Hybrid oxide-polymer layer formed on Ti-15Mo alloy surface enhancing antibacterial and osseointegration functions

    Hybrid oxide-polymer layer formed on the Ti-15Mo alloy surface enhancing antibacterial and osseointegration functions Alicja Kazek-Kesik, Joanna Jaworska, Małgorzata Krok-Borkowicz, Monika Głoda-Cepa, Małgorzata Pastusiak, Monika Brzychczy-Włoch, El˙zbieta Pamuła, Andrzej Kotarba, Wojciech Simka PII: DOI: Reference:

S0257-8972(16)30457-1 doi: 10.1016/j.surfcoat.2016.05.073 SCT 21228

To appear in:

Surface & Coatings Technology

Received date: Revised date: Accepted date:

7 March 2016 24 May 2016 25 May 2016

Please cite this article as: Alicja Kazek-Kesik, Joanna Jaworska, Malgorzata KrokBorkowicz, Monika Gloda-Cepa, Malgorzata Pastusiak, Monika Brzychczy-Wloch, El˙zbieta Pamula, Andrzej Kotarba, Wojciech Simka, Hybrid oxide-polymer layer formed on the Ti-15Mo alloy surface enhancing antibacterial and osseointegration functions, Surface & Coatings Technology (2016), doi: 10.1016/j.surfcoat.2016.05.073

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ACCEPTED MANUSCRIPT Hybrid oxide-polymer layer formed on the Ti-15Mo alloy surface enhancing antibacterial and osseointegration functions

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Alicja Kazek-Kęsika)*, Joanna Jaworskab), Małgorzata Krok-Borkowiczc),

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Monika Głoda-Cępad), Małgorzata Pastusiakb), Monika Brzychczy-Włoche),

a)

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Elżbieta Pamułac), Andrzej Kotarbad), Wojciech Simkaa) Faculty of Chemistry, Silesian University of Technology, B. Krzywoustego Street 6, 44-100 Gliwice, Poland

Center of Polymer and Carbon Materials, Polish Academy of Sciences, M. Curie-

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b)

c)

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Sklodowskiej Street 34, 41-819 Zabrze, Poland Faculty of Materials Science and Ceramics, AGH University of Science and Technology, Mickiewicza Street 30, 30-059 Krakow, Poland

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Faculty of Chemistry, Jagiellonian University, Ingardena Street 3, 30-060 Krakow, Poland e)

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Department of Microbiology, Jagiellonian University Medical College, Czysta 18 Street,

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31-121 Krakow, Poland

Corresponding author: Alicja Kazek-Kęsik, e-mail: [email protected]

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Abstract: Poly(D,L-lactide-co-glycolide) 50/50 layer was formed on the previously anodized Ti-15Mo alloy surface and evaluated in terms of morphology (SEM), electrochemical analysis and biocompatibility (cell adhesion and biofilm formation tests). The polymer uniformly covered the porous oxide layer and filled the pores. The molar ratio of the comonomers in the copolymer was chosen to obtain the material, with degradation time of 4 weeks. The degradation process of polymer-oxide coatings was evaluated using 1H and

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C NMR. The

oxide and polymer-oxide layers were cytocompatible, and osteoblast-like MG-63 cells were well adhered to the modified surfaces. The hybrid gentamicin-PLGA loaded coatings formed on the titanium alloy surface exhibited antimicrobial activity against Staphylococcus aureus. The coatings improved the corrosion resistance of the substrate in Ringer solution. Formation

ACCEPTED MANUSCRIPT of the hybrid polymer-oxide layer on the Ti-15Mo alloy surface is a promising approach to obtain 'smart materials'.

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Key words: titanium alloy, implant, plasma electrolytic oxidation, PLGA, cytocompatibility,

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Staphylococcus aureus

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1. Introduction

Titanium alloys are considered as a future, long-term materials for orthopaedic devices. Beta-phase titanium alloys e.g. Ti-xMo, Ti-xNb-xSn, Ti-xNb-xZr are characterized by

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mechanical properties close to natural bone [1-3]. Various techniques are applied to modify

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and include specific functions on the titanium alloys surface. Depending on the applied methods the coatings are composed of ceramics (anodizing, plasma electrolytic oxidation, solgel, electrophoretic deposition, physical or chemical vapour deposition) or polymers (dip

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coatings, electrophoretic deposition). Hybrid coatings are formed on metal surface when two or more techniques are coupled or using advanced ultrafine dry powder coatings technology

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to obtain polymer-ceramic layers [4-9]. The coatings may be enriched with biologically active substances to reduce possibility of bacteria adhesion and biofilm formation. Failure of the

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implanted materials e.g. dental implants, is associated with infections caused by selected bacteria [10].Usually, the active substances like drugs cannot be incorporated into the ceramic layer using a single technique. Solvents and high temperature applied during surface modification may influence biological stability of the drugs. Thus, functionalization of the metal surface with well-adhered and biologically active substance is challenging. Nowadays, biodegradable and bioresorbable devices are used in medical practice. They are made of aliphatic polyesters obtained by the ring-opening polymerization of lactones (ROP). ROP technique makes it possible to create materials of a specified microstructure and properties, tailored to further medical application. Poly(D,L-lactide-co-glycolide) (PLGA) is highly biocompatible polymer and the products of its degradation are non-toxic for living

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ACCEPTED MANUSCRIPT organisms [11]. The active chemical substance is easy to blend with PLGA to form the polymer layer on the titanium alloy surface. The adhesion of polymer layer to the metal

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surface is inferior as compared to adhesion of converse coating such as oxide layer. Thus, the

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anodization of valve metal and then coating of its surface with polymer containing active

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substance allows to obtain multifunctional material for bone tissue regeneration. Materials are defined as 'smart biomaterials' when they are able to respond to a particular stimulus in the surrounding tissue: temperature, pH, ionic strength, and magnetic field [12,13]. The oxide

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layer obtained by anodization or plasma electrolytic oxidation improves the corrosion

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resistance and enhances the bioactivity of the substrate (due to presence in the coatings of selected compounds, incorporated from solution). Kinetics of drug release from polymers depends on their chemical composition, morphology and polymer layer thickness. The hybrid

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oxide-polymer coatings may be classified as a next-generation advanced bone implants. The poly(lactic-co-glycolic acid) (PLGA) is a biodegradable, biocompatible polymer, and its

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degradation time depends on proportions of the comonomers and polymerization conditions. PLGA is widely used to produce scaffolds for tissue engineering, as well as nano- and

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microspheres for protein and drug delivery. The active substance can be melted or encapsulated with this polymer. For example, the PLGA/gelatin [14], PLGA/chitin [15], PLGA-iron oxide [16] PLGA-dexamethasone [17] microspheres have been studied as potential materials for tissue regeneration. The bacteriostatic properties of orthopaedic implants are especially important during the first hours after implantation. Biomaterial Assosciated Infections (BAIs) are usually caused by microorganisms adhered on the implanted material. Bacteria grow and live usually in biofilm exhibiting high resistance to common antibiotics [18]. The bacteria and host osteoblast cells compete to adhere on the protein layer that forms on the implant surface. Deposition of an additional polymer layer filling in the pores in the oxide layer may decrease

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ACCEPTED MANUSCRIPT the possibility of bacteria adhesion. The polymer layers also change the amount and conformation of adsorbed proteins (due to their different chemical composition, surface

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charge, wettability, and roughness) thus influence the adhesion of both types of cells:

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osteoblasts and bacteria [19].

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The aim of this study was to design and fabricate the hybrid oxide-polymer coating with antibacterial function on β-phase titanium alloy implant material. The Ti-15Mo alloy surface was modified by plasma electrolytic oxidation in suspension with bioactive wollatostonite

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(CaSiO3) which is known to enhance osseointegration [20]. Then, the layer of biodegradable

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PLGA containing antibiotic (gentamycin) was formed on the anodized surface. PLGA layer degraded during 4 weeks and released the active substance into simulated body fluid. The main goal of the hybrid coatings was to prevent infection, reduce bacteria adhesion and

2. Materials and methods

Surface modifications and microstructure characterization

Surface treatment

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2.1.

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proliferation, and stimulate osteointegration after material implantation into bone tissue.

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The composition of the titanium-molybdenum (Ti-15Mo) alloy (BIMO Metals, Wrocław, Poland) used in this investigation was as follows: 14.73-14.98 wt.% Mo, 0.016 wt.% N, 0.06 wt.% Fe, 0.08 wt.% C, 0.15 wt.% O, 0.01 wt.% H, Ti balance. The sample possessed cylindrical shape. The surface area of the specimens was equal to 0.8 cm2. Before anodization process, the surface of the samples was ground with 600 and 1000 grit SiC abrasive paper, then etched in 1 M HF with 4 M H2SO4 solution and rinsed ultrasonically in deionized water for 5 min. The procedure of the titanium alloys surface preparation prior to anodization was the same like that described in our previous papers [21]. The Ti-15Mo alloy samples were anodized in solution 0.1 M Ca(H2PO2)2 (Alfa Aesar, Germany) with 50 g/dm3 of CaSiO3 (Carl Jäger, Germany). The process was performed in a

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ACCEPTED MANUSCRIPT cooled electrolyzer at an initial current density of 100 mA/cm2 for 5 min using a DC power supply (PWR800H Kikusui, Japan). The voltage limit during the anodic reaction was 300 V.

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Treated specimens was used anode and a titanium mesh was served as cathode. After the

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surface modifications the samples were rinsed ultrasonically in deionized water for 5 min.

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Polymer coating

Poly(D,L-lactide-co-glycolide) PLGA(50/50) was used to prepare biodegradable polymer layer. The synthesis was performed according to the following procedure:

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Monomers and initiators

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D,L-lactide and glycolide were provided by Corbion Purac BV (The Netherlands). The monomers were purified by recrystallization from dried ethyl acetate and then dried in a vacuum oven at room temperature. A low toxic initiator, zirconium(IV) acetylacetonate

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(Sigma-Aldrich) Zr(acac)4 was used as received. Methanol (99.8%) and chloroform (98.5%) were supplied by Avantor Performance Materials Poland S.A.

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Copolymerization procedure

Copolymerization of D,L-lactide with glycolide was performed in bulk using Zr(acac)4 as

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initiator. Reaction was carried out in a glass two-necked flask (50 cm3) equipped with a magnetic stirrer, under an inert argon atmosphere. The selected monomers and initiator were introduced under argon into the flask. The flask was immersed in a thermostatically controlled oil bath at 130°C. After the specified reaction time (about 72 hours), the flask was cooled to room temperature. The resulting copolymer, for purification, was dissolved in chloroform and precipitated in cold methanol. Finally, the purified material was dried in a vacuum at room temperature. Polymer coatings on anodized titanium alloys samples Implants were coated with poly(D,L-lactide-co-glycolide) by dip-coating method. PLGA solutions (10% w/v) were prepared by dissolving the polymer in dichloromethane (Sigma

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ACCEPTED MANUSCRIPT Aldrich, Poland). The gentamicin (Sigma Aldrich, Poland) (5% w/v polymer) was dissolved in dichloromethane and then blended with PLGA solution. The anodized titanium alloy

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samples were immersed into PLGA solution for 30 s and then air-dried for 24 h.

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The sample labels and treatment conditions for each sample were presented in Table 1.

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SEM investigations

The morphology of the layers formed on the titanium alloys surfaces were examined using a scanning electron microscope (SEM, Hitachi TM-3000, accelerating voltage of 15 kV). Biological investigations

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2.2.

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The samples (without gentamicin) were sterilized in ethyl alcohol (Avantor Poland) during 1 h and transferred into 24-well plates. The osteoblast-like cells were cultured at an initial density of 8.0 x 103 cells per well in 1 cm3 of EMEM cell culture medium (ATTC,

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USA) supplemented with 10% FBS, 1% penicillin/streptomycin and 0.1% amino acids and sodium pyruvate (PAA, Germany) at 37°C under a humidified atmosphere with 5.0% CO2.

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Cell metabolic activity was evaluated using Alamar Blue reagent (In Vitro Toxicology Assay Kit, Resazurin based). 0.1 cm3 Alamar Blue reagent was added and the cells were

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incubated for 4 h at 37°C. Reduction of Alamar Blue was measured fluorescently (excitation wavelength 530 nm, emission wavelength 590 nm) (FLUOstar Omega, BMG labtech) and calculated according to the following formula: %Reduction of Alamar Blue =

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100%

where: Sx – fluorescence of samples Scontrol – fluorescence of medium without cells S100%reduced – fluorescence of reagent reduced in 100% (reagent with medium was placed in The result of this measurement is the reduction ratio of the reagent (the higher the reduction, the more cells). Measurements were performed in triplicate. 6

ACCEPTED MANUSCRIPT The results of the number of metabolically active cells were calculated based on a calibration curve, and they are expressed as the average and standard deviation (S.D.) from

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three independent samples. Student’s t-test was used for statistical analysis of the differences

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between the studied groups. Statistically significant differences (p < 0.05) from Ti-15Mo are

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indicated by *, and differences between samples are marked by #.

The viability, attachment and distribution of the adhered cells were evaluated using live/dead staining. First, 1 cm3 of a phosphate buffered saline (PBS) solution supplemented

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with 2 μl (1 mg/cm3) of calcein AM (Sigma) and 2 μl (1 mg/cm3) of propidium iodide

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(Sigma) were added to each well followed by incubation for 10 min at room temperature. After that, the samples were washed twice in PBS and examined under a fluorescence microscope (Zeiss Axiovert 40, Carl Zeiss, Germany).

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2.3. The polymer layer degradation

Changes in the comonomeric unit composition as well as changes in the chain

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microstructure of the copolymer during degradation were monitored on the basis of 1H- and C-NMR spectroscopy. Spectra were recorded with NMR Bruker AVANCETM II 600 MHz

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Spectrometer with UltraShieldTM Plus magnet spectrometer operating at 600 MHz (1H) and 150 MHz (13 C) using CDCl3 as deuterated solvent. 1H-NMR spectra were obtained with 32 scans, 11 µs pulse width, and 2.65 s acquisition time. 13C-NMR spectra were obtained with 20,000 scans, 9.4 µs pulse width, and 0.9 s acquisition time. The molar mass and molar mass distribution of the polymers were determined by GPC with a Physics SP 8800 chromatograph (eluent: tetrahydrofuran, flow rate: 1 cm3/min,

detector: Shodex SE 61, calibration:

polystyrene standards, Styragel columns). Degradation study was performed in 20 dm3 of Ringer solution at 37 ± 1C for 28 days. 2.4. Bacteria adhesion test

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ACCEPTED MANUSCRIPT The reference strain used in the study was S. aureus DSM 24167 (Deutsche Sammlung von Mikroorganismen und Zellkulturen). The inoculate was incubated in 50 cm3 of Bacto™

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Tryptic Soy Broth (TSB) (Becton Dickinson) for 18 h at 37°C. The method, applied to assess

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bacterial adhesion to the examined discs, was the static adhesion assay. The experiments were

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performed in 24-well tissue culture plates, where each disc was placed and incubated for 4 h in 1 cm3 of specific bacteria strain suspension in TSB (~1 × 106 CFU/cm3) (Colony Forming Units — CFU). As a control, wells with TSB of 24-well tissue culture plate were used. After

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this time, discs were carefully washed 3 times with PBS (Institute of Immunology and

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Experimental Therapy, PAN) in order to remove non-attached bacteria cells from the surface. Within each experiment incubations were carried out in triplicates. For fluorescence microscopy observation bacteria were stained using LIVE/DEAD BacLight Bacterial

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Viability Kit (Life Technologies) according to the manufacturer’s recommendations. The observations were carried out on Olympus IX51 fluorescence microscope, equipped with a

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XC10 camera.

2.5. Electrochemical analysis

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The electrochemical analysis of the titanium alloy samples was investigated using Ringer solution, which was composed of 8.6 g/dm3 NaCl, 0.3 g/dm3 KCl, and 0.48 g/dm3 CaCl2·6H2O (Baxter, USA). The apparatus included a standard two-chamber electrolysis cell with three electrodes: a working electrode, a platinum auxiliary electrode and a Haber-Luggin capillary with a reference electrode (saturated calomel electrode – SCE). The electrolysis cell was powered by a potentiostat (PARSTAT 4000, AMETEK) equipped with Versa Studio software. The investigations included the following measurements: (a) recording the opencircuit potential (EOCP) as a function of time and (b) determining the log j=f(E) curve over a potential range of EOCP −20 mV to EOCP +20 mV (dE/dt= 1 mV/s), which provides information regarding: (1) the corrosion potential, ECORR, in mV; (2) the corrosion current

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ACCEPTED MANUSCRIPT density, jCORR, in A/cm2; (3) the polarization resistance, Rp, in Ω cm2 and (4) the cyclic polarization curve (CV) over the potential range from EOCP - 0.1 V to 3 V (dE/dt= 10 mV/s).

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The corrosion characteristics, such as the corrosion potential (ECORR), polarization resistance

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(Rp) and corrosion current density (jCORR), were extracted from the curves and calculated via

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Stearn-Geary method: (1)

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where B is the Stern–Geary constant (V), Tafel slopes (anodic βa and cathodic βc) were calculated from Tafel extrapolation (V/decade) and Rp of a corroding metal is defined as the

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slope of a potential versus current density plot at j=0: (2)

Results and discussion

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For each sample, three independent measurements were performed.

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Surface characterization and polymer degradation The surface of the Ti-15Mo alloy was anodised in suspensions contained CaSiO3, and the

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PEO-coating was characterized in detail in our previous work [21]. The aim of this work was to obtain the additional polymer layer on the previously anodized Ti-15Mo alloy surface. The plasma electrolytic oxidation (PEO) is an easy method for the titanium alloys surface modification. The advantages of the conversion oxide layers were characterized in many papers [22-25]. The oxide layer is well adhered to the substrate, and the layer thickness and chemical composition depend on the electrochemical parameters and solution used during the anodization. In our case, the oxide layer was formed in the suspension containing calcium hypophosphite (Ca(H2PO2)2) and wollastonite (CaSiO3) to enhance the cytocompatibility and bioactivity of the Ti-15Mo alloy. The oxide layer thickness was determined in our previous work [21] and for the TM-W sample it was between 3.40–4.60 μm. However, during the anodization biologicaly active substance like antibiotics or drugs could not be incorporated. 9

ACCEPTED MANUSCRIPT Thus, the biodegradable polymer coating was formed on the porous oxide layer, as a layer which can be enriched with the active substances. The SEM images of the oxide and hybrid

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oxide-polymer coatings formed on the titanium alloy surface are presented in Fig. 1.

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The polymer layer filled the pores of the oxide layer. The PLGA layer was thin and the

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structure of the porous oxide layer was still visible. The cracks present on the polymer surface were formed during observation via scanning electron microcopy. The morphology of porous oxide layer containing macropores, chemical composition and several other factors were

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found to enhance osteoblasts adhesion in vitro and osteointegration process in vivo [26,27].

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Poly(D,L-lactide-co-glycolide) was analyzed by means of NMR since this technique provides useful information on chain microstructure of the polymers. Since drug delivery systems (DDS) may contain polymeric matrix which serves as a reservoir for the drug

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(polymer-based drug delivery depots) its degradation in such cases can be controlled by the chain microstructure [28]. Thus, it is of a great importance to design appropriate material

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which will release drug in a defined manner. Observed signals in the 1H NMR (Fig. 2) and

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C NMR (Fig. 4) spectra of PLGA layers

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were assigned to the appropriate groups present in the polymer chain. Spectral region of the methine and methylene protons (1H NMR) as well as spectral region of the carbonyl carbon groups (13C NMR) is helpful in analyzing polymeric chain microstructure and its changes during degradation process. In the 1H NMR spectra recorded during degradation of the PLGA layers, next to the peaks assigned to the main chain of the polymer, signals corresponding to the short chains of oligomers were also noted (Fig. 3), what suggests advanced degradation. The presence of the oligomers was also confirmed by the GPC measurement - the dispersity (Mw/Mn) of the samples increased from 1.5 to around 5 during degradation. Although, PLGA layers were highly degraded after 28 days (Mn decreased from 15 kDa to around 2 kDa), changes in the copolymer chain microstructure were not rapid what suggests 1H NMR and 13C

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ACCEPTED MANUSCRIPT NMR analysis (Fig. 3 and 4). Changes in the average length of the lactidyl and glycolidyl microblocks present in the main polymeric chain were insignificant. This suggests even way

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of degradation, which is very advantageous for medical implants.

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NMR analysis revealed a slight, expected decrease in glycolidyl units content in the

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polymeric chain during 28 days of incubation in Ringers solution (Fig. 5). Implant failure can be classified into biological, mechanical, iatrogenic and inadequate patient adaptation failures [29]. A biological failure can be classified as early (or primary) or

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late (or secondary). The early failures occur before creation of load-bearing union of the

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implant with the bone, and the late failures occur after the prosthetic rehabilitation. Preimplant diseases, peri-mucositis and peri-implantitis influence the osseointegration. Thus, the PLGA is often blended with various active substances. In our case, the polymer coatings

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degraded during 4 weeks. The fast degradation of the polymer may result in release of the biologically active substance in a very short time. The concentration of the active substance

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should be safe for the living organism. Prevention of the infection during the first hours after implantation may improve the osseointegration process.

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MG-63 osteoblast-like cells cytocompatibility and bacterial adhesion The viability of MG-63 osteoblast-like cells after 1, 3 and 7 days on the control Ti-15Mo alloy and modified alloy samples is shown in Fig. 6. After 1 day of cell culture, the numbers of cell cultured on the TM-W and TM-W-P surfaces were similar. A slightly higher percent of Alamar Blue reduction was determined for the TM samples. After 3 and 7 days of the culture the number of cells increased on all the samples, so the percent of Alamar Blue reduction was also higher. The TM-W-P-2 was the sample, in which the coating was degraded in Ringer solution before in vitro investigation. On the TM-W-P-2 sample surface, the products of the polymer degradation were observed. After 1 and 3 days, the percent of Alamar Blue reduction was lower as compared to the other samples. However, after 7 days the percent of reduction

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ACCEPTED MANUSCRIPT increased. It means, that products of coating degradation did not significantly influence the cell culture. On the TM-W-P sample also the products of polymer degradation occurred, but

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the polymer degraded in culture medium during 7 days. On the TM-W-P significant increase

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in cells proliferation was observed. The PLGA is used to form composite scaffolds with

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hydrogels, bioglass, bioceramics [30,31]. The products of PLGA degradation are non-toxic, and for example medical product composed of PLGA (SYNERGY stent) has been approved for clinical application by US Food and Drug Administration (FDA) and European Medical

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Agency [32]. The biocompatility of PLGA-based scaffolds or microspheres [33,34] or PLGA

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membranes, nanofibres, loaded by various drugs or nanoparticles [35-38] was investigated. Depending on the final form of PLGA and presence of inert or bioactive additives, the product may be used in nerve, cornea, cartilage, ligament, ligament/tendon tissue engineering

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[32]. Thus, in our case cytocompatibility of the degradation products of PLGA were not analyzed. PLGA may be regarded as an appropriate biodegradable polymer to form hybrid

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coatings with bioactive substance on previously anodized titanium alloy surface. Fig. 7 presents the osteoblast-like MG-63 cell after 1, 3 and 7 days of culture on the

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samples modified by PEO technique and PLGA layers and on reference TCPS. The adhesion of osteoblast-like MG-63 cells on TM-W and TM samples were presented in our previous paper [39], so in this paper we present only the results for the samples modified by PEO and coated with PLGA. The fluorescence images showing live (green) and red (dead) cells correspond to the results of cell viability evaluated by Alamar Blue test. The cells were better spread and more homogenously distributed on TM-W-P than on TM-W-P-2, after 24 h and 3 days of the culture. On all of the samples no dead cell were observed after 24 h of culture. After 7 days the number of the cells significantly increased. The cells covered the surfaces of both samples. No dead cell were observed on day 3 and 7 of culture. Cytotoxicity of PLGA

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ACCEPTED MANUSCRIPT was evaluated using various cell lines, and the results confirmed that the polymer and product of its degradation were not cytotoxic [40].

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To evaluate the impact of modified titanium alloy surface resistance on Staphylococcus

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aureus adhesion, simple biological test was conducted. Fig. 8 presents representative images

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obtained for S. aureus strain after 4 h of culture on the investigated samples. On the investigated samples no dead bacteria were observed. The adhesion area of bacteria in control experiment, was 2670 μm2/image, which is 0.46% of the available surface (Fig. 8a). For the

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TM-W-P samples, slightly higher adhesion area of S. aureus was observed (Fig. 8b). The

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average surface coverage by bacteria was 9367 μm2/image which is 1.4% of the available surface. For the samples with gentamicin-loaded PLGA (TM-W-PG), no adhesion of bacteria was observed (Fig. 8c).

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Combining structural and functional materials, is an effective way to provide antimicrobial properties to the developed biomaterials. Majority of studies report that S.

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aureus, E. coli strains and coagulase-negative straphyloccoci are the most common infection pathogens in North America, Latin America and Europe [41,42]. Implant-associated bacterial

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infections are harmful for tissue regeneration. Biofilm formed on the surface of the implant is difficult to eradicate and resistant to majority of known antibiotics. Thus, many investigations are carried out to design implants with appropriate surface morphology, roughness, chemical composition and antibacterial properties [43,44]. The synthesis of antibacterial composite (PLGA/TiO2) was proposed by J-Y. Wu et al [45]. TiO2 powder (anatase) was prepared by sol-gel method and blended with PLGA. The investigation showed that composite film containing 10% TiO2 nanoparticles had an effective antibacterial properties against S. aureus and E. coli stains. Additionally, the composites were not cytotoxic for human keratinocytes and mouse fibroblasts cell (L929) [41]. Electrochemical analysis

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ACCEPTED MANUSCRIPT The results of the electrochemical analysis of the Ti-15Mo alloy samples are shown in Table 2. The open circuit potential (EOCP) for the samples was measured for 60 min in Ringer

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solution at 37°C (Fig. 9a). The surface modification caused that the EOCP values were much

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higher than those of the as-ground sample. The EOCP for the samples coated with PLGA was

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smaller compared to the only anodized sample. The polymer degraded in Ringer solution during measurement, so the EOCP curve was not well stabilized compared to the other samples. The Ringer solution penetrated the polymer layer, the EOCP was strongly changed after 2400 s

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of measurement. The highest open circuit potential was determined for the TM-W sample. It

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is well known, that on the titanium alloys the oxide layer is spontaneously formed in aerated solutions. Depending on chemical composition of the titanium alloys, the corrosion resistance is different. Anodizing process enhanced the resistant properties of the substrate in Ringer

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solution. The oxide layer formed on the TM-W was porous, and some cracks were observed on the coating [21]. The additional polymer layer may act as a glue of the oxide layer to the

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alloy surface. However the degradation of the polymer during measuring caused that finally the EOCP was lower as compared to EOCP of the only anodized sample.

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The polarization resistance and corrosion current density (jCORR) were calculated according to the Stern-Geary method. The representative polarization curves are presented in Fig. 9b. The measured corrosion potentials (ECORR) for all of the investigated samples were slightly lower as compared to the open-circuit potentials EOCP, due to polarization of the samples. The calculated jCORR was similar for the TM and TM-W sample, and slightly higher for the TM-W-P sample. The polar resistance was lower for the modified samples compared to the TM. The surface areas of the modified samples were higher due to the presence of pores, and cracks thus, the polarization and current density of the modified samples could be lower than the as-ground sample. The chemical composition of the TM-W-P sample was continuously changed, so the corrosion resistance of the sample was different. The main goal

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ACCEPTED MANUSCRIPT of the polymer-oxide layer, was to explore the possibility to enrich the polymer in active substance to be released after implantation. Thus, the polymer should also degrade fast. The

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oxide layer under the polymer layer reduced electrolytic conductivity. The anodization is

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widely used to delay degradation of biodegradable materials, such as magnesium and its

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alloys. The oxide layer was found to decrease the evolution of hydrogen, improve the corrosion resistance and increase biocompatibility of the substrate [46-49]. The polymer layers e.g. polycaprolactone, polyvinylpyrrolidone are considered as a

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barrier preventing the reaction between the magnesium or titanium alloy and corrosive ions

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[50-52]. When the magnesium alloys are coated with polymer layer, the corrosion rate is slower. However, the biodegradable polymers on the titanium alloys surface function rather as an additional layer to enhance the functionality of the material. The corrosion is blocked

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especially by oxide layer formed on the titanium alloys surface. In our case, the polymer was used as a reservoir of antimicrobial drug to prevent infection in the implant vicinity area. Conclusions

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The biodegradable PLGA polymer was deposited on the porous oxide layer formed on the

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β-phase titanium alloy. The hybrid coatings were not cytotoxic for osteoblast-like MG-63 cells. Also, the coatings after 2-week degradation significantly influenced proliferation of the cells. The surface modification resulted in antibacterial properties toward S. aureus. The electrochemical analysis showed that the polymer layer degraded uniformly, and the surface modification improved corrosion resistance of the substrate. The novel hybrid oxide-polymer coating formed on the titanium alloy surface clearly indicates a promising approach to obtain multifunctional biomaterials for bone tissue engineering. The polymer layer may be enriched in biologically active substance, preventing development of infection after material implantation.

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ACCEPTED MANUSCRIPT References [1] V.G. Pina, A. Dalmau, F. Devesa, V. Amigó, A.I. Muñoz, Tribocorrosion behavior of beta titanium biomedical alloys in phosphate buffer saline solution, J. Mech. Behav.

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Biomed. Mater. 46 (2015) 59-68.

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Table 1. The sample labels and treatment conditions.

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surface modification grinding anodization

PLGA layer

+ + +

+ +

+

TM-W-P-2

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TM-W-PG

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notes

sample after 2 weeks degradation in Ringer solution -

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gentamycinloaded PLGA -

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sample

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jCORR

mV

mV

Ω cm2

A/cm2

-317.4±5.1 190.5±10.8 -155.7±9.3

-315.6±6.1 178.7±8.1 -168.8±4.3

1.2±0.2ˑ105 6.4±0.2ˑ105 1.3±0.2ˑ104

3.2±0.1ˑ10-6 6.7±0.4ˑ10-6 3.5±0.1ˑ10-5

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Tafel slopes, V βa 0.18 0.07 0.10

βc 0.20 0.40 0.19

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TM TM-W TM-W-P

EOCP

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Figure 1. SEM images of the anodized Ti-15Mo alloy surface before (A) and after covering

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Figure 2. 1H NMR spectra of poly(D,L-lactide-co-glycolide) 50/50.

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Figure 3. 1H NMR spectra of poly(D,L-lactide-co-glycolide) 50/50 (0) before degradation and

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after one, two, three and four (1-4) weeks of degradation in Ringer solution.

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C NMR spectra of poly(D,L-lactide-co-glycolide) 50/50 (0) before degradation

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units content [%]

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40 30

%GG

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2 3 degradation time [week]

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Figure 6. MG-63 osteoblast-like cell viability after 24 h, 3 days and 7 days of culture on

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control Ti-15Mo alloy and modified alloy samples. The results are expressed as average ± S.D. Statistically significant differences from the control (Ti-15Mo) samples (p < 0.05) are

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indicated by # and from the TCPS by *.

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Figure 7. Fluorescence microscope images of cells after 1, 3 and 7 days of culture on modified titanium alloy samples, and reference (TCPS) samples, bar = 100 µm. Live/dead

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Figure 8. Representative images of a) control, b), TM-W-P c) and TM-W-PG gentamicin-

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Figure 9. Variation of the OCP with time (A), polarization curves (B) recorded for the ground and modified titanium alloy surfaces.

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Graphical abstract

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ACCEPTED MANUSCRIPT Highlights: 1. The biodegradable PLGA polymer was deposited on the porous oxide layer formed on the β-phase titanium alloy.

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development of infection after material implantation.

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2. The polymer layer may be enriched in biologically active substance, preventing

3. The surface modification resulted in antibacterial properties toward Staphylococcus

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