Metal ion-dependent tailored antibacterial activity and biological properties of polydopamine-coated titanium implants

Surface & Coatings Technology xxx (xxxx) xxxx

Contents lists available at ScienceDirect

Surface & Coatings Technology journal homepage: www.elsevier.com/locate/surfcoat

Metal ion-dependent tailored antibacterial activity and biological properties of polydopamine-coated titanium implants Hsiang Kaoa, Chun-Cheng Chenb,c, Yun-Ru Huanga, Ying-Hung Chua, Attila Csíkd,∗, Shinn-Jyh Dinga,c,∗∗ a

Institute of Oral Science, Chung Shan Medical University, Taichung City, 402, Taiwan School of Dentistry, Chung Shan Medical University, Taichung City, 402, Taiwan c Department of Stomatology, Chung Shan Medical University Hospital, Taichung City, 402, Taiwan d Institute for Nuclear Research (ATOMKI), Hungarian Academy of Sciences, Bem tér 18/c, H-4026, Debrecen, Hungary b

A R T I C LE I N FO

A B S T R A C T

Keywords: Titanium Metal ion Antibacterial activity Osseointegration

There is a great need to develop new approaches for preventing bacterial adhesion and promoting cell growth on the surface of Ti implants. To this end, various metal ions (Ag+, Cu2+, Sr2+ or Zn2+) were sprayed onto the surface of the PDA (polydopamine)-immobilized Ti. The antibacterial activity of metal ion coatings against Gram-negative E. coli and Gram-positive S. aureus were examined. The content of reactive oxygen species (ROS) production from the metal ions in bacteria was also determined. L929, RAW 264.7 and human mesenchymal stem cells (hMSCs) were used to analyze biological function of the coatings. The results of phase composition and microstructure showed that metal ions were successfully coated on PDA-immobilized Ti surface. Three divalent metal ion coatings effectively inhibited bacterial growth in a concentration-dependent manner. Samples of Cu, Sr and Zn ions in E. coli and S. aureus produced 3–8 times more ROS than the negative control. The presence of Cu and Ag remarkably reduced the cell function, as evidenced by poor cell viability and differentiation. In contrast, high content of Sr promoted hMSC proliferation and differentiation. It is eventually concluded that, taking antibacterial ability and osteogenic activity into account, 10% Sr-coated Ti had potential for implant applications.

1. Introduction

been utilized to reduce the risk of bacterial infection and/or enhance the osseointegration [9–12], where the surface properties such as chemical composition and structure are improved [13–15]. According to the literature [10,14,16–18], the individual osteogenesis or antibacterial ability of the implant does not meet all clinical needs. Although considerable advances have been made in the development of implant systems using various surface modification techniques, the need to design and manufacture new implant surfaces as an integral part of advanced implants remains a focus of research [17–19]. There is a significant need for a broad-spectrum antimicrobial agent that can prevent bacterial colonization on biomaterials, minimize the development of bacterial resistance, and especially display long-term stability. The incorporation of an antimicrobial agent on Ti surfaces is one of the effective modifications. Some in vitro studies have indicated that metal ions such as silver (Ag), zinc (Zn), and copper (Cu) on the surfaces of the implants could effectively minimize initial bacterial adhesion [4,20,21]. The unique advantages of metal ions in therapeutic

Metal implants such as titanium (Ti) in dental and orthopedic prostheses have received considerable attention because of their superior mechanical properties. Early failure of Ti implants is due to the peri-prosthetic infection, extensive inflammation, and poor osseointegration [1–3]. The implant surface is always designed to produce osseointegration that refers to a direct structural and functional connection between living bone and the implant surface, but this feature may also induce bacterial adhesion [2,4]. Therefore, preventing bacterial adhesion and colonization on the surface of bone implants remains a challenge [5–7]. The use of antibiotics is a recognized treatment for the prevention of bacterial infection [8]. However, the resistance to new antibiotics has emerged in microbial populations within a few years of the introduction of a new therapeutic drug. In addition, bacteria binding on the surface of implant may result in the formation of antibiotic-resistant biofilms [5]. Surface modification of Ti implants has

∗

Corresponding author. Corresponding author. Institute of Oral Science, Chung Shan Medical University, Taichung City, 402, Taiwan. E-mail addresses: [email protected] (A. Csík), [email protected] (S.-J. Ding).

∗∗

https://doi.org/10.1016/j.surfcoat.2019.124998 Received 26 June 2019; Received in revised form 7 September 2019; Accepted 14 September 2019 0257-8972/ © 2019 Elsevier B.V. All rights reserved.

Please cite this article as: Hsiang Kao, et al., Surface & Coatings Technology, https://doi.org/10.1016/j.surfcoat.2019.124998

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

2.2. PDA bond coat

applications include reduced cost, improved stability, and possibly less risk compared to recombinant proteins or genetic engineering techniques [21]. Among the antimicrobial agents, Ag+ ions and Ag nanoparticles are widely used, but they are toxic to humans to a large extent [22]. A small amount of Cu and Zn ions is recorded to be essential for various metabolic processes of most organisms, but a higher amount may induce toxicity [13,23]. Zinc not only plays an important role in stimulating bone formation, but also shows antibacterial ability [9,24]. The Zn/Ag co-implanted titanium provides excellent osteogenic activity and antibacterial ability [19]. In the case of Cu, Jia et al. [25] found that the regenerated cellulose films coated with Cu nanoparticles had efficient antibacterial activity against S. aureus and E. coli. Strontium (Sr) has the dual effects on stimulation of osteoblast differentiation and inhibition of osteoclast formation and resorption in vitro [26–28]. In a study by Kokubo's group [27], Sr-treated Ti implants produced high cell viability and promoted osteoblast differentiation in vitro. More importantly, in rabbit studies, Sr–Ti implant had stronger biomechanical strength and bone-implant contact than the control Ti implants. Conversely, the available information on the antibacterial activity of Sr is limited. Dabsie and others [29] found that reduction to one tenth of its control value against bacteria was observed at a Sr concentration of 1.11 mol/L. One of the challenges in designing metal-based drugs is to balance the potential toxicity of an active formulation with the positive efficacy of therapeutic agents [30]. Therefore, strict tailoring of the used concentration of metal ions is essential to compromise the effects of Ti implant surfaces on microbial and osteoblasts [9,12,30]. In addition, an ideal coating technology should be cost-effective, suitable for any implant geometry, and maintain the shape and efficacy of the implants. To solve these problems, this study loaded metal ions such as Cu2+, Zn2+ or Sr2+ into the PDA-coated Ti implant surface by using simple and effective coating technology. Many materials, such as metals, ceramics and plastics, can be induced a self-assembled layer of adhesive PDA on their surfaces without the need for further modification after soaking in a dopamine solution [15,31]. In this study, the effects of concentration and types of metal ions on the antibacterial effect and osteogenic activity of Ti implants were systematically studied. The research hypothesis was that the correct dose of the ionic coating had no toxic effect on the cells, but can inhibit the growth of bacteria. Two different bacterial species, represented by different bacterial types, were used to examine the efficacy of antimicrobial agents. E. coli is a widespread intestinal parasite of mammals and has been commonly used as a Gramnegative model organism, while the Gram-positive S. aureus bacteria is one of the major causes of community-acquired and hospital-acquired infections [7,32]. The biological behavior of coated implants was understood by using mouse fibroblast cell line L929, mouse macrophage cell line RAW 264.7 cells and hMSCs.

To prepare 2 mg/mL dopamine deposition solutions, 20 mg of dopamine hydrochloride (Sigma-Aldrich, St Louis, MO, USA) was dissolved in 10 mL of 10 mM trishydroxymethyl aminomethane buffer (Tris–buffer; pH 8.5) [15]. The Ti substrates were immersed in the deposition solution that was stirred at 60 °C using a magnetic stirrer at 500 rpm for 1 day. The coatings were then ultrasonically cleaned in ultrapure water to remove the weakly bonded PDA, and dried at 60 °C. 2.3. Metal ion coating To avoid the anion effect, the metal nitrates were adopted. Copper (II) nitrate trihydrate (Sigma-Aldrich), strontium (II) nitrate (SigmaAldrich), zinc nitrate hexahydrate (JT Baker, Phillipsburg, NJ, USA) and silver nitrate (JT Baker) are the sources of Cu2+, Sr2+, Zn2+ and Ag+ ions, respectively, without further purification. Copper ions were obtained by dissolving powders in distilled water in three different concentrations (0.2, 0.5 and 1%), while 2, 5 and 10% zinc (or Sr) ion concentrations were prepared. The 0.02% silver nitrate solution in distilled water was regarded as a positive control. A 100 μL aliquot was cast on the PDA-treated substrates by spin coating at 500 rpm for 1 min and then stored at 40 °C for 1 day. Afterwards, the coating were ultrasonically cleaned in water and dried in the oven at 40 °C for 1 day. For the sample code, “Cu0.2” represented the PDA-treated Ti implants coated with a 0.2% Cu(NO3)2 solution, while “Sr5” stood for the use of a 5% Sr(NO3)2 solution. 2.4. Composition and morphology Phase composition of sample surfaces with and without ion coatings was examined using a grazing incidence X-ray diffractometer (GIXRD; Bruker D8 SSS, Bruker Corporation, Karlsruhe, Germany) with Ni-filtered CuKα radiation operating at 40 kV and 100 mA. GIXRD measurements are conducted at an incident angle of 1° or less to maximize the signal from the thin film, which are different from the conventional X-ray diffractometer with high penetration deep into a material [33]. An incident angle at 0.5° and a scanning speed of 1°/min were used in this study. To further understand the functionalities of the coating surface, the chemical composition was analysed with an X-ray photoelectron spectroscopy system (XPS; PHI 5000 VersaProbe, ULVAC-PHI Inc., Osaka, Japan) equipped with an Al-Kα X-ray source (excitation energy: 1486.6 eV) that had a 24.4 W source energy at the anode. Surface morphologies of various samples were coated with gold and observed under a field emission scanning electron microscope (SEM; JEOL JSM-6700F, JEOL Ltd., Tokyo, Japan). A secondary neutral mass spectrometry (SNMS) system (INA-X, SPECS GmbH, Berlin, Germany) was used for the depth profiling of the samples, which was employed to calculate the thickness of the coating layer [34]. The sample surface is bombarded by an ion beam and the sputtered neutral particles, atoms and atomic clusters are detected by a mass spectrometer after electron gas post-ionization [35].

2. Materials and methods 2.1. Preparation of substrate

2.5. Corrosion measurement

Grade 2 commercially available 1 mm-thick titanium plates (99.6 at %, Grade 2, Spemet Co., Taipei, Taiwan) of 10 × 10 mm2 were selected as the substrate materials. Before coating, the surface of the substrate was mechanically polished to #1000 grit level. The substrates were then etched in 30% HNO3 for 30 min at room temperature, followed by ultrasonic cleaning in ethanol for 30 min, rinsing with distilled water for 30 min and air drying. Afterwards, an experimental procedure including the preparation of a PDA layer followed by a metal ion coating, physicochemical determination, in vitro antibacterial examination, and cell function evaluation were schematically illustrated in Fig. 1, while was described in detail below.

Using the open-circuit potential (OCP) time method, the CHI 660A electrochemical system (CH Instrument, Austin, TX, USA) was used to detect the corrosion resistance of coated samples in a simulated body fluid (SBF) solution. Two electrodes including working electrode and reference electrode (saturated calomel reference electrode; SCE) were involved. The SBF solution, the ionic composition of which is similar to that of human blood plasma, consisted of 7.9949 g NaCl, 0.3528 g NaHCO3, 0.2235 g KCl, 0.147 g K2HPO4, 0.305 g MgCl2·6H2O, 0.2775 g CaCl2 and 0.071 g Na2SO4 in 1000 mL distilled H2O and was buffered to pH 7.4 with hydrochloric acid (HCl) and tris-hydroxymethyl aminomethane (Tris, (CH2OH)3CNH2) [36]. All chemicals used were of 2

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

Fig. 1. Schematic diagram of the experimental procedure and purpose.

on stubs, coated with gold layer, and viewed using SEM.

reagent grade and used as obtained. The evaluation of the samples was started after soaking in SBF for 1 h.

2.8. ROS detection 2.6. In vitro release The production of intracellular ROS was examined with a cell-permeant fluorescent probe, 5-(and-6)-carboxy-2′,7′-dichlorodihydrofluorescein diacetate (carboxy-H2DCFDA (C400), ThermoFisher, Eugene, Oregon, USA) according to the manufacturer's guideline. The probe solution (5 mM) was prepared in 100% ethanol. 1 mL of bacteria at a density of 107 CFU/mL in Bacto tryptic soy broth (Becton Dickinson, Sparks, MD, USA) were cultured for 1 day and then were loaded with 1 μL of C400 at a working concentration of 5 μM for 30 min in the dark. Bacterial cells were collected by centrifugation at 10,600 rpm for 1 min and washed by PBS. The dye freely penetrated cells and was then hydrolyzed by intracellular esterases to deacetylated H2DCF, and trapped inside the cells. Bacteria were then treated with the antibacterial agents including Ag, Cu, Zn and Sr ions for 0.5, 1 and 4 h. The 10% H2O2 was used as a positive control [37], while the PBS solution without the addition of an antibacterial agent was regarded as a negative control. The formation of the highly fluorescent DCF is due to the oxidation of deacetylated H2DCF with various ROS like hydroxyl radicals, hydrogen peroxide, and superoxide anions [38]. After reaction, a 200 μL of solution was transferred to the fresh 96-well of microtiter plate and then was detected at an excitation wavelength of 488 nm and an emission wavelength of 525 nm using a CLARIOstar® high performance microplate reader (BMG Labtech, Offenburg, Germany). ROS generation of the metal ion samples was expressed as the fold of the negative control without treatment in terms of fluorescence intensity. All tests were conducted at least six separate experiments.

To assess the release of metal ions, a Tris-HCl solution (pH 5.0) containing 10 mM Tris was prepared. Sample surface-to-solution volume ratio of 0.1 cm−1 (10 mL solution) was used. The samples were soaked in Tris-HCl at 37 °C for 6 h, 12 h, 1 day and 3 days. The concentrations of Cu, Sr and Zn in the collected solutions were determined via an inductively coupled plasma with atomic emission spectrometry (ICP-AES; Perkin-Elmer OPT 1MA 3000DV, CT, USA) system. The surface morphology of the 3-day-soaked samples was also observed by SEM. 2.7. Antibacterial activity The antibacterial activity of samples was evaluated using E. coli (ATCC 8739 Hsinchu, Taiwan) and S. aureus (ATCC 25923, Hsinchu, Taiwan). Before seeding of bacteria, all samples were sterilized with 75% ethanol and then exposure to ultraviolet (UV) light overnight. After washing three times with phosphate buffer solution (PBS, pH 7.4), samples were placed in 24-well culture plates and 1 mL of bacteria were seeded at a density of 107 CFU/mL in Luria Bertani Broth (HiMedia, Mumbai, India) for a culture period of 3, 6, 12, 24 and 48 h. Afterwards, the antibacterial activity of samples was examined by an alamarBlue (Invitrogen, Grand Island, NY, USA) assay. At the end of the culture period, the samples were washed with PBS two times to remove loosely adherent bacteria from the sample surfaces. Each well was filled with 500 μL of solution at a ratio of 1:10 of alamarBlue to broth and then incubated at 37 °C for 30 min. Subsequently, 150 μL of the solution from each well was transferred to a new 96-well plate and read in a BioTek Epoch spectrophotometer (Winooski, VT, USA) at 570 nm with reference wavelength of 600 nm. The antibacterial efficacy of the samples was expressed as the bacterial viability normalized to that of the Ti control in terms of optical density. To further observe the amounts of bacterial colony, the samples were washed three times with PBS and fixed in 2% glutaraldehyde (Sigma, St. Louis, MO, USA) after 24 h seeding. After that, the samples were dehydrated using a graded ethanol series for 20 min at each concentration and then were mounted

2.9. Cell culture The mouse fibroblast cell line L929 (BCRC RM60091, Hsinchu, Taiwan), mouse macrophage cell line RAW 264.7 cells (BCRC 60001) and hMSCs (Cell Engineering Technologies, Coralville, IA, USA) were used to evaluate biological behaviours of the coating implants. The cells were suspended in Dulbecco's modified Eagle medium (DMEM; Gibco, Langley, OK) containing 10% fetal bovine serum (FBS) (Gibco) and 1% penicillin/streptomycin solution (Gibco) in 5% CO2 at 37 °C. Before cell incubation, specimens were sterilized by washing in 75% ethanol and 3

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

2.12.3. Extracellular matrix mineralization Mineralized matrix synthesis was analysed using Alizarin Red S staining, which identifies calcium deposits. After culturing for 14 and 21 days, hMSC cells were washed with PBS and fixed in 4% paraformaldehyde (Sigma-Aldrich) for 10 min at 4 °C. It was then stained for 10 min at room temperature in 0.5% Alizarin Red S (Sigma-Aldrich) in PBS. The stained cells were completely washed with PBS to reduce nonspecific Alizarin Red S staining. To quantify matrix mineralization, the calcium mineral precipitate was destained with 10% cetylpyridinium chloride (Sigma-Aldrich) in PBS for 30 min at room temperature. The absorbance of Alizarin Red S extracts was measured at 560 nm using a BioTek Epoch microplate reader. The analysis was performed in three separate experiments.

exposure to UV light overnight. 2.10. L929 cytotoxicity The cytotoxicity assay of the various implants was carried out according to ISO 10993–5. L929 cells were incubated on the specimens for 12, 24 and 48 h. One mL of Cell suspensions (104 cells per well) were seeded over each of the samples in a 24-well plate. 100 μL of 10% dimethylsulfoxide (DMSO; Sigma-Aldrich) was used as a positive reference material, and Ti plate was used as the negative reference material. After the established L929-cell incubation period, the cytotoxicity was assayed using the MTT (3-(4,5-dimethylthiazol-2-yl)2,5-diphenyltetrazolium bromide; Sigma-Aldrich) assay, in which tetrazolium salt was reduced to formazan crystals by the mitochondrial dehydrogenase of living cells. Briefly, 3 h before the end of the incubation period, 100 μL of MTT solution and 900 μL of DMEM containing 1% penicillin/streptomycin were added to each well. Upon removal of the MTT solution, 500 μL of DMSO were also added to each well. The plates were then shaken until the formazan crystals had dissolved, and 150 μL of the solution from each well was transferred to a new 96-well plate. Plates were read using a BioTek Epoch microplate reader at 570 nm, with a reference wavelength of 650 nm. The absorbance results were recorded for six independent measurements. The viability data (%) were obtained by normalizing to the Ti control in terms of absorbance.

2.13. Statistical analysis One-way analysis of variance (ANOVA) was used to evaluate significant differences between means in the measured data. The significance of standard deviation in the measured data of each specimen under different experimental conditions was determined by using Scheffe's multiple comparison testing. In all cases, the results were considered statistically significant, with p-values below 0.05. 3. Results 3.1. Phase and chemical composition

2.11. RAW 264.7 cell viability Fig. 2a shows the GIXRD patterns of the Ti surfaces before and after ion coating on the PDA layer. Two characteristic peaks located at around 38.4° and 40.1° were attributed to the (002) and (101) crystal faces of the Ti implants [39]. It appears that neither PDA nor Ag coating greatly affected the peak intensities of the Ti control. In contrast, peaks of (002) and (101) reflections were found to decrease on the ion coated surfaces with high coating concentrations such as Cu1, Sr10 and Zn10. Undoubtedly, these divalent metal ion coatings exhibited corresponding characteristic peaks from the metal ion component. For example, the Zn10 coating showed clear diffraction peaks at 2θ = 27.0° and 34.4°, while the Sr10 and Cu1 coatings had a peak at 32.5° and 25.8°, respectively. To further identify the presence of metal ions, XPS spectra survey of ion-coated samples on the Ti surfaces were examined. In addition to C 1s and O 1s, the PDA coating showed the appearance of N 1s at 399.6 eV to some extent, as shown in Fig. 2b. As expected, the Ag was found on the Ag-coated surface, which included the binding energy at 368.6 eV (3d), 572.6 (3p3/2) and 603.6 (3p1/2). Cu 3p and Cu 2p took place at the 78.0 and 936.0 eV, respectively, while Sr 3d and 3p were at a binding energy of 134.5 and 270.5 eV, respectively. The binding energy values at 92.3, 1026.3, and 1048.3 eV were ascribed to Zn 3p, 2p3/ 2 and 2p1/2, respectively.

RAW 264.7 macrophage cells were used to examine the osteoclastic responses to the coating implants using the MTT assay as described previously. Cell suspensions (104 cells per well) were seeded on sample surfaces in a 24-well plate for 1, 3 and 7 days. Data were expressed as percentages normalized to the Ti control from six independent measurements. 2.12. hMSCs responses 2.12.1. Cell attachment and proliferation The osteogenic properties of the coating samples were evaluated by incubation with hMSCs at passage 3–6. hMSCs were seeded onto the sterilized samples at a density of 104 cells/well in a 24-well plate. The cells were grown in the osteogenic induction medium consisting of DMEM supplemented with 10% fetal bovine serum, 1% penicillin/ streptomycin, 10 nM dexamethasone, 10 mM β-glycerophosphate, and 0.05 mM ascorbic acid. To assess attachment, cells were cultured for 12 h and 24 h. Proliferation was assessed at days 3 and 7. Medium was exchanged at 2-day intervals. Cells cultured on the Ti group as a control. The MTT assay was used to detect the absorbance. Data were the average of three independent measurements. 2.12.2. ALP activity To evaluate early cell differentiation, the alkaline phosphatase (ALP) activity of hMSC at a density of 104 cells per well on the various coating samples was examined after 7 and 14 days of incubation. ALP catalysed the hydrolysis of the colourless organic phosphate ester substrate (p-nitrophenyl phosphate; pNPP) to p-nitrophenol (a yellow product) and phosphate. ALP activity was measured using the TRACP & ALP assay kit (Takara, Shiga, Japan) according to the manufacturer's instructions. To perform the assay, after incubation, cells were washed with 0.9% NaCl and lysed with 300 μL of 1% NP-40 (Sigma-Aldrich) in 0.9% NaCl. For measurement purposes, 300 μL of the substrate solution (20 mM Tris-HCl, 1 mM MgCl2, 12.5 mM p-nitrophenyl phosphate, pH 9.5) was added to each well and allowed to react at 37 °C for 30 min. The reaction was stopped by adding 300 μL of 0.9 N NaOH and read at 405 nm using a BioTek Epoch microplate reader. Each value represented an average of three runs.

3.2. Morphology and coating thickness Fig. 3 indicates surface SEM images of various coatings. It can be clearly seen that numerous granules built up the coating structure on PDA-coated surface (Fig. 3b) compared with a smooth surface of the polished Ti substrate (Fig. 3a). The addition of Ag ion to the PDA bond coat produced white dots (Fig. 3c). In sharp contrast, the Cu, Sr, and Zn incorporation resulted in a distinct morphology covering the PDAcoated Ti surface. Cu-coated surface (Fig. 3d) exhibited a film-covered morphology, while Sr (Fig. 3e) and Zn surfaces (Fig. 3f) showed a particle-aggregate structure. SNMS showed the element depth profiles of some coating samples (Fig. 4). Since the PDA acted as an adhesive layer on the Ti surface, C and Ti elements were used as reference elements. It is reasonable to take into account that the Ti concentration on the surface of all coating groups gradually increased and reached the saturation, while the C 4

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

Fig. 2. (a) GIXRD patterns and (b) XPS spectra survey of various samples.

3.3. Electrochemical corrosion To examine the corrosion behavior of the Ti implants with and without ion modification, the change in OCP of the samples in SBF over time is demonstrated in Fig. 5. The OCP of all samples reached steady state after soaking in SBF for 3 h. During soaking, the OCP of the Ti control shifted to a more negative value, indicating its tendency to corrosion. Compared with the Ti control, all coating samples showed higher potential values after 3 h of testing, exhibiting excellent corrosion resistance. The PDA coating on the Ti surface exhibited the highest corrosion-resistant ability at the early stage of soaking. Interestingly, the OCP value of the Ag-incorporated coating increased with the increase in the measurement time and then reached a higher OCP value than other coatings, possibly due to the stable passive film formation. It appeared that higher ion concentrations resulted in lower OCP values. At the same concentration, the difference in OCP between Sr and Zn coatings was not obvious.

3.4. In vitro release To further understand the acidic pH value-induced dissolution process, the in vitro release profiles of various metal ions from the coating samples after soaking in Tris-HCl with pH 5.0 for the predetermined periods of time are plotted in Fig. 6. For Ag coating, the ultra-low release amount was less than the detection limit of ICP-AES, thus resulting in the unavailable data. It should be noted that except Ag, all coatings could attest the burst release at the first 6 h of soaking in pH 5.0. The Cu1 surface released 34 ± 2 ppm after 3 days, while Cu02 was 8 ± 1 ppm. The Zn-coated surface had a similar release profile to the Sr-coated surface at the same coating concentration of 2 wt%. However, it can be clearly seen that during the soaking time intervals, the release amount of the 10 wt% Sr-coated surface (533 ± 10 ppm) was twice the release amount of the 10 wt% Zn surface (284 ± 23 ppm). The use of higher coating concentrations on PDA-modified Ti surfaces would reasonably result in higher release levels. Broad face SEM micrographs of the ion coatings after soaking in a Tris-HCl solution for 3 days are shown in Fig. 7. The ion coatings such as Cu02 (Fig. 7a), Sr2 (Fig. 7d), and Zn2 (Fig. 7g) at low coating concentrations had a similar structure to the PDA layer. However, the 0.5 wt% (Fig. 7b) and 1 wt% Cu (Fig. 7c) coating showed residual particles. In the case of Sr coatings, the surface structure had a significant change compared with the original form over the range of coating concentrations used (Fig. 7d–f), demonstrating an almost dissolution of Sr. With regard to the surface morphology of the 5 wt% Zn

Fig. 3. Surface SEM images of (a) Ti control, (b) PDA bond coating, (c) Ag coating, (d) Cu1 coating, (e) Sr10 coating and (f) Zn10 coating.

element would approach a decreased tendency from the surfaces. Depth profiling of the PDA-coated samples with (Fig. 4c) and without (Fig. 4b) Ag incorporation showed a modified layer thickness of about 250 nm which was higher than the oxide layer of Ti control (50 nm) (Fig. 4a). Thus, the coating thickness of the PDA layer alone was around 200 nm, while that the thickness of a native Ti oxide layer was not excluded. SNMS also identified the presence of Cu (Fig. 4d), Sr (Fig. 4e) and Zn (Fig. 4f), indicating a similar average depth of about 350 nm. It can be deduced to be approximately 100 nm for the thickness of the divalent metal ion coatings. For the same ion resource, the thickness of the ion coating did not increase with increasing used concentration.

5

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

Fig. 4. SNMS depth profiles of (a) Ti control, (b) PDA bond coating, (c) Ag coating, (d) Cu1 coating, (e) Sr10 coating and (f) Zn10 coating to show the coating thickness.

Fig. 5. Open circuit potential vs. time curves of various samples in SBF. Fig. 6. Release profiles of the Cu, Sr and Zn from PDA-coated Ti surfaces as a function of soaking time in a Tris-HCl with pH 5.0.

6

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

Fig. 7. Surface SEM images of (a) Cu02, (b) Cu05, (c) Cu1, (d) Sr2, (e) Sr5, (f) Sr10, (g) Zn2, (h) Zn5 and (i) Zn10 coatings after soaking for 3 days in Tris-HCl solution at pH 5.0.

3.6. ROS production

coating, a significant change was found with the presence of an etchinginduced band structure (Fig. 7h), particularly for a 10 wt% Zn coated surface (Fig. 7i).

To more fully evaluate the role of metal ions in ROS generation, first, the quantification of the fluorescence intensity of the C400 probe without bacterial seeding should be performed as a blank test. The results showed that the intensity of the blank tests in all tested agents was similar due to the lack of bacteria (not shown), confirming that the metal ion itself did not cause the ROS production. In contrast, exposure of various metal ions to E. coli or S. aureus bacteria induced a significant increase in the intracellular ROS in a dose-dependent manner (Fig. 11). At various reaction time intervals, the degree of ROS generation within the two bacterial strains caused by 10 wt% H2O2 and Ag ion was much higher than that obtained from the other ions. Interestingly, as the reaction time increased, the levels of ROS treated with H2O2 and Ag gradually decreased relative to the negative control. Regardless of the reaction time, the Cu, Sr and Zn ion samples revealed that the ROS data generated from E. coli and S. aureus were within 3–8 times of the negative control. The higher ion concentration produced higher ROS level. Of note, the use of a low Cu concentration, such as 1 wt%, could result in a similar amount of ROS to the 10 wt% Sr or 10 wt% Zn.

3.5. Antibacterial activity When compared with the Ti control, the viability (%) of E. coli on the pure PDA coating was slightly decreased throughout the culture time (Fig. 8a). As expected, E. coli cultured on the ion coatings had a significant lower viability than those on the PDA coatings at all culture time points. At the first 3 h of seeding, the bacterial survival percentage on the surfaces of all ion coatings was less than 15%, indicating a relatively effective antibacterial activity at the initial stage. After culture for 48 h, the Ag, Cu02, Zn2, and Zn5 coatings showed values of about 30%. It is worth noting that the Cu05, Cu1, Sr5, Sr10 and Zn10 coatings caused a lower bacterial survival than the Ag coating. In the case of S. aureus, a similar trend in the survival percentage was also found (Fig. 8b). As a result, the antimicrobial activity of Cu02, Sr5 and Zn2 coating was almost equivalent that of the Ag coating during the culture time periods. Not surprisingly, it also showed a clear concentrationdependent antibacterial ability of the ion coatings against S. aureus. To further demonstrate the antibacterial ability of the ion-coated samples, bacterial colonies on the surfaces were observed by SEM. SEM images of E. coli and S. aureus bacterial adhesion on the surfaces of the samples after seeding bacterial species for 24 h are shown in Figs. 9 and 10, respectively. Gram-negative E. coli (Fig. 9a) colonies spread and grew well on the surface of the PDA-coated Ti substrate. The number of rod-shaped E. coli bacteria adhering to the ion coatings was appreciably reduced in comparison with that of the PDA coating, regardless of which ion was used (Fig. 9b–h). Similarly, spherical Gram-positive S. aureus bacteria on the PDA coating aggregated into grape-like colonies (Fig. 10a); however, few bacteria adhered to the ion coatings (Fig. 10b–h).

3.7. L929 cytotoxicity The cytotoxicity of the coating samples should be examined prior to the use in vivo. The viability results of the MTT assay for L929 with respect to Ti control are shown in Fig. 12. As the incubation time increased, the positive control (DMSO) exhibited a high degree of cytotoxicity as viability was reduced to 29% from 82%, whereas the PDAcoated sample was slightly above 100% viability. Significant depression of cellular activity took place when L929 cells were seeded on Ag and Cu-coated implants. Cu coatings resulted in cytotoxicity similar to the Ag coating and the cell viability on the both coatings decreased with increase in the culture time, like DMSO. It is evident that the Ag coating had a viability value of less than 30% after 48 h of culture. The Zn coating onto the PDA layer also resulted in lower cell viability. The Cu and Zn coatings elicited a concentration-dependent decrease in cell 7

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

Fig. 8. (a) E. coli and (b) S. aureus bacterial viability cultured on the various coating surfaces for various time points. The Ti control was as a reference.

viability at all culture time points. In contrast, all Sr coatings showed no signs of cytotoxicity at all culture time points. Of particular note is that L929 cells seeded on the Sr10 coating exhibited a viability value of about 120% after 48 h of culture.

Fig. 9. SEM images of E. coli bacterial adhesion on the surfaces of (a) PDA, (b) Ag, (c) Cu02, (d) Cu1, (e) Sr2, (f) Sr10, (g) Zn2 and (h) Zn10 coatings after culture for 24 h. The arrows indicate E. coli bacteria.

surfaces at all culture time points, while PDA-modified layer did not adversely affect the cell attachment (Fig. 14a). Evaluation of cell proliferation was also shown in Fig. 14b after culture for 3 and 7 days. It was found that the absorbance value increased steadily with increasing in the culture time point, revealing increasing numbers of viable cells, except the Ag coating. The presence of Cu and Zn resulted in the significantly (p < 0.05) lower cell attachment and proliferation on the coating surfaces compared with the Ti control (Fig. 14a–b). The increased content of Cu and Zn ions in the coating gave rise to the lower hMSC cell growth. Contrary to the findings, cells seeded on the surface of the Sr-containing sample had higher metabolic activity than the other metal ion coatings, which was comparable to the Ti control.

3.8. Macrophage viability After 1 day of culture, RAW 264.7 cells seeded on PDA-coated surfaces with and without Sr ion maintained a higher viability than the Ti control, while the positive control (DMSO) and Ag ion coating showed extremely lower survival (Fig. 13). The Cu and Zn coatings significantly caused RAW 264.7 toxicity, and the viability decreased as the ion concentration increased. When cultured for 3 and 7 days, the number of viable macrophages seeded with metal ions was found to be significantly reduced compared to the PDA coating alone. There was a similar trend in viability for Cu and Zn coatings. At the same coating concentration, the Sr ion coating had a higher viability than the Zn ion coating.

3.9.2. ALP activity Fig. 14c shows the amount of intracellular ALP in hMSC cells cultured on the surfaces of various coatings for 7 and 14 days. The ALP level of cells on PDA coating surfaces was similar to that of the Ti control, while the Ag-coating group significantly (p < 0.05) reduced the ALP level at the two culture time points. After 7 days of culture, all Cu coatings showed lower ALP activity, indicating lower cell differentiation at the earlier stage of culture. For all Sr-containing and Zncontaining surfaces, higher ALP levels than the Ti control was determined. As the culture time was extended to 14 days, all other ion

3.9. hMSC responses 3.9.1. Cell attachment and proliferation To elucidate the effects of various metal ions on osteogenic activities, the biological function of hMSC cells cultured on coating samples were determined. It can be clearly seen that the attachment of hMSC cells cultured on the surface containing Ag was significantly (p < 0.05) lower than that on the surface of the Ti control and PDA-modified 8

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

Fig. 12. Cytotoxicity of various samples seeded with L929 cells at various timepoints. DMSO solution was as a positive control. Cell viability was normalized to Ti control.

3.9.3. Mineralization To further explore the effects of different ion-coated samples on cell mineralization, quantification of extracellular matrix mineralization was performed by the Alizarin Red S assay. As culture time increased, Ca deposit expression increased (Fig. 14d), except for the Ag sample. After 14 and 21 days of culture, the Ca contents on the Ag and Cu surfaces were significantly (p < 0.05) lower than those obtained from the Ti control and PDA surfaces. Similar to the findings of the ALP activity, hMSCs grown on the surfaces containing lower Zn and higher Sr had a greater content of Ca deposits compared to the Ti control after culture for 21 days. It was noted that the promotion of mineral deposition by Zn and Sr elicited different trends with the culture time. After culture for 21 days, 10 wt% Sr showed significant higher Ca deposits than the control, which was possibly due to the positive osteogenic stimulation of Sr. Fig. 10. SEM images of S. aureus bacterial adhesion on the surfaces of (a) PDA, (b) Ag, (c) Cu02, (d) Cu1, (e) Sr2, (f) Sr10, (g) Zn2 and (h) Zn10 coatings after culture for 24 h. The arrows indicate S. aureus bacteria.

4. Discussion A variety of surface coatings on the Ti implants are designed to inhibit initial adhesion of bacteria to the Ti surfaces because the native Ti surfaces do not prevent bacterial colonization. Antimicrobial agents such as metal ions, instead of antibiotics, can be applied to the Ti surface by simple spin coating. The PDA layer is easily applied to large or complex implant surface and adheres well to the substrate [31,40], which was also used as bonding glue to ions. The ion-coated Ti implants may leach these ions to produce positive antibacterial efficacy and/or biological responses to provide effective use [11,19]. There are few

samples exhibited an increased ALP amount except for the Ag sample; however, the time-dependent trends of the three coatings were different. In the case of Cu and Zn, the greater ion content in the coating resulted in the lower ALP levels. Contrary to this finding, the Sr-containing coatings depicted a positive correlation between ion concentration and ALP amount. The 10 wt% Sr-containing (Sr10) sample had the greatest ALP expression.

Fig. 11. Determination of the ROS production within the (a) E. coli and (b) S. aureus bacteria treated with the different concentrations of metal ions for different reaction time. 9

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

XPS and SIMS results consistently demonstrated the introduction of metal ions into the surface of the PDA-coated Ti substrates, in agreement with the changes in the coating morphology. The higher concentration used in the preparation of coating, the greater characteristic peaks in X-ray diffraction patterns of the ion coatings were found. Regarding the coating thickness, the PDA layer alone had the same thickness with the Ag-coated sample because of an extremely low Ag concentration such as 0.02% used. On the contrary, the addition of Cu, Sr and Zn ions to PDA surfaces made the film thickness to increase from 250 nm to about 350 nm. The similar film thickness of the divalent metal ion coating was possibly due to the same coating volume used during the preparation. On the other hand, it is no doubt that the presence of a PDA layer with and without ion coatings could enhance the corrosion-resistant ability of the Ti implants. Bacterial adhesion is the first step of bacterial colonization, which is linked to the characteristics of microbial cell surfaces, the surface properties of the implant materials, and the biological bathing fluid [41,42]. To ascertain the current ion coatings as promising antibacterial implant materials, the antibacterial ability of the samples should be examined, which is one of the crucial issues in their clinical use. Bacterial viability directly responds to the antibacterial ability of the metal ion. The PDA-coated surface had a little bactericidal effect. In contrast, the coating with 0.02% Ag ions showed high bactericidal efficiency against the two bacterial species. SEM images also clearly showed the inhibition of E. coli and S. aureus growth on the surface of the Ag coating. Not surprisingly, Ag ions were found to have a strong inhibitory and bactericidal effect against a broad spectrum of bacteria [7,13,24]. The antibacterial activity of Ag ions depends on the inhibition of respiratory chain and cell pathways, the generation of ROS promoting oxidative stress, and damage of DNA and RNA [43]. Metal ions, such as Cu and Zn, are used as an alternative to Ag in the coating due to antibacterial effect. This study found that although the concentration of Cu, Sr and Zn ions used was higher than that of the Ag

Fig. 13. RAW 264.7 cell viability after culture on various samples for different time-points. The data as percentages were obtained by calculation related to Ti control.

reports on the relative antibacterial activity of different metal ions at the same concentration or order of magnitude. In this study, we used a simple and versatile method to apply metal ions to PDA-coated surfaces for antibacterial and osteogenic efficacy. Typically, the surface of Ti implant is pre-treated by sandblasting with a large grit and acid etching (SLA) for biomechanical interlocking to allow the bone ongrowth [2]. However, a relatively smooth sample surface was used throughout the study to greatly reduce the interference of the rough substrate and to effectively evaluate the properties of the metal ion coatings. Before elucidating in vitro biological responses, it was necessary to account for the evolution of phase composition and microstructure. The GIXRD,

Fig. 14. (a) Cell attachment, (b) proliferation, (c) ALP activity, and matrix mineralization of hMSCs cultured on various samples surfaces at various culture timepoints. Asterisk represents statistically significant difference (p < 0.05) from Ti. 10

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

osteogenic activity are highly desirable in clinical applications, as schematically outlined in Fig. 15. Increasing investigations have been devoted to loading antibacterial agents on the surface of implants to inhibit bacterial adhesion or to improve osteointegration at the implant sites. Generally, it is difficult to simultaneously achieve the requirements of inhibiting bacteria adhesion and promoting osteogenic activity on the implant surfaces [10–12,15]. Ions could effectively eradicate bacteria; however, under certain conditions, it may appear to be detrimental to cellular functions [55,56]. Actually, attention should be paid to the cytotoxicity of metal ions. The cytotoxicity of metal compounds is known to depend on their chemical species and concentration [56]. As an example, Ag ions can cause high antibacterial activity and extreme cytotoxicity even at low concentration (for example, 0.02% in this study). When the concentration of Ag exceeds a certain threshold, Ag has certain toxicity to human cells [56]. The cell function on the surface of the implant is closely related to the physical, chemical and biological characteristics of the materials used [7,57]. To understand the biological function of metal ions, L929 and RAW 264.7 cytotoxicity assay and in vitro osteogenesis of hMSCs cultured on the coating surfaces were evaluated. The results of the three cell cultures consistently indicated that Ag significantly reduced cell viability and osteogenesis. According to ISO 10993-5 standard, a viability of more than 70% is considered to be non-cytotoxic. A lower viability value elucidated a higher cytotoxic potential of the test sample. Ag causes DNA damage and increases the number of apoptotic and necrotic cells [58], which led to higher L929 cytotoxicity than Sr and Zn. Although the essential metal ions such as Cu and Zn are used as cofactors in the cytoplasm, their high concentration can cause cytotoxic reactions [23]. Moreover, the cytotoxicity of Cu and Zn ions is dose dependent [59,60]. In contrast, Sr showed a positive enhancement of L929 viability. The RAW 264.7 cell line is widely used to detect immunological activity and innate immune response to biomaterials because it exhibits a stable and mature adherent macrophage phenotype [61]. As a result, the Ag showed a significant negative impact on the macrophage viability. Three divalent metal ions had a dose-dependent toxic effect on RAW 264.7 cells. Li et al. [62] found Cu inhibited the attachment and activation of macrophages. Moonga and Dempster [63] reported that Zn highly inhibited osteoclastic bone resorption in vitro. Petrochenko et al. [64] mentioned that ZnO coatings produce gradual ion release and were toxic to RAW 264.7 macrophages after a certain level. Sr released from bioactive glass could inhibit RAW264.7 activity in a dosedependent manner [65], confirming the current findings. In the present study, hMSCs were also seeded onto the surface of the coating to examine the in vitro osteogenic activity of all ion-coated implant materials. In fact, there were significant differences between the numbers of cells detected in the ion coating samples. The antibacterial agents such as Ag and Cu on the surface of the implant largely reduced the numbers of cells, while the Sr coating did not exhibit the same trend. In vitro osteogenesis analysis consistently indicated significant higher ALP and calcium content in cells cultured on Zn and Sr coatings than on Ag and Cu coatings, although these coatings had a comparable antibacterial ability. The membrane-bound exoenzyme ALP is involved in the bone-regeneration process at the early stage of osteogenic differentiation, and the ability of cells to produce mineralized matrices in materials is important for bone regeneration [28]. On the other hand, there was a linear positive correlation between the released Sr concentration and osteoblastic function, while Zn elucidated a negative correlation. It is believed that an appropriate concentration of Zn ions is a factor that promotes the function of osteoblast and has a partly antibacterial effect on bacteria without causing undesired side effects [9]. Studies have shown that the influence of Zn ions on osteogenic differentiation of mouse primary bone marrow stromal cells depends on Zn ion concentration and incubation time [66]. Not surprisingly, cell proliferation, differentiation, and matrix mineralization were reduced with higher Zn content. Sr ions are also used to improve the

ion, the use of these divalent metal ions as a coating material can significantly reduce the bacterial viability. The bacterial viability was inversely correlated with the ion concentration gradient, indicating that the antibacterial activity of the three metal ions was dose-dependent, which was consistent with the previous studies [10,44–46]. Studies have showed that, with the increase of Zn concentration, the bacterial adhesion of S. mutans on Zn-implanted Ti surfaces was significantly reduced [45]. Liu et al. [46] reported that the growth inhibition percentage of P. gingivalis and A. actinomycetemcomitans increased with the concentration of SrCl2·6H2O solution. The current findings also corresponded to SEM observation in which the number of rod-shaped E. coli and spherical S. aureus bacteria adherent to these ion coatings was obviously reduced in comparison with those on the PDA-coated surface. Regarding the type of metal ions, there was no significant difference in the antibacterial activity of Sr2+ and Zn2+ coatings at the same concentrations, but the Cu2+ ions could achieve the same ability at a concentration one order of magnitude lower than the other two metal ions. This comparison clearly pointed out that the antibacterial activity of Cu2+ ions was superior to that of Sr2+ or Zn2+ ions. Although the exact mechanism of the antibacterial activity of these divalent metal ions was not fully understood, the biocidal effects were possibly because of damage to their cytoplasmic membrane, proteins denaturation or DNA damage [13,47,48]. For example, copper ions can inactivate proteins by disrupting Fe–S clusters in cytoplasmic hydratases [48]. Zinc ions have been shown to inhibit a variety of bacterial activities, such as transmembrane proton translocation, glycolysis and acid tolerance [49]. In addition, electrostatic attraction between positively charged metal ions and negatively charged bacterial membranes may cause the leakage of proteins and other intracellular constituents [42,44]. However, a complementary study such as ROS was required to elucidate the mechanism by which metal ions acted on the bacteria. ROS include superoxide anions, hydroxyl radicals, singlet oxygen, and hydrogen peroxide and are byproducts of cellular oxidative metabolism [50]. ROS are important for physiological processes, such as intracellular signal transduction, and against pathogens, but excess can impair cell function by destroying lipids, proteins, RNA and DNA [51]. Based on the current results, the trend of intracellular ROS induced by divalent metal ions was quite consistent with the antibacterial ability, indicating that higher ion concentration resulted in higher ROS level, causing the greater antibacterial activity. Therefore, it can be speculated that the metal ions used in this study may have a significant biocidal effects due to the generation of ROS. Moreover, type-dependent ROS levels of metal ions were occurred, which indicated 1% Cu solution produced a ROS value similar to 10%Sr and 10% Zn solutions. The use of ions may intensify the antibacterial activity because it readily penetrates cell membranes and oxidizes cell contents through ROS [52]. Ma et al. [53] reported that Cu(II)-promoted ROS overproduction and mitochondrial disruption might be the main cause of increased cell apoptosis. Applerot et al. [32] advocated that ZnO exhibited antibacterial activity against E. coli and S. aureus due to a significant increase in the oxidative stress originating from the generation of ROS. Antibacterial activity of ZnO nanoparticles was found to be dose-dependent with ROS production [54]. On the other hand, to date, the effect of Sr ions on the antibacterial function has been relatively less noticed, and the possible mechanism by which ROS generation contributed to Sr-induced antibacterial activity has not been directly explored. Brauer et al. [1] demonstrated the use of Sr2+ ions in the SiO2–CaO–CaF2–MgO-based cement because it inhibited bacterial growth and impeded cytoplasmic membrane permeability, cell wall synthesis and bacterial chromosome replication. According to the current ROS results, we believed that the antibacterial activity of Sr ions may be caused by ROS production. The incorporation of Sr ions into the implant materials can allow for the dual action of Sr: inhibiting bacteria and enhancing bone formation, the latter of which will be discussed below. Ti implants with dual functions of antibacterial efficacy and 11

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

Fig. 15. Schematic representation of Sr ion-coated Ti implants with dual functions of antibacterial efficacy and osteogenic activity in this work.

because of the higher release in an acidic environment. It necessitates to exploring the loading of metal ions on coatings to minimize cytotoxicity and maximize bactericidal efficacy. Yamamoto et al. [55] found that the 5 ppm CuCl2, 12 ppm ZnCl2, and 1817 ppm SrCl2 caused L929 cell death to reach 50% (inhibitive concentration; IC50) after 1 day of culture, which meant that the order of cytotoxicity was Cu2+ > Zn2+ > Sr2+. In the current study, the release of three metal ions from the coatings was measured by ICP-AES. After soaking for 1 day in Tris-HCl (pH 5.0), the Cu02 coating released about 6 ppm, which resulted in a 1-day viability of L929 cells of about 50%. The Zn2 and Zn10 coatings released Zn concentration of 70 and 176 ppm, respectively, and L929 viability was about 90% and 67%, respectively, at 1 day of incubation. The difference in the Zn release concentration between the current result and the previous reports [55,59] was attributed to the different soaking environments, such as solution pH. Nevertheless, it is reasonable to observe greater release and cytotoxicity with higher coating concentrations of the Zn coating. Regarding the Sr coating, the highest concentration of Sr10 was 533 ppm after soaking for 3 days, which was much lower than IC50 derived from Yamamoto et al. [55]. It is interesting to note that the release rate of the ion coating was Sr2+ > Zn2+ at the same concentration and soaking time. This result may be interpreted by the chemical binding of metal ions with PDA. Since Zn is a transition metal ion, a strong Zn-PDA coordination bonds can be formed [70], which led to a lower release rate than that obtained with the Sr coating. The current results provided the evidence that 2% Zn coating and 10% Sr coating not only had no toxic effect on cells, but also effectively inhibit bacterial growth of bacteria. More importantly, from the unexpected results of Sr against bacterial species, the combined antimicrobial and osteogenic properties of Sr may be harnessed for a variety of medical applications (Fig. 15). Even a high Sr ion concentration gave rise to a substantial reduction in bacterial growth that was also conducive to cell growth. However, the antibacterial mechanism of Sr ions should be further studied as the potential material was used to prevent microbial adhesion.

osteogenesis of the biomaterials. Indeed, matrix mineralization of hMSCs cultured for 21 days was significantly enhanced on Sr samples, particularly on the Sr10 group. Previous studies have reported that the Sr-containing materials can up-regulate osteoblastic cell viability, ALP activity, and stimulate bone formation [67], as found in our current study. Bonnelye et al. [26] reported that osteoblast differentiation and number of bone nodules were enhanced with only 0.1 mM Sr ions. Therefore, optimization of the Sr content on the surface of the implant could be beneficial to improve the antibacterial efficacy and osteogenic ability of the implant. In summary, a 10 wt% Sr coating may be suitable for surface modification of Ti implant in terms of antibacterial efficacy and osteoblast function compared to the Cu or Zn coatings. The antibacterial material is placed at a site of an infection in order to controllably release antibacterial ions as the material degrades. The duration and effectiveness of the antibacterial efficacy of these coating depends on the loading and release kinetics, in addition to the coating thickness. Current results from quantitative antimicrobial tests demonstrated that all metal ion-coated Ti samples clearly resulted in reduced viable bacteria and different cell responses. It should be recognized that since high doses of metal ions may be toxic, the kinetics of the ion release from coating materials must be tailored, and the metal ion concentration can be easily adjusted by varying the concentration used in the coating process. Actually, as the amount of Cu, Sr and Zn on the PDA-modified Ti surface increased, a statistically significant (p < 0.05) increase in the antimicrobial activity was found. To investigate the release behaviour, the amount of Cu, Sr and Zn released from the ion-incorporated coatings were measured after soaking in TrisHCl at pH 5.0. Due to bacterial induced local metabolic acidosis or tissue inflammation, the bone lesion environment may have varying pH from a neutral pH of 7.4 to an acidic pH as low as 5.0 [68]. The current in vitro release profile clarified the high initial release of ions followed by a lower sustained release profile. Because of the highest risk of infection in the perioperative periods, a high initial release rate of metal ions is required to prevent bacterial colonization [69]. Indeed, the currently used coating techniques ensured ions on the surface of the Ti substrate and accelerated the release of biocidal ions into the surrounding medium. Moreover, rapid degradation would occur under acidic conditions, such as pH 5.0, but slow degradation under neutral or basic conditions, which may be advantageous to inhibit bacteria

5. Conclusions When applied to clinical implantation, it is essential to design new 12

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

materials with bactericidal efficacy and osteogenic activity. This study supported the hypothesis that the release of some antibacterial agents not only inhibited the bacterial attachment, but also it enhanced biologic properties of the implants. From the current results, it seems to be a feasible way to incorporate Zn or Sn ion into PDA coatings by a simple spin coating, which promoted the rapid release of ions. Within the limits of this study, ion-tailored coatings displayed the concentrationdependent bactericidal efficacy. The Cu-, Sr- or Zn-incorporated PDA coatings were highly effective in inhibiting adhesion of E. coli and S. aureus strains, possibly due to the ROS generation. Biological investigation confirmed that high content of Sr promoted the proliferation and differentiation of hMSC, and low content of Zn had similar effects. Taken together, 10 wt% Sr coating on the PDA-immobilized Ti surface can be considered a high potential implant candidate. However, in addition to prolonging release behavior by means of covalent binding of ions to PDA-modified Ti implant and increasing the coating thickness, the underlying mechanism of Sr-induced antibacterial activity needs further investigation.

Med. 9 (1998) 129–134. [21] V. Mouriño, J.P. Cattalini, A.R. Boccaccini, Metallic ions as therapeutic agents in tissue engineering scaffolds: an overview of their biological applications and strategies for new developments, J. R. Soc. Interface 9 (2012) 401–419. [22] S. Eckhardt, P.S. Brunetto, J. Gagnon, M. Priebe, B. Giese, K.M. Fromm, Nanobio silver: its interactions with peptides and bacteria, and its uses in medicine, Chem. Rev. 113 (2013) 4708–4754. [23] L.A. Finney, Transition metal speciation in the cell: insights from the chemistry of metal ion receptors, Science 300 (2003) 931–936. [24] O.M. Goudouri, E. Kontonasaki, U. Lohbauer, A.R. Boccaccini, Antibacterial properties of metal and metalloid ions in chronic periodontitis and peri-implantitis therapy, Acta Biomater. 10 (2014) 3795–3810. [25] B. Jia, Y. Mei, L. Cheng, J. Zhou, L. Zhang, Preparation of copper nanoparticles coated cellulose films with antibacterial properties through one-step reduction, ACS Appl. Mater. Interfaces 4 (2012) 2897–2902. [26] E. Bonnelye, A. Chabadel, F. Saltel, P. Jurdic, Dual effect of strontium ranelate: stimulation of osteoblast differentiation and inhibition of osteoclast formation and resorption in vitro, Bone 42 (2008) 129–138. [27] Y. Okuzu, S. Fujibayashi, S. Yamaguchi, K. Yamamoto, T. Shimizu, T. Sono, K. Goto, B. Otsuki, T. Matsushita, T. Kokubo, S. Matsuda, Strontium and magnesium ions released from bioactive titanium metal promote early bone bonding in a rabbit implant model, Acta Biomater. 63 (2017) 383–392. [28] I.T. Wu, T.Y. Chiang, C.C. Chen, Y.C. Chen, S.J. Ding, Dopant-dependent tailoring of physicochemical and biological properties of calcium silicate bone cements, Bio Med. Mater. Eng. 29 (2018) 773–785. [29] F. Dabsie, G. Gregoire, M. Sixou, P. Sharrock, Does strontium play a role in the cariostaticactivity of glass ionomer? Strontium diffusion andantibacterial activity, J. Dent. 37 (2009) 554–559. [30] K.H. Thompson, C. Orvig, Boon and bane of metal ions in medicine, Science 300 (2003) 936–939. [31] H. Lee, S.M. Dellatore, W.M. Miller, P.B. Messersmith, Mussel-inspired surface chemistry for multifunctional coatings, Science 318 (2007) 426–430. [32] G. Applerot, A. Lipovsky, R. Dror, N. Perkas, Y. Nitzan, R. Lubart, A. Gedanken, Enhanced antibacterial activity of nanocrystalline ZnO due to increased ROSmediated cell injury, Adv. Funct. Mater. 19 (2009) 842–852. [33] J. Dudognon, M. Vayer, A. Pineau, R. Erre, Grazing incidence X-ray diffraction spectra analysis of expanded austenite for implanted stainless steel, Surf. Coat. Technol. 202 (2008) 5048–5054. [34] C. Buga, M. Hunyadi, Z. Gácsi, C. Hegedűs, J. Hakl, U. Schmidt, S.J. Ding, A. Csík, Calcium silicate layer on titanium fabricated by electrospray deposition, Mater. Sci. Eng. C 98 (2019) 401–408. [35] K. Vad, A. Csik, G.A. Langer, Secondary neutral mass spectrometry: a powerful technique for quantitative elemental and depth profiling analyses of nanostructures, Spectrosc. Eur. 21 (2009) 13–16. [36] S.J. Ding, Y.H. Chu, D.Y. Wang, Enhanced properties of novel zirconia-based osteoimplant systems, Appl. Mater. Today 9 (2017) 622–632. [37] M. Oparka, J. Walczak, D. Malinska, L.M.P.E. van Oppen, J. Szczepanowska, W.J.H. Koopman, M.R. Wieckowski, Quantifying ROS levels using CM-H2DCFDA and HyPer, Methods 109 (2016) 3–11. [38] A. Gomes, E. Fernandes, J.L.F.C. Lima, Fluorescence probes used for detection of reactive oxygen species, J. Biochem. Biophys. Methods 65 (2005) 45–80. [39] C. Hegedűs, C.C. Ho, A. Csík, S. Biri, S.J. Ding, Enhanced physicochemical and biological properties of ion-implanted titanium using electron cyclotron resonance ion sources, Materials 9 (2016) 25. [40] C.C. Ho, S.J. Ding, Structure, properties and applications of mussel-inspired polydopamine, J. Biomed. Nanotechnol. 10 (2014) 3063–3084. [41] H. Van de Belt, D. Neut, W. Schenk, J.M. Van Horn, H.C. Van der Mei, H.J. Busscher, Infection of orthopedic implants and the use of antibiotic-loaded bone cements: a review, Acta Orthop. Scand. 72 (2001) 557–571. [42] F. Song, H. Koo, D. Ren, Effects of material properties on bacterial adhesion and biofilm formation, J. Dent. Res. 94 (2015) 1027–1034. [43] W.K. Jung, H.C. Koo, K.W. Kim, S. Shin, S.H. Kim, Y.H. Park, Antibacterial activity and mechanism of action of the silver ion in Staphylococcus aureus and Escherichia coli, Appl. Environ. Microbiol. 74 (2008) 2171–2178. [44] M. Raffi, S. Mehrwan, T.M. Bhatti, J.I. Akhter, A. Hameed, W. Yawar, M.U. Hasan, Investigations into the antibacterial behavior of copper nanoparticles against Escherichia coli, Ann. Microbiol. 60 (2010) 75–80. [45] J. Xu, G. Ding, J. Li, S. Yang, B. Fang, H. Sun, Y. Zhou, Zinc-ion implanted and deposited titanium surfaces reduce adhesion of Streptococccus mutans, Appl. Surf. Sci. 256 (2010) 7540–7544. [46] J. Liu, S.C.F. Rawlinson, R.G. Hill, F. Fortune, Strontium-substituted bioactive glasses in vitro osteogenic and antibacterial effects, Dent. Mater. 32 (2016) 412–422. [47] Y. Ohsumi, K. Kitamoto, Y. Anraku, Changes induced in the permeability barrier of the yeast plasma membrane by cupric ion, J. Bacteriol. 170 (1988) 2676–2682. [48] L. Macomber, J.A. Imlay, The iron-sulfur clusters of dehydratases are primary intracellular targets of copper toxicity, Proc. Natl. Acad. Sci. U.S.A. 106 (2009) 8344–8349. [49] T.N. Phan, T. Buckner, J. Sheng, J.D. Baldeck, R.E. Marquis, Physiologic actions of zinc related to inhibition of acid and alkaliproduction by oral streptococci in suspensions and biofilms, Oral Microbiol. Immunol. 19 (2004) 31–38. [50] P.P. Fu, Q. Xia, H.M. Hwang, P.C. Ray, H. Yu, Mechanisms of nanotoxicity: generation of reactive oxygen species, J. Food Drug Anal. 22 (2014) 64–75. [51] K. Ito, T. Suda, Metabolic requirements for the maintenance of self-renewing stem cells, Nat. Rev. Mol. Cell Biol. 15 (2014) 243–256. [52] T.Y. Wang, M.D.J. Libardo, A.M. Angeles-Boza, J.P. Pellois, Membrane oxidation in

Acknowledgement The work is supported by Ministry of Science and Technology, Taiwan under contract no. MOST 107-2221-E-040-002. References [1] D.S. Brauer, N. Karpukhina, G. Kedia, A. Bhat, R.V. Law, I. Radecka, R.G. Hill, Bactericidal strontium-releasing injectable bone cements based on bioactive glasses, J. R. Soc. Interface 10 (2013) 20120647. [2] C.J. Chen, S.J. Ding, C.C. Chen, Effects of surface conditions of titanium dental implants on bacterial adhesion, Photomed. Laser Surg. 34 (2016) 379–388. [3] T.C. Huang, C.J. Chen, S.J. Ding, C.C. Chen, Antimicrobial efficacy of methylene blue-mediated photodynamic therapy on titanium alloy surfaces in vitro, Photodiagn. Photodyn. Ther. 25 (2019) 7–16. [4] L. Grenho, F.J. Monteiro, M.P. Ferraz, Antibacterial effect of nanoHA-ZnO composites, J. Biomed. Mater. Res. A 102 (2014) 3726–3733. [5] C.A. Fux, J.W. Costerton, P.S. Stewart, P. Stoodley, Survival strategies of infectious biofilms, Trends Microbiol. 13 (2005) 34–40. [6] M. Zlowodzki, J.S. Prakash, N.K. Aggarwal, External fixation of complex femoral shaft fractures, Int. Orthop. 31 (2007) 409–413. [7] C.K. Wei, S.J. Ding, Dual-functional bone implants with antibacterial ability and osteogenic activity, J. Mater. Chem. B 5 (2017) 1943–1953. [8] L. Zhao, P.K. Chu, Y. Zhang, Z. Wu, Antibacterial coatings on titanium implants, J. Biomed. Mater. Res. B 91 (2009) 470–480. [9] G. Jin, H. Cao, Y. Qiao, F. Meng, H. Zhu, X. Liu, Osteogenic activity and antibacterial effect of zinc ion implanted titanium, Colloids Surf., B 117 (2014) 158–165. [10] S. Kalaivani, R.K. Singh, V. Ganesan, S. Kannan, Effect of copper (Cu2+) inclusion on the bioactivity and antibacterial behavior of calcium silicate coatings on titanium metal, J. Mater. Chem. B 2 (2014) 846–858. [11] M. Li, Q. Liu, Z. Ji, X. Xu, Y. Shi, Y. Cheng, Y. Zheng, Polydopamine-induced nanocomposite Ag/CaP coatings on the surface of titania nanotubes for antibacterial and osteointegration functions, J. Mater. Chem. B 3 (2015) 8796–8805. [12] Y. Zhao, H. Cao, H. Qin, T. Cheng, S. Qian, M. Cheng, X. Peng, J. Wang, Y. Zhang, G. Jin, X. Zhang, X. Liu, P. K Chu, Balancing the osteogenic and antibacterial properties of titanium by codoping of Mg and Ag: an in vitro and in vivo study, ACS Appl. Mater. Interfaces 7 (2015) 17826–17836. [13] F. Bir, H. Khireddine, A. Touati, D. Sidane, S. Yala, H. Oudadesse, Electrochemical depositions of fluorohydroxyapatite doped by Cu2+, Zn2+, Ag+ on stainless steel substrates, Appl. Surf. Sci. 258 (2012) 7021–7030. [14] Q. Wu, J. Li, W. Zhang, H. Qian, W. She, H. Pan, J. Wen, X. Zhang, X. Liu, X. Jiang, Antibacterial property, angiogenic and osteogenic activity of Cu-incorporated TiO2 coating, J. Mater. Chem. B 2 (2014) 6738–6748. [15] C.C. Ho, S.J. Ding, Novel SiO2/PDA hybrid coatings to promote osteoblast-like cell expression on titanium implants, J. Mater. Chem. B 3 (2015) 2698–2707. [16] S.B. Goodman, Z. Yao, M. Keeney, F. Yang, The future of biologic coatings for orthopaedic implants, Biomaterials 34 (2013) 3174–3183. [17] J. Raphel, M. Holodniy, S.B. Goodman, S.C. Heilshorn, Multifunctional coatings to simultaneously promote osseointegration and prevent infection of orthopaedic implants, Biomaterials 84 (2016) 301–314. [18] J. Qiu, L. Liu, B. Chen, Y. Qiao, H. Cao, H. Zhu, X. Liu, Graphene oxide as a dual Zn/ Mg ion carrier and release platform: enhanced osteogenic activity and antibacterial properties, J. Mater. Chem. B 6 (2018) 2004–2012. [19] G. Jin, H. Qin, H. Cao, S. Qian, Y. Zhao, X. Peng, X. Zhang, X. Liu, P.K. Chu, Synergistic effects of dual Zn/Ag ion implantation in osteogenic activity and antibacterial ability of titanium, Biomaterials 35 (2014) 7699–7713. [20] T.N. Kim, Q.L. Feng, J.O. Kim, J. Wu, H. Wang, G.C. Chen, F.Z. Cui, Antimicrobial effects of metal ions (Ag+, Cu2+, Zn2+) in hydroxyapatite, J. Mater. Sci. Mater.

13

Surface & Coatings Technology xxx (xxxx) xxxx

H. Kao, et al.

cell delivery and cell killing applications, ACS Chem. Biol. 12 (2017) 1170–1182. [53] L. Ma, X. Li, Y. Wang, W. Zheng, T. Chen, Cu(II) inhibits hIAPP fibrillation and promotes hIAPP-induced beta cell apoptosis through induction of ROS-mediated mitochondrial dysfunction, J. Inorg. Biochem. 140 (2014) 143–152. [54] M. Shoeb, B.R. Singh, J.A. Khan, W. Khan, B.N. Singh, H.B. Singh, A.H. Naqvi, ROSdependent anticandidal activity of zinc oxide nanoparticles synthesized by using egg albumen as a biotemplate, Adv. Nat. Sci. Nanosci. Nanotechnol. 4 (2013) 035015. [55] A. Yamamoto, R. Honma, M. Sumita, Cytotoxicity evaluation of 43 metal salts using murine fibroblasts and osteoblastic cells, J. Biomed. Mater. Res. 39 (1998) 331–340. [56] F. Heidenau, W. Mittelmeier, R. Detsch, M. Haenle, F. Stenzel, G. Ziegler, H. Gollwitzer, A novel antibacterial titania coating: metal ion toxicity and in vitro surface colonization, J. Mater. Sci. Mater. Med. 16 (2005) 883–888. [57] T.R. Su, Y.H. Chu, H.W. Yang, Y.F. Huang, S.J. Ding, Component effects of bioactive glass on corrosion resistance and in vitro biological properties of apatite-matrix coatings, Bio Med. Mater. Eng. 30 (2019) 207–218. [58] J. Jiravova, K.B. Tomankova, M. Harvanova, L. Malin, J. Malohlava, L. Luhova, A. Panacek, B. Manisova, H. Kolarova, The effect of silver nanoparticles and silver ions on mammalian and plant cells in vitro, Food Chem. Toxicol. 96 (2016) 50–61. [59] W. Song, J. Zhang, J. Guo, J. Zhang, F. Ding, L. Li, Z. Sun, Role of the dissolved zinc ion and reactive oxygen species in cytotoxicity of ZnO nanoparticles, Toxicol. Lett. 199 (2010) 389–397. [60] B. Cao, Y. Zheng, T. Xi, C. Zhang, W. Song, K. Burugapalli, H. Yang, Y. Ma, Concentration-dependent cytotoxicity of copper ions on mouse fibroblasts in vitro: effects of copper ion release from TCu380A vs TCu220C intra-uterine devices, Biomed, Microdevices 4 (2012) 709–720. [61] D.M. Higgins, R.J. Basaraba, A.C. Hohnbaum, E.J. Lee, D.W. Grainger, M. Gonzalez-

[62]

[63] [64]

[65]

[66]

[67]

[68]

[69] [70]

14

Juarrero, Localized immunosuppressive environment in the foreign body response to implanted biomaterials, Am. J. Pathol. 175 (2009) 161–170. L. Li, Y. Xu, Z. Zhou, J. Chen, P. Yang, Y. Yang, J. Li, N. Huang, The effects of Cudoped TiO2 thin films on hyperplasia, inflammation and bacterial infection, Appl. Sci. 5 (2015) 1016–1032. B.S. Moonga, D.W. Dempster, Zinc is a potent inhibitor of osteoclastic bone resorption in vitro, J. Bone Miner. Res. 10 (1995) 453–457. P.E. Petrochenko, Q. Zhang, R. Bayati, S.A. Skoog, K.S. Phillips, G. Kumar, R.J. Narayan, P.L. Goering, Cytotoxic evaluation of nanostructured zinc oxide (ZnO) thin films and leachates, Toxicol. In Vitro 28 (2014) 1144–1152. E. Gentleman, Y.C. Fredholm, G. Jell, N. Lotfibakhshaiesh, M.D. O'Donnell, R.G. Hill, M.M. Stevens, The effects of strontium-substituted bioactive glasses on osteoblasts and osteoclasts in vitro, Biomaterials 31 (2010) 3949–3956. T. Wang, J.C. Zhang, Y. Chen, P.G. Xiao, M.S. Yang, Effect of zinc ion on the osteogenic and adipogenic differentiation of mouse primary bone marrow stromal cells and the adipocytic trans-differentiation of mouse primary osteoblasts, J. Trace Elem. Med. Biol. 21 (2007) 84–91. K. Lin, X. Wang, N. Zhang, Y. Shen, Strontium (Sr) strengthens the silicon (Si) upon osteoblast proliferation, osteogenic differentiation and angiogenic factor expression, J. Mater. Chem. B 4 (2016) 3632–3638. C.K. Wei, S.J. Ding, Acid-resistant calcium silicate-based composite implants with high-strength as load-bearing bone graft substitutes and fracture fixation devices, J. Mech. Behav. Biomed. Mater. 62 (2016) 366–383. S. Jaiswal, P. McHale, B. Duffy, Preparation and rapid analysis of antibacterial silver, copper and zinc doped sol–gel surfaces, Colloids Surf., B 94 (2012) 170–176. H. Wu, J.M. Ang, J. Kong, C. Zhao, Y. Dud, X. Lu, One-pot synthesis of polydopamine–Zn complex antifouling coatings on membranes for ultrafiltration under harsh conditions, RSC Adv. 6 (2016) 103390–103398.