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Nano-crystalline pulsed laser deposition hydroxyapatite thin films on Ti substrate for biomedical application.

Abstract Hydroxyapatite (HA) thin coating has been coated on titanium substrates by pulsed laser deposi -- tion (PLD). Structural and morphological studies by transmission electron microscopy (TEM) and X-ray diffraction (XRD) were performed. The HA film is polycrystalline and irregular with a range size of particulates from 100 nm to 1 Adhesion of the HA films to the Ti substrates was excellent as observed by cross-scanning electron microscopy (X-SEM) and transmission electron microscopy (XTEM). The resid -- ual stresses state determined on uncoated Ti substrates was compressive. After deposition of thin PLD HA on titanium substrates at 400 [degrees] C the residual stress state was in a very low tensile state.

Keywords Nanostructures, Residual/internal stress, Electron microscopy

Introduction

Thin coatings are an important part of highly developed medical prostheses, bestowing improved functionalities to these devices. Many research works in the field of biomaterials are focused on forming bioactive coatings on metallic and non-metallic substrates which are implanted into human body. The purpose of this effort is to utilize the mechanical properties of the substrate and the bioactivity of the coatings. The bioceramic thin films [e.g., calcium phosphates, in particular hydroxyapatite (HA)] of metallic implants have been proposed as a solution for combining the mechanical properties of the metallic material with the bioactive character of the ceramic layer, leading to a better integration of the entire implant with the newly remodeled bone.

Titanium materials (commercially pure titanium ASTM Grades 1-4 or Ti-based alloys) are considered to be the most biologically compatible materials to vital tissue(1) because they have high mechanical resistance, low modulus of elasticity, high corrosion resistance, excellent general biocompatibility, and atoxicity.(2)

Hydroxyapatite (HA, ([Ca.sub.10] ([PO.sub.4]).sub.6]([OH].sub.2)), the major component of human hard tissues (65-70 wt (degrees), is characterized by high osteoconductivity, its ability to enhance bone in-growth(3),(4) and fairly good attachment to implant surfaces especially at early critical stages of the implantation.(5-12) HA is extensively used for many applications, including coatings of orthopedic and dental implants, and scaffolds for bone growth. HA implant coatings showed an improved bone apposition as compared to uncoated implants in the first several weeks after operation.(13) In the last 20 years, physical and chemical techniques for coating calcium phosphate compounds on titanium for biomedical application have been developed: plasma spraying,(14),(15) dipping,(16) electrocodeposition,(17) pulsed-layer deposition (18), sputtering,(19) and sol-gel-derived coating.(20)The plasma-spray (PS) technique is currently a commercially available method for coating implant devices with HA. Although the PS technique is currently employed to produce coatings used clinically, the long-term stability of the coating/implant is questionable. PS coating implant devices with HA present disadvantages that affect the long-term stability and lifetime of implant. In spite of successful microstructural and structural results, these investigations of PS ceramic coatings on metal substrates have shown failure at the interface.(21),(22)

Pulsed laser deposition (PLD) has attracted interest due to its versatility and controllability as well as its ability to synthesize and deposit uniform films with an accurate control of the stoichiometry and crystallinjty (18),(23) jn a typjcaj pld process, a pulsed laser beam superficially vaporizes the surface of the target (a process usually described as ablation), and the vapor condenses on a substrate. One of the most important issues that can determine the success or failure of the deposition process is the film adhesion.

PLD allows for the control of the interface layer between the substrate material and the thin film, which in turn can be used to substantially improve the film adhesion to the substrate. Knowledge of the residual stresses (RS) induced in the coating and in the substrate during deposition is important in predicting the location of failures.(24-27) Moreover, it is difficult to obtain precise RS measurements on thin layer (1 urn thickness) coatings using laboratory X-ray diffraction. It is due to the high surface roughness of the metal substrate -- necessary for a good osteointegration -- compared to the HA thickness. As a matter of fact the surface topography strongly affects the bone response, and roughness can exert a major influence on the interaction between cells and a biomateriaFs surface. There is considerable evidence from in vitro and in vivo experiments that the cell biology, that is, shape and function, and the bone-implant interfacial shear strength are influenced by the substrate surface microtopography.(28-30) In some recent works of the author, the effect of HA coated by [sol-gel.sup.31] and by plasma [sprayed.sup.21] on a titanium surface was shown. This article gives a first mechanical approach on RS in titanium substrate before and after deposition on PLD HA. Therefore, only the state of residual stress in the substrates without and with HA coating has been determined by the [sin.sup.2] [phi] method by the X-ray laboratory. Structural and morphological studies were performed on both the uncoated and coated Ti substrates. No biological tests have been performed.

Materials and methods

Preparation of PLD HA/Ti

Some characteristics of titanium and HA are shown in Table 1. Ti (wt%: C 0.10 max, Fe 0.20 max, H 0.015 max, N 0.03 max, O 0.18 max, Ti rem) disks ([PHI] = 15 mm, thickness = 2 mm) were prepared with a final polishing by silicon carbide sandpaper (1200#) and finally treated chemically. The chemical etching consisted in a pretreatment by specimen immersion in 1 M sodium hydroxide (NaOH) and 0.5 M hydrogen peroxide ([H.sup.2] [O.sup.2]) at 75 [degrees] C for 10 min for cleaning and decontaminating the titanium surface from embedded particles and machining impurities. Treatment in 0.2 M oxalic acid ([H.sup.2] [C.sup.2] [O.sup.2]) at 85 [degrees] C for 10 min was performed in order to produce a microporous surface and an immersion in nitric acid was done for final passivation.
Table 1: Characteristic data of titanium and hydroxyapatite:
crystalline structure, space group, lattice parameters, coefficient
of thermal expansion (a), elastic modulus ([pounds sterling]),
Poisson's ratio (v), and radio crystallographic constants (Si
and S2)

    Structure  Space            a = b    c(nm)     a (p        E (MPa)
               group            (nm)                pm[degrees]
                                                     C)

HA  hcp        [P.sub.p]6mmc  0.94075  0.68775           14    117,000
Ti  Hcp        [P.sub.p]m     0.29505  0.46826          8.5    110,000

       V      [S.sub.1]=[(v/e).sub.hkl]   1/2[S.sub.2]=[(1+v/E
                  ([(MPa.sup.-1]) )            sub.hkl]
                                            ([(MPa.sub.-1])

HA  0.28                  2.39                  10.94
Ti  0.34                 -3.09                  12.18


A UV KrF * laser source ([lambda] = 248 nm, [iota] = 25 ns) was used which was placed outside the irradiation chamber (Fig. 1). The deposition conditions are collected in Table 2. To avoid target piercing during multipulse ablation, a rotational movement was applied. After deposition, the obtained structures were treated in a water vapor enriched atmosphere at 400 [degrees] C for 6 h to improve their crystalline state. (32)
Table 2: Experimental conditions for PLD coatings

Target                             HA
Substrate                          Ti etched Ti
Fluency, J/c[m.sup.2]              2.6
Dynamical pressure, Pa             30 Pa [H.sub.2] O vapors
Distance target-substrate, mm      40
Pulse repetition rate, Hz          2
Number of pulses 20,000
Substrate temperature, [degrees]C  400
Coating thickness, [mu]m           1.2


Characterization and analysis of PLD HA/Ti

Phase analysis of HA coatings was performed using grazing incidence ([iota] = 3 [degrees]) X-ray diffraction (GIXRD) with Cu /Ca radiation ([lambda] = 0.15418 nm). X-ray diffractograms were recorded step-by-step with a step size of 0.02 [degrees] at 40 kV, 40 mA, with an acquisition time of 20 s. The primary beam was delimited using a 3-mm diameter collimator. The limitation of X-ray divergence was achieved using 0.3 [degrees] vertical Soller slits mounted before the detector.

For structural investigation of HA, selected area electron diffraction (SAED) was performed with a Topcon EM 002B electron microscope operating at 200 kV with a 0.18 nm point-to-point resolution, equipped with a low dose camera and an Si/Li detector for analysis. A cross-section of HA/Ti interface was prepared by ion millin[g.sup.33] by cross-section transmission electron microscopy (XTEM). Morphological characterization of HA coating and HA/Ti interface were observed by field emission scanning electron microscopy (FESEM) using a JEOL JSM-6700F.

RS tensors were determined in the [alpha] - phase of titanium using the si [n.sup.2] [phi] metho[d.sup.34] without taking into account texture in the titanium substrate.

[FIGURE 1 OMITTED]

Diffraction scans were run with a generator tension of 30 kV and a generator current of 10 mA. Each diffraction scan was done at [phi] angles (0[degrees], [+ or -] 16.8[degrees]. [+ or -]24 [degrees], [+ or -]30 [degrees], [+ or -]35.2 [degrees], [+ or -]40.2 [degrees], [+ or -]45 [degrees]), and 0 angles (0 [degrees], 60 [degrees], 120 [degrees], 180 [degrees], 240 [degrees], 300 [degrees]) with a step size of 0.02 [degrees] and an acquisition time of 300 s. The titanium radiation ([lambda] = 0.27496 nm) and diffracting planes {110} were selected to perform the strain measurements on Ti substrate with and without PLD HA coating.

Results and discussion

Crystalline phase of HA by GIXRD (hep crystal symmetry) was detected and the evolution of the background noise shows an amorphous phase that can be identified as an amorphous HA phase. No impurity phases such as a-tricalcium phosphate (C [a.sup.4] [P.sup.2] P [O.sup.9],) (sup.2) TCP) or tetracalcium phosphate (C [a.sup.4] [p.sup.2] [O.sup.9], TTCP), etc. were identified, [alpha] (hep) and [beta] (bcc) phases of titanium substrate were also determined and confirmed by the XRD pattern (Fig. 2a). The SAED pattern (Fig. 2b) proves the crystallinity of HA with the characteristic (211) and (002) crystallographic planes of HA.

Under these deposition conditions (Table 2), in addition to atoms and ions, in most cases some droplets of target material are also deposited on the substrate surface. In most systems, the formation of large droplets can be reduced by using dense and smooth targets.36 However, the ablation of smaller droplets originating from the fast heating and cooling processes of the target cannot be completely avoided. The surface is compact and well-crystallized and exhibits a quite irregular morphology (Fig. 3). Some grain-like particles and droplets -- partially covered with a thin film and finally embedded in a bigger particle -- were observed by FESEM on the surface of the film.

[FIGURE 2 OMITTED]

Size of HA particulates ranges from 100 nm to 1 [mu] m and this is characteristic of PLD coatings. For biomedical applications, the presence of irregular morphology and structure enhances the cells' proliferation due to a surface extension. Furthermore, these surfaces are expected to be more susceptible to the natural remodeling processes when they are implanted in a living body. Biological responses to HA surfaces are influenced by the size, morphology, and structure of HA particles. (38,39) The nano-sized biological apatite found in bones has been reported to be involved in the natural bone remodeling process rather than being phagocytised. (40) As such, the synthesized nano-HA was suggested to be biocompatible. It has been observed that nano-HA supported the attachment and the spread of human osteoblast cells. (41) The surface roughness of the substrate is particularly important for HA coatings not only because a rough surface can provide increased wettability of the HA solution on the substrate, but also because mechanical interlocking between the HA-coated layer and substrate may be enhanced. (42) When bone cells are attached to a solid substrate their behavior and function depend on the physico-chemical and morphological properties of the biomaterial surface. (43) The biocompatibility of bioma-terials is very closely related to cell behavior on contact with them and particularly to cell adhesion to their surface. Surface characteristics of materials -- whether their topography, chemistry, or surface energy -- play an essential part in osteoblast adhesion on biomateri-als. Thus, attachment, adhesion, and spreading belong to the first phase of cell/material interactions and the quality of this first phase will influence the cell's capacity to proliferate and to differentiate itself on contact with the implant. Thereafter, the quality of this adhesion will influence their morphology, and their capacity for proliferation and differentiation. (44)

[FIGURE 3 OMITTED]

In Fig. 4 is shown an frtiSfcM micrograpn 01 me interface HA/Ti cross-section that provides an overview of the possible repercussion of the morphology in the mechanical properties of the coatings and offer information.

In applications for bioceramic coatings, bond coats should reduce the release of metal ions from the metallic substrate to the surrounding living tissue of mice (45) and human osteoblasts. (46)

A good bonding prevents a steep gradient in the coefficients of thermal expansion between Ti and HA that promotes the formation of strong tensile forces in the coating, giving rise to crack generation, chipping, and/or delamination. Moreover, it may be cushion damage (47) by cracking and delamination of the coating initiated by cyclic micromotions of the implant during movement of the patient in the initial phase of the healing process. (48) Thus, it is highly desirable to engineer the substrate/HA coating interface in such a way that by application of a suitable thin biocompatible bond coat layer the advantages addressed above can be realized.

[FIGURE 4 OMITTED]

Figure 5a shows the microstructure of HA/Ti film obtained by XTEM (dark field) prepared by ion milling. Columnar-like growth of the HA interlayer was observed, characteristic of directional deposition methods. This evidence was confirmed by a bright field image (Fig. 5b). HA is well defined, dense, and it has a compact block-like shape and HA/Ti is smooth (Fig. 5a).

The mechanical properties of the substrate HA coatings was studied by the non-destructive XRD residual stress evaluation by the [sin.sub.2] [psi] method. Figure 6 shows the compressive RS state ([sigma] [sub.11] = [sigma] [sub.22] = -329 [+ or -] 40 MPa) for the uncoated Ti substrates. The compressive state is linked to the mechanical and chemical polishing of the substrate. The errors on stress components were given for 95% of confidence ([+ or -]2 SD). Normal stress ([sigma] [sub.733]) was assumed to be equal to zero due to no lateral stress gradient along longitudinal and transverse directions. No shear stresses were observed.

[FIGURE 5 OMITTED]

[FIGURE 6 OMITTED]

The stress state on coated substrate is very low ([sigma] * 25 [+ or -] 18 MPa and [sigma] 22 = 20 [+ or -] 17 MPa). During the deposition of the HA, the RS are generated in the coating and in the substrate because of different reasons depending on the deposition technique, such as the thermal conditions, the growth of coatings, and structural mismatches. (49) One of the reasons is the superposition of the RS state of titanium (compressive state before coating) and the local mechanical equilibrium. The compressive stress on the surface of the substrate before coating is linked to the RS due to surface preparation by successive mechanical polishing and chemical etching.

In the PLD deposition process, the thermal stresses originated during the process. In addition, the difference in the coefficients of thermal expansion between the coating and the substrate leads to RS between these two partners by cooling down the sample from process temperature to room temperature. Furthermore, stresses are induced by the growth of the coating driven by capillary stresses during the processes. (49) The difference in crystal structure between the coating and the substrate is another reason leading to RS by structural mismatch (Table 1). (50) This change of stress in the substrate before and after coating could be due to the relaxation in the substrate under temperature.

Conclusions

The irregular surface of the PLD coating is a crystalline phase of HA. The found nano-HA can be involved in the natural bone remodeling process when implanted in the living body. A representative cross-section image has shown that the HA/Ti interface is significantly adhesive. The residual stress in the prepared substrate changed from a compressive state in the surface to close to zero after PLD coating.

Acknowledgments Sincere thanks are due Eng. Jacques Faerber IPCMS, Strasbourg for SEM measurements.

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A. Carrado Institut de Physique et Chimie des Materiaux de Strasbourg,

UMR 7504 UDS-CNRS, 23 rue du Loess, BP 43, 67034 Strasbourg cedex 2, France e-mail: adele.carrado@ipcms.u-strasbg.fr

DOI 10.1007/S11998-011-9355-9
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