Electrochemically deposited antimicrobial hydroxyapatite coatings
Mokabber, Taraneh
DOI:
10.33612/diss.132596200
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Publication date: 2020
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Mokabber, T. (2020). Electrochemically deposited antimicrobial hydroxyapatite coatings. University of Groningen. https://doi.org/10.33612/diss.132596200
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Bioactive and antimicrobial
coatings – a literature review
Summary
This chapter is a literature review to capture preliminary insight into the silver-containing calcium phosphate coatings applied on metallic implants to improve their biocompatibility and antimicrobial properties. First, it provides an introduction to the titanium implants and the necessity of applying bioactive and antimicrobial coatings on them. Next, biocompatible and bioactive calcium phosphate coatings are introduced and the electrochemical deposition as one of the most appropriate deposition methods for these coatings is discussed. Finally, the antimicrobial mechanism of silver species and the biocompatibility and antimicrobial properties of silver-containing calcium phosphate coatings are discussed.
2.1 Titanium implants
The use of titanium implants for orthopedic and dental replacements is on the rise because of their developing success and the growing aging population. In general, the majority of implants is extremely successful, however approximately 5-11% of dental implants and 10% of knee and hip replacements fail due to biological or biomechanical factors [1,2]. Recently, critical challenges of titanium implants have been the improvement of bioactivity and preventing bacterial infection and biofilm formation, which is closely related to the interaction of the body and the implanted material [1]. The implantation of a biomaterial evokes specific responses from the host tissue, which is called foreign body response. The different steps of foreign body response to an implant are illustrated in Figure 2-1. These responses are described as follows: initial blood-material interactions occur upon implantation (protein adsorption), followed by the formation of a provisional matrix, then acute inflammation, chronic inflammation, granular tissue development, and, finally, fibrous capsule development [3].
Figure 2-1 Schematic representation of the host response at different points of time after implantation of a biomaterial [3, adapted with permission].
Although, some of the implanted materials evoke foreign body response and are encapsulated by fibrous, still their performance is acceptable. However, in some cases, foreign body response leads to serious problems such as complete rejection of
implant by the body. In the case of orthopedic metallic implants the general situation is that such encapsulation does not occur at the bone-metal interface for certain materials such as stainless steel and titanium, however there are contradictive reports that these materials may also cause host responses [1,3]. In general, titanium biomaterials can be implanted inside the body successfully as long as the implant is in bulk form, has sufficient mechanical stability and does not become infected. However, if one of these conditions is not achieved, the implant may be associated with an acute/chronic inflammatory reaction, osteolysis, loosening and finally failure [1].
In order to enhance the interaction between the host bone and the metallic implant, different strategies have been studied in literature, for example tailoring protein adsorption [4], promoting macrophage polarization [5], surface topographical patterns [6], biomimetic coatings [7], inorganic bioactive coatings [8,9], and chemical surface treatment [10]. As far as a bioactive coating is concerned, inorganic calcium phosphate coatings are the most favored to improve the bioactivity of the titanium implants because of its compositional similarity to natural bone. Moreover, the methods that are used to deposit calcium phosphate coatings on the metallic substrates allow achieving a uniform and reproducible coatings with sufficient mechanical adhesion to the substrate [1]. As a consequent, coating titanium with bioactive calcium phosphate components is the most impressive way to improve the bioactivity of the implant and is expected to overcome the disadvantages of titanium.
Bacterial contamination is another issue, which can restrict the long-term application of the titanium implants. Indeed, implanted biomaterials are susceptible to bacterial infections; therefore, in order to obtain biomaterials with optimal resistance to the infection, the interactions between bacteria and biomaterials have to be understood. In general, bacteria contamination occurs in two forms: dry state and wet conditions. In dry state, contamination takes place by direct transfer of bacteria through contaminated objects and by airborne bacteria. However, in wet condition, contamination occurs by the exposure to physiological fluids. Dry state contamination mainly occurs before the implantation, while contamination in wet condition takes place after the implantation. Dry state contamination can be controlled by adequate sterilization of operating room and surgical supplies and disinfecting the whole procedures of manipulation. In the case of aqueous solutions, bacterial adhesion on biomaterial surfaces is significantly affected by different variables including the surface properties, the type of pathogen and the nature of physiological fluids (Table 2-1) [11].
Table 2-1 Variables controlling bacterial adhesion [11]. Variables influencing bacterial adhesion and colonization on biomaterial surfaces Surface morphometry Macroporosity
Microporosity Micro-roughness Nano-roughness Physico-chemical properties Surface energy
Hydrophylicity/superhydrophylicity Hydrophobicity/superhydrophobicity Hydrophobic functional groups Polar functional groups Charged functional groups
Functional groups with specific activities Degree of hydration
Environmental conditions Electrolytes pH
Temperature
Host proteins/host adhesions Shear rate/fluid viscosity Fluid flow rate
Pathogen Gram-positive/Gram-negative Genus/Species
Bacterial shape Surface energy
Strain type and specific set of expressed adhesions
In general, implant-associated infections are the result of bacterial adhesion and colonization on the implant surface. The attachment and proliferation of bacteria on surfaces were first recognized in the 1930s. Over the past decades, the formation of the biofilm has been extensively studied. Although the complex mechanism of biofilm formation is still under investigation, in vitro experimental models suggest that biofilm formation has four steps (Figure 2-2): 1) initial
attachment of bacterial cells to the biomaterial surface; 2) cell colonization and accumulation in multiple cell layers; 3) biofilm formation and 4) detachment of bacterial cells from the biofilm into a planktonic state to start a new cycle of biofilm formation elsewhere [11–13].
Figure 2-2 The cycle of biofilm formation [13, adapted with permission].
The development of bacterial biofilms on the implant surface critically restricts the functionality and performance of the implant, and also exposes the patient at serious risk of systemic infections. It is very difficult to treat the bacterial cells that grow within the biofilm, since they exhibit increased resistance to antibiotics. Especially if the wrong antibiotics or low doses of relevant antibiotics are used, the problem becomes much severer and implant removal and replacement often represents the only chance to eradicate the infection. Consequently, in order to prevent bacterial infections, using synthesized antimicrobial material is more appropriate than utilizing medications [14–16].
Over the past decade, intensive efforts have been focused on the fabrication of antimicrobial surfaces or infection-resistant materials to eliminate or reduce the attachment of bacteria and formation of a biofilm on the biomaterial’s surfaces. Achievement of these purposes can be based on different strategies:
• Modify the surface of the biomaterial to give anti-adhesive properties (coatings based on hydrophilic polymeric brushes, polyamidoamine dendrimers, and biosurfactants);
• Dope the material with antimicrobial substances (antibiotic loaded biomaterials, loading with disinfectants and bactericidal substances, biofilm-disaggregating agents, grafted chitosan, hydrophobic polycationic coatings, immobilized antimicrobial peptides, and titanium oxide (TiO2) and
silver oxide (Ag2O) nanoparticles);
• Combine anti-adhesive and antimicrobial coatings (multilayer film constructed by assembling layer by layer heparin and chitosan, and covalent conjugation of antimicrobial peptides immobilized onto a hydrophilic polymer);
• Develop multifunctional materials, which are able to prevent biofilm formation as well as support tissue integration (silver containing hydroxyapatite coatings, poly(L-lysine)-grafted-poly(ethylene glycol) functionalized with adhesive peptides such as RGD, and bioglasses doped with gold nanoparticles) [13].
Nowadays, in orthopedics, the main approaches are focused on multifunctional surfaces, which are able to simultaneously improve the bioactivity of titanium implants and prevent the bacterial infection. Multifunctional bioactive and antimicrobial surfaces can be achieved via different strategies, which are summarized in Figure 2-3 [1]. Thick coatings are mainly attributed to a thick HA coating doped with different inorganic antimicrobial agents such as Ag, Cu, Zn, and Ce as metallic ions or nanoparticles. In addition, bioactive glasses in combination with ZnO or Ag nanoparticles, and bioactive TiO2 thick layers enriched by Zn or Cu
have been studied as thick bioactive coatings with antimicrobial properties. Ag-incorporating TiO2 thin films and Ag-doped HA thin films may be considered as
multifunctional thin coatings. These coatings are synthesized mainly by the use of sputtering-base deposition processes. Another type of multifunctional coatings are TiO2 nanotubes enriched by Ag nanoparticles which are useful to improve cell
adhesion and infection inhibition. Chemical surface treatments are cost-effective methods to develop multifunctional surfaces on titanium. Different solutions such as acidic solutions (HCl, H2SO4, HNO3, HF) and alkali solutions (NaOH, KOH) are
used in chemical treatments. Most of the surfaces formed via chemical treatments have large capacity to ion exchange which enable incorporation of various types of functional metal ions and induce antimicrobial properties. Incorporation of antibiotics in bioactive coatings is another strategy to provide multifunctional coatings [1].
Figure 2-3 Strategies for imparting bioactive and antimicrobial properties to Ti surfaces through different types of surface coatings or modifications [1, adapted with permission].
Regarding all the above mentioned strategies, particularly in orthopedic and dental implants, silver containing calcium phosphate multifunctional coatings possess enormous advantages since calcium phosphate can enhance osteoconductivity, while silver prevents the bacterial infection.
2.2 Calcium phosphate coatings and synthetic methods
2.2.1 Calcium phosphate minerals
The major inorganic constituent of dentin and bone is apatite, which belongs to the family of calcium phosphates. As shown in Figure 2-4, the structure of the bone can be considered as natural anisotropic composites which consists of biominerals (apatite) embedded in a protein matrix (collagen) [17]. Apatite comprises approximately 65-70 wt.% of bone tissue. Hence, calcium phosphate ceramics are beneficial to improve the bioactivity of metallic implants. Indeed, calcium phosphate coatings are able to promote direct chemical bonding at the implant/bone interface. Because of their compositional similarity to natural bone mineral and excellent biocompatibility, they can stimulate the formation of bone tissue through the osteoconduction mechanism without causing any toxicity, inflammation or foreign
body response. When a calcium phosphate based ceramic is implanted in the body, a layer of carbonated apatite, which is free of fibrous tissue, forms on its surface and leads to a bonding between the implant and the living bone, resulting in better fixation of the implant to the surrounding tissues [17–20].
Figure 2-4 The hierarchical structure of typical bone at various length scales. The microstructure of cortical or compact bone consists of Haversian systems (circles in cross section and microscopic view) with osteonic canals and lamellae; at the nanoscale, the structural framework is collagen fibers composed of bundles of mineralized collagen fibrils [17, adapted with permission].
In different environments and conditions calcium phosphate components include different phases, which are classified according to the ratio of CaO to P2O5.
Biologically relevant calcium phosphates are hydroxyapatite (HA), octa-calcium phosphate (OCP), tri-calcium phosphate (TCP), and brushite (dicalcium phosphate dihydrate, DCPD). Each of the various calcium phosphates has a different crystallographic structure and properties. For instance, the solubility product constants (Ksp) of HA, OCP, TCP, and DCPD in aqueous solution are 2.34×10-59,
2×10-49, 2.83×10-30, and 2.32×10-7, respectively [21–24]. According to the
application of the biomaterial, different phases can be applied on the implant. For example, if an increased amount of calcium and phosphate ions are required near the tissue–implant interface in early stages of implantation, DCPD could be a good option because it is relatively well soluble. In contrast, if the long-term stability of the coating is needed, phases with lower solubility such as hydroxyapatite are preferred [23]. For long-term applications, the specifications of hydroxyapatite coatings based on the Food and Drug Administration guidelines and also ISO standards are shown in Table 2-2 [25].
Table 2-2 specifications of hydroxyapatite coatings [25]. Property Value
Thickness Not spesific
Phase purity ˃95% for chemical stability Crystallinity 62% for resorption resistance Tensile strength ˃50.8 MPa
Shear strength ˃22 MPa Ca/P ratio 1.67-1.76 Density 2.98 g/cm3
Heavy metals ˂50 ppm
2.2.2 Electrochemical deposition of calcium phosphate coatings
Over the past decade, various deposition methods such as plasma spraying [26], hydrothermal [27], pulsed laser and electron depositions [28,29], magnetron sputtering [30], sol-gel process [31], electrophoretic deposition [32], and electrochemical deposition [33] have been used to deposit calcium phosphate coatings on metal substrates. Among these techniques, electrochemical deposition has received more attention because of its advantages such as comparably low operation temperature, sufficiently low cost, the possibility to obtain stoichiometric composition, high purity, simplicity of performance, and rapid deposition. This method allows highly irregularly shaped objects, including porous implants, to be coated uniformly [34,35]. During the electrochemical deposition of calcium phosphate coatings, an electric current is passed through two electrodes (the anode is usually a platinum plate and the cathode is the target substrate) immersed in an aqueous electrolyte that contains Ca2+ and PO
4
ions. Due to the electric current the reduction of water takes place and results in a local increase in the pH at the surface of cathode to a level where the calcium phosphate species become insoluble and then nucleate and deposit as calcium phosphate crystals on the cathodic substrate [36].
In electrochemical deposition, the process parameters can significantly influence the properties of the deposited calcium phosphate coatings. The type and the amount of applied current density or voltage, the composition of the electrolyte, deposition temperature and time, and finally the amount of additives are some of the parameters that can significantly affect the chemical composition, phase, crystallinity, morphology, and adhesion strength of the deposited coatings. Therefore, in order to control the properties of the deposited coatings,
understanding the mechanism of electrochemical deposition is of great importance. The mechanism of electrochemical deposition of calcium phosphate coatings on metallic substrates can be described by three main reactions [23,25,37]:
1. An electrochemical reaction (hydrogen evolution; generation of hydroxyl ions);
2. An acid–base reaction (formation of different phosphate ions such as HPO4
2-and PO4
ions);
3. Precipitation reactions (nucleation and growth of calcium phosphates crystals as coating).
One of the most effective parameters during the electrochemical deposition is the applied current density or voltage, which consequently affects the pH value in the vicinity of cathode surface. Different phosphates (e.g. H2PO4-, HPO42- and PO43-)
are stable in different pH ranges. As a result, the value of pH near the cathode directly affects the composition of the deposited calcium phosphates [23]. In order to have a better understanding of the influence of pH, the reactions that take place during the electrodeposition process are studied in more details.
There are different reactions that may occur during electrodeposition of calcium phosphates by applying current density [23,37–39]:
O + 2H O + 4e → 4OH (2-1) 2H O + 2e → H ↑ +2OH (2-2) 2H + 2e → H ↑ (2-3) NO + H O + 2e → NO + 2OH (2-4) H PO + H O + 2e → H PO + 2OH (2-5) H PO + H O + 2e → HPO + H + OH (2-6) HPO + H O + 2e → PO + H + OH (2-7) H PO + e → HPO + 1 2⁄ H (2-8) H PO + 2e → PO + H (2-9) 2HPO + 2e → 2PO + H (2-10)
Through the reactions (2-1) to (2-7), hydroxyl ions are produced and hydrogen ions are consumed. Consequently, pH in the vicinity of the cathode surface is locally enhanced. As the concentration of the hydroxyl ions increases in the solution, the concentration of HPO42- and PO43- ions also increases by the reactions [37]:
H PO + OH → HPO + H O (2-11)
HPO + OH → PO + H O (2-12)
The HPO42- and PO43- ions can also be supplied by the deprotonation reaction of
H2PO4- according to Eqs. (2-13)-(2-15) that illustrate the dissociation constants of
phosphoric acid depend on the temperature [37]:
H PO → H PO + H , p 1 = −4.5535 + 0.01349 + 799.3/ (2-13) H PO → HPO + H , p 2 = −5.3541 + 0.01984 + 1979.5/ (2-14)
HPO → PO + H , p 3 = −76.17 − 0.033574 + 134.05/ + 29.658log (2-15) The dissociation constants for the reactions (2-13)–(2-15) are pK1=2.509, pK2=7.278, and pK3=12.078, respectively, at the temperature of 80 °C. Figure 2-5 shows the speciation curves of phosphoric acid at 80 °C as a function of pH. This figure indicates that in the case of acidic electrolyte, when the pH is 2.509 or less, deprotonation almost does not occur. By increasing the pH to the value of 2.509-7.278, reaction (2-13) takes place and the predominant species is H2PO4-. In the pH
range of 7.278-12.078, HPO4
becomes more dominant via reaction (2-14), whereas when the pH is higher than 12.078, PO43- is the predominant species [37].
Consequently, the local pH in the close vicinity of the cathode is an important parameter that directly affects the local concentration of acid phosphate groups.
Figure 2-5 The distribution of phosphate species as a function of pH at 80 °C [37, adapted with permission].
By applying the current density, the calcium ions migrate toward the cathode, and according to the type of the predominant phosphate species in the solution, react with different phosphate groups and hydroxyl ions, and several kinds of calcium phosphate phases deposit on the surface of the cathode via the following reactions:
Ca + HPO + 2H O → CaHPO . 2H O (DCPD) (2-16) 4Ca + HPO + 2PO + 2.5H O → Ca (HPO )(PO ) ∙ 2.5H O (OCP) (2-17) 10Ca + 6PO + 2OH → Ca (PO ) (OH) (HA) (2-18) Therefore, it can be concluded that in moderate pH (7.278-12.078) where HPO42- is
more dominant, OCP and DCPD phases form, whereas if the pH is higher than 12.078 and the PO4
is the predominant species the deposited phase is hydroxyapatite [23,37–39].
2.2.3 The effect of deposition parameters
In the electrochemical deposition, the properties of the coatings can be remarkably influenced by the deposition parameters, among which the applied current density or voltage is the most effective one, since it directly affects the pH value in the vicinity of the cathode surface. Some reports show that by increasing the current density the pH increased and consequently the deposited layer consists of hydroxyapatite. Kuo et al. [40] studied the different calcium phosphate phases, which deposited at different current densities of 1 - 20 mA/cm2. Figure 2-6 illustrates
Figure 2-6 XRD diagrams of specimens deposited at current densities of 1, 5, 10, 15, and 20 mA/cm2 for 30 min, respectively [40, adapted with permission].
the X-ray diffraction patterns of these phases: at a low current density of 1 mA/cm2,
DCPD is the major phase and HA is the minor deposited phase. By increasing the current density to 5 mA/cm2, the major structure is still DCPD, however HA
increases relatively. When the applied current density is higher than 10 mA/cm2, the
peaks corresponding to DCPD disappear completely and HA becomes the main phase in the coating. It can be concluded that at low current densities the pH is low and the concentration of hydroxyl ions is insufficient to produce PO4
3-, via reaction (2-11) and (2-12), and only at a higher current density such as 10-20 mA/cm2, the
concentration of hydroxyl ions is high enough to provide sufficient PO43- which is
essential for hydroxyapatite deposition [37,40].
On the other hand, applying a high current density results in more hydrogen evolution. As discussed previously, in the electrochemical reactions (Eqs. 6) to (2-10)) hydrogen bubbles are generated, which practically prevent the deposition of calcium phosphate coatings on the cathode and/or decrease the adhesion of coatings to the substrate surface. One approach to limit this problem is adding hydrogen peroxide (H2O2) to the electrolyte, which increases the amount of OHˉ ions by Eq.
(2-19):
H O + 2e → 2OH (2-19)
Therefore, addition of H2O2 to the electrolyte can influence the mechanism of
the electrochemical process through providing an alternative source of OHˉ. More OHˉ ions means more PO43- ions production, thus the deposited layer would be
hydroxyapatite with strong adhesion to the substrate [36,41].
The electrolyte temperature is another influential factor and it can affect the deposition of calcium phosphate coatings in several ways:
1. Accelerating the diffusion rate of ions and changing the reaction rate as well, a higher deposition temperature encourages the formation of more hydroxyapatite coatings;
2. Decreasing the solubility of hydroxyapatite, accelerating particle nucleation rate and stimulating the deposition of the film;
3. Reducing hydrogen bubbles attachment to the substrate surface during the electrodeposition process, resulting in a less damaged hydroxyapatite layer, with better adhesion, and more homogeneous coatings [36,41].
Similar results were reported by Zogbi Jr et al. [39]. They deposited calcium phosphate crystals at two different temperatures of 40 and 70 °C on vertically aligned multi-walled carbon nanotubes, which were grown on Ti substrates, and investigated the morphology of the coatings (Figure 2-7) as well as the atomic ratio of Ca/P (Table
2-3) in the deposited layers. As shown in Figure 2-7, the coatings present considerably different morphologies. Figure 2-7 (a) illustrates globular-like crystals of calcium phosphate when deposition occurs at 40 °C. However, Figure 2-7 (b) indicates that crystals deposited at 70 °C have a plate-like morphology.
Figure 2-7 Scanning electron microscopy (SEM) images from hydroxyapatite electrodeposited for 3600 s maintaining the temperature constant at (a) 40 °C, forming globular-like crystals and (b) 70 °C, forming plate-like crystals [39, adapted with permission].
Table 2-3 indicates the atomic percentage of different elements as well as the Ca/P ratio in the layers deposited at (A) 40 and (B) 70 °C. The results show that the layer deposited at 40 °C with globular-like crystals has an average Ca/P atomic ratio of 1.55 which is typical for TCP. However, the average Ca/P atomic ratio of plate-like crystal deposited at 70 °C is 1.67 (typical of stoichiometric HA). It can be concluded that, the deposition temperature can significantly affect the morphology of calcium phosphate coatings. Increasing deposition temperature to 70 °C leads to an increased calcium content in the composite and the possibility of hydroxyapatite formation enhances [39].
Table 2-3 Average calcium and phosphorus elemental content and Ca/P atomic ratio in electrodeposited hydroxyapatite coatings deposited in (A) 40 and (B) 70 °C [39].
(A) Element Atomic % (B) Element Atomic %
O K 67.81 O K 75.66
P K 12.58 P K 9.11
Ca K 119.61 Ca K 15.24
2.2.4 Adhesion strength of electrodeposited calcium phosphate coatings
The long-term performance of implants coated with hydroxyapatite is extensively influenced by the physical and chemical properties of the coating such as chemical purity and phase stability, as well as the mechanical properties like hardness and adhesion strength of the coatings. In general, the adhesion strength of the coatings is a function of both adhesive strength (coating to the substrate) and cohesive strength (within the coating layers themselves). The cohesive strength is controlled by coating structure, i.e. crystallinity, porosity and structure uniformity, however the adhesive strength is mostly dependent on the coating properties such as the work of adhesion, residual stress and surface roughness of the substrate. So far, a number of methods have been developed to increase the adhesion strength of the hydroxyapatite coatings on metal substrates [36,42,43].
In order to enhance the cohesive strength, coatings must be uniform with low porosity. Some researchers claim that hydroxyapatite coatings deposited by the traditional electrochemical method (applying a static potential) cannot provide these properties. As mentioned previously, hydrogen bubbles, which are generated during the electrochemical process, adhere to the surface of the cathode and prevent uniform deposition of the coatings. Moreover, in the electrochemical deposition a concentration gradient occurs because of the low mobility of ions diffusing from the main body of the solution to the surface of the substrate. These two phenomena lead to porous coatings with low adhesion strength.
To overcome these problems, two solutions have been suggested: to prevent hydrogen bubble evolution, as discussed before, H2O2 is added into the electrolyte to
provide more hydroxyl ions without excessive an increase of the current density; and to decrease the concentration gradient, it is suggested to apply pulsed current instead of direct continuous current. During the pulsed current, as shown in Figure 2-8, the current density is turned on and off for a specific duration resulting in a relaxation time. Consequently, dense and uniform coatings can be deposited with improved adhesion strength [44–46].
Figure 2-9 shows the schematics of pulsed electrodeposition of hydroxyapatite coatings on 316L stainless steel. As shown in Figure 2-9 (a), once pulsed current applied, OHˉ ions appear in the vicinity of the cathode due to the reduction of water and H2O2, and lead to increasing pH at the interface between the substrate and the
electrolyte. By increasing the pH, Ca2+ and PO 4
ions in the vicinity of the substrate surface react with OHˉ ions to form a hydroxyapatite layer according to Eq. (2-18) (Figure 2-9 (b)). Formation of hydroxyapatite consumes PO4
and Ca2+ ions, so the
concentration of these ions decreases near the electrode, which causes difference of concentration of PO4
and Ca2+ ions between the bulk solution and the vicinity of the
electrode as shown in Figure 2-9 (b). This difference results in the concentration gradient of the substrate. The relaxation time (pulse off time) gives a chance to Ca2+
and PO4
ions to diffuse from the bulk solution to the vicinity of the substrate (Figure 2-9 (c)), which reduces the concentration gradient. By the successive current pulse, Ca2+ and PO
4
ions react with OHˉ ions to form again hydroxyapatite layer, as
Figure 2-9 The schematic illustration for the mechanism of pulsed electrodeposition of hydroxyapatite. The application of pulsed on time (a) and (b), application of pulsed off time (c) and again the application of pulsed on time (d) [45, adapted with permission].
represented in Figure 2-9 (d). Nevertheless, during direct continuous current, Ca2+
and PO43- ions do not have enough time to diffuse from the bulk solution and get to
the cathode surface, so concentration gradient prevents the super saturation needed for deposition of hydroxyapatite. Therefore, deposited coating becomes porous with poor adhesion strength. In conclusion, pulsed electrodeposition retains the concentration gradient and improves the physico-chemical properties of the coatings by depositing more uniform, compact and denser coatings [45].
In order to improve the adhesive strength of the coating (coating to substrate), several pre-treatments on the substrate have been suggested, such as treatment in NaOH solution and anodizing the substrate [36,47–49]. The surface roughness of the substrate is highly important, since a rough surface can increase the wettability of the hydroxyapatite solution on the substrate, and also provide mechanical interlocking between the hydroxyapatite layer and the substrate [48]. One approach to enhance the surface roughness of the titanium substrate is to anodize the surface to produce a TiO2 layer. Figure 2-10 illustrates the typical morphology of titanium
substrate after anodic oxidation. The oxide film has porous microstructure with rounded nanopores. This rough surface can be helpful in forming a mechanical interlock between the titanium substrate and the hydroxyapatite coating and improve the adhesion strength of the hydroxyapatite coating [48,49].
Figure 2-10 SEM micrograph of anodic TiO2 oxide layer formed on Ti–6Al–4V alloy surface [48, adapted with permission].
It has been shown that further treatment of anodized titanium in an alkaline solution, such as NaOH, can even improve the adhesion strength of hydroxyapatite
coating [47]. Kar et al. [47] investigated the adhesion strength of the hydroxyapatite coating on the nanotubular TiO2 surface in different experimental conditions. The
adhesion strength of the coatings increases from 19 to 25 MPa by alkaline treatment. This is attributed to the fact that immersing the samples in a NaOH solution increases the pH inside the TiO2 nanopores, which results in a superior nucleation
of the hydroxyapatite crystals during electrodeposition.
2.3 Silver-containing calcium phosphate coatings
2.3.1 Silver as an antimicrobial agent
In vitro investigations have indicated that metal ions such as silver, copper and zinc, as well as silver and gold nanoparticles can be used as antimicrobial agents to prevent the bacteria adhesion [50,51]. In the past few years, silver has gained more attention since it is an effective bactericide over a wide range of bacteria including antibiotic-resistant bacteria. In addition, it possesses many advantages such as antifungal and antiviral properties, strong inhibitory, a broad spectrum antimicrobial activity, high thermal stability and non-toxic to human cells at low concentrations. Silver nanoparticles have been shown to be effective against a wide range of bacteria such as Escherichia coli, Staphylococcus aureus, Staphylococcus epidermis, Leuconostoc mesenteroides, Bacillus subtilis, Klebsiella mobilis, and Klebsiella pneumonia and also be effective against different fungi such as Aspergillus niger, Candida albicans, Saccharomyces cerevisia, Trichophyton mentagrophytes, and Penicillium citrinum [15,16,52].
As an antimicrobial material, silver has been studied to identify the effects of silver species (ionic state (Ag+) versus metallic state as nanoparticles
(AgNPs));
particles size (μm or nm);
combination with different matrix (Ti, HA, TiO2 and Wollastonite);
synthetic methods (plasma spraying, thermal spraying, pulse laser deposition, RF magnetron sputtering, wet chemical precipitation, electrochemical deposition, and sol gel) [53].
All of the factors mentioned above can affect the properties of silver-containing materials, and consequently influence their antimicrobial efficiency. For instance, the synthetic method has a direct influence on the properties of the material. Consequently, according to the application of the biomaterial, it is necessary to choose a proper synthetic process to develop an effective material with excellent biocompatibility and antimicrobial properties. In addition, the type of the silver (ionic silver or metallic silver) in the silver-containing materials remarkably affects
the antimicrobial properties and the biocompatibility of the coatings due to the fact that different species induce different killing mechanisms [53–55].
So far, according to the silver type, different killing mechanisms have been proposed [16,53,55]. Figure 2-11 illustrates the different steps of the killing mechanism for silver ions [16]. It is known that silver ions bind and react with electron-donating groups such as amino, carboxyl and thiol groups in the cellular proteins on the membrane and weaken the stability of the outer membrane. Consequently, the accumulation of the silver ions on the membrane can damage it and increase its permeability which leads to cytosolic leakage (step (1) in Figure 2-11). Once the silver ions penetrate the bacteria through the damaged membrane, they combine with nucleoid and inactivate the respiratory chain as well as the tricarboxylic acid (TCA) cycle enzymes and disrupt the ATP production (step (2) in Figure 2-11). This step is the most effective method by which the silver ions cause bacterial death. Furthermore, silver ions can bind to bacterial DNA and RNA and destroy their structure. The destruction of DNA structure inhibits the bacterial replication and finally leads to bacterial cell distortion and death (step (3) in Figure 2-11) [12,16,50].
Figure 2-11 Illustration for the proposed bacteria killing mechanism caused by silver ions released from Ag/HA bioceramics [16, adapted with permission].
In the case of silver nanoparticles, bacteria can be killed via different mechanisms. Firstly, bacteria may directly contact with silver nanoparticles, which can attach to the cell wall of the bacteria, form pits in the cell membrane, penetrate
the cytoplasm and eventually cause cell death [53,55,56]. It is also possible that silver nanoparticles release silver ions, consequently the killing mechanism of silver nanoparticles would follow the above mentioned steps (Figure 2-11) [15,57]. Another mechanism, which can be attributed to both silver ions and silver nanoparticles, is the generation of reactive oxygen-containing species (free radicals), such as superoxide radicals (O2- and H2O2). If the reactive oxygen species increase to toxic
level, the cells would experience a very high oxidative stress which causes the cellular inactivation [50,54]. In conclusion, although various type of the silver species exhibit different killing mechanisms, silver as an effective antimicrobial agent can be extensively used in biomedical industry to prevent implant-associated infection.
2.3.2 Synthesizing silver-containing calcium phosphate coatings
In recent years, several methods have been investigated to introduce silver into the bioactive coatings, such as biomimetic deposition, plasma spraying, magnetron sputtering, sol-gel process, and electrochemical deposition. Electrochemical deposition has many advantages compared to the other methods to synthesize ceramic coatings. Incorporation of the silver into the calcium phosphate coatings can be applied through different approaches, for example calcium ions in the crystal structure can be substituted by silver ions. This structure is called silver-doped calcium phosphate component [16,58]. Furthermore, silver and calcium phosphate components can be synthesized as separate phases and built up as a composite material [59–61].
The crystallographic structures of the calcium phosphates have several different sites that allow for atomic substitution, hence various elements with various ionic charges can be replaced in those positions. For instance, hydroxyapatite has a hexagonal structure in which the calcium ions occupy two different sites called I and II. Ca(I) are in columns parallel to the c axis, while Ca(II) form two equilateral triangles along the c axis, which are called anion channel and mainly occupied by monovalent ions such as OHˉ. Phosphate ions (PO4
3-) with tetrahedral geometry, where each P atom is surrounded by four oxygen atoms, complete the structure (Figure 2-12). Theoretically, calcium ions can be replaced by a large number of bivalent and trivalent metal cations; in practice, however, it is not possible to substitute a lot of the calcium cations [62,63]. Substitution with bivalent cations such as Mg2+, Zn2+ and Sr2+ is more feasible comparing with monovalent cations such as
Na+, K+ and Ag+, since monovalent substitution causes a charge imbalance.
Moreover, the cationic radius of the metal that substitutes the calcium ions must be taken into consideration, larger cationic radius differences mean more structural distortion [20,62].
Figure 2-12 View of the atomic structure of hydroxyapatite as a stand-in for bioapatite. Viewed (A) down the c axis, and (B) perpendicular to the c axis [62, adapted with permission].
In general, the incorporation of different ions inside the hydroxyapatite structure leads to some changes in the crystal structure, the surface charge, the solubility, and the crystallinity, which consequently influence the overall biological properties and performance upon implantation. For example, the incorporation of antimicrobial cations such as silver induces antimicrobial properties to the calcium phosphate coatings. However, in order to incorporate silver ions into the calcium phosphate structure, it is very crucial to choose an appropriate method, since the radius of silver ions is larger than that of calcium ions [58,62]. Electrochemical deposition is one of the techniques that can be used to synthesize silver-doped calcium phosphate coatings. For example, Yan et al. [58] synthesized silver-doped hydroxyapatite coatings via electrochemical deposition method by adding 0.1 mmol/l of AgNO3 to the conventional calcium phosphate electrolyte. During the
electrochemical deposition, calcium and silver ions react with HPO4
and a coating of Ca1-xAgx HPO4.2H2O deposits on the substrate following the reaction:
HPO + (1 − )Ca + Ag + 2H O → Ca Ag HPO ∙ 2.5H O (2-20) By alkaline treatment in 1 M NaOH for 2h, the deposited coating is converted to silver-doped hydroxyapatite through a hydrolysis reaction:
10Ca Ag HPO ∙ 2.5H O → Ca Ag (PO ) (OH) + 4H PO + 18H O (2-21) By using different analytical methods such as X-ray diffraction, scanning electron microscopy equipped with energy dispersive spectroscopy and X-ray photoelectron spectroscopy, these authors proved the substitution of the calcium ions with silver ions in the hydroxyapatite structure [58]. They also claimed silver-doped hydroxyapatite coatings deposited via electrochemical deposition with silver concentration of 2.03 wt.% not only exhibit significant antimicrobial properties, but also show a remarkable biocompatibility [58].
Another approach of synthesizing silver-containing calcium phosphate coatings is the deposition of calcium phosphate with silver nanoparticles as composite coatings. For instance, Xie et al. [61] utilized a conventional calcium phosphate electrolyte containing 5 mM Ca(NO3)2 and 3 mM NH4H2PO4 together with 0.05 g/l
AgNO3 and 0.22 g/l chitosan (CS) to deposit hydroxyapatite coatings incorporating
silver nanoparticles. Chitosan was used as a suitable stabilizing agent to provide a uniform distribution of silver nanoparticles in the coating. Silver ions in the electrolyte solution are chelated by two amino groups of chitosan. In this case, the deposition was conducted by pulsed electrochemical deposition with a pulse width of 100 s for 1 h at -1.3 V on pure Ti substrates. After deposition, bone morphology protein-2 (BMP-2) along with a heparin solution was absorbed on the coatings to enhance the biocompatibility of the deposited coatings. Figure 2-13 shows the synthesizing steps of the CS/Ag/HA composite coating [61].
Figure 2-13 (a) Structure of chitosan. (b) Chitosan chelates Ag+ in electrolytes. (c) CS/Ag/HA coatings are prepared by co-electrodeposition. (d) CS/Ag/HA coatings are formed. (e) 2 and heparin are mixed. (f) BMP/CS/Ag/HA coatings are formed by adsorbing BMP-2/heparin mixed solution on CS/Ag/HA coatings [61, adapted with permission].
The morphology of the deposited coatings is illustrated in Figure 2-14. SEM micrographs show that in the presence of chitosan (Figure 2-14 (a)) silver nanoparticles are homogenously distributed in the calcium phosphate coating. The size of the nanoparticles ranges from 100 to 200 nm. Without chitosan the deposited nanoparticles agglomerate and form large clusters, as seen in Figure 2-14 (b).
Moreover, in vitro investigations demonstrate that CS/Ag/HA coatings have high antimicrobial properties along with good osteoinductivity. In summary, electrochemical deposition is a proper method to synthesize silver-containing calcium phosphate coatings with superior properties [61].
Figure 2-14 SEM micrographs of (a) CS/Ag/HA coatings, (b) Ag/HA coatings, and (c) HA coatings [61].
2.3.3 Biological properties of silver-containing calcium phosphate coatings
As soon as a silver-containing biomaterial is subjected to the aqueous environment of the body, the surrounding cells will be affected either by direct contact with the silver-containing biomaterial or by silver ions that are released from the biomaterial. If the concentration of silver ions in human blood exceeds 300 ppb, it can induce toxicity to some human cells as well as side-effects in the form of leukopenia (white blood cells), liver and kidney damage. In addition, long-time exposure to high concentrations of silver leads to argyria which is a type of skin disease caused by excessive exposure to chemical compounds of silver. In contrast, some studies have demonstrated that if a low concentration of silver is incorporated into the biomaterial, it retains its biocompatibility [64–66]. Therefore, in order to guarantee superior antimicrobial properties along with excellent biocompatibility, it is of utmost importance to optimize the silver concentration in the biomedical coatings. Although, many studies have been done to optimize the silver concentration in the coatings according to their biocompatibility and antimicrobial activity, further investigations are still required because the effects greatly depend on the form of silver and the synthetic approach by which it is incorporated.
The biological properties of silver-containing calcium phosphate coatings are evaluated through two different aspects:
• Evaluating the antimicrobial properties of the coatings (the effect of the silver on bacteria cells);
• Evaluating the biocompatibility of the coatings (the effect of the silver on human cells).
Depending on the application of the biomaterial, different methods are used to evaluate the antimicrobial properties. One of the most commonly used methods to evaluate the antimicrobial properties of the coatings is the disc diffusion test method (commonly known as the Kirby-Bauer disc diffusion method), in which a zone of inhibition of bacterial growth around the disc is examined. Additionally, the type of the bacterial strain, which is chosen to evaluate the antimicrobial properties, is very determinative, since various bacterial strains exhibit different sensitivity when exposed to antimicrobial materials [67–69].
Mocanu et al. [50] evaluated the antimicrobial properties of silver-containing hydroxyapatite composites against five different bacteria strains, namely Escherichia coli, Staphylococcus aureus, Staphylococcus spp, Bacillus cereus and Candida albicans by using the disc diffusion test. Figure 2-15 illustrates the inhibition zone for hydroxyapatite composites with different silver content exposed to different bacteria. As can be seen, silver-containing hydroxyapatite composites are effective against all of the five microbial species, while no effect is observed for the control samples. The diameter of the inhibition zone depends on the silver content of the composites. In all of the six groups, the largest inhibition zone is related to the highest content of silver (5.4%). Staphylococcus aureus shows the highest sensitivity against silver nanoparticles and has the zones of inhibition around 11–16 mm (Figure 2-15 b). Among all the species, Staphylococcus spp and Bacillus cereus are less sensitive [50]. However, among the wide variety of bacteria that can cause infection on an implant, Staphylococcus aureus and Staphylococcus epidermidis are the most common pathogens, which cause almost 70% of the infections [70].
It has been reported that the minimum concentration of silver ions that is required to induce the bactericidal effect is 0.1 ppb. This is far below the safe value for human cell, since eukaryotic cells are able to withstand 10 ppm exposure. Nevertheless, the maximum silver ion concentration that eukaryotic cells can withstand without suffering cell death, is not clear [12]. Therefore, it can be concluded that the human cells are not as sensitive as bacteria when exposed to silver ions. This may be due to the different structure and size of bacteria and human cells. For instance, a fibroblast cell has a diameter of 10-50 μm, whereas the diameter of Staphylococcus epidermidis cell is between 0.5-1.5 μm. It means mammalian cells are much larger than bacterial cells. As a result, for mammalian cells the selective silver concentration relative to the surface area is remarkably lower. In addition, mammalian cells have a much more complex structure than bacteria cells. In mammalian cells, the core, mitochondria and endoplasmic reticulum protect the cell against the silver ions attack because they can serve as additional passage barriers against silver ions. Thus, a silver-containing material with optimum content of silver can be non-toxic to the tissue cells, while largely inhibiting bacterial infection [16,71].
Figure 2-15 Inhibition zones for Escherichia coli (a), Staphylococcus aureus (b), Staphylococcus spp (c), Bacillus cereus, (d), Candida albicans (e), and after removing the disks for Staphylococcus aureus (f). Disks are numbered according to their silver nanoparticles content: disk 1=2%; disk 2=2.5%; disk 3=3%; disk 4=3.6%; disk 5=4.5%; disk 6=5.4%; W stands for the control samples [50, adapted with permission].
The first phase of the cell-implant interaction is cell adhesion and this phenomenon can extensively affect the other cellular functions such as proliferation and differentiation [16]. Shi et al. [16] fabricated 0.04, 0.27, 2.2 and 197 ppm doped HA bioceramics and investigated whether the release of silver ions from Ag-HA bioceramics affects osteoblast adhesion. They seeded mouse embryonic osteoblast MC3T3-E1 cells onto the bioceramic surface. According to their investigation (Figure 2-16), it can be concluded that Ag-HA bioceramics facilitate osteoblast adhesion with plentiful active filopodia formation. The bioceramics doped with low amount of Ag (0.04 ppm – 2.2 ppm) show better cell adhesion compared with the bioceramics doped with higher concentration. Meanwhile, Ag-doped HA bioceramics with 2.2 ppm silver concentration exhibit bacteria reduction of over 80% against E. coli and S. aureus bacteria [16].
Figure 2-16 Confocal microscopy images of MC3T3-E1 cells cultured for (a) 1 h, (b) 4 h and (c) 24 h on HA and Ag-HA bioceramics with a cell density of 3.0 × 104 cells per mL. The F-actin was stained with FITC (green) and the nucleus was stained with DAPI (blue) [16, adapted with permission].
Similar results were reported by Herkendell et al. [71]. They optimized the content of silver nanoparticles and multiwalled carbon nanotubes (CNT) as reinforcement of mechanical properties in hydroxyapatite composites. Figure 2-17
illustrates a comparative presentation of bacteria and osteoblast cell activity on CNT-Ag composites with different concentration of silver nanoparticles. The HA-CNT-Ag composite with 4 wt.% CNT and 1 wt.% silver (C4A1) has excellent biocompatibility with significant bacteria reduction (specially against S. epidermidis) as compared to pure HA. However, once the concentration of the silver nanoparticles is increased to 5-10 wt.%, the bacteria reduction improves, but the cell density slightly decreases [71].
Figure 2-17 Variation of density of fibroblast cells, S. epidermidis and E. coli with silver content in HA-CNT-Ag composites. The rectangular box shows silver content for optimized biological properties [71, adapted with permission].
An efficient implant must provide a good interaction with the human body in terms of biocompatibility and bioactivity, along with the prevention of infection. Titanium implants coated with silver-containing calcium phosphate coatings possess all these properties. Nevertheless, the synthetic method, the type and concentration of the silver species inside the coating remarkably influence the properties of the coating. Therefore, gaining deeper insight into the physical, chemical and biological properties of silver-containing calcium phosphate coatings is of paramount importance for achieving an optimal coating for future bone implants.
2.4 Conclusion
Titanium implants coated with calcium phosphate can be widely used in biomedical applications, since the coating provides bioactivity and biocompatibility, while the titanium provides mechanical strength. However, an advanced implant not
only needs to fulfill the biocompatibility and bioactivity, but also needs to prevent bacterial infection. One approach to enhance the resistance of the coating against the bacterial contamination is to incorporate silver into the synthetic coating. Among the different methods, electrochemical deposition technique is favorable to deposit silver-containing calcium phosphate coating because it allows to coat complex shapes and porous surfaces. It was shown that applying different deposition parameters significantly influences the properties of the coating such as chemical composition, thickness and morphology, which are the characteristics that strongly affect the biological properties as well.
Despite the extensive work carried out on the silver-containing calcium phosphate coatings, a deeper insight into the physical, chemical and biological properties of the coatings is still lacking. The deposition parameters play an important role in determining the properties of the deposited coatings. However, the relationship between the deposition parameters and the mechanical/biological properties of the coatings and also their effects on modulating cellular response is not clearly understood. Moreover, there has been no systematic study of the main mechanism responsible for morphological changes and of the growth mechanism in electrochemical deposition of calcium phosphate coatings. There is also a remarkable variation in the observed antimicrobial mechanisms of silver-containing materials. A more comprehensive investigation of the properties of silver-containing calcium phosphate coatings can be helpful to better understand the antimicrobial activity of the coatings. Determining the role of the silver type in the antimicrobial properties and biocompatibility of silver-containing coatings is of outstanding importance, which can provide a great opportunity to improve the bactericidal coatings for future biomedical implants.
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