A functional coating to enhance antibacterial and bioactivity properties of titanium implants and its performance in vitro

Abstract

Bacterial infections and lack of osseointegration may negatively affect the success of titanium (Ti) implants. In the present study, a functional coating composed of chitosan (CS) microspheres and nano hydroxyapatite (nHA) was prepared to obtain antimicrobial Ti implants with enhanced bioactivity. First, the chitosan microspheres were fixed to Ti surfaces activated by alkali and heat treatment, then nHA coatings were precipitated onto these surfaces. Ciprofloxacin was loaded into the microspheres using two different procedures; encapsulation and diffusion. Scanning electron microscopy micrographs of the modified Ti surfaces showed that the coating was successfully deposited onto the Ti surfaces and stable for 30 days in PBS. The drug was completely released from free microspheres loaded by encapsulation in 21 days whereas only 89% release was observed after immobilization. The burst release also decreased from ca. 55% to ca. 35%. The release was further reduced following the nHA precipitation. The modified Ti surfaces showed antimicrobial activity based on the bacterial time-kill assay using S. aureus, but the efficiency was affected by both nHA precipitation and drug loading strategy. Highest antimicrobial activity was seen in the samples without nHA layer, and when the drug was loaded by diffusion. Fourier transform infrared spectroscopy and X-ray diffraction analyses revealed that nHA on the surface enhanced HA growth in simulated body fluid for 3 weeks, showing increased osseointegration potential. Therefore, the proposed coating may be used to prevent Ti implant failure originated from bacterial infection and/or low bioactivity.

Keywords

Titanium chitosan hydroxyapatite infection local drug release antimicrobial activity

Introduction

Titanium (Ti) and its alloys are commonly used materials in the fabrication of hard tissue implants due to their excellent mechanical strength, biocompatibility, and anti-corrosion properties.1,2 However, factors such as bioactivity and post-surgery infections limit the success of Ti implants in biomedical applications.3 –5 For instance, the osseointegration properties of the pristine Ti surfaces are limited due to their bioinertness, and poor adhesion to surrounding tissues. The weak adherence leads to displacement of the implant and fibrous tissue formation at the implant/tissue interface causing implant failure.6,7 Another important cause for the failure of Ti implants is post-surgery bacterial infections. Adhesion of bacteria onto the implant surfaces causes biofilm formation that reduces bone generation on these surfaces.8,9 The biofilm also reduces the efficiency of systematic antibiotic treatment following surgical operations and protects the bacteria from the immune system.10 In all these cases, second surgery becomes inevitable which means increased pain and cost for patients. Therefore, functional surfaces are needed both to prevent infections and to increase the osseointegration.

Staphylococcus aureus (S. aureus) is by far the most common pathogen causing osteomyelitis11 and antibiotics such as gentamicin and ciprofloxacin are used in its treatment.12 These antibiotics are administrated to patients intravenously or orally at regular periods. However, the treatment with systemic drug delivery has some drawbacks such as the generation of super bacteria and systemic toxicity because of the high concentrations of antibiotic intake.13 Local treatment of bacterial infections by implant materials with antimicrobial properties rather than systemic drug delivery may reduce these problems. The strategies to obtain antimicrobial implant materials may be mainly classified into two categories; bacteria-repelling and bacteria-killing.14,15 Bacteria-repelling approaches aim to prevent bacterial adhesion onto surfaces through different methods such as the formation of different surface topographies or altering surface charge/hydrophilicity.14,16 –18 Nevertheless, these approaches are insufficient in long term applications and unable to kill all bacterial strains.19 In contrast, bacteria-killing approaches aim to eradicate the bacteria using biocides such as nanoparticles, antimicrobial peptides, and antibiotics.20 –23 When these substances delivered locally, the required amount is directly administered to where it is needed, and administration frequency, total drug concentration, and side effects are reduced.24,25 Besides, the blood flow to surgical sites is significantly reduced in orthopedic applications, and these sites become prone to infection due to the reduction in the concentration of carried immune-responsive agents.24 Local drug delivery systems also address this problem. Although bacteria-killing approaches relying on local drug delivery seem attractive, they may also have some negative effects on osteogenic activities such as cell attachment, growth, and normal cell expression.8,26,27 Therefore, local drug delivery systems with enhanced osteogenic activity are highly demanded in the treatment of bacterial infections.

Both degradable and nondegradable materials have been used to design drug carrier systems for local drug delivery. Polymethylmethacrylate (PMMA) is the most common nondegradable material for local delivery of therapeutic agents, but it provides a substratum for bacterial colonization and may cause allergy.28 In contrast, degradable carrier systems, e.g. chitosan (CS), hyaluronic acid, gelatin, collagen, polyglycolic acid, polylactic acid (PLA), poly (lactic-co-glycolic acid), polyethylene glycol, hydroxyapatite (HA), degrade in time and leave no substratum for bacterial colonization following the complete release of the therapeutics.28,29 CS is good candidate to be used in drug carrier systems since apart being biodegradable, it is also non-toxic, and has anti-bacterial effect against many pathogens.30,31 However, it is majorly used in composite form or mixed with different materials such as multi-walled carbon nanotubes, graphene, alginate, gelatin, collagen, and hydroxyapatite because of its weak mechanical properties, and limited bone regeneration potency.32

HA is an ideal material to design bioactive drug carrier systems and coatings on implant surfaces due to its osteogenic activity.33 –36 Nano sized HA (nHA) had also been shown to have antimicrobial properties, so it can be used in the design of antimicrobial surfaces.36 For example, Lian et al, fabricated vancomycin-loaded nano HA (nHA)/collagen/PLA bone graft and in subcutaneous implantation, the composite graft showed high level of bacterial inhibition (> 99%) and exhibited good adhesion to the trauma site without any inflammation responses.37 In another study, drug loaded nHA structures were deposited onto the Ti substrates as antimicrobial and bioactive coating using electrochemical deposition, and enhanced bacterial inhibition and HA crystal precipitation were detected.9 HA is also suitable to load inorganic antibacterial agents and nanoparticles (i.e Ag, Cu, Zn, and F).38 For instance; Surmeneva et al, designed a coating based on Ag release from Ti surfaces on which Ag nanoparticles were stacked between HA layers39 while Jeyachandran et al, produced fluoride and zinc modified HA films by sol–gel method.40 In order to prepare antibacterial and bioactive coatings on Ti surfaces, different strategies have been suggested such as releasing antimicrobial agents from nanotubular (NT) structures41 or chemically treated surfaces,42 and construction of hybrid coatings on the surfaces using organic/inorganic materials including collagen mimetic proteins, antimicrobial peptides, polymers, HA, etc.38,43,44

In the present study, we aim to develop a functional coating to obtain both antimicrobial activity and enhanced bioactivity for Ti-based implant materials. CS and nHA were chosen due to their antimicrobial properties and nHA has also known to enhance osseointegration. Moreover, ciprofloxacin was chosen as model biocide because of its efficiency in the treatment of wide-range of infections. The performance of the proposed coating was evaluated in vitro by drug release studies, assessment of antimicrobial activity using S. aureus, and bioactivity studies in simulated body fluid (SBF).

Materials and methods

Materials

Chitosan from shrimp shells (low-viscosity), mineral oil (for molecular biology), ciprofloxacin (≥ 98.0%, HPLC), span® 80 (viscosity 1000–2000 mPa.s), gluteraldehyde solution (grade II, 25% in H₂O), (3-Aminopropyl)triethoxysilane (≥98%, APTES), hexane (laboratory reagent, ≥ 95%), toluene and sodium phosphate dibasic (Na₂HPO₄) were purchased from Sigma. Calcium chloride dihydrate (CaCl₂), acetic acid (glacial) 100% anhydrous, sodium hydroxide (≥ 99%, NaOH) were purchased from Isolab Chemicals. Titanium (Ti, grade II) was purchased from Bag-san, Turkey.

Methods

Synthesis of the chitosan microspheres

Microspheres were prepared by using the emulsion cross-linking method.45 Firstly, 0.1 g CS was dissolved in 5 ml acetic acid (0.1 M) and then 1 ml ciprofloxacin (500 µg/ml) was added into the polymer solution. Meanwhile, 50 ml mineral oil and 1000 µl Span 80 were mixed with a mechanical stirrer for 10 minutes at room temperature. The prepared polymer-drug solution was added dropwise into the oil-span80 solution and mixed for 30 min. For crosslinking, glutaraldehyde (GA) was added (400 µl, 25%) to the oil-polymer solution, and the mixture was stirred for another hour. The microspheres were collected following the incubation of the mixture at 60 °C for 1.5 hours and washed at least three times with hexane using centrifuge (4000 rpm, 5 min) to remove the oil phase. Finally, the microspheres were dried at 37 °C for 24 h and stored under vacuum atmosphere until use. Different stirring rates (200 – 500 rpm) were used to optimize the size of the microspheres. The same procedure was applied to prepare drug free microspheres for control studies.

The shape and diameter of the prepared microspheres were determined by measuring the diameter of the randomly selected microspheres (100 microspheres) under light microscopy (Olympus BX60). The microspheres were also examined with scanning electron microscopy (SEM) (FEI Quanta- 600 FEG) to determine their surface morphologies.

Preparation of the chemically activated titanium surfaces

Ti plates (1.5 × 1.5 cm²) were chemically cleaned using etching acid mixture composed of hydrofluoric acid (40 ml), nitric acid (90 ml), and dH₂O (70 ml) for 4–5 min. Following the cleaning process, the samples were washed three times using dH₂O in ultrasonic bath. The samples were then incubated with 5 M NaOH solution (24 h, 60 °C), rinsed three times using dH₂O to remove excess NaOH, and dried at room temperature. Following the alkali treatment, the samples were heat-treated at 600 °C (1 h) and were exposed to APTES solution (5%) for silanization (24 h). Finally, the samples were washed with toluene, ethyl alcohol, and distilled water, and dried at 37 °C. Surface topographies of the Ti surfaces after the alkali and heat treatment, and silanization were analyzed with SEM. Phases and chemical groups formed on the Ti surfaces were determined with X-ray diffraction spectroscopy (XRD; Philips PW3710) and Fourier transform infrared spectroscopy (FTIR; Perkin-Elmer Spectrum One FT-IR). Pristine and alkali treated Ti samples were used as control in the characterization studies.

Immobilization of the microspheres onto the activated titanium surfaces

The Ti samples were exposed to 8% glutaraldehyde (GA) to form aldehyde groups on the surfaces (1 h) following the alkali and heat treatment. Then, a homogeneous solution composed of dissolved CS (2%) and microspheres was prepared and spread on the activated surfaces at room temperature (30 min). Finally, the samples were kept at -20°C for 30 min and dried in freeze-drier overnight. The stability of the coatings was determined by SEM analyses of Ti samples before and after incubation in PBS for 30 days.

In order to obtain full and homogeneous microsphere distribution on the surface, optimization was done using different amounts and ratio of CS solution and microsphere (100:5, 50:10, 50:5; µl:mg). The microspheres were sieved (vwr, 45 µm) before immobilization to narrow down the size distribution.

Synthesis of the nHA on the coated titanium surfaces

The prepared microspheres were put into 20 ml CaCl₂ solution (1.25 mM, pH 10) at 37 °C and stirred for an hour. Na₂HPO₄ (20 ml; 0.75 mM) solution was then added dropwise, and pH was adjusted to 10. The mixture was stirred for another hour and kept at 37 °C overnight. Finally, microspheres were washed with dH₂O and dried at 37 °C. The same procedure was used to synthesize nHA on the surface of microsphere coated Ti substrates.nHA coatings on the free microspheres and Ti surfaces were visualized using SEM. The phases, and chemical groups on the surfaces after the nHA deposition were evaluated using XRD and FTIR, respectively.

Drug loading and release studies

Two different drug loading methods were tried to determine the effect of loading procedure on efficiency and release. In the first one, the drug (ciprofloxacin) was added during microsphere synthesis, so loading was done by encapsulation as explained in ‘Methods’ section. These microspheres were also used to prepare uncoated and nHA coated Ti surfaces (50 mg on each surface). The amount of the drug used during the microsphere synthesis was 12.5 µg. Drug loading efficiency was determined according to equation (1), using drug release profiles of the samples.

drug loading efficiency (%) = \frac{released drug}{total drug} \times 100

(1)

In the second method, first uncoated and nHA coated microsphere containing Ti surfaces (50 mg on each surface) were prepared and the drugs were then loaded to these samples by diffusion. For this, the prepared samples were incubated with drug solution (12.5 µg) at 37 °C for 24 h and the unloaded drug amount was determined after removing the Ti samples from PBS. The free ciprofloxacin was determined at 278 nm using spectrophotometer (Shimadzu, UV-1601). Drug loading efficiency was then calculated according to equation (2).

drug loading efficiency (%) = \frac{total drug - free drug}{total drug} \times 100

(2)

In order to determine the drug release profiles, free microspheres and coated Ti samples were incubated in PBS solution (3 ml, pH 7.4) at 37 °C. Samples were then taken from the solution at different time intervals (2–500 h), fresh PBS was added, and the samples were analyzed at 278 nm.

Antimicrobial tests

Time-Kill assay was used to evaluate the antibacterial activity of the modified Ti surfaces. S. aureus (ATTC 29213) was chosen as a standard strain for antimicrobial susceptibility testing. Firstly, the samples were sterilized by UV (2 h) and placed into 5 ml Mueller-Hinton (MH) broth containing ca. 1×10⁶ CFU/ml bacteria from overnight culture in tubes. To determine the initial bacterial concentration, 100 microliter sample (t = 0 sample) was taken, and the tubes were then incubated in a shaker (100 rpm) at 37 °C. The amount of bacteria in each tube was determined at 3, 6, 12, and 24 hours by taking 100 μl of culture samples. Serial dilutions (10⁻¹–10⁻⁶) were prepared using sterile saline solution (0.85% NaCl) and then, 100 μl sample taken from each dilution was spread to Luria Bertani (LB) agar plates. After incubating at 37 °C for 16 hours, colonies on agar plates were counted and the number of live bacteria was calculated according to equation (2).

\frac{CFU}{m l} = \frac{Number o f Colonies \times Dilution Factor}{Volume o f culture (m l)}

(3)

Serial dilutions were performed in triplicate. Only bacteria and drug free surfaces, nHA coated and uncoated microspheres were used as negative controls while drug loaded microsphere coated surfaces and pure drug solution (5.5 µg) was used as positive control. The performance of drug loaded surfaces containing nHA coated and uncoated microspheres were evaluated in comparison with these controls. Three parallel samples of each group were used for the antibacterial test.

To determine the effect of the drug loading procedure on the antimicrobial activity, samples prepared using two different procedures (encapsulation and diffusion) as given in ‘Drug loading and release studies’ section were compared.

Determination of the bioactivity of the prepared Ti surfaces in vitro

Simulated body fluid (SBF) was used to evaluate the bioactivity of the prepared Ti surfaces in vitro. For this purpose; SBF, (1X) was prepared according to our previous protocol.46 Briefly, 0.319 g NaCl, 0.013 g NaHCO₃, 0.009 g KCl, 0.009 g K₂HPO₄ .3H₂ O, 0.01 g MgCl₂ .6H₂O, 1.48 ml HCl (1 M), 0.01 g CaCl₂, 0.003 g Na₂SO₄ and 0.24 g (CH₂OH)₃CNH₂ were dissolved in 40 ml distilled water sequentially, and the pH was adjusted to 7.4. The Ti samples were then soaked in the prepared SBF solution for 3 weeks at 37 °C, and the solution was replaced by fresh one every 2 days. Microsphere coated Ti samples were also soaked in SBF and used as control group. SEM analyses were carried out on the Ti surfaces to determine HA growth following the SBF studies. Elemental composition of the surfaces and formed chemical groups on these surfaces were studied using Energy-Dispersive X-ray Spectroscopy (EDS) (JEOL JSM 7000 F) and FTIR, respectively.

Statistical analysis

Statistical analyses were done to compare antimicrobial activates of the coated Ti surfaces. It was performed using ANOVA test. The p values < 0.05 were considered as statistically significant.

Results

Size distribution of the synthesized microspheres

The size of the synthesized microspheres was gradually decreased with increasing stirring rate as shown in light microscopy analyses (Figure 1(a)). The diameter of the microspheres was reduced 5–fold (from 100 ± 4.4 µm to 21.9 ± 4.3 µm) when the stirring rate was increased from 200 rpm to 500 rpm. More narrow size distribution was obtained with increasing stirring rate but the variations can still be observed at light microscopy images and SEM micrographs (Figure 1(b) and (c)).

Characterization of the activated Ti surfaces

A porous network structure was observed on Ti surfaces after alkali and heat treatment (Figure 2(a)). XRD analyses revealed that new Na titanate phases were formed after alkali treatment and rutile-TiO₂ phases were formed following the heat treatment on Ti surfaces (Figure 2(b)). When the alkali and heat treated surfaces were exposed to the APTES solution, the porous network started to disappear and a dense surface coating appeared (Figure 2(c)).

In accordance with the FTIR analyses, Si-O-Si groups (1034 cm⁻¹) and Si-OH groups (1121 cm⁻¹) were formed on the Ti surfaces after silanization (Figure 2(d)). Besides, amine groups (−NH₂; N-H stretching at 1,64,01,576 cm⁻¹) and methylene groups (-CH₂; -Si–CH₂- stretching at 2933 and 1340 cm⁻¹) were detected on these surfaces while only oxide peaks (651, 908, 1638, and 3361 cm⁻¹) were seen on the alkali and heat-treated surfaces.

Characterization of the coated titanium surfaces

Thick CS film layer was observed in SEM micrographs of the surfaces prepared using high amount of CS solution (100 µl) and low amount of microsphere (5 mg) (Figure 3(a)). In addition, the microspheres were not homogeneously spread over these surfaces and the coatings were detached after incubation with PBS for 2 days. When the CS to microsphere ratio was set to 50:10 (µl:mg), it was seen that the surface was densely covered with microspheres. However, local microsphere detachments were still detected after PBS incubation (Figure 3(b)). The Ti samples were homogeneously coated without any detachment when CS to microsphere ratio was set to 50:5 (µl:mg) (Figure 3(c)). On these surfaces, a thin-film was seen to cover both the microspheres and the microsphere free area (Figure 3(c), insert). Moreover, SEM analysis showed that the coatings were undisrupted even after 30 days of PBS incubation (Figure 3(d)).

Drug release profiles of the free and Ti-immobilized microspheres

The amount of total released drug was determined as 10.5 ± 0.4 µg for the free microspheres loaded by encapsulation (Figure 4). The drug encapsulation efficiency was calculated as 83.8% when this amount was accepted as total loaded drug amount. The drug release rate slowed down when the microspheres were immobilized onto the Ti surfaces. The majority of the loaded drug (89%; 9.4 µg) was released after 21 days from the microsphere coated Ti surfaces (Figure 4, immobilized microsphere (encapsulation)). Besides, the burst release decreased to ca 35% from ca 55% after immobilization.

The amount of the loaded drug onto the microsphere coated Ti surfaces by diffusion was 9.3 ± 0.2 µg and encapsulation efficiency was calculated as 75%. The majority of the loaded drug (88.6%; 8.3 µg) was released after 21 days (Figure 4(b)). The burst release was slightly higher (38%) compared to the Ti surfaces prepared with the encapsulation method.

Characterization of the nHA coatings

SEM micrographs showed that the surfaces of the microspheres were smooth prior to the nHA synthesis (Figure 5(a)). Ca-P deposits were determined in SEM micrographs of both free microspheres (Figure 5(b)) and microsphere coated Ti surfaces (Figure 5(c)) following the nHA synthesis. EDS analysis showed that Ca and P amounts were not homogenous on nHA coated surfaces (Table 1). For example, Ca and P amounts were 19.3 at % and 14.9 at %, respectively on dense nHA containing parts (point 3). However, they were 1.2 at % and 3.4 at % on the thin nHA containing areas (point 2). SEM micrographs also demonstrated that nHA coated microspheres and the CS film were still on the surfaces after incubation with PBS for 30 days (Figure 5(d)).

Table 1.

Ca-P amounts of the surfaces following the nHA precipitation.

Point	Ca (at % )	P (at % )	Ca:P
Free microspheres
1	23.4	30.4	0.8
2	35.1	35.3	1.0
3	54.2	38.4	1.4
Mean	37.6 ± 12.7	34.7 ± 3.3	1.1 ± 0.3
Microsphere coated Ti surface
1	5.2	6.4	0.8
2	1.2	3.4	0.4
3	19.3	14.9	1.3
Mean	8.6 ± 7.8	8.2 ± 4.9	0.8 ± 0.4

The FTIR spectra of uncoated and nHA coated free microspheres were shown in Figure 6(a). For uncoated free microspheres, the peaks at 889 and 1559 cm⁻¹ are indicating saccharide units and amino groups in CS. The narrow adsorption peak at 1034 cm⁻¹ is attributed to C–O stretching mode of CH₂–OH of saccharide in the GA crosslinked CS. Hydroxyl and amino group of CS, and amide I peak of N-acylated CS are assigned to 875 and 1416 cm⁻¹, respectively. The other peaks are attributed to –CH₂ bending (at 1377 and 1405 cm⁻¹) and C–H stretching (at 2868 and 2933 cm⁻¹) in the CS FTIR spectrum. In contrast, new peaks attributed to phosphate groups (563, 601, and 964 cm⁻¹) and carbonate groups (at 875 and 1416 cm⁻¹) were detected following the nHA deposition onto microspheres. Furthermore, the peak at 1034 cm⁻¹ for uncoated microspheres shifted to 1023 cm⁻¹, and adsorption peak of the hydroxyl group at 3328 cm⁻¹ broadened compared to uncoated microspheres.

Only the Ti peak was detected in the XRD patterns of the microsphere coated surfaces prior to the nHA precipitation (Figure 6(b)). Following the nHA precipitation, peaks attributed to Ca-P phases were also obtained on these surfaces (Figure 6(b)).

Effect of the nHA coating on drug release rate

The drug release rates of microspheres immobilized on Ti surfaces were analyzed before and after nHA deposition (Figure 7). For both loading methods, the release rate was slowed down after nHA deposition and the difference was more pronounced at the early stages of the release but started to diminish after 10 days for the encapsulated drug containing samples (Figure 7(a)) and disappear after 12 days for diffusion-loaded samples (Figure 7(b)). The burst release was also considerably decreased for both method, e.g. from ca. 35% to ca. 15% for encapsulated drug containing surfaces (Figure 7, inserts). For all samples, the majority (80–95%) of the drug was released in 21 days.

Antimicrobial activity of the surfaces

Bacterial density reached to its maximum after 6 hours (Figure 8, only bacteria), and to its minimum upon 24 hours incubation with the free drug (Figure 8, drug). In control groups, Drug-free microsphere and drug free microsphere + nHA containing surfaces showed no antimicrobial activity (Figure 8), while drug loaded microspheres showed antimicrobial activity (Supplementary Figure 2). nHA coated surfaces showed different antimicrobial activity that was strongly dependent on procedure (Figure 8).

When the drug was loaded during the microsphere synthesis, a notable reduction was detected in antimicrobial activity (Figure 8, microsphere (encapsulation) + nHA) in comparison to drug loading after nHA precipitation (Figure 8, microsphere (diffusion) + nHA). When in vitro drug release profiles for the first three hourswere considered, the released drug amount was ca. 1.9 µg for the nHA coated surfaces which was almost half of the drug released (ca 3.6 µg) from the only microsphere coated Ti surfaces.

Bioactivity tests

SEM image of the nHA deposits on the Ti surfaces before soaking into the SBF was shown in Figure 9(a). Following the SBF incubation for 3 weeks, enhanced HA mineralization was observed on nHA-deposited surfaces (Figure 9(b)) compared to nHA-free surfaces (Supplementary Figure 4). In accordance with the EDS analyses, the amounts of the Ca (10.4 ± 0.5 at %) and P (8.8 ± 0.5 at %) were increased after incubation in SBF for 3 weeks. They were 3.2 ± 0.2 at % and 3.1 ± 0.6 at % respectively prior to the immersing into the SBF.

Different elements such as Mg²⁺, Na⁺, and Cl^- ions were also detected on these surfaces (Figure 9(c)). In FTIR analysis, new peaks indicating P–O bending (at 46,75,17,570 cm⁻¹) and stretching (at 1052 cm⁻¹) modes of the phosphate, and –OH groups (at 630 cm⁻¹) were observed (Figure 9(d)). Furthermore, the broad adsorption peak at 3390 cm⁻¹ generated by the -OH and –NH₂ groups of CS shifted to 3422 cm⁻¹ after 3 weeks.

SEM micrograph of the interlayer between the coating and Ti substrate was shown in Figure 10. Thick HA growth was observed on the back of the coating and on the surface of the Ti below the coating after incubation in SBF for 3 weeks. It can be concluded that HA structures covered the whole surface from the outermost layer of the deposited coating to the Ti substrate.

Discussion

In the presented study, we aimed to modify the Ti surfaces using drug loaded microspheres and nHA coatings both to improve its antimicrobial activity and bioactivity. CS based drug carrier system was prepared with the emulsification-crosslinking method.45 The reduction in the size of the microspheres with the increasing stirring rate (Figure 1(a)) was probably due to increased shear stress in the emulsification technique.47 In addition, the size distribution of the microspheres become narrower with increasing rate (Figure 1(b) and (c)). To further reduce the variations in sizes of the microspheres, the samples were sieved before the immobilization onto the alkali and heat treated Ti surfaces.

A porous network like surface texture was formed on Ti surfaces following the alkali and heat treatment (Figure 2(a)) as reported by many researchers.48 –50 When the Ti surfaces are exposed to an alkali solution, hydroxyl groups partially dissolve the passive TiO₂ layer while hydrated TiO₂ structures are formed. The hydroxyl group further attacks the hydrated TiO₂ and negatively charged hydrates are formed on the surface of the substrates. The interaction between negatively charges species and alkali ions in the aqueous solution leads to the formation of an amorphous alkali hydrogel titanate layer.51 This amorphous structure may be transformed into a stable crystalline alkali titanate layer with the heat treatment as the amorphous gel layer is not mechanically stable and the surface layer can be detached from its substrate if no heat treatment was done.52 Kim et al. revealed the conversion of the amorphous sodium titanate layer to crystalline alkaline titanate layers due to the dehydration and densifying of the hydrogel layer at 600 °C.52 In accordance with the XRD patterns of the alkali-treated and dried (60 °C) Ti surfaces, Ti peaks and broad Na titanate peaks with low intensity were indexed (Figure 2(b)). Therefore, it could be said that only amorphous titanate phases were formed on the surfaces prior to the heat treatment. Rutile peaks indicating crystalline alkali titanate formation were only obtained after the heat treatment at 600 °C (Figure 2(b)) showing the importance of heat treatment to transform amorphous titanate phases to crystalline ones.

Microspheres were deposited and fixed on the Ti surfaces successfully with the help of CS thin film (Figure 3(c)). The coating stability was based on two mechanisms; covalent bonding and entrapment. Amino-silanes such as APTES have been commonly used to chemically immobilize molecules onto the titanium-based surfaces because of their ability to bind to hydroxyl groups on oxide surface.53 A covalent bond (Si-O-Ti) is formed between hydroxyl groups of the substrates and the ethoxy groups of APTES. Functional free amino groups that are desirable for surface modification of biomaterials are also formed on these surfaces.54 –56 In this study, free amino groups formed on Ti surfaces after APTES exposure (Figure 2(d)) were used to form covalent bonds with amino groups found in CS microspheres and the film.57 CS formed a continuous thin film around the microspheres and entrapped them on the surfaces further increasing the stability of microspheres on the surfaces. Thanks to these two mechanisms, the microspheres were still on the surfaces even after 30 days of incubation in PBS (Figure 3(c) and (d)).

Both release rate and burst release from microspheres were reduced after the immobilization (Figure 4). These changes might be originated from the additional coating layer which can act as diffusion barrier. Similar results were obtained in a study conducted by Wang et al, in which poly(lactic-co-glycolic acid) coatings with different thicknesses (1–10 layers) were deposited onto the titania nanotubes.58 Overall release time was extended from 5 to 40 days with the increase in polymer thickness while the burst release was reduced from 84% to 27%. Similarly, Li et al, developed a localized drug delivery system consisting of an antibiotic-loaded PEG hydrogel film covalently bound to Ti implants to prevent implant-associated infections and showed that the coating slowed down the drug release rate.59

CS has free amino and hydroxyl groups to be used as nucleation center during the nHA growth. Ca²⁺ ions are attracted by the –NH₂ group of the CS through coordination bonding or by –OH group of CS via electrostatic interactions. nHA structures are then grown on the CS in the presence of phosphate (PO₄⁻³) ions. Li et al, demonstrated in situ synthesis of spindle-like HA nanoparticles on pristine and CS modified graphene oxide (GO) sheets based on these interactions.60 In our study, free –NH₂ and –OH groups were available on the coated Ti surfaces due to the presence of CS, so the growth mechanism of nHA on Ti surfaces (Figure 5(b) to (d)) is thought to be the same.

It has been reported that HA affects the early bone regeneration process in a positive manner by increasing the new bond formation at implant-tissue interface and accelerating cell proliferation.61 Therefore, nHA deposits formed on the Ti surfaces have the potency to increase the bond strength and promote the new bone generation. Furthermore, nHA coating may be used to control the release rates of biomolecules from drug carrier systems, because it decreased both drug release rate and burst release (Figure 7). This change in the release profiles probably occurred due to the increased thickness following nHA deposition as in the case of CS film (Figure 4).

Based on antimicrobial activity results, both nHA coated and uncoated samples may be used against bacterial infections. However, the antimicrobial activity of these surfaces may differ greatly depending on the nHA presence and the drug loading strategy. For example, when the drug was loaded during microsphere synthesis, the antimicrobial activity was reduced significantly following the nHA precipitation (Figure 8, microsphere (encapsulation) + nHA). In contrast, it was higher when the drug was loaded after nHA precipitation (Figure 8, microsphere (diffusion) + nHA). This difference might be occurred due to the differences in drug release rate, loaded drug amount and type, surface properties affecting bacterial attachment, and position of the drug on the surfaces. In vitro drug release studies showed that the released drug amount from microsphere containing surfaces decreased ca 2 fold after nHA precipitation and the antimicrobial activity was almost lost. This explains the higher antimicrobial activity observed for the latter (Supplementary Figure 2). Loaded drug amounts were probably reduced due to loss of drug during nHA synthesis process in encapsulation procedure and increased thickness on the surfaces following the nHA synthesis in diffusion procedure. Therefore, the decrease in loaded drug amounts may be the cause of reduction in antimicrobial activity.

Moreover, the drug release profile can be different in bacterial growth media due to the attachment of bacteria, and the adsorption of proteins onto the surfaces. It has been reported that nHA deposits may greatly affect bacterial attachment and protein adsorption.36 Besides, attachment of the bacteria may be different for surfaces that were prepared with the adsorption of different amounts of protein. Hence, nHA coated surfaces might have enhanced both bacterial attachment and protein adsorption that may cause a reduction in drug release compared to nHA free surfaces. The location of the drug might also affect the antimicrobial activity since it changes the release profile. When the drug was loaded by diffusion, the majority of the drug may be located near the surface so showed its effect more rapidly. This can explain the increase in antimicrobial activity of nHA coated Ti surfaces loaded with diffusion.

It was seen that nHA presence enhanced the HA mineralization (Figure 9) compared to uncoated ones (Supplementary Figure 4). The positive (Ca²⁺) and the negative (PO₄⁻³ and OH) sites on the nHA coatings along with the free –NH₂ and –OH groups of CS acted as nucleation centers for HA growth and caused enhanced HA crystal precipitation (Figure 9(a)). The nucleation centers may attract different ions such as Na⁺, Cl^-, Mg²⁺ from SBF solution similar to the in vivo biomineralization during HA growth. The presence of various ions in HA structures in SBF was shown in EDS analysis (Figure 9(c)). In contrast, HA crystal precipitation was fewer on the nHA free surfaces (Supplementary Figure 4 b) probably due to the decrease in number of the nucleation center. On these surfaces, only the negative (-OH) and free -NH2 groups of CS act as nucleation centers for HA precipitation. FTIR analyses revealed that the number of peaks that are attributed to P–O bending and stretching mode peaks of phosphate ions were higher on nHA coated surfaces (Figure 9(d)), and this may indicate the crystalline HA precipitation. Besides, -OH peak at 630 cm⁻¹ is a characteristic peak for stoichiometric HA,62 and it was not seen on the uncoated surfaces. The shift of 3390 cm⁻¹ peak toward 3422 cm ⁻¹ might also indicate that the amount of -OH groups on the surfaces increased due to the HA growth since adsorption peaks of the –OH groups in HA mostly appear about 3450–3650 cm⁻¹.62

Conclusion

In this study, a functional coating composed of drug-loaded CS microspheres and nHA were prepared on Ti implants, and its antimicrobial activity and bioactivity were evaluated in vitro. It was shown that all samples released the drug in a controlled manner but, the antimicrobial activity of the surfaces following the nHA deposition was altered depending on the procedure used for drug loading. When the drug was loaded by diffusion, the surfaces showed a better antimicrobial activity against S.aureus. Therefore, the proposed coatings may be used to prevent S. aureus related infections efficiently, but to better understand its antimicrobial activity, further tests using different bacterial species needs to be done. Furthermore, nHA coating on Ti surfaces improved the bioactivity of the surfaces by enhancing HA crystal growth. In conclusion, the proposed coating provided both antimicrobial activity and enhanced bioactivity for Ti surfaces. The application of the coating is not limited to the release of antibiotics for the antimicrobial property, it may also be used to release different biomolecules to control cellular activities such as differentiation and proliferation at the implant-tissue interface.

Supplemental Material

sj-pdf-1-jba-10.1177_0885328220977765 - Supplemental material for A functional coating to enhance antibacterial and bioactivity properties of titanium implants and its performance in vitro

Supplemental material, sj-pdf-1-jba-10.1177_0885328220977765 for A functional coating to enhance antibacterial and bioactivity properties of titanium implants and its performance in vitro by Burcu Doymus, Gizem Kerem, Ayten Yazgan Karatas, Fatma Nese Kok and Sakip Önder in Journal of Biomaterials Applications

Footnotes

Declaration of conflicting interests

The author(s) declared no potential conflicts of interest with respect to the research, authorship, and/or publication of this article.

Funding

The author(s) disclosed receipt of the following financial support for the research, authorship, and/or publication of this article: This work was supported by “The Scientific and Research Council of Turkey” [TUBITAK –Project #217M220].

ORCID iD

Sakip Önder

Supplemental material

Supplementary material for this article is available online.

References

Zhang

Chen

LY.

A review on biomedical titanium alloys: recent progress and prospect. Adv Eng Mater 2019; 21: 1801215.

Rack

Qazi

JI.

Titanium alloys for biomedical applications. Mater Sci Eng C 2006; 26: 1269–1277.

Neoh

Zheng

, et al. Balancing osteoblast functions and bacterial adhesion on functionalized titanium surfaces. Biomaterials 2012; 33: 2813–2822.

Cochis

Azzimonti

Della Valle

, et al. Biofilm formation on titanium implants counteracted by grafting gallium and silver ions. J Biomed Mater Res A 2015; 103: 1176–1187.

Priyadarshini

Rama

Chetan Vijayalakshmi

Bioactive coating as a surface modification technique for biocompatible metallic implants: a review. J Asian Ceram Soc 2019; 7: 397–406.

Duyck

Vandamme

Geris

, et al. The influence of micro-motion on the tissue differentiation around immediately loaded cylindrical turned titanium implants. Arch Oral Biol 2006; 51: 1–9.

Bragdon

Burke

Lowenstein

, et al. Differences in stiffness of the interface between a cementless porous implant and cancellous bone in vivo in dogs due to varying amounts of implant motion. J Arthroplasty 1996; 11: 945–951.

Costerton

Stewart

Greenberg

EP.

Bacterial biofilms: a common cause of persistent infections. Science 1999; 284: 1318–1322.

Geuli

Metoki

Zada

, et al. Synthesis, coating, and drug-release of hydroxyapatite nanoparticles loaded with antibiotics. J Mater Chem B 2017; 5: 7819–7830.

10.

Sirivisoot

Pareta

Webster

TJ.

A conductive nanostructured polymer electrodeposited on titanium as a controllable, local drug delivery platform. J Biomed Mater Res A 2011; 99: 586–597.

11.

Del Pozo

Collazos

Cartón

, et al. V. Bacterial osteomyelitis: microbiological, clinical, therapeutic, and evolutive characteristics of 344 episodes. Rev Esp Quimioter 2018; 31: 217–225.

12.

Rayner

Munckhof

WJ.

Antibiotics currently used in the treatment of infections caused by Staphylococcus aureus. Intern Med J 2005; 35 Suppl 2: S3–16.

13.

Fathi

Akbari

Taheriazam

Antibiotics drug release controlling and osteoblast adhesion from titania nanotubes arrays using silk fibroin coating. Mater Sci Eng C Mater Biol Appl 2019; 103: 109743.

14.

Kaur

Liu

Antibacterial surface design – contact kill. Prog Surf Sci 2016; 91: 136–153. Aug

15.

Siedenbiedel

Tiller

JC.

Antimicrobial polymers in solution and on surfaces: overview and functional principles. Polymers (Basel) 2012; 4: 46–71.

16.

Rzhepishevska

Hakobyan

Ruhal

, et al. The surface charge of anti-bacterial coatings alters motility and biofilm architecture. Biomater Sci 2013; 1: 589–602.

17.

Zhang

Wang

Sun

, et al. Superhydrophobic surface with hierarchical architecture and bimetallic composition for enhanced antibacterial activity. ACS Appl Mater Interfaces 2014; 6: 22108–22115.

18.

Tran

Le Thi

Hoang Thi

, et al. Graphene oxide immobilized surfaces facilitate the sustained release of doxycycline for the prevention of implant related infection. Colloids Surf B Biointerfaces 2019; 181: 576–584.

19.

Hasan

Crawford

Ivanova

EP.

Antibacterial surfaces: the quest for a new generation of biomaterials. Trends Biotechnol 2013; 31: 295–304.

20.

Yeh

Y-C

Huang

T-H

Yang

S-C

, et al. Nano-based drug delivery or targeting to eradicate bacteria for infection mitigation: a review of recent advances. Front Chem 2020; 8: 1--22, www.frontiersin.org/article/10.3389/fchem.2020.00286/full (accessed 26 November 2020).

21.

d'Angelo

Casciaro

Miro

, et al. Overcoming barriers in Pseudomonas aeruginosa lung infections: engineered nanoparticles for local delivery of a cationic antimicrobial peptide. Colloids Surf B Biointerfaces 2015; 135: 717–725.

22.

Stigter

Bezemer

De Groot

, et al. Incorporation of different antibiotics into carbonated hydroxyapatite coatings on titanium implants, release and antibiotic efficacy. J Control Release 2004; 99: 127–137.

23.

Back

Bormann

Calafi

, et al. Testing of antibiotic releasing implant coatings to fight bacteria in combat-associated osteomyelitis – an in-vitro study. Int Orthop 2016; 40: 1039–1047.

24.

Swanson

Cheng

Friedrich

Development of chitosan-vancomycin antimicrobial coatings on titanium implants. J Biomed Mater Res A 2011; 97A: 167–176.

25.

Grainger

DW.

Drug/device combinations for local drug therapies and infection prophylaxis. Biomaterials 2006; 27: 2450–2467.

26.

Ince

Schütze

Karl

, et al. Gentamicin negatively influenced osteogenic function in vitro. Int Orthop 2007; 31: 223–228.

27.

Rathbone

Cross

Brown

, et al. Effect of various concentrations of antibiotics on osteogenic cell viability and activity. J Orthop Res 2011; 29: 1070–1074.

28.

Kluin

van der Mei

Busscher

, et al. Biodegradable vs non-biodegradable antibiotic delivery devices in the treatment of osteomyelitis. Expert Opin Drug Deliv 2013; 10: 341–351.

29.

Tiwari

Bannerjee

, et al. Drug delivery systems: an updated review. Int J Pharm Investig 2012; 2: 2–11.

30.

Ali

Ahmed

A review on chitosan and its nanocomposites in drug delivery. Int J Biol Macromol 2018; 109: 273–286.

31.

Peptu

Humelnicu

Rotaru

, et al. Chitosan‐based drug delivery systems. In: Lambertus A M, van den B, Carmen GB (eds) Chitin and chitosan. Hoboken: Wiley, 2019, pp.259–289.

32.

Zarghami

Ghorbani

Pooshang Bagheri

, et al. Prolongation of bactericidal efficiency of chitosan – bioactive glass coating by drug controlled release. Prog Org Coatings 2020; 139: 105440.

33.

Saber-Samandari

Nezafati

Saber-Samandari

The effective role of hydroxyapatite-based composites in anticancer drug- delivery systems. Crit Rev Ther Drug Carrier Syst 2016; 33: 41–75.

34.

Onder

Calikoglu-Koyuncu

Kazmanli

, et al. Behavior of mammalian cells on magnesium substituted bare and hydroxyapatite deposited (Ti,Mg)N coatings. N Biotechnol 2015; 32: 747–755.

35.

Onder

Calikoglu-Koyuncu

Kazmanli

, et al. Magnesium doping on TiN coatings affects mesenchymal stem cell differentiation and proliferation positively in a dose-dependent manner. Biomed Mater Eng 2018; 29: 427–438.

36.

Bhardwaj

Yazici

Webster

TJ.

Reducing bacteria and macrophage density on nanophase hydroxyapatite coated onto titanium surfaces without releasing pharmaceutical agents. Nanoscale 2015; 7: 8416–8427.

37.

Lian

Liu

Wang

, et al. Antibacterial and biocompatible properties of vancomycin-loaded nano-hydroxyapatite/collagen/poly (lactic acid) bone substitute. Prog Nat Sci Mater Int 2013; 23: 549–556.

38.

Spriano

Yamaguchi

Baino

, et al. A critical review of multifunctional titanium surfaces: new frontiers for improving osseointegration and host response, avoiding bacteria contamination. Acta Biomater 2018; 79: 1–22.

39.

Surmeneva

Sharonova

Chernousova

, et al. Incorporation of silver nanoparticles into magnetron-sputtered calcium phosphate layers on titanium as an antibacterial coating. Colloids Surf B Biointerfaces 2017; 156: 104–113.

40.

Jeyachandran

Narayandass

Mangalaraj

, et al. A study on bacterial attachment on titanium and hydroxyapatite based films. Surf Coatings Technol 2006; 201: 3462–3474.

41.

Zong

Bai

Liu

, et al. Antibacterial ability and angiogenic activity of Cu-Ti-O nanotube arrays. Mater Sci Eng C Mater Biol Appl 2017; 71: 93–99.

42.

Yamaguchi

Nath

Sugawara

, et al. Two-in-One biointerfaces – antimicrobial and bioactive nanoporous gallium titanate layers for titanium implants. Nanomaterials 2017; 7: 229.

43.

Xie

Wang

Pulse electrochemical synthesis of spherical hydroxyapatite and silver nanoparticles mediated by the polymerization of polypyrrole on metallic implants for biomedical applications. Part Part Syst Charact 2015; 32: 630–635.

44.

Xie

C-M

Wang

K-F

, et al. Silver nanoparticles and growth factors incorporated hydroxyapatite coatings on metallic implant surfaces for enhancement of osteoinductivity and antibacterial properties. ACS Appl Mater Interfaces 2014; 6: 8580–8589.

45.

Ozdabak‐Sert

Sen

Kok

FN.

Construction of a sandwich‐type wound dressing with pain‐reliever and pH‐responsive antibiotic delivery system. J Appl Polym Sci 2019; 136: 48252.

46.

Onder

Kok

Kazmanli

, et al. Magnesium substituted hydroxyapatite formation on (Ti,Mg)N coatings produced by cathodic arc PVD technique. Mater Sci Eng C Mater Biol Appl 2013; 33: 4337–4342.

47.

Lee

S-J

Jeong

J-R

Shin

S-C

, et al. Nanoparticles of magnetic ferric oxides encapsulated with poly(D,L latide-co-glycolide) and their applications to magnetic resonance imaging contrast agent. J Magn Magn Mater 2004; 272–276: 2432–2433.

48.

Liang

Zhou

Wang

Apatite formation on porous titanium by alkali and heat-treatment. Surf Coatings Technol 2003; 165: 133–139.

49.

Gemelli

Resende

De Almeida Soares

GD.

Nucleation and growth of octacalcium phosphate on treated titanium by immersion in a simplified simulated body fluid. J Mater Sci Mater Med 2010; 21: 2035–2047.

50.

Camargo

Takemoto

Hoekstra

, et al. Effect of surface alkali-based treatment of titanium implants on ability to promote in vitro mineralization and in vivo bone formation. Acta Biomater 2017; 57: 511–523.

51.

Lee

Yoo

Large-area sodium titanate nanorods formed on titanium surface via NaOH alkali treatment. Arch Metall Mater 2015; 60: 1371–1374.

52.

Kim

Miyaji

Kokubo

, et al. Effect of heat treatment on apatite-forming ability of Ti metal induced by alkali treatment. J Mater Sci Mater Med 1997; 8: 341–347.

53.

Tan

Zhang

Ning

, et al. Preparation and characterization of APTES films on modification titanium by SAMs. Thin Solid Films 2011; 519: 4997–5001.

54.

Meroni

Lo Presti

Di Liberto

, et al. A close look at the structure of the TiO2-APTES interface in hybrid nanomaterials and its degradation pathway: an experimental and theoretical study. J Phys Chem C Nanomater Interfaces 2017; 121: 430–440.

55.

Renoud

Toury

Benayoun

, et al. Functionalization of titanium with chitosan via silanation: evaluation of biological and mechanical performances. PLoS One 2012; 7: e39367.

56.

Nie

Long

, et al. A comparative analysis of antibacterial properties and inflammatory responses for the KR-12 peptide on titanium and PEGylated titanium surfaces. RSC Adv 2017; 7: 34321–34330.

57.

Chuah

Kuddannaya

Lee

MHA

, et al. The effects of poly(dimethylsiloxane) surface silanization on the mesenchymal stem cell fate. Biomater Sci 2015; 3: 383–390.

58.

Wang

Weng

Liu

, et al. Controlled release and biocompatibility of polymer/titania nanotube array system on titanium implants. Bioact Mater 2017; 2: 44–50.

59.

Fan

, et al. The immobilization of antibiotic-loaded polymeric coatings on osteoarticular Ti implants for the prevention of bone infections. Biomater Sci 2017; 5: 2337–2346.

60.

Wang

Liu

, et al. In situ synthesis and biocompatibility of nano hydroxyapatite on pristine and chitosan functionalized graphene oxide. J Mater Chem B 2013; 1: 475–484.

61.

Jiang

Han

Zheng

, et al. Characterization and biocompatibility study of hydroxyapatite coating on the surface of titanium alloy. Surf Coatings Technol 2019; 375: 645–651.

62.

Wang

Jia

, et al. Gradient structural bone-like apatite induced by chitosan hydrogel via ion assembly. J Biomater Sci Polym Ed 2011; 22: 505–517.

Supplementary Material

Please find the following supplemental material available below.

For Open Access articles published under a Creative Commons License, all supplemental material carries the same license as the article it is associated with.

For non-Open Access articles published, all supplemental material carries a non-exclusive license, and permission requests for re-use of supplemental material or any part of supplemental material shall be sent directly to the copyright owner as specified in the copyright notice associated with the article.

0.00 MB

1.70 MB