Antibacterial and biological characteristics of plasma sprayed silver and strontium doped hydroxyapatite coatings

Abstract

Infection in primary total joint prostheses is estimated to occur in up to 3% of all surgeries. As a measure to improve the antimicrobial properties of implant materials, silver (Ag) was incorporated into plasma sprayed hydroxyapatite (HA) coatings. To offset potential cytotoxic effects of Ag in the coatings, strontium (Sr) was also added as a binary dopant. HA powder were doped with 2.0 wt% Ag₂O, 1.0 wt% SrO and the powder was then heat treated at 800° C. Titanium substrates were coated using a 30 kW plasma spray system equipped with a supersonic nozzle. X-ray diffraction (XRD) confirmed the phase purity and high crystallinity of the coatings. Samples were evaluated for mechanical stability by adhesive bond strength testing. Results show that the addition of dopants did not affect the overall bond strength of the coatings. The antibacterial efficacies of the coatings were tested against Pseudomonas aeruginosa. Samples that contained the Ag₂O dopant were found to be highly effective against the bacterial colonization. In vitro cell-material interactions using human fetal osteoblast (hFOB) cells were characterized by 3-(4,5-Dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assay for cell viability, field emission scanning electron microscopy (FESEM) for cell morphology and confocal imaging for the important differentiation marker alkaline phosphatase (ALP). Our results showed evidence of cytotoxic effects in the Ag-HA coatings, characterized by poor cellular morphology and cell death and nearly complete impediment of functional ALP activity. The addition of SrO to Ag-HA coatings was able to effectively offset these negative effects and improve the performance when compared to pure HA coated samples.

1. Introduction

Hydroxyapatite (HA) has been used commercially as a coating on metallic implants since the 1980’s. It has excellent biocompatibility due to its compositional similarity to natural bone and exhibits surface chemistry that supports bone in growth[1–3]. Currently in the orthopedic setting, HA coatings are used in hip and knee applications as an alternative to cemented implants or uncoated press fit implants. Cemented implants are usually recommended for patients that are less likely to put stresses on the cement that could lead to fatigue fractures, such as older patients or younger patients with compromised bone health. Uncemented, porous coated implants are generally recommended for patients that are likely to lead more active lifestyles. The use of HA as a coating material has been highly criticized in the past for fear of two possible failure mechanisms: 1) Delamination of the coating causing aseptic loosening of the implant and 2) Due to the natural dissolution process of HA, free particles or grit may become a 3^rd party wear accelerant between the femoral head component and the acetabular cup component of the implant. Because of the relative newness of HA coated implants in clinical use, advocates of the coating have been hard pressed to allay the concerns of skeptics until recently. Several long term clinical follow-ups have shown that implants that have been coated with HA perform just as well or outperform their cemented or cementless uncoated counterparts with 98% survival after 10 years and an estimated 96% stem survival at 20 years [4–7]. Of several extensively researched coating methods[8], plasma spray deposition is regarded as the most efficient and economical and is the only method currently used in industry.

The current surgical strategy for preventing infection is to minimize contamination during surgery and to administer peri-operative antibiotic prophylaxis. The estimated risk for infection of an implant after total joint replacement surgery is fairly low, 0.5% – 5%, but the consequences are very serious [9]. For severe infections the standard protocols include implant removal, surgical debridement and long-term treatment with full spectrum antibiotics. It has been estimated that the treatment costs for a single occurrence of infection exceeds $50,000 [10]. Because infections can occur from many different causes and can happen at several stages of the implant lifetime[11], preventative treatment such as peri-operative prophylaxis and decontamination procedures may be considered an incomplete regimen. There is a significant need for a solution that can fight infection over time locally. Recent studies have demonstrated that antibiotics may be easily loaded as a coating on implant surfaces with great short term effects [12–14]. There is much left to be desired, however. The release kinetics are not able to satisfy long term requirements. Also, because antibiotics are fragile compounds, there is concern of decomposition during implant sterilization. Antibiotics would have to be loaded in a sterile environment just before surgery. Furthermore, there are concerns over the efficacy of antibiotics in the face of the growing problem of antibiotic resistant bacterial strains. Due to these inadequacies, researchers have recently begun to investigate the antimicrobial effects of cations incorporated into implant materials.

When compared to other heavy metal ions, silver has demonstrated a high affinity for antimicrobial activity, while maintaining a relatively low cytotoxicity [15–17]. The silver ion is a highly active ion that binds strongly to electron donor groups containing sulfur, oxygen or nitrogen. In the case of biological molecules, components such as thio-, amino-, imidazole-, carboxylate- and phosphate groups contain these electron donors [18]. Studies have shown that Ag⁺ ions are able to penetrate through the bacterial cell wall and cause DNA to transform into a condensed form which reacts with the thiol group proteins and results in cell death[19]. It was also found that silver ions are able to interfere with the replication process [20]. It has been demonstrated that the higher levels of silver incorporated into a material, the better the antimicrobial effect is, but it comes at the cost of increasing cytotoxicity [21]. It is, therefore, a good idea to incorporate a secondary chemical to alleviate potential negative effects while maintaining optimal antimicrobial properties of Ag.

Metal ion substitution into calcium phosphate (CaP) materials has been studied by our lab and others as a way to influence the mechanical and biological properties. Different combinations of dopants as well as varying concentrations can have significant effects on the overall mechanical strength, strength degradation profiles and cell materials interactions. [22–25]. Strontium is a non-essential element that has bone seeking behavior. The mechanism by which strontium acts is by way of displacing Ca²⁺ ions in osteoblastic calcium-mediated processes. More specifically, researchers have identified that strontium stimulates bone formation by activating the calcium sensing receptor [26,27], while inhibiting bone resorption by increasing osteoprotegerin (OPG) and preventing receptor activator of nuclear factor kappa B ligand (RANKL) expression [28]. OPG is a protein produced by osteoblasts that inhibits RANKL induced osteoclastogenesis by operating as a decoy receptor for RANKL [29]. The OPG/RANKL ratio, then, is an important balance for regulating bone resorption and osteoclastogenesis.

In this study we have incorporated Ag₂O and SrO into plasma sprayed hydroxyapatite coatings for improved antimicrobial activity and cell material interactions. The Ag₂O and SrO doped HA coatings were prepared using a 30 kW inductively coupled RF induction plasma spray equipped with supersonic plasma nozzle. The coatings were characterized for phase purity and mechanical stability. Antimicrobial effectiveness of HA, Sr-HA, Ag-HA and Sr/Ag-HA coatings were determined using Pseudomonas aeruginosa. This study also investigates the cytocompatibility of the coatings using human fetal osteoblast cells (hFOB).

2. Experimental

2.1 Coating preparation

HA powder of 150–212 μm was used to coat 2-mm thick grade 2 commercially pure titanium (Cp-Ti) substrate (President Titanium, MA). Before coating, the substrates were sandblasted, ultrasonically cleaned in de-ionized water, and then cleaned with acetone to remove the organic materials. Four compositions of HA were used in the present study: HA, HA doped with 2 wt % Ag₂O (Ag-HA), 1 wt % SrO (Sr-HA) and a combination of 2 wt % Ag₂O and 1 wt %SrO (Sr/Ag – HA). The powders were prepared by mixing HA with silver oxide (Ag₂O) and/or strontium oxide (SrO) in 250 mL polypropylene Nalgene bottles containing 75 mL of anhydrous ethanol and 100 g of 5 mm diameter zirconia milling media. The mixtures were then milled for 6 h at 70 rpm followed by drying in an oven at 60° C for 72 h. Ball milled powder was heat treated at 800° C for 6 h to maintain precursor powder size characteristics. A 30 kW inductively coupled RF plasma spray system (Tekna Plasma Systems, Canada), equipped with axial powder feeding system was used for the coating preparation. In this study, coatings were prepared at 25 kW plate power and at 110 mm working distance using supersonic plasma nozzle. Argon (Ar) was used as the central, and carrier gas with flow rate of 25 and 10 standard liters per minute (slpm). A mixture of 60 slpm Ar and 6 slpm hydrogen (H₂) was used as sheath gas. The chamber pressure was maintained at 5 pound-force per square inch gauge (psig). The parameters were selected from previous optimization study where details of the supersonic plasma nozzle were explained [30]. In the supersonic plasma nozzle, HA particles are introduced in the lower region of the plasma torch with a velocity of 510 m/s, much higher than that of the conventionally used plasma nozzle. Particles introduced with a traditional plasma nozzle reside in plasma for 5 ms, while particles introduced into the plasma with a supersonic nozzle reside for 290 μs. The supersonic plasma nozzle employment, then, results in reduced heat treatment of HA particles compared to conventional plasma nozzle.

2.2 Phase analysis

Siemens D500 Krystalloflex X-ray diffractometer using Cu Kα radiation at 35 kV and 30 mA at room temperature was used to determine different phases in the coating with a Ni-filter over the 2θ range between 20° and 60°, at a step size of 0.02° and a count time of 0.5 seconds per step.

2.3 Mechanical properties

The bond strength of the as sprayed HA coatings was evaluated using a standard tensile adhesion test (ASTM C633), where 5 replicates were used [26]. The counter Ti substrate was sand blasted and attached to the surface of the HA coating using quick set epoxy resin as an adhesive glue. After curing in an oven at 120 °C for 2h, the fixtures were subjected to a tensile test at a constant cross head speed of 0.0013 cm/sec until failure. The adhesive bond strength was calculated as: failure load/ sample area (A=5.06 cm2). The data were reported as mean ± standard deviation.

2.4 Silver ion release

In order to determine the silver ion release from the coatings, samples were immersed in phosphate buffered saline (PBS) at pH 7.4. The experiments were performed in an oven maintained at a temperature of 36.5± 0.5 °C and subjected to a constant shaking at 150 rpm. 5 ml of media was collected at each time point starting at 0.5 h. The dissolution tubes were refilled with PBS at each time. The collected media was analyzed for Ag⁺ concentration using a Shimadzu AA-6800 Atomic Absorption Spectrophotometer (Shimadzu, Kyoto, Japan). Ag⁺ and ionization buffer standards were purchased from High-Purity Standards (Charleston, SC, USA). The media was diluted to 1:100 using deionized water and 10 ml of the solution was transferred to a cylindrical tube for the measurement.

2.5 In vitro bone cell-materials interaction

All samples were sterilized by autoclaving at 121 °C for 20 min. In this study established human osteoblast cell line hFOB 1.19 (ATCC, Manassas, VA) were used. Cells were seeded onto the samples in 24-well plates at a density of 10⁴ cells/mL. The base medium for this cell line was a 1:1 mixture of Ham’s F12 Medium and Dulbecco’s Modified Eagle’s Medium (DMEM/F12, Sigma, St. Louis, MO), with 2.5 mM L-glutamine (without phenol red). The medium was supplemented with 10% fetal bovine serum (HyClone, Logan, UT) and 0.3 mg/ml G418 (Sigma, St. Louis, MO). Cultures were maintained at 34 °C under an atmosphere of 5% CO₂ as recommended by ATCC for this particular cell line. Medium was changed every 2 days for the duration of the experiment.

2.5.1 Cell morphology

Samples for testing were removed from culture after 3 days of incubation. All samples for SEM observation were fixed with 2% paraformaldehyde/2% glutaraldehyde in 0.1M cacodylate buffer overnight at 4 °C. Post-fixation was performed with 2% osmium tetroxide (OsO₄) for 2 h at room temperature. The fixed samples were then dehydrated in an ethanol series (30%, 50%, 70%, 95% and 100% three times), followed by a hexamethyldisilane (HMDS) drying procedure. After gold coating, the samples were observed under field emission scanning electron microscope (FESEM) (FEI 200F, FEI Inc., OR, USA) for cell morphologies.

2.5.2 MTT assay

MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide) assay was used to evaluate cell proliferation. The MTT (Sigma, St. Louis, MO) solution of 5 mg/ml was prepared by dissolving MTT in sterile filtered PBS. 10% MTT solution was then added to each sample in 24-well plates. After 2 h of incubation, 1 ml of solubilization solution made up of 10% Triton X-100, 0.1N HCl and isopropanol was added to dissolve the formazan crystals. 100 μl of the resulting supernatant was transferred into a 96-well plate, and read by a plate reader at 570 nm. Statistical analysis was performed using a one way ANOVA and p<0.05 was considered statistically significant. Triplicate samples were used in MTT assay experiments to insure reproducibility. Data are presented as mean ± standard deviation.

2.5.3 Immunohistochemistry and confocal microscopy

Samples were fixed in 3.7% paraformaldehyde/ phosphate buffered solution, pH 7.4 and at room temperature for 10 min. Samples were then washed in PBS 3 times (5 min each) and cells were permeabilized with 0.1% Triton X-100 (in PBS) for 4 min at room temperature. Next, samples were rinsed in TBST3 times (5 minutes each) and incubated in TBST-BSA (Tris-buffered saline with 1% bovine serum albumin, 250 mMNaCl, pH 8.3) blocking solution for 1h at room temperature. Primary antibody against alkaline phosphate (ALP) (Sigma, St. Louis, MO) was added at a 1:100 dilution and incubated at room temperature for 2 h and kept at 4 °C overnight. Samples were then washed with TBST 3 times (10 min each). The secondary antibody, goat anti-mouse (GAM) Oregon green (Molecular Probes, Eugene, OR), was diluted 1:100 in TBST and was used to incubate the cells for 1h. After rinsing three times for 10 minutes each with TBST, the samples were then mounted on glass coverslips with Vectashield mounting medium (Vector Labs, Burlingame, CA) with propidium iodide (PI) and kept at 4 °C for future confocal laser scanning microscope (CLSM) imaging. Micrographs were then taken on a Zeiss 510 laser scanning microscope (LSM 510 META, Carl Zeiss MicroImaging, Inc., NY, USA).

2.6 Antimicrobial activity

Antimicrobial activity of doped HA coated samples was determined by challenging with Pseudomonas aeruginosa. Before bacterial culture, the coated samples were sterilized by autoclaving at 121 °C for 20 min. The bacterial stock solution was prepared by overnight growth of Pseudomonas aeruginosa in tryptic soy broth (TSB) at room temperature with constant stirring. The bacterial cell density was measured by optical methods at 530 nm and a concentration of 105 cells/ml was used. All the samples were vertically placed in a specially designed bioreactor with 0.22 μm filtered air ventilation. A total of 500 ml TSB culture media was poured in the bioreactor with appropriate amount of bacteria. The reactor was kept at 37 °C with constant stirring. After 24 h, each sample was taken out of the bioreactor and washed with PBS. In order to determine the viable and adherent bacteria on the sample surfaces, the cells were stained with LIVE/DEAD Backlight kit (Invitrogen, USA) for 15 min in dark and immediately observed under Zeiss 510 confocal laser scanning microscope.

3. Results

3.1 Physico-chemical analysis

X-ray diffraction patterns of the pure and doped HA coatings are given in Figure 1. The major phase in all samples was identified as HA (JCPDS no. 09-0432). Secondary phases such as alpha-tricalcium phosphate (α-TCP, JCPDS no. 09-0348) and beta-tricalcium phosphate (β-TCP JCPDS no. 09-0169) were also found in the coatings to some small extent. Phases that are undesired, such as tetratricalcium phosphate (TTCP) and calcium oxide (CaO) were absent from the results. In samples that contained Ag₂O, a characteristic Ag₂O (101) peak was observed (JCPDS no. 01-1041) as well as peak shifts, indicating effective incorporation of Ag₂O into the apatite structure. Visual inspection also confirmed the presence of Ag₂O as the color of the coating was brown. Quantitative analysis of the XRD data, completed according to ASTM F2024, showed that the fraction of α/β TCP was less than 2% of the coating in all samples.

Figure 1.

X-ray diffraction spectra of HA and Sr/Ag doped HA.

3.2 Mechanical properties

The tensile adhesive bond strength of coatings was determined according to ASTM standard C633. Adhesive strength (Figure 2) of HA, Sr-HA, Sr/Ag-HA and Ag-HA were found to be 17.4 ± 2.9, 18.21 ± 4.8, 16.3 ± 0.26, and 16.94 ± 0.41 MPa. All of the coatings demonstrated cohesive failure, with the exception of the Ag-HA samples. They exhibited mixed modes of failure with evidence for both adhesive and cohesive, but the strength was not significantly affected. Adhesive failure is described as a failure at the coating-substrate interface, while cohesive failure is a failure within the coating. Because no sample exhibited pure adhesive failure mode, it is indicated that the bonding of the coating to the substrate was strong in all samples.

Figure 2.

Adhesive strength of coatings to substrate.

3.3 Silver ion release

Figure 3 shows the cumulative Ag⁺ release over the course of seven days from the Ag-HA and Sr/Ag-HA coatings in PBS at 37 ° C. There was no significant difference between the two samples at any time point, which asserts that the addition of Sr to the Ag doped coating did not have a large effect on the degradation kinetics of the coating. The release kinetics for both samples were slightly slower over the course of the first 12h, but reached near steady state afterwards, indicating a long term sustainable release.

Figure 3.

Cumulative Ag⁺ ion release in PBS as a function of time.

3.4 In vitro cell-materials interactions

3.4.1 MTT assay

Cellular proliferation of the hFOB cell line, as determined by the MTT assay, is shown in Figure 4. At day 3 and day 11, the Sr-HA and Sr/Ag-HA samples had significantly higher cell density than the Ag-HA samples. Although not significant at the α = 0.05 level, there was a clear trend that the Ag-HA samples were not allowing proliferation at the rates of all other samples and the Sr-HA and Sr/Ag-HA samples were facilitating proliferation better than the pure HA samples. All samples showed increase in cell density between days 3 and 11, with samples containing Ag to a lesser extent, indicating possible cytotoxic effects on the osteoblast cells.

Figure 4.

Optical density measurements for hFOB cell proliferation on HA, Sr-HA, Sr/Ag-HA and Ag-HA coated samples after 3, 7 and 11 days of culture. (*, p < 0.05, n=3)

3.4.2 Cellular morphology

Cellular attachment and growth of the hFOB cells on the pure and doped HA coatings were analyzed by FESEM and are shown in Figure 5. The HA, Sr-HA and Sr/Ag-HA all show normal signs of healthy cell growth. The cells are elongated with many filopodia and pseudopodia extensions. The Ag-HA samples do show signs of poor attachment qualities and maturation denoted by their round and stunted shape. There is also evidence of cell death or premature apoptosis revealing more evidence of possible cytotoxic effects.

Figure 5.

FESEM images depicting hFOB cell morphology after 3 days in culture (a) HA, (b) Sr-HA, (c) Ag-HA, (d) Sr/Ag-HA

3.4.3 Alkaline phosphatase activity

ALP is a key enzyme that is expressed during the differentiation process of osteoblast cells. Expression of ALP was evaluated for hFOB cells cultured on the HA, Sr-HA, Sr/Ag-HA and Ag-HA. Results are shown in Figure 6. The nuclei of the cells fluoresce a blue color, while green indicates ALP expression. At day 5, the Sr-HA and Sr/Ag-HA samples showed the most activity while the pure HA coating showed moderate activity. The Ag-HA samples showed almost no activity, indicating a negative effect of Ag on differentiation of the osteoblast cells. At day 11, the trends were similar with the exception of a slight decline in activity in all samples evidencing idle differentiation activity.

Figure 6.

Confocal micrographs showing ALP expression in hFOB cells at 5 days and 11 days. Green fluorescence indicates active ALP; blue fluorescence indicates PI bound to nucleus of cells.

3.5 Antimicrobial activity

The effectiveness of the coatings against bacterial propagation was measured by live/dead fluorescent staining and is shown in Figure 7. The green fluorescence indicates live bacterial cells or colonies, while the red depicts dead bacteria. After 24 hours of culture, the HA and Sr-HA samples had large quantities of live bacteria present and a small number of dead as shown in Figure 7a and Figure 7b. Conversely, the Sr/Ag-HA and Ag-HA samples had large quantities of dead bacteria present and a very small amount of live cells/colonies, suggesting highly effective bactericidal properties in the samples containing Ag₂O.

Figure 7.

Live/dead confocal images of bacteria after 24 hours of exposure to (a) HA, (b) Sr-HA, (c) Ag-HA and (d) Sr/Ag-HA. Dead bacteria appear red, while live bacteria appear green.

4. Discussion

One of the major challenges in plasma sprayed HA coatings is the retention of phase purity and crystallinity. Decomposition of HA and the formation of amorphous phases are a common occurrence in traditional plasma spray systems due to the high temperature of the plasma arc and the rapid cooling on the relatively cool substrate. The melt ratio (the ratio of the mass of not melted portion of a HA particle to the melted portion) is a key parameter in retaining phase purity and crystallinity. A higher melting ratio usually results in greater amounts of decomposed phases and amorphous phases. Research has shown that with increasing amounts of decomposed HA and amorphousness, the less stable HA is considered [31,32]. The plasma machine in this study was equipped with a supersonic nozzle. The HA particles are discharged at a high speed, 510 m/s, in the lower region of the plasma arc where temperatures are relatively low and therefore, phase decomposition and amorphous phase formation is restricted due to reduced exposure to the plasma flame. Our previous studies with this nozzle have demonstrated the ability to retain high phase purity and crystallinity[25,30,33]. The XRD data, shown in Figure 1, indicates very little phase decomposition and amorphous phase formation in all samples. This is evident by the sharp and distinct HA peaks. The addition of SrO, Ag₂O and the combination of SrO/ Ag₂O had insignificant effects on the physical properties of the coatings. In naturally occurring HA (biological apatite), the calcium ion is highly substituted with trace cations. Research has shown that HA is an ideal platform for cationic substitution as the addition of dopants to HA may alter basic physico-chemical properties such as long rang order and crystallite mean size, while maintaining the apatite nature of the structure [34]. Samples doped with Ag₂O and the combination of SrO/ Ag₂O demonstrated peak shifts in the XRD analysis indicating significant incorporation of the dopants into the HA lattice structure. While no peak shift is observed for SrO doped HA, it is still likely that Sr was incorporated into the HA lattice structure as no separate SrO phase peaks were detected in the diffractogram. Mole ratios of Sr:HA, Ag:HA and Sr/Ag:HA were 1:18, 1:11 and 1:6.45 respectively. Samples that contained Ag₂O and the combination of SrO/Ag₂O had nearly two and three times higher ratios of dopant:HA than SrO alone and could explain why peak shifts were observed only in these samples. Also, the ionic radii of Ca²⁺, Ag⁺ and Sr²⁺ ions are 114 pm, 114–116 pm and 132 pm respectively, for their lowest co-ordination numbers. Thus, it is much easier for Ag⁺ to replace Ca²⁺ in the HA lattice structure. Powders were pretreated to ensure comparable particle size before coating the substrate materials, therefore both the doped and pure HA particles were exposed to similar heat treatments in the plasma. Because of the comparable heat treatments and stability of ionic substitutions, it was expected that the coatings displayed similar behavior in terms of phase decomposition and amorphous phase formation. Since phase decomposition into α-TCP and β-TCP was less than 2% of the entire coating, the coating is also expected to retain its mechanical properties for long periods of time. α-TCP and β-TCP degrade quicker than HA, therefore larger percentages in the coating may result in the formation voids which lead to long term instability of the coating.

The adhesive strengths of the coatings (Figure 2) showed no significant difference with the incorporation of SrO, Ag₂O or SrO/ Ag₂O. Visual analysis of the failure modes of the samples indicated a strong adhesive bond to the substrate, where the common failure mode was cohesive. All samples showed adhesive strength with an average above 16 MPa which is a requirement for these coatings. It was not expected that the addition of dopants would affect the bonding of the coating to the substrate, as the coating method is more mechanical in nature than chemical [30]. Cohesive failure mode is the preferable failure mode (as opposed to adhesive failure mode) because it can effectively demonstrate that the bond strength of the HA or doped HA coating to the Ti substrate is stronger than the tensile strength of the HA material itself. In other words, the HA material will likely fail before delamination of the coating from the implant can occur.

Two of the most important factors in utilizing Ag as an antimicrobial agent in coatings are the ability to have sustained action for long term effects while minimizing the potential cytotoxic effects of the cation. In our previous work, we have reported that the presence of Ag on the surface of LENS^™ processed TCP coatings as well as electrophoretically deposited Ag onto stainless steel orthopedic nails significantly enhanced the antimicrobial activity, but did not have long term efficacy [35,36]. In this study, the antimicrobial Ag is incorporated directly into the HA, rather than just a surface modification. Our group and others have shown that the addition of dopants to CaP materials can directly affect the degradation profiles of the ceramic and, in turn, the ionic release kinetics of the dopants[15,23,37,38]. Previously, the concentration of silver incorporated into HA was optimized to meet the minimum inhibitory concentrations (MIC) needed for the antimicrobial activity, while mitigating the risk of cytotoxic effects [39]. The Ag₂O concentration was found to be optimum at 2%. With the addition of Sr as a binary dopant in the system, it was important to investigate the possibility that the Ag⁺ release could be affected. The results from AAS analysis (Figure 3) show that Sr did not have any noticeable effect on the release of Ag⁺ from the samples. Furthermore, the release is consistent with a controlled long term profile. HA not only degrades very slowly over time, but it is also broken down in the remodeling process by osteoclasts. Because the Ag is incorporated directly into the apatite structure and coating, it is predicted that the release kinetics should remain consistent over an extended period of time making it an effective process for controlling late stage bacteriological colonization. It is important to note, however, that the Ag release rate from the coating is quite slow. Ag is not known to play a major role in any metabolic pathways, and there is always concern of build-up and potential damage to surrounding tissues. At the level of release measured, it is likely that the Ag will be excreted from the body naturally without concern for widespread toxic effect [40], while still maintaining a MIC at the microenvironment level making biofilm formation on the implant surface difficult. A further and more in depth long term study is required to more fully understand the release kinetics and the

Although Ag as an antimicrobial agent has been used extensively in clinical settings, there are still concerns of cytocompatibility [41–43]. MTT results (Figure 4) show an insignificant negative effect on cellular proliferation on samples incorporating Ag alone when compared to pure HA coated samples. Samples that incorporated SrO (Sr-HA and Sr/Ag-HA) showed significantly greater cellular viability at days 3 and 11 when compared to Ag-HA samples indicating a strong positive effect on hFOB cells due to Sr. These results are consistent with other research. It was shown that Sr may directly interact with the calcium sensing receptor in osteoblast cells to trigger mitogenic signals in the protein kinase c/d signaling pathways resulting in increased cell division activity [44]. FESEM micrographs taken after 3 days of culture (Figure 5) showed morphological characteristics of the osteoblast cells. Cells on the surface of pure HA, Sr-HA and Sr/Ag HA show excellent health, while some cells on Ag-HA samples showed dysfunctional characteristics such as premature apoptosis, delayed differentiation characterized by spherical shape and cell death. The above mentioned results indicate that Ag is having some amount of cytotoxic effects on the cells and that Sr is able to offset these effects. It is possible that Sr is acting as a competitor to Ag for binding sites specific to cellular function. This is further clarified by the results for ALP immunohistochemistry (Figure 6). The micrographs show moderate activity for HA coated samples and extensively upregulated activity for Sr-HA and Sr/Ag-HA samples. Cultures on Ag-HA samples showed nearly complete inhibition of active ALP production. Alkaline phosphatases are glycosyl-phosphatidylinositol anchored, Zn²⁺ metallated glycoproteins that catalyze the hydrolysis of phosphomonoesters into inorganic phosphates. They are an essential marker of osteoblastic differentiation that helps to create an alkaline environment that favors the mineralization of inorganic phosphates. It is important to note that Zn²⁺ is the native metal ion to ALP and plays an important role in its structural/functional stabilization, but it can easily be substituted by competing ions, in this case Sr²⁺ and Ag⁺ [45]. It is possible that Ag⁺ ions released in the Ag-HA samples are binding to ALP’s high affinity metal ion sites causing functional destabilization, while in the binary system the same sites are showing a greater preference to Sr²⁺ and having very little negative effect. In this situation, it is a combinatorial effect of solubility, ion preferability and the availability of Zn²⁺ in the cell media. In the case where Ag₂O is the only dopant, noticeable effect is seen on cell health and active ALP activity. This is likely because of the limited availability of Zn²⁺ and other important divalent cations in the cell media. The metal binding sites on various proteins are searching for something to bind to and if there is no competitive advantage to more preferable cations (i.e., lack of availability), then the excess silver ions will be the driving force for binding. In the case of the binary dopant system, we see very different results. Not only is the size and charge of Sr²⁺ more preferable than Ag⁺ in forming stable proteins, but the aqueous solubility product of SrO is much greater than that of Ag₂O. In essence, in Sr²⁺ we have a cation that is able to form a more stable “active” from of various proteins and there is also a much greater abundance of Sr²⁺ ions available due to the greater solubility of SrO. Strontium has been extensively researched as an enhancer to osteoblastic differentiation and has been shown to have positive effects on Runx2 expression, osteocalcin expression and bone sialoprotein [46–49].

Results from the live/dead staining (Figure 7) show excellent antimicrobial properties of the coatings containing Ag. The control samples of HA and Sr-HA had significant live colonies of bacteria after 24 hours, while the Ag-HA samples and Sr/Ag-HA samples demonstrated nearly complete eradication of the bacteria. These results confirm that in the immediate microenvironment, the MIH concentration was effectively reached and was not affected by the addition of Sr in the binary system.

5. Conclusions

HA coating materials have been well studied as an osteogenic enhancing material and are effectively used in several medical and dental applications. In this study, by incorporating Ag₂O into the coating we were able to effectively enhance the coating with sustainable long term antimicrobial properties while minimizing negative cytoplasmic effects on osteoblast cells. With the addition of SrO as a binary dopant, any evidence of negative effects on osteoblastic cells was not only reversed, but also showed significant positive enhancement of cell proliferation and differentiation activity when compared to pure HA coated samples. The addition of dopants to the HA coating did not significantly alter the phase purity, crystallinity or adhesive strength of the coatings. All samples showed excellent adhesive strength with an average above 16 MPa.