Assessment of the Addition of Fluorapatite in Hydroxyapatite Coatings: Implementation Prosthetics/Bone in Vivo

Abstract

Hydroxyapatite (Hap: Ca10 (PO4)6 OH2)-Fluorapatite (Fap: Ca10 (PO4)6F2) composite coating on 316 L stainless steel, using the High-Velocity Oxy-Fuel Spray (SHVOF), was investigated. This work is an evaluation of the bioactivity of bone/Hap–Fap composite coatings implanted in the tibia of the rabbit. A small amount of Fap (6.68, 13.26 and 26.52 w% Fap attributed to 0.25, 0.5 and 1% fluor) was introduced into Hap. The fluorapatite provides more stable and adherent deposits. The characteristics of the coatings were investigated with various instruments including Scanning Electron Microscopy (SEM) and X-Ray Diffraction (XRD) and biological (in-vivo and in-vitro) tests. Hap−Fap coating showed excellent behavior in vitro and in vivo tests revealing that the Fap is effective in improving biocompatibility and bioactivity. This study draws inspiration from technological and biological selection solutions adopted by evolution, to transpose the principles and processes of human engineering.

1. Introduction

Research has never ceased trying to improve materials for bone and dental implants and the techniques used to synthesize them. To begin with, metals were the most widely adopted materials for implants. However, they were not without problems. Indeed, Schulze et al. assessed the effect of In-Vivo exposure to metallic nanoparticles on bone marrow In-Vitro, revealing significant alterations in cell biology [1]. Similarly, in Total hip arthroplasty ‘THA’, the problem of implant loosening due to aseptic osteolysis, was observed to be triggered by wear particles from the implant articulating surfaces. The overall In-Vivo performance of metallic hip implants turned out to be comparably poor and caused some manufacturers to recall their products. Thus, several researchers turned to the study of various bioactive ceramics, such as Hydroxyapatite (Hap), Tricalcium phosphate (TCP) and Bioglass, developed to activate bone regeneration [2]. Given its similar chemical and physical characteristics to bone, Hap was widely used as bone substitute material in orthopedic domain [3, 4]. However, when dispersed as nanoparticles Hap can cause inflammation by activating monocytes and neutrophils [4, 5, 6].

Within this prospect, Fluorapatite (Fap) was initially chosen for this purpose because of its fairly good biological and mechanical properties compared to β-TCP and hydroxyapatite (Hap) when used alone. Nevertheless, the long-term stability of bioactive ceramic implants was criticized for at least two shortcomings: the presence of low solubility of the coating and the absence of high adhesion strength between coating and substrate [7]. Hence, as a solution to these problems, Dhert and Cheng proposed to maintain F-concentration down to the minimum level in order to decrease the solubility of fluorohydroxyapatite (FHa) [8, 9]. In a similar attempt to decrease the solubility of the coating, Wolke et al. recommended the use of a fluoride coating containing Hap [10]. Indeed, Hap and Fap constitute the inorganic compound of the human hard tissue. Unfortunately, Baltag et al. and Overgaard et al. reported that the high degradation rate of Hap coating in biological environments is a serious concern, which might be harmful to adhesion properties, resulting in undesirable debris and even delaminating, which eventually leads to the failure of the implant [11, 12]. Furthermore, the stability of Hap−Fap composites was also a source of concern because dissolution or reprecipitation influences cell behavior [13]. Indeed, when Hap−Fap composites are suspended in SBF liquid, they can dissolve and precipitate rapidly, leading to ion release, change in the pH, size, and morphology [14, 15].

As a solution to these problems, some studies [16, 17], presented Fap as potential replacement for Hap in implants because of the higher degradation of the latter material in biological environments and its lower adherence. For these reasons, other researchers worked on developing the mechanical properties of coatings from Fap and Hap together to increase their efficiency. Bahandag et al. [17] studied the influence of Fap on the properties of Hap coating using thermal projection. These scholars revealed that Fap increases Hap coating crystallinity. Furthermore, the slow release of Fap reduces the delivering particles leading to the improvement of its osteointegration. Similarly, Chang et al. [18] showed that the fluorine ions improve the osteoblastic cells’ proliferation and differentiation. Hence, scholars such as Fraz and Telle and Somrani opted for Fap as an additive to Hap and a replacement for pure Hap coating on metallic implants because of their chemical composition which is similar to the bone mineral [19, 20]. Mark and Brown confirmed this finding and explained that these materials presented good biocompatibility [21]. Particularly in dental prostheses, fluor is effective in inhibiting caries [22] and proves to be compatible with the human bone that contains approximately 1 wt% of fluor [23, 24, 25].

The second issue related to implants was the technique of synthesizing the chosen materials. Thermally sprayed bio-ceramic on metallic substrate was widely used in orthopedic prostheses given its great potential in bone regeneration activity In-Vivo. Among the adopted Ceramic coating techniques are plasma spraying, flame spraying and high-velocity oxy-fuel (HVOF) spray [26]. The Plasma spray process is the most commercially preferred technique for clinical applications. The use of bond coats to improve the surface properties of metallic substrates was also studied extensively in the thermal spray literature [27]. However, despite the abundance of research on Hap coating and on the mechanical properties and bioactivity of Fap, there is still a serious need to improve the synthesis, characterization and application of Fap−Hap coating composite. For this reason, this work intended to investigate the Fap (6.68; 13.26 and 26.52 w%)-Hap coating on the 316 L stainless steel using the SHVOF technique. The remainder of this paper is divided into three sections. Section two will present the characterization of raw materials and of as-sprayed apatite coating. Section three will exhibit the study of the bioactive behavior of samples using the simulated body fluid (SBF). Finally, section four will discuss the adhesion and biocompatibility of Hap and Hap-Fap coating on rabbit bone cells as applied through the SHVOF technique.

2. Materials and methods

2.1 Materials

The Fap powder was synthesized using a wet-chemical method [20]. A calcium nitrate solution was slowly poured into a boiling solution containing di-ammonium hydrogen-phosphate, and a 28% NH₄OH solution was added to the mixture in order to adjust the pH to 9. The precipitate was filtered, washed, dried at 70°C for 12 h and calcinated at 500°C. The Hap powder with Ca/P ratio of 1.66−1.71 (Medicoat, HA-15-183, 95%) was used as base material. The Hap−Fap powders were mixed using a dried-mechanical method.

2.2 Methods

The Hap and Hap−Fap coatings were carried out on 316 L stainless steel discs by High-velocity oxy-fuel spray (SHVOF) using a Sulzer-Metco F4-MB (Switzerland). Before applying the coating, degreasing and grit blasting were carried out to make the substrate surface coarse and clean. The thermal sprayed process parameters are listed in Table 1. The coating thickness was evaluated using a micrometer (200 μm).

Starting powders	Ch4flow rate (L/min)	N2 flow rate (L/min)	O2 flow rate (L/min)	Spray flow rate (ml/min)	Spray distance (mm)
Hap	135	20	270	100	130
Hap-Fap	135	20	270	100	130

Table 1.

Spray process parameters.

The simulated body fluid (SBF) which has ionic concentrations very similar to those of human plasma was used to study the bioactive behavior of samples. For this, a commonly used SBF solution of pH 7.4 was prepared according to the procedure recently described by Kokubo et al. [28]. Table 2 presents the ionic composition of the as-prepared SBF and compares it with that of human plasma. The Hap and Hap−Fap coatings were cut in parallelepipedic slices of size 2*6*12 mm and then cleaned before being immersed in 100 ml of SBF. The temperature was maintained at 37°C.

Ion concentration (mmol/l)	SBF	Blood plasma
Na^+	142	142
K^+	5.0	5.0
Mg^2+	1.5	1.5
Ca^2+	2.5	2.5
Cl^−	147.8	103.0
HCO^−	4.2	4.2
HPO^2−4	1.0	1.0
SO^2−4	0.5	0.5
pH	7.25−7.42	7.24−7.40

Table 2.

Ion concentrations in supersaturated SBF solution prepared in the present study and in human blood plasma.

The phase compositions of the coated samples were examined with an X-ray diffractometer (PHILIPS PANALYTICAL) and with a scanning electron microscope (JEOL JSM 5800LV). The roughness of the surfaces of both substrate and coatings was measured using Mitutoyo (ISO 1997).

Prosthesis implantation, for an “In Vivo” study in rabbits, was carried out in a laboratory of animal experimentation and approved by the faculty of medicine of Sfax and by the ethical committee of the Habib-Bourguiba University Hospital, Sfax; Tunisia. The surgical operation was performed on the rabbit, by an appropriate medical team with respect for all the rules of asepsis. The shape and dimensions of prostheses are shown in Figure 1. They were carefully chosen after several graft tests in order to perfectly fill in the medullar canal of the Rabbit tibia.

Figure 1.

Prostheses dimensions for implantation: Great diameter: 5 mm; small diameter: 4 mm; degree of inclination: 30° and length: 37 mm.

These prostheses were then sterilized by ⁶⁰CO gamma irradiation (Equinox, UK).

Five adult white rabbits, aged from 5 to 10 months, were used for the experiments. All surgical procedures were done under strict aseptic protocol. The animals were premedicated and anesthetized with a xylasine/ketamine mixture (10 mg/kg). The animal tibialis anterior face was shaved. After the injection, the animal was left at rest with its eyes closed for about 20 minutes. The skin was disinfected with a povidone-iodine solution of 10% (Betadine, Medapharma). Local anesthesia (Unicaine 2%) was employed in the anterior tibialis face. Four cylindrical bars with a length of 37 mm and diameter of 4–5 mm (Figure 1) were placed in each animal. Surgical preparations for the cylinders were done using first a pilot drill and then a 2 mm twist drill. Careful drilling was done with a low rotary handpiece. The skin was sutured with interrupted threaded sutures. Skin was again soaked with a povidone solution. After 28 days the rabbits were sacrificed, and the implants extracted. The samples were stored in a formalin-based solution called BB’S and then included in methacrylic resin to be used for radiological testing of bone tissue. X-ray radiography was performed using a Faxitron X-ray system (Edimex, Angers, France) equipped with a camera (5X5 CCD). The coatings were characterized by Scanning Electron Microscopy (SEM) and X-Ray Diffraction (XRD). Moreover, the bonding strength of the as-sprayed coatings was measured using a universal testing system. The microstructure of the detached surfaces was examined using SEM.

3. Results and discussion

3.1 Characterization of the different powders

Figure 2 shows the XRD patterns for Hap and Hap-synthesized Fap composite. This observation can have four implications. Firstly, Hap-diffraction peaks shown in Figure 2a can be indexed as typical hexagonal phases (ICDD 77−0120). These peaks situated at approximately 25.9, 32, 33.1, 40.1 and 46.7 correspond well to (002), (211), (300), (212) and (222) lattice planes of the classic hexagonal phase Hap-diffraction peak (a) can be indexed as typical hexagonal phase (ICDD 77−0120). Secondly, it was clearly observed that the addition of Fap in Hap did not affect the diffraction peaks location shown in Figure 2b–d. However, the increase of Fap quantity in Hap powders shown in Figure 2b–d slightly increased the diffraction peaks intensity. This observation, in total agreement with the standard requirements, indicates that the crystallinity was well preserved. Thirdly, this observation confirms Feki-Ghorbel et al.’s previous claim that the crystalline stability of Fap is greater than that of Hap [29]. Finally, this observation shows that there were no impurity diffraction peaks or phases in the XRD patterns of the Hap and Hap−Fap.

Figure 2.

XRD patterns of powders: (a) hap; (b) hap−Fap (0.25 w% F); (c) hap−Fap (0.5 w% F) and (d) hap−Fap (1 w% F).

3.2 Characterization of as-sprayed apatite coating

3.2.1 SEM analysis

The SEM micrograph shown in Figure 3 reveals a characteristic lamellar microstructure of thermal spray coatings. A typical surface microstructure HVOF sprayed Hap coatings are generally porous. The porous structure might be beneficial to the biomedical application involving the mechanical fixation by bone in growth. Firstly, indicate where you can see the microstructures in terms of Figure 3a–c.

Figure 3.

Surface microstructures of hap (a); hap−Fap (0.25% F) (b); hap−Fap (0.5% F) (c) and hap−Fap (1% F) (d).

Secondly, these observations were not the result of the same quantity of Fap. Therefore, the differences in microstructures can be the result of the quantity of Fap. This should be clearly explained. The microstructure of the Hap and Hap−Fap coatings consists of particles of different shapes and sizes: fully molten splats which are merged in each other (white circles in the photos), some small globular particles in the matrix give a dense appearance and others probably are the un-melted particles. No crack can be observed.

3.2.2 XRD analysis

Figure 4 shows the XRD patterns for Hap and Hap−Fap as-sprayed coating. We can observe the presence of the pattern peaks attributed to Fap and Hap. Diffraction peaks of Hap−Fap coating showed more intensity than that of Hap coating. This can be interpreted as an indication that the addition of Fap increases the crystalline phase of Hap coating. Hence, our XRD analysis confirms once more Feki-Ghorbel et al.’s previous claim that the crystalline stability of Fap is greater than that of Hap [29]. In addition, this result suggests that the HVOF spray parameters were optimized. Finally, the observation of a little amorphous phase implies that Fap can be chosen as a potential partial substitute for Hap.

Figure 4.

XRD patterns of powders: (a) hap; (b) hap−Fap (0.25 w% F); (c) hap−Fap (1 w%F).

3.2.3 Surface roughness

The surface properties of implants are of vital importance for implant tissue interaction which further influences the biocompatibility for clinical use [30, 31]. Particularly, surface roughness alters osteoblastic attachment proliferation of bone cells and their differentiation and matrix production [32, 33]. The surface roughness parameters (Ra, Rq and Rz) for blasted stainless steels and for 316 L substrates coated by Hap and Hap−Fap are shown in Figure 5. The HVOF spray treatment clearly modified the surface roughness of the samples (Ra). In fact, the average surface roughness of blasted steel substrate increased from1.5 to 2.9 ± 0.2 μm. Moreover, the addition of Fap improved the average surface roughness which reached 4.7 ± 0.2 μm for the Hap−Fap (w% 1 fluor). Such results are relevant for the clinical application of Hap−Fap coated implants. As was reported by authors [34, 35], the modification of the surface roughness of an implant significantly influences In-Vitro osteoblastic response. Furthermore, a better long-term In-Vivo response of the implant is achieved when the surface roughness is increased as the amount of bone in direct contact with the implant surface as well as the loads and torques required for extracting the implant from bone growth increase [28].

Figure 5.

Surface roughness measurements for the SHVOF sprayed hap; hap-Fap-coated 316 L along with that of the grit blasted uncoated 316.

The comparison between the surface roughness results obtained in this work and those of the reported HVOF-sprayed and the flame-sprayed Hap coating [36, 37, 38] shows that the Hap−Fap coatings presented a comparative surface roughness while the addition of Fap yielded a higher roughness of the surface.

3.3 In-vitro tests

3.3.1 XRD analysis

Figure 6 presents the XRD patterns of SHVOF Hap−Fap (0.25 and 1 w% F) composite coatings in the function of immersion time in SBF. After each immersion time in SBF (0, 3 and 28 days), the examination of the coating surface shows the expansion of the diffraction lines group around 37° (2θ) corresponding to Hap phase. After a 28-day immersion in SBF, the XRD pattern presented a diffraction halo situated between 35° and 40° (2θ). The mineralogical phase analysis, using the software “High score”, allowed identifying the transition phases appearing on the Hap−Fap surface. The diffraction patterns can be attributed to a hydrated carbonated apatite layer. These patterns included lines of diffraction which can be mainly attributed to a carbonated apatite [00–012-0529]. An intensity peak around 34.5° which can be attributed to a.

Figure 6.

XRD patterns of SHVOF hap−Fap composite coating in function immersion time in SBF: 0.25 w% F (a) and 1w% F (b).

Hydrated Apatite [01–077-0128] was also detected. The crystalline structure modification after a 28-day immersion in SBF produced a formation of a new carbonated apatite precipitate covering all the Hap−Fap (0.25 w% and 1 w% F) surface.

3.3.2 SEM morphologies

The bioactivity and biocompatibility of the Hap coating and Hap−Fap composite coatings were evaluated by immersing the samples in SBF for 0; 3; 7 and 8 days.

As shown in Figure 7, spherical-shaped particles were observed on the Hap surface of the coated material after three days of immersion in SBF. Figures 8 and 9 confirmed with certainty the appearance of the new crystals on the Fap phase of the Hap−Fap composite on Hap−Fap at 0.25 and 1w% composite. As the immersion time increased, the crystals seemed to grow in size and to form an Apatite layer as shown after A 28-day immersion in SBF. This layer appeared to consist of many nano-sized crystallites with spherical morphology, as was observed by previous authors [39]. Feki-Ghorbel et al. [29] revealed that the incorporation of 1 w% of Fap into Al₂O₃ promoted the apatite layer formation on the coating surface when soaked in SBF. According to Kokubo [28], the bone-bonding ability of a material depends on the ability of apatite to form on its surface in SBF. In light of this idea, we can say with much certainty that since we observed the appearance of the apatite crystals on the Fap phase of the Hap−Fap composite surface, then we expect to reach very encouraging results in the in vivo tests.

Figure 7.

SEM observation of hap surface after immersion in SBF solution.

Figure 8.

SEM observation of hap−Fap (0.25 w% F) surface after immersion in SBF solution.

Figure 9.

SEM observation of hap−Fap (1w% F) surface after immersion in SBF solution.

Figure 10 presents the micrographs of the SHVOF coating/bone interface following sacrifice of the animal 28 days after the implantation.

Figure 10.

SEM observation of a. hap/bone interface (a); hap−Fap (0.25w%F)/bone interface (b) and hap-Fap (1 w% F) /bone interface (c).

The Hap (Figure 10a) and Hap−Fap (Figure 10b, c) composite coating exhibits a good adhesion with bone. The addition of Fap encourages the presence of crystalline forms interconnected and constituting an array at the Hap−Fap/Bone interface.

A Cross-section observation also confirms the presence of a thin apatite layer on the Hap-composite surface as shown by In-vitro tests. In this work, we have observed that the incorporation of a little amount of Fap (1w%) into Hap promoted the apatite layer formation on the coating surface and could achieve bone-bonding when implanted in bone tissues, as do bioactive ceramics.

The understanding of the biological mechanisms involved in osteoconduction seems complex. In fact, the correlation between the microstructure and the biological activity can intervene to be able to interpret the results. Nevertheless, two facts should be kept in mind when dealing with calcium phosphate films. Firstly, as was previously stated [40, 41], the bioactivity of the calcium phosphates coating depends on the capacity of their surface nucleated crystallized carbonated apatite, like the osseous mineral from the biological fluids. Secondly, the calcium phosphate films are crystallized under conditions such as the formation of the random nucleated spherulites, which can be considered as only a stage essential to the mechanism of the osteo-conduction. Hence, it can be concluded that the structural composition of the biomaterial’s surface has an influence on bioactivity. In this study, the addition of the various percentages of Fap in the Hap coating acted on the surface’s physical state and so increased its bioactivity with nucleated apatite spherulites on the biomaterial surface. As regards Hap/Fap composite coatings, after 28 days of immersion in the SBF and in contents raised in Fap (1% F), the observed transition phase can prove the presence of an amorphous neo-formed structure on the surface justifying the samples bioactivity.

Indeed, the important solubility of the amorphous phase increases the SBF oversaturation when in contact with the sample surface. Therefore, it triggers the nucleation of a neo-formed film similar to the osseous mineral. The SEM study of the surface film morphology after 3, 7 and 28 days of incubation showed a large amount of extra-cellular matrix proliferation.

This matrix stems from the crystallized apatite layer which triggered the osseous regrowth. It is important to note that this new apatite phase will result in osteoconduction. The physicochemical reactions occurring on the surface of bioactive materials could be decomposed into several stages described in the in vitro analysis above. Thus, the development of a layer of amorphous calcium phosphate occurred after a stage consisting of relegated ions forming a network of modifiers present in the deposits of the Hap−Fap matrix.

In chemical terms, this implies that Ca^{2 +} ions present on the bone surface are quickly exchanged with the ions H+ present in the SBF. Hence, the amorphous precipitates of phosphate of calcium grow on the surface. In addition, the migration of the ions Ca^{2 +} on the bone surface, the presence of PO₄³⁻ in the Hap−Fap matrix and the existence of soluble phosphates in the SBF promote the bioactive phase formation.

At this stage, the amorphous calcium phosphate layer covers the coating surface containing an important amount of bioactive Fap. The calcium phosphates are crystallized thanks to the presence of the ions OH-, Mg²⁺ and carbonates CO₃²⁻ in the middle. Then, the apatite carbonated phase occurs close to the bone mineral phase.

A Previous study, [16, 17] presented Fap as potential replacement for Hap in implants because of the higher degradation of the latter material in biological environments and its lower adherence.

4. Conclusions

To conclude, this investigation resulted in four important findings.

Firstly, this study demonstrated that the addition of Fap would produce a higher roughness at the bone surface which would certainly enable the osteoblastic cells to grow more easily. This would be very likely due to the fact that the more the surface of the bone is rough, the more it provides direct contact with the implant surface and the more it requires loads and torques for extracting the implant.

Secondly, the SEM observation of In Vitro tests confirms the formation of an Apatite layer on Hap and Hap−Fap coating after a 28-day immersion in SBF. The addition of 0.5w% and 1 w% Fap in the Hap coating increases the bone surface bioactivity through the growth of nucleated apatite spherulites on the biomaterial surface. In addition, it reduces the Hap phase solubility and leads to a stable amorphous layer.

Thirdly, this study revealed that the exchange of the calcium ions “ Ca^{2 +} ” and the phosphate ions “ PO4³⁻ ” between the Hap-Fap matrix and the SBF promotes the bioactive apatite phase formation. This new apatite phase leads to osteo-conduction.

Fourthly, In Vivo tests showed partial resorption of the implant in the form of zones less dense than the bone. The presence of this area would indicate the new bone growth accompanied by a slight osteointegration.

Our results suggest that the addition of Fap into Hap is more suitable for implantation of prostheses and for the study of bone substitutes. We can conclude with certainty that it represents the solution of choice the principles and processes of human engineering.

The bioactive behavior comparison of the two composites Hap-Fap with 0.25 wt% F and 1 wt% F revealed that the increased Fap level until 1 wt% F is presented as a better choice because it allowed the implant to develop a better integration which will guarantee a more durable life.

Acknowledgments

We would like to thank the medical team at the University hospital of Habib Bourguiba, Sfax, Tunisia for their valuable help in conducting the in vivo implantation on the rabbit in their laboratory.