface. As immersion proceeds, the smooth and compact surface of Mg is significantl transformed, with many large, deep pits forming on the surface of corroding Mg. In view of Mg being considered for a multitude of primaril load-bearing applications, e.g. as coronary stent or orthopaedic fixation material, such corrosion is highly undesirable, as it may result in local stress concentration and rupture [37].

[0190] Untimely degradation of Mg can be controlled b limiting th direct contact between the surface of the metal and the physiological fluid. Polymer coatings can provide such a barrier. The degradation onset and/or degradation rate can be controlled by tailoring the material properties of the coating, by changing the chemistry and/or. the degree of cross-linking of the polymer coating, and/or by using multiple coatings with different material properties.

[0191] Example: Samples of pure Mg (0.5mm thick foil; -~0.05g ± O.OOSg in weight} were coated with terpinene-4-ό! using RF plasma polymerisation at 25W. A single coating of approximately 200nm was used on each side of Mg substrate. Samples were placed into 24-well culturing plates. A human cell line (THP-1) was suspended in RPM1-1640 culture medium supplemented with. 100 IU/mL penicillin, 100 lJ/mL streptomycin, with 2 % (v/v) 15 pg mL L-ghitamine, 10 % (v/v) Hi-Foetal Bovine Serum, and 2 % (v/v) HEPES. The THP-1 cells were then incubated at 37 °C with 5 % CO2 in humidified atmosphere. Once sufficient cell density was achieved, the THP-1 ceils were harvested, centrifuged and re-suspended in the fresh media for culture. The ceils were seeded at density of I * 10^s ceils per ml in 24-well plates (NUNC, Thermo Fisher Scientific, Australia). Wells with RPM 640 medium alone was used as negative control, with untreated THP-1 cells used as a positive control. The plates were incubated at 37°C and 5% CO2 for 14 days. The media was changed regularly.

[0192] Surfaces of unmodified Mg samples were demonstrated to degrade more readily when exposed to THP-1 cell-containing media. The amount of detached corrosion products was less in the case of polymer-coated samples compared to unmodified Mg. The corrosion products of Mg in simulated body fluid are known to be composed of high concentrations of O, Ca, P, and Mg [37]. These degradation products have been shown to slow down the corrosion process, whereas high concentration of CI ions in the physiological environment positively contributes to the corrosion of Mg, by reactin with Mg(OHh corrosion product to form a highly soluble MgCla [29]. The increase in solution alkalinity is known to contribute to the formation of hydroxyapatite on the surfaces of Mg(OH)? and the polymer coating. Hydroxyapatite crystals can be clearly observed on the surfaces of both uncoated and polymer-coated metal surfaces. The crystals are larger and more numerous on the surfaces of uncoated Mg samples, suggesting the degradation is more advanced in these surfaces. Overall the surfaces of polymer-coated samples appear smoother, with the difference particularly apparent after 14 days of incubation (FIG. 5).

[01 3] Example: Human aortic smooth muscle cells were placed into flasks containing control (no substrate; uncoated Mg substrate) and polymer-coated Mg samples. Polymer-coated Mg samples were fabricated by coating pure Mg pieces (0.5mm thick foil; 0.05g ± 0.005g in weight) with 200nm thick polymer* using RF plasma polymerisation at .25 W and terpinene-4-ol as a precursor. Samples were placed into individual flasks, which were then filled with 5ml of working medium (Dulbecco's modified Eagle's medium, 0% foetal bovine serum) and seeded with I ^x 10⁵ cells. The flasks were incubated at 37°C and 5% CO_?, for 14 days. Medium was refreshed once on day 7 of incubation.

[0194] Similar to samples incubated in the presence of THP-1, FIG. 6 shows the surfaces of the samples degraded as a result of exposure to cell -containing media. The degradation was more pronounced, even after 3 days of incubation, with crystals developing on both surfaces. The crystals were more abundant, denser, longer and thinner on uncoated Mg surfaces, with thicker and sparser crystals observed on the polymer-coated samples. By day 14 of the incubation, larger crystals also developed on the surfaces of uncoated samples, whereas the crystal distribution and size remained relatively stable on polymer-coated surfaces throughout the incubation. The foniiation of the hydroxvapatite on the surfaces of the degrading magnesium and the polymer coating may further contribute to the stability of the material in the solution.

[0 5 ] Biocompaiibility of polymer-coated Mg ^' vivo; The evolution of hydrogen as a result of magnesium corrosion in vivo can be too rapid to be absorbed by the surrounding tissues, leading to the formation of a hydrogen balloon. Such balloon can prevent the cells from forming a direct contact with the metal surface, thus hindering tissue and organ regeneration and healing. Further to gas liberation, Mg alloy corrosion leads to local alkalization of the peri-implant milieu, which ma be detrimental to the well-being of the cells. Addition of polymer coating should slow down the metal corrosion sufficiently to sustain the removal of evolved hydrogen by the surrounding tissues and maintain the pH level at a reasonable level.

[0196] Example; Uncoated and polymer-coated Mg substrates were implanted sabcutaneously into thirty 20-week-old male C57bl6 mice. Sheets of pure Mg (0.5mm thickness) were cut into 3mm <5mm pieces and mechanically polished to remove any oxide layer. The deposition of the coating was performed using RF plasma polymerization at 25W and room temperature, with terpeiie as a monomer. Both sides of Mg samples were coated with approximately 200nm thick film. Animals were maintained strictly in accordance with guidelines for the care and use of lab animals. The animals were kept for 35 days, with the serum being collected at day 0, 7, 21, 35. On day 35, the animals were euthanized by means of carbon monoxide. Serum samples were stored at -80°C -until, analysis. ELISA was used to assay the collected serum samples for inflammation marker interleuMn (lL)-6.

[01 7] Polymer-coated samples did not induce any acute toxicity over the examination period of 35 days, with all of the test mice displaying good health post-implantation. There was no mortality or systemic toxicity in either group; overall, the test animals responded similarly to the controls. The pocket filled with hydrogen was smaller in the case of polym er-coated Mg samples, indi cating a slower rate of degradation as compared to uncontrolled Mg samples. One mouse implanted with an uncoated Mg sample developed an abscess.

[0198] FIG. 7 shows that immediately following tire implantation, the 1L-6 levels are increased for all animals. The levels the reduced over the next week to baseline levels, with mice implanted with polymer-coated Mg recovering quicker as compared to those animals implanted with uncoated Mg samples. Slower recovery in the case of the latter is likely to the earlier onset of implant degradation and higher rate of corrosion observed in uncoated Mg. Statistical evaluation of the serum data found no statistical difference in the pro-inflammatory response to the two sample types.

[0199] Biodegradation of polymer-coated Mg in vivo: The degradation of materials in vivo often differs markedly from that observed under in vitro conditions. Single-layer and multi-layer polymer coatings can delay the onset of Mg degradation in vivo and slow down the rate of Mg degradation, particularly at the early stages of implantation.

[0200] Example: Uncoated and polymer-coated Mg substrates were implanted subcutaneotisly into thirty 20-week-old male€57bl6 mice. Animals were maintained strictly n accordance with guidelines for the care and use of lab animals. The animals were kept for 35 days, then euthanized by means of carbon monoxide. Implants were surgically removed from the euthanized mice, dried and stored under ambient conditions until they were required for SEM visualisation.

[0201] After 35 days of implantation, the surfaces of polymer- coated samples were relatively smooth and displayed limited surface degradation (FIG. 8). Polymer coatin provided a sufficient barrier between the corrosive in vivo environment and the underlying substrate. It is likely that a limited amount of hydroxyapatite formed on the surface of the polymer coating during early stages of corrosion aided in stabilising the surface of the implant, further slo wing down the degradation.

[0202] Drug-loading: The use of drug delivery systems to administer drug intervention in a controlled manner directly and discretely at the site of the implantation offers a potentially safer, more effective alternative to systemic medication, improving treatment outcomes for the patient,

[0203] The protective coating can be directly loaded with at least one biologically active pharmaceutical ingredient, for example an anti-bacterial; anti-inflammatory, antiproliferative, anti-angiogenic, antirestenotic or anti-thrombogenic agent. The onset of drug release and the rate of drug release can be controlled by controlling the degree of cross-linking in the coating and/or by using a multilayer coating structure.

[0204] The degree of cross-linking is directly related to the rate at which the polymer coating will degrade, with more cross-linked films degrading at a slower rate. In the latter case, an additional polymer layer on top of drug-containing layer may temporarily protect the underlying dmg-containing layer, delaying the onset of drug release.

[0205] The properties of this layer, such as the degree of cross -linking, chemical composition, and layer thickness, can also be controlled independentl from the underlying layers, so as to the degradation rates of the layers can be different from one another. In a similar fashion, the drug-containing layer can be coated on top of the protective layer to have a more complex sequence of degradation events.

[0206] Example: A multilayer polymer structure was fabricated from terpinene-4-ol using RF plasma polymerisation, in the manner stent polymer layer (7W)/rhOPG

layer (10W)/polymer layer (50W). Stent was commercial polyethylene vascular tubing. A test compound, recombinant human osteoprotegerin (rhOPG), selected based upon known cellular effects in vitro. In vitro, sustained release of rhOPG was observed from the device into medium maintained at 37°C over a 1 8-hour (7 day) period (FIG. 9). The biological activity of the eluted protein was demonstrated by its stimulation of IL-6 production in THP-I cells (FIG. 10).

[0207] FIGS, 1 1 -13 show the attachment of osteoblast-type cells on the polyterpenol coatings fabricated under different conditions on MTi substrate: at 10W for lOmins (10/10, ~200mn thick coating); at 10W for 30mms (10/30, -600nm thick coating); at 25W for lOmins (25/10, ~200nm thick coating); at 25W for 30mm (25/30, ~600nm thick coating). Alloys such as Ti can be potentially cytotoxic to cells, resulting in poor attachment to the metal substrate. Coating NiTi with polyterpenol resulted in the improved attachment of osteoblast-type cells similar to that observed on the culture dish control. This demonstrates the potential suitability of the present invention for applications to bone regeneration. These results also show that the present invention can improve the biocorapatibility of non -resorbable metallic substrates, including those that are mechanical flexible. Being an organic film, polyterpenol coating is highly flexible and does not break as a result of mechanical deformation.

[0208] FIG. 11 shows^■polyterpenoi-coated NiTi; MG63 bone cells cultured for 14 days on uncoated and polyterpenoi-coated commercial pure Ti, control - tissue culture plastic. 10/10 = polymer deposition at 10 W for 10 rains; 10/30 - polymer deposition at 10 W for 30 mitts; etc.

[0209] FIG. 2 shows polyterpenoi-coated cpTi: MG63 bone cells cultured for 14 days on uncoated and polyterpenoi-coated commercial pure Ti, control - tissue culture plastic. 10/10 = polymer depositio at 10 W for 10 rains; 10/30 = polymer deposition at 10 W for 30 mins; etc.

[0210] FIG. 13 shows polyterpenoi-coated NiTi; MG63 bone cells cultured for 14 days display healthy size and morphology. Polyterpenol film deposited at 25 W for I Gmins on cp-Ti (A) and 30mins on NiTi (B).

[021.1] FIG. 14 shows polyterpenol coated onto two types of magnesium - pure Mg and mechanically processed (equal channel angular pressing) Mg. Equal channel angular pressing (ECAP) is a technique to substantiall enhance the strength of bulk metallic materials by the formation of a sub-micron or nano-sca!e grain structure.

[0212] Corrosion test in 0.1M NaCl solution, exposed area ; 0.785 cut (Diameter = 1cm), the remainder of the sample was covered by Epoxy-resin, amount of medium : 300ml [0213] Both types of Mg degraded more slowly when coated with 50nm thick polyterpenol coatings (B) than when not coated (A). ECAP processed samples showed better corrosion resistance than as-received Mg, indicating that ECAP process had a positive influence on corrosion resistance. This shows that a combination of mechanical processing and polyterpenol coating can significantly improve the degradation stability of Mg-based implants.

CONCLUSIONS

[0214] Amorphous monoterpene alcohol-based thin films were successfully deposited on the surface of pure magnesium using plasma enhanced chemical vapour deposition. The in vitro degradation testing indicated that polymer-coated Mg samples had a notably lower degradation rate compared to the unmodified or argon plasma etched surfaces. The polyme coating was demonstrated to be cytocompatible with THP-i cell line and murine macrophages. Based on the outcomes of this preliminary study, it could be suggested that the amorphous ultra polymer thin coatings deposited using F plasma polymerisation is a potential candidate for surface biomodification of resorbable implantable metals.

[0215] Clinical utility of biodegradable magnesium implants is undermined by the untimely degradation of these materials in vivo. Their high corrosion rate leads to loss of mechanical integrity, peri-implant aikalization and localised accumuiation of hydrogen gas. Biodegradable coatings were produced on pure magnesium using RF plasma polymerisation. A monoterpene alcohol with known an ti --inflammatory and antibacterial properties was used as a polymer precursor. The addition of the polymeric layer was found to reduce the degradation rate of magnesium in simulated body fluid. The In vitro studies indicated good eytocompatibility of non-adherent THP-1 cells and mouse macrophage cells with the polymer, and the polymer coated sample. The viability of THP-1 cells was significantly improved when in contact with polymer encapsulated .magnesium compared to unmodified samples. Collectively, these results suggest plasma enhanced polymer encapsulation of magnesium as a suitable method to control degradation kinetics of this biomaterial.

[0216] The present invention thereby had the significant effect of eliminating or at least reducing the release of large amounts of Mg" , localised hydrogen gas (¾) accumulation and aikalization. and to an untimely loss of mechanical strength of the implanted material that is associated with rapid degradation. As such, clinical application of the present invention may overcome such prior art problems have occurred in absorbable Mg stents in human coronaries which indicated the loss of the radial force and consequent early recoil as a main contributor for restenosis at 4 months as shown by intravascular ultrasound imaging of.

[0217] Hie present invention of nanoscale surface modification using 'cold' plasma polymerisation is particularly attractive, as these facilitate surface modification with little detriment to bulk properties of Mg. Plasma polymerised coatings are very thin (tens to hundreds of nanometres) uniform and defect-free; they adhere well to many types of substrates, and their properties ca be optimised for a given application by controlling processing conditions.

[0218] Advantageously, the addition of the smooth polymeric layer was found to reduce the degradation rate of magnesium in simulated body fluid, with encapsulated samples showing good cytocompatibility with non-adherent THP-l cells and mouse macrophage cells.

[0219] One of ordinary skill in the art will appreciate that materials and methods, other than those specifically exemplified can be employed in the practice of the inventio without resort to undue experimentation. All art-known functional equivalents, of any such materials and methods are intended to be included in this invention. The terms and expressions which have been employed are used as terms of description and not of limitation, and there is no intention that in the use of such terms and expressions of excluding any eq uivalents of the features shown and described or portions thereof, but it is recognized that various modifications are possible within the scope of the invention claimed. Thus, it should be understood that although the present invention has been specifically disclosed by examples, preferred embodiments and optional features, modification and variation of the concepts herein disclosed may be resorted to by those skilled in the art, and that such modifications and variations are considered to be within the scope of this invention as defined by the appended claims.

[0220] The present invention is not to be limited in scope by the specific embodiments and Examples described herein, which are intended as single illustrations of individual aspects of the invention, and functionally equivalent methods and components are within the scope of the in vention. Indeed, various modifications of the invention, in addition to those shown and described herein will become apparent to those skilled in. the art from the foregoing descriptio and accompanying Example. Such modifications are intended to fall within the scope of the present invention.

[0221] Further patent applications may be filed in Australia or overseas on the basis of, or claiming priority from, the present application. The following example claims are provided by way of example only and are not intended to limit the scope of what may or may not be claimed in any future application. It should also be noted that features may be added to, and/or removed from, the claims in order to further define the claimed invention.

[0222] All patent and scientific literature referred to herein and included in the "References" section is incorporated herein by reference.

[0223] TABLES

Table 1. Coraparative evaluation of physico-chemical attributed of uiiraoditied, plasma- etched and polymer thin film-coated Mg surfaces.

Polymer thin

Plasma-

Unmodified film

treated*

coated*⁵*

Contact angle Θ ° ± SD

9 w 71.8 ± 3.3 47.4 ± 1.5 60.3 ± 1.6

0 _m 7-1.4 ± 0.7 54.4 ·: 2.6 51.8 ± 1.5

Θ 54.7 ± 4.2 47.3 ± 3.6 35.7 ± 2.2

Surface free energy, mj n/

γ 20.8 27.1, 36.8

y^LW 31.7 35.9 41.6

γ 1 Ϊ 0.4 0.2

27.8 51.6 28.1

Surface roughness parameters, , ski ? JO imxlfyim

R_ff(a& nm 406.4 300.4 405.

R_A nm 43.5 19.6 41.1

R._?. nm 55.7 26.7 53 9

R** 0-2 0.6 0.2

Rto 0.4 3.7 0.6

* Argon plasma treatment for 30 tnin at input RF power of 50 W; ** Polymer coating deposited at an input power of 25 W; *** The components of surface free energies (7101) (mJ/m²): Lifshitz van der Waal component (γ ^Λ ), acid/base component (γ^ΑΒ), electron acceptor (γ^*) and electro donor (y^*); **** Roughness parameters: maximum peak height (R_milx), average roughness (Ra), root mean square (R_(J), surface skewness (R_Sk), coefficient of kurtosis O )- Table 2. THP 1 cell viability and macrophage attachment to unmodified and polymer- coated glass samples after incubation at 37 °C and 5% C02.

THP- 1 cell Macrophage

Number x 10 ^<! cells/ml Viability % ^'Number * 10 ² ceils/area

24 h 48 h 24 h 48 h 24 48 h LPS

7 W 1.13 ± 0.18 1.84 ± 0.56 96.9 ± 0.5 96.2 ± 0.7 2.2 ± 0.2 2.1 ± 0.4 2.0 ± 0.2 10 W 1.03 ± 0.24 1.89 ± 0.22 96.9 ± 1.0 95.6 ± 0.9 2.1 ± 0.3 2.1 ± 0.2 1.8 ± 0.2 25 W 0.77 ± 0.06 1.45 ± 0,17 95.6 ± 1.8 95.6 ± 1.0 2.3 ± 0.3 2,0 ± 0.2 1.9 ± 03

* initial cell count 0.5 * 10 ' celts/ml, viability 95 %.

Table 3. Roughness parameters of samples after 72 hours of immersion in simulated body fluid, as derived from 10 μηι ^x 10 pm scans.

Unmodified Plasma-etched Polymer-coated

R_ffi;¾ rim 704.4 367.3 479.7

.:,. ntn 72.6 41.6 50.1

R_(?, nm 94.6 52.0 65.1

0.003 -0.04 0.3

0.5 -0. 1 0,8

Table 4. Viability and proliferation of TMP-l cells after incubation in the presence of unmodified, plasma-etched, and polymer-coated Mg for 24 and 48 hours at 37 °C and 5% CQ₂.

Ceil number ^x 1 ⁵ cells/nil Viability %

24 h 48 h 24 h 48 h

Unmodified 7.03 ± 0.92 9.14 ± 1 ,33 83.7 ± 0.5 29.8 ± 1.1

Plasma-etched 25 W 7.57 ± 1.13 8.83 ± 0.96 84.1 ± 0.7 30.3 ± 0.9

Plasma-etched SOW 6.33 ± 0.71 9.70 ± 0.78 77.5 ± 0.4 20.8 ± 1.2

Polymer-coated 10 W 6.80 ± 0.47 12.57 ± 1.01 90.9 ± 0.9 823 ± 0.9

Polymer— coated 25 W 7.33 ± 0.81 1.4.97 ± 1.08 89.2 ± 1.0 78.4 ± 1.5

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Claims

1. A coating for a biomateriai., the coating comprising monomelic units comprising a terpene or terpens analogue or derivative polymerised by plasma polymerisation.

2. A method, of coating a biomateriai, the method comprising polymerising monomelic units comprising a terpene or terpene analogue or derivative with plasma polymerisation to form a polymer and depositing the polymer as a coating on a biomateriai substrate to thereby coat the biomateriai.

3. A biomateriai comprising a substrate and a polymeric coating, the polymeric coating comprising monomelic units comprising a terpene or terpene analogue or derivative polymerised by plasma polymerisation and deposited on the substrate,

4. A method of making a biomateriai comprising polymerising monomelic units comprising a terpene or terpene analogue or derivative with a plasma polymerisation to form a polymer and depositing the polymer as a coating on a biomateriai substrate to coat the biomateriai and thereby make the resorbable biomateriai.

5. A method of reducing the degradation rate of biomateriai, the method comprising polymerising monomerie units comprising a terpene or terpene analogue or deri vative with plasma polymerisation to form a polymer and depositing the polymer as a coating on a biomateriai substrate to coat the biomateriai and thereby reduce the degradation rate of the biomateriai.

6. The coating, method or biomateriai according to any above Claim wherein the coating is when used or for use as a coating for a biomateriai substrate or optionally a resorbable biomateriai substrate.

7. The coating, method or biomateriai according to any above Claim wherein the resorbable biomateriai substrate comprises or consists of a resorbable metal such as one or more of a magnesium; a maganesium alloy; an iron; an iron alloy; a zinc; and a zinc alloy.

8. A coating, method or biomateriai according to any above Claim wherein the monomelic units are polymerised and deposited by plasma-enhanced chemical vapour deposition.

9. A coating, method or biomateriai according to any above Claim wherein a source of the terpene or terpene analogue or derivative comprises a essential oil or a constituent of an essential oil such as, tea tree oil, lavender oil, eucalyptus oil, d- lemonene, and or a-pinene.

10. A coating, method or biomaterial according to any above Claim wherein the coating comprises an antibacterial or antimicrobial coating.

11. A coating, method or biomaterial according to any above Claim wherein the plasma-enhanced chemical vapour deposition comprises radio frequency plasma- enhanced chemical vapour deposition,

12. A coating, method or biomaterial according to any above Claim wherei the plasma-enhanced chemical vapour deposition comprises a radio frequency between 13 - 14 MHz; between 13,4 - 1.3.7; or between 13.5 - 13.6.

13. A coating, method or biomaterial according to any above Claim wherein when radio frequency plasma-enhanced chemical vapour deposition is used an input power may be between 5 and 100 RF power; 5 and 50 RF or between 7 and 25 RF power.

14. A coating, method or biomaterial according to any above Claim wherein the input power is controlled or varied to increase or decrease the amount of cross- linking in the p lymer.

15. A coating, method or biomaterial according to any above Claim wherein the hydrophilicity and/or hydrophobic! ty of the coating may be varied.

16. The coating, method or biomaterial according to Claim 15 wherein die hydrophobfcity is varied by increasing or adding a carrier gas and/or the hydrophilicity is increased by controlling carrier gas amount and/or controlling surface roughness.

37. A coating, method or biomaterial according to any above Claim wherein the when radio frequency plasma-enhanced chemical vapour deposition is used polymerisation is performed at a pressure between 25 - 400 mTorr; between 30 - 100 mTorr; or between 40 - 60 mTorr.

18. A coating, method or biomaterial according to any above Claim wherein the radio frequency plasma-enhanced chemical vapour deposition comprises depositing the coating at a pressure between 50 - 400 mTorr; between 100 - 300 mTorr; or betwee 125 - 250 mTorr.

19. A coating, method or biomaterial according to any above Claim wherein the monome source may be introduced at a rate of between 0.5 - 5 cm^'Vrnm; 1 - 4 cm min; or 1.5 - 2.5 cm in.

20. A coating, method or biomaterial according to any above Claim wherein the coating comprises a degradable coating.

21. A coating, method or biomaterial according to any above Claim wherein the coating comprises a thin film optionally comprising a thickness between 5 - 2000 mil

22. A coating, method or biomaterial according to any above Claim wherein the thickness of the coating is controlled by var ing one or more of time, monomer flow, monomer concentration, pressure, radio frequency, input power or temperature.

23. A coating, metliod or biomaterial according to any above Claim wherein the coating comprises an amorphous, highly branched and/or cross-linked coating.

24. A coating, method or biomaterial according to any above Claim wherein the degree of cross-linking may increase with the deposition power.

25. A coating, method or biomaterial according to any above Claim wherein the coating comprises a smooth, dense and/or uniform coating.

26. A coating, method or biomaterial according to any above Claim wherein the coating does not alter or significantly alter the nanotopography of the biomaterial substrate.

27. A coating, metliod or biomaterial according to any above Claim wherein the coating degrades slower than the substrate.

28. A coating, method or biomaterial according to any above Claim wherein the coating comprises a eytocompatible coating.

29. A coating, metliod or biomaterial according to any above Claim wherein the biomaterial comprises a fixation device; a stent; or a soft-tissue suture.