- Email: [email protected]
Cytocompatibility of titanium metal injection molding with various anodic oxidation post-treatments
Materials Science and Engineering C 32 (2012) 1919–1925
Contents lists available at SciVerse ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Cytocompatibility of titanium metal injection molding with various anodic oxidation post-treatments Clémence Demangel a, Delphine Auzène a, Muriel Vayssade b, Jean-Luc Duval b,⁎, Pascale Vigneron b, Marie-Danièle Nagel b, Jean-Claude Puippe c a b c
CRITT-MDTS, ZHT du Moulin Leblanc, 3, Bd Jean Delautre 08000 Charleville-Mézières, France Université de Technologie de Compiègne, UMR 7338 Biomécanique-Bioingénierie, BP 20529 60205 Compiègne Cedex, France Steiger Galvanotechnique, Route de Pra de Plan, 18 CH-1618 Châtel-St-Denis, Switzerland
a r t i c l e
i n f o
Article history: Received 20 September 2011 Received in revised form 26 April 2012 Accepted 22 May 2012 Available online 27 May 2012 Keywords: Metal injection molding Anodic oxidation Biocompatibility Titanium Implant
a b s t r a c t Metal injection molding (MIM) is a near net shape manufacturing method that allows for the production of components of small to moderate size and complex shape. MIM is a cost-effective and flexible manufacturing technique that provides a large innovative potential over existing methods for the industry of implantable devices. Commercially pure titanium (CP-Ti) samples were machined to the same shape as a composite feedstock with titanium and polyoxymethylene, and these metals were injected, debinded and sintered to assess comparative biological properties. Moreover, we treated MIM-Ti parts with BIOCOAT®, BIODIZE® and BIOCER®, three different anodic oxidation techniques that treat titanium using acid, alkaline and anion enriched electrolytes, respectively. Cytocompatibility as well as morphological and chemical features of surfaces was comparatively assessed on each sample, and the results revealed that MIM-Ti compared to CP-Ti demonstrated a specific surface topography with a higher roughness. MIM-Ti and BIOCER® samples significantly enhanced cell proliferation, cell adhesion and cell differentiation compared to CP-Ti. Interestingly, in the anodization post-treatment established in this study, we demonstrated the ability to improve osseointegration through anionic modification treatment. The excellent biological response we observed with MIM parts using the injection molding process represents a promising manufacturing method for the future implantable devices in direct contact with bones. © 2012 Elsevier B.V. All rights reserved.
1. Introduction Titanium and its alloys are widely used in the medical devices industry, due to a unique combination of properties including low density, high strength, corrosion resistance and biocompatibility. Titanium provides a good interface with bone, and as a consequence, titanium implants offer improved recovery and rehabilitation from injury and contribute to increased patient comfort. However, titanium production is hampered by the high cost of traditional manufacturing processes, and poor workability for complex shape production. Therefore, more efficient processing techniques such as metal injection molding (MIM) have been developed [1]. This technique was inspired from the methods developed for plastic injection molding, and is particularly relevant for complex, small size, near-net shape mass manufacturing. Furthermore it is cost-effective and enhances design flexibility without damaging mechanical properties, and therefore has relevant potential for biomedical applications. The MIM technique is carried out in four steps: (a) feedstock pellets are obtained by mixing a polymer binder and a metal powder, (b)
⁎ Corresponding author. Tel.: + 33 33 44 23 44 21; fax: + 33 33 44 23 79 42. E-mail address: [email protected] (J.-L. Duval). 0928-4931/$ – see front matter © 2012 Elsevier B.V. All rights reserved. doi:10.1016/j.msec.2012.05.037
feedstock is injected into a mold with the specific shape of the components, (c) binder is removed by catalytic and thermic methods, and (d) components are sintered in a furnace for grain cohesion and densification. In the literature, the biocompatibility of the MIM technique has only been tested for stainless steel (316L) injection molding [2]. In this study, we focused on titanium, the most frequently used metal in the biomaterial field. We evaluated the impact of the titanium MIM process and anodic oxidation post-treatments on cytocompatibility, with a particular focus on surface aspects. Since the MIM products reach a chemical composition close to raw materials, it should not affect cytocompatibility. Nevertheless, other features of MIM processing could have an impact on biological response. Because it is considered as a net-shape technique, special attention was paid to aspects of the component surface properties. Indeed it has been shown that surface roughness has a direct impact on cellular response [3]. Both chemical and structural parameters play a major role. It has been established that surface energy of a coating significantly promotes cell adhesion [4]. As a consequence, recent studies aim to distinguish the influence of the surface chemistry and topography on the cellular response [5]. The objective of our work was to compare the MIM technique to the existing approach, to fully assess its bio-application potentials and target its possible use in orthopedic medical devices. To this
1920
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
aim, we performed micrometric and nanometric topographies combined with chemical investigations of the surface. Then, we used bone explant cultures to study cytocompatibility by comparing cell behavior (migration, colonization and adhesion) between the different manufacturing techniques, which are conventional machining and metal injection molding. In addition, we assessed the ability of MC-3T3-E1 pre-osteoblasts to differentiate these conditions. All along, titanium surface modification with three different types of anodic oxidation treatments was compared to determine the optimal surface for biocompatibility. 2. Materials and methods 2.1. Surface preparation and characterization Commercially pure titanium (CP-Ti) (ACNIS, Villeurbanne, France) grade 4 bar (following surgical implants standards ISO 5832‐2 and ASTM F 67) was machined. Same dimension plates (11 mm× 55 mm× 3 mm) were injected and sintered from Catamold® Ti BASF (Ludwigshafen, Germany). Feedstock pellets composed of titanium spherical particles were embedded in plastic binder, predominantly polyoxymethylene (POM). Injection molding of the samples was performed on an Arburg (Lossburg, Germany) 320C 600–100 machine at a barrel temperature of 160 to 170 °C for the nozzle and a mold temperature of 130 °C. Material in the screw consisted of a plasticizing cylinder and mold which were specifically made to resist abrasion and corrosion due to the use of feedstock. The debinding stage was done in two steps: first we used catalytic binder removal with pure nitric acid >98% under nitrogen atmosphere at 120 °C. Treatment duration depended on sample thickness: about 1 h/mm. Second, a thermal debinding was performed at 600 °C for 3 h under argon (65 mbar) in the sintering furnace in order to remove the binder entirely. Sintering was carried out at 1300 °C under argon partial pressure (65 mbar) for 200 mn. Samples were set on zirconia plates into the furnace. Catalytic and sintering equipment were obtained from Elnik Systems (Cedar Grove, USA). Then MIM-Ti as well as CP-Ti samples underwent the same preparation steps: degreasing in an alkaline detergent, blasting with zirconia at 2 bar pressure, etching with a mixture of nitric and hydrofluoric acids (BIOETCH), and passivation in nitric acid. Three different surface treatments were applied on the sintered parts: a color anodization (BIOCOAT®), an alkaline anodization (BIODIZE®) and a glow discharge anodization (BIOCER®). These surface treatments are described in the literature [6–10] and their characteristics are summarized hereafter. The final rinsing occurred in biologically controlled water (b1 mesophylic aerobic germ per mL). The samples were dried in hot air and packaged in clean room (ISO Class 7). Sterilization with gamma rays (35 kGy) was then applied. The color anodization BIOCOAT® was essentially a titanium oxide layer TiO2 with evenly distributed thickness around the coated part. The layer was very adherent to the substrate and was generated by a transformation of the titanium metallic surface into an oxide phase. For this, the titanium part was immersed in an electrolyte consisting typically of diluted sulphuric acid or phosphoric acid and polarized anodically. Typical values of the coating thickness were between 30 and 300 nm, corresponding to 2–3 nm per volt, as described in Table 8.2 of Titanium in Medicine [11]. The coating acted as interference filter and its color was directly related to its thickness. The coating BIODIZE® consisted of an alkaline anodization according to the standard AMS 2488c. A titanium oxide TiO2 was formed by the anodic transformation of the titanium surface. The oxide content was higher at the top and decreased inwards to gradually approach the composition of the substrate. The thickness of the layer showed a gradient of oxide about 3 μm. The main characteristics of the alkaline anodizing consisted of good wear resistance, prevention and reduction of fretting wear, as well as improved resistance to fatigue by 15 to 20%. The glow discharge anodization BIOCER® was a treatment
under high voltage polarization where intense sparking occurred allowing the formation of titanium oxide. The anions present in the electrolyte were incorporated into the deposit in significant quantities, particularly phosphate (~20%), calcium (~20%) and magnesium (~7%). The typical coating thickness was between 5 and 10 μm. The deposit was amorphous and its topography presented a large amount of pores around 2 μm. MIM-Ti and CP-Ti sample features relative to surface micro and nano topography, microstructure and chemical composition were assessed with scanning electron microscopy with energy dispersive X-ray spectroscopy (SEM/EDX — Carl Zeiss, Nanterre, France), a topographic optical plateform Altisurf 500 (ALTIMET, Thonon-Les-Bains, France), an atomic force microscopy (AFM — CSM Instruments, Peseux, Switzerland), optical microscope, gas analyzers, and wettability apparatus. 2.2. Bone explant culture and assays Chick embryo organotypic culture technique developed in our laboratory [12,13] was used in this study. Bone tissue samples were obtained from the tibias of 14-day-old chick embryos. They were cut in 2 mm 3 explants and cultivated in semi-solid medium. The culture medium was Dulbecco's modified Eagle's medium (Gibco Invitrogen, Cergy-Pontoise, France) supplemented with 40% fetal bovine serum (FBS), 2% L-glutamine 200 mM, 0.15% penicillin and streptomycin mixture (100 UI/mL and 100 μg/mL). This complete medium was mixed v/v with buffered agar (2% Bacto-Agar Difco, Detroit, USA in Gey's solution). Each explant was cultured in contact with 1 cm 2 (10 mm × 10 mm square) of either Ti material surfaces to be tested or Thermanox® used as controls. Explants were grown for 14 days in a 100% humidified incubator at 37 °C with 5% CO2. Forty explants were cultured for each experimental condition. Two independent experiments were carried out. All cultured explants were stained with neutral red (2%) (Sigma Aldrich, St Quentin Fallavier, France) and cell-layers developed around the explants were measured using a stereo-microscope coupled with a camera. Cell migration area was calculated from the pictures using Image J analysis software [14]. Cell density and cell adhesion were assessed after removing the cells and counting them in a Coulter® Multisizer (Beckman-Coulter, Villepinte, France). Explants were removed from their substrata and cell-layers treated with 0.25% trypsin–EDTA (Gibco, Invitrogen, CergyPontoise, France). The cells that detached from the sample surface after incubation for 5, 10, 20, 30 and 60 min were harvested and counted. The samples were then placed in (0.25%) trypsin–EDTA for 15 min to remove any residual attached cells. The cell adhesion is expressed from an index that represents the area between the curve of cell dissociation (% of detached cells vs time) and the X axis. This value generally between 2000 and 6000 is inversely related to the cell adhesion. This is illustrated through a diagram producing three areas: strong adhesion (A b 3000), moderate adhesion (3000 b A b 4500) and weak adhesion (A > 4500). A histogram presents the migration surface in mm2. 2.3. Pre-osteoblast cultures and assays The pre-osteoblasts MC-3T3 have been widely used to assess the biocompatibility of titanium [15,16]. Pre-osteoblasts MC-3T3-E1 subclones 24 and 4 (ATCC CRL2595 and CRL2593) were cultured in minimum essential medium alpha without ascorbic acid (MEMalpha w/o ascorbic acid, Gibco Invitrogen, Cergy-Pontoise, France) supplemented with 10% FBS, L-glutamine 2 mM, penicillin (100 UI/mL) and streptomycin (100 μg/mL). 1×104 cells/cm2 were seeded on Ti material samples or on tissue culture polystyrene dishes (TCPS) (Nunc®, D. Dutscher, Brumath, France) used as a reference material. After 3 days in culture, the medium was supplemented with ascorbic acid (50 μM),
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
dexamethasone (10 − 8 M) and β glycerophosphate (10 mM) and changed every 2–3 days (all products were purchased from Invitrogen). To assess alkaline phosphatase (ALP) activity, after 10 days in culture, cells were washed briefly with Phosphate Buffer Saline (PBS) and fixed in 70% ethanol for 30 min at 4 °C. Cells were incubated for 1 h at 37 °C in a filtered staining solution made of 20 mg naphthol AS-BI phosphate dissolved in 1 mL N.N.dimethylformamide, 10 mL 200 mM Tris-200 mM maleic acid pH 8.5, 8 mL 160 mM NaCl and 6 mg fast red violet LB salt (all products were from Sigma Aldrich, St Quentin Fallavier, France). Positive red cells were observed in situ with a light microscope. In addition, ALP + cells were counted after detachment with 0.25% trypsin–EDTA treatment and cytospinning on glass slides. An alizarin red‐based assay of mineralization and immunodetection of Type I Collagen by adherent cells was performed 17 day postseeding. Some cell layers were rinsed with NaCl 0.9%, fixed in ethanol 70% for 30 min and after washing stained with alizarin red S 2% (Sigma) for 10 min. Observations were made using a light microscope. Other cell layers were fixed with paraformaldehyde 4% in PBS for 10 min, washed, and permeabilized with Triton X-100 0.5%. Cells were then incubated for 1 h with BSA 1% and then with polyclonal anticollagen I antibody (Chemicon AB749, 25 μg/mL), washed and incubated with rabbit secondary Cy3-conjugated IgG antibody (Jackson Immunoresearch, 111–165–144, 15 μg/mL) and stained with DAPI (1 μg/mL). Observations were made using Zeiss LSM 710 confocal microscope. 2.4. Statistical evaluation All statistical evaluations were performed by ANOVA using GraphPad InStat software and all the reported values are expressed as the mean± standard deviation. Tukey–Kramer multiple comparisons test and Mann–Whitney non-parametric test were used respectively for bone explants and MC-3T3-E1 cell results. A value of Pb 0.05 was considered statistically significant. 3. Results 3.1. Material and surface characterization We first investigated the BASF feedstock by SEM (20 kV), which revealed that particle diameters range from 2 μm to 35 μm. The gas characterization of the feedstock performed on gas analyzers showed that carbon content was 0.01% and oxygen rate was 0.12%. After sintering, the MIM material density was controlled and considered the primary determinant for the efficiency of the whole technique. The raw material density is typically around 4.50 g/cm 3, while in our experiment the sintered parts reached 4.32 g/cm 3 (as compared with BASF data requirement which is 4.20 g/cm 3). This density corresponded to a porosity of 4%, and pores were mostly localized on the surface of the material, with a denser and less porous core. On the other hand, MIM-Ti exhibited a specific surface aspect with larger pores, and few pores smaller than 10 μm in size on SEM micrograph (Fig. 1). MIM-Ti properties were assessed according to standard ASTM F 67 for medical devices, presenting a monophase alpha microstructure with a grain size No. 5.0 (corresponding to an average diameter of 63.5 μm). The carbon content after sintering had a range between 0.09 and 0.11% when oxygen content was 0.23%. The surface chemistry was assessed by energy dispersive X-ray spectroscopy. On each sample, titanium was the main element while carbon, oxygen and nitrogen were also detected. Calcium, phosphate and magnesium were distributed homogeneously on the surface of BIODIZE®. A notable amount of zirconium and aluminum was noticed on both CP-Ti due to grit-blasting residue. This contrasted with the absence of residue contamination on the MIM processing samples.
1921
Fig. 1. SEM micrograph on MIM-Ti surface.
No difference of contact angle was detected by wettability measurements on CP-Ti and MIM-Ti. As shown in our laboratory (unpublished study), BIOCER® treatment was characterized by a high surface energy recognizable by its high wettability when immersed in deionized water (positive meniscus). During the same test, pure titanium presented a low wettability (negative meniscus). Topographical surface examinations revealed significant differences between CP-Ti and MIM-Ti samples. At a micrometer level, the MIM-Ti surface presented a higher roughness than CP-Ti (Table 1) and peak dimension for MIM-Ti was higher and more irregular than CP-Ti (Fig. 2). These results were confirmed by the comparative Rz (average maximum height of the profile) values, which showed that peak height was about 5 μm on CP-Ti, whereas the profile of MIM-Ti had an amplitude of 20 μm and could reach amplitudes up to 80 μm in more irregular areas. At a nanometer level, the roughness was also higher for MIM-Ti by Ra (roughness average)=0.35 μm and Rms (root mean square)=0.59 μm whereas for CP-Ti, Ra=0.27 μm and Rms=0.44 μm. The AFM topography of MIM-Ti appeared deeper but smoother than CP-Ti (Fig. 3). On the CP-Ti surface, some alterations are present. Indeed, as titanium alloys present a poor workability, the grinding operation should have caused tearing and plastic deformation on the surface.
3.2. Post treatment effect on surfaces We next examined whether the specific post-treatments impacted the surface morphology. Anodic oxidation BIOCOAT® treatment with acid electrolyte did not significantly change the surface shape but coloration was observed. The oxide film, about 0.2 μm was too thin to be observed by SEM on cross section. Changes were also observed in the specific surface aspects treated with BIODIZE® and BIOCER® (Fig. 4). For instance, on surfaces treated with alkaline anodic oxidation BIODIZE®, we noted the appearance of an oxide film with micro-porosities, and surface roughness was slightly increased. Surfaces that received the BIOCER® treatment, however, developed an oxide layer that was nearly 10 μm thick with numerous porosities
Table 1 Roughness parameters on the five studied surfaces (scan length: 4.8 mm — cut-off 0. 8 mm). (μm)
MIM
CP-Ti
BIOCOAT®
BIODIZE®
BIOCER®
Ra Rz
1.7 11.6
0.7 5.0
1.8 12.2
2.0 15.5
1.8 11.3
1922
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
Fig. 2. Topographic 3D view of CP-Ti (a) and MIM-Ti (b) surfaces, 1 × 1 mm.
about 2 μm diameter, and there was an increase in surface roughness of the specimen.
Cell adhesion shown in Fig. 6b was weak on Thermanox® while it was in the mid range on CP-Ti, MIM-Ti, BIODIZE® and BIOCOAT®. In contrast, a strong adhesion was observed on BIOCER®.
3.3. Bone explant behavior 3.4. MC-3T3-E1 pre-osteoblast behavior Explants (≈2 mm 3) from tibias of 14 day-old chick embryos were cultured in semi-solid medium for 14 days in contact with Ti samples or Thermanox® control. Depending on the surface properties of the substratum, cells could escape from each bone explant, spread, migrate and proliferate, then develop a cell layer around the explant. The area of each cell layer was measured and cells were removed and counted. Cell migration areas are presented in Fig. 5a. Cells migrated similarly on all the Ti sample surfaces, but migration was significantly increased on Thermanox® control (P b 0.001). The mean number of cells per explant is shown in Fig. 5b. Cell numbers were higher on Thermanox® control than on all Ti sample surfaces (P b 0.001), and all the MIM-treated samples except for BIODIZE® showed a higher number of cells per explant than CP-Ti (P b 0.001). Fig. 6a presents the cell colonization/mm 2 of material surfaces: Thermanox® (948 cells/mm²) = MIM-Ti (1035 cells/mm²) > BIOCER® (846 cells/mm²) (P b 0.001) > BIOCOAT® (746 cells/mm²) (Pb 0.05)>BIODIZE® (635 cells/mm²) (P>0.05)>CP-Ti (411 cells/mm²) (P b 0.001). We also observed that MIM process significantly increased the cell density (P b 0.001).
Cell capability to differentiate to osteoblasts was assessed through different complementary assays carried out on Ti material samples and TCPS used as a control. In each case, when ALP activity was detected in situ after 8 days in culture (results not shown), MC-3T3E1 (subclone 24) cell layers showed ALP positive cells. Fig. 7a illustrates the quantitative evaluation assessed from cytospun cells 10 days post seeding. ALP + cell percentage in decreasing order was: MIM-Ti = BIOCER®> BIODIZE® = BIOCOAT® (P b 0.01) > CP-Ti = TCPS (P b 0.001). All layers of MC-3T3-E1 (subclone 4) exhibited calcium deposition after 17 days in culture. Alizarin red staining was especially well developed and showed numerous foci of concentration on BIOCER® (Fig. 7b). Moreover, collagen I synthesis was detected from all the experimental and control samples. Fig. 7c illustrates immunodetection of collagen I on BIOCER®. 4. Discussion This study highlights the comparison of the specific surface roughness of MIM parts, to determine the optimal surface for biocompatibility
Fig. 3. Atomic force microscopy 20 μm × 20 μm on CP-Ti (a) and MIM-Ti (b) (performed by CSM Instruments).
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
1923
Fig. 4. SEM micrograph of BIODIZE® (a) and BIOCER® (b) surfaces.
of titanium parts. One factor that is known to increase the roughness of titanium components is the initial powder particle size; however, besides increasing roughness, it has been reported that higher particle size could decrease carbon contamination [17], and does not have a direct influence on mechanical properties.
Fig. 5. (a) Migration area measurements (mm²) on the different biomaterials. ***P b 0.001 compared to Thermanox®. (b) Number of cells per explant calculation on the different biomaterials. The results are the mean ± SD of two independent experiments (n = 36). ***P b 0.001 significantly different of Thermanox®, ###P b 0.001 significantly different of CP-Ti.
Furthermore, as titanium is a reactive material and binders employed are commonly composed with organic elements, particular attention should be given to the carbon and oxygen contamination that can damage the mechanical properties of the material [18]. Moreover, commercial feedstocks revealed relative high carbon and oxygen content as also stated by Nyberg et al. [19]. In recent studies alternative binders and
Fig. 6. (a) Cell density calculation on the different biomaterials. (b) Cell adhesion assessment on the different biomaterials. The results are the mean ± SD of two independent experiments (n = 36). ***P b 0.001.
1924
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
Fig. 7. (a) Quantitative analysis of ALP expression onto the different biomaterials. **P b 0.01, ***P b 0.001. (b) Calcium deposition onto BIOCER®. (c) Collagen synthesized by the cells on BIOCER® (nuclei are stained with DAPI).
powders have been investigated to avoid interstitial contamination [20–22]. While the oxygen content in material rises, mechanical stiffness is increased at the expense of the ductility. Therefore, to preserve ductility, an oxygen rate below 0.25% is advised [23]. Other parameters such as time, temperature, vacuum level or even gas atmosphere used during sintering or removal binder steps have a significant influence on the component contamination, highlighting the importance of choosing the optimal parameters to reach the desired properties of MIM-Ti materials. Therefore, we investigated the mechanical properties of MIM-Ti under a range of manufacturing and treatment modalities. Parameters used in this present study were determined previously, and ASTM F 67 requirements were fulfilled. It has been previously reported in the literature that porosity rate of MIM-Ti is commonly about 3 to 5% [24], and that expected density is between 94 and 98% [18]. Thus, in this study, the MIM green body has been fully debinded and sintered, and MIM pore size of less than 10 μm was in accordance with previous reports [23]. Hence, MIM technique seemed to be very promising but more work was needed to reach the optimal level of decreasing grain size and gas contamination. Considering biological response, two complementary approaches were used to assess the cytocompatibility of Ti material samples. Bone explant cultures were performed to compare the ability of bone cells to migrate, to colonize the surfaces and to adhere, depending on the treatment applied to the Ti [26]. We used organotypic culture for this study because it gives the possibility to test the organ (bone) that will be impacted at the implantation site, and will be in direct contact with the material. This organotypic technique is consistent with the results obtained in cell culture [25].
Our results indicate that bone cell migration and surface colonization were promoted on Thermanox® (control), which is in accordance with other previous studies [26,27]. Cell migration was roughly the same on the different Ti material samples regardless of the treatment applied, however cell colonization was improved by the MIM process (MIM-Ti, BIOCER®, BIODIZE®, BIOCOAT®)(Fig. 5b and 6a). It has been demonstrated that MIM process does not alter L929 fibroblast behavior [28] and it is well known that Ti implant surface topography influences osteoblastic proliferation, differentiation and extracellular matrix protein expression [29]. Cell adhesion was poor on Thermanox® control, much improved on MIM-Ti, BIODIZE® and BIOCOAT®, and dramatically increased on BIOCER®. From bone explant cultures, we can therefore conclude that MIM process improves the cytocompatibility of Ti, and that among the Ti material samples tested, BIOCER® appears to be ideally suited to favor bone cell adhesion. The second approach consisted of verifying the capability of preosteoblasts to support formation, secretion and mineralization of extracellular bone matrix when seeded on the different Ti surfaces, with TCPS used as a control. For that purpose, we first checked ALP positivity as an early marker of the osteoblast differentiation. A significant increase of ALP + cells was observed from MIM-Ti and BIOCER®. Moreover, the previous observations in favor of BIOCER® were reinforced by an increased mineralization by calcium deposition. Increased collagen I synthesis reinforced the observation of efficient osteoblast differentiation. Such observations are consistent with that described by Whiteside et al. [30] who reported that titanium-based coatings produced by plasma electrolytic oxidation with cratered topography enhance human primary osteoblast proliferation, as
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
well as cell adhesion and synthesis of collagen. Osteoblast behavior might, at least in part, depend on plasma protein adsorption on TiO2 [31]; further studies will be necessary to compare protein adsorption onto the surfaces of our Ti material samples. As the chemistry and wettability of CP-Ti and MIM-Ti samples are nearly the same (except some ceramic residues embedded in raw material), the improved biological response of MIM-Ti could be explained in a large part by the increase in specific surface features, namely the higher roughness. Indeed, it has been largely reported that surface architecture has a positive impact on the biological reaction [32–34]. Several parameters are known to influence the roughness in the MIM technique including: the ratio of binder in the feedstock [35], the size distribution of the metallic powder, and the surface aspect of the mold cavities. For BIOCER® treatment, both topographic and chemical considerations are involved, which demonstrates the potential of increasing biocompatibility by adding a surface modification treatment. In conclusion, the cytocompatibility studies suggest that MIM and BIOCER® techniques improve both osteoblast adhesion and function. BIOCER® deposition is also particularly well suited to the incorporation of bioactive substances by post-processing and represents very promising strategies to improve Ti implant bone integration. 5. Conclusion The cytocompatibility study on MIM processing pure titanium has demonstrated that this manufacturing process is suitable for the implantable device industry by enhancing biological properties in terms of adhesion, proliferation and differentiation. It has also been demonstrated that MIM technique respects standard requirements for mechanical strength, chemical composition and microstructure. Some processing parameters could be adapted in the future to find the best compromise and optimization between all these properties, including the optimization of surface parameters to favor cytocompatibility. Moreover, surface modification treatments on titanium by anodic oxidation processes have demonstrated a similar achievement to sintered material; especially for BIOCER® treatment that enriches the surface with mineral elements and clearly promotes cell adhesion. Generally this study highlights how surface aspects like roughness and chemical issue could affect cell activity. As they obtained good biological response, MIM-Ti and BIOCER® treatment would have a potential benefit in the orthopedic and dental devices manufacturing field. In this way, MIM technique presents economical advantages for such small, complex products, manufactured in high volume. Finally MIM technique versatility may be an improvement over techniques for larger purposes like sintering of new titanium alloys. Acknowledgments To CSM Instruments SA, Switzerland — Dr Guillaume Berthout, for their implication in this work with the AFM measurements for nanotopography characterizations.
1925
Appendix A. Supplementary data Supplementary data to this article can be found online at http:// dx.doi.org/10.1016/j.msec.2012.05.037. References [1] Y. Itoh, H. Miura, T. Uematsu, K. Sato, J. Solid Mech. Mater. Eng. 3 (12) (2009) 1297–1305. [2] S.H. Zhu, G.H. Wang, Y.Z. Zhao, Y.M. Li, K.C. Zhou, B.Y. Huang, J. Cent. South Univ. Technol. 12 (1) (2005) 9–11. [3] L. Le Guéhennec, A. Soueidan, P. Layrolle, Y. Amouriq, Dent. Mater. 23 (2007) 844–854. [4] Y.M. Zhang, P. Bataillon-Linez, P. Huang, Y.M. Zhao, Y. Han, M. Traisnel, K.W. Xu, H.F. Hildebrand, J. Biomed. Mater. Res. A 68A (3) (2004) 383–391. [5] D.M. Dohan Ehrenfest, P.G. Coelho, B.S. Kang, Y.T. Sul, T. Albrektsson, Trends Biotechnol. 28 (4) (2010) 198–206. [6] J.C. Puippe, Ulmer Gespräch, Galvanotechnik, 95(12), Eugen G. Leuze Verlag, 2004, pp. 2908–2915. [7] N.K. Kuromoto, R.A. Simão, G.A. Soares, Mater. Charact. 58 (2) (2007) 114–121. [8] D. Capek, M.P. Gigandet, M. Masmoudi, M. Wery, O. Banakh, Surf. Coat. Technol. 202 (8) (2008) 1379–1384. [9] Y.T. Sul, C.B. Johansson, Y. Jeong, T. Albrektsson, Med. Eng. Phys. 23 (5) (2001) 329–346. [10] X. Zhu, K.H. Kim, Y. Jeong, Biomaterials 22 (16) (2001) 2199–2206. [11] D.M. Brunette, P. Tengvall, M. Textor, P. Thomsen, Titanium in Medicine, 8, Springer-Verlag, ISBN: 3-540-66936-1, 2001, pp. 233–266. [12] J.L. Duval, M. Letort, M.F. Sigot-Luizard, Biomaterials 9 (1988) 155–161. [13] E. Leclerc, J.L. Duval, I. Pezron, F. Nadaud, Mater. Sci. Eng., C 29 (2009) 861–868. [14] M.D. Abramoff, P.J. Magalhaes, S.J. Ram, Biophotonics Int. 11 (7) (2004) 36–42. [15] B.G. Keselowsky, D.M. Collard, A.J. Garcia, Proc. Natl. Acad. Sci. U. S. A. 102 (2005) 5953–5957. [16] V. Muhonen, C. Fauveaux, G. Olivera, P. Vigneron, A. Danilov, M.D. Nagel, J. Tuukkanen, J. Biomed. Mater. Res. 88A (2009) 787–796. [17] R.M. German, PIM Int. 3 (4) (2009) 21–37. [18] K.S. Weil, E. Nyberg, K. Simmons, J. Mater. Process. Technol. 176 (1–3) (2006) 205–209. [19] E. Nyberg, M. Miller, K. Simmons, K. Weil, Met. Powder Rep. 60 (10) (2005) 8–13. [20] W. Limberg, E. Aust, T. Ebel, R. Gerling, B. Oger, Met. Powder Rep. 60 (3) (2005) 22–27. [21] G. Shibo, Q. Xuanhui, H. Xinbo, Z. Ting, D. Bohua, J. Mater. Process. Technol. 173 (3) (2006) 310–314. [22] E. Nyberg, M. Miller, K. Simmons, K.S. Weil, Mater. Sci. Eng., C 25 (3) (2005) 336–342. [23] T. Ebel, PIM Int. 2 (2) (2008) 21–30. [24] Y. Itoh, H. Miura, T. Uematsu, K. Sato, M. Niinomi, T. Ozawa, Met. Powder Rep. 64 (5) (2009) 17–20. [25] J.L. Duval, R. Warocquier-Clérout, M.F. Sigot-Luizard, Cells Mater. 2 (3) (1992) 179–192. [26] C. Bussy, R. Verhoef, A. Haeger, M. Morra, J.L. Duval, P. Vigneron, A. Bensoussan, E. Velzenberger, G. Cascardo, C. Cassinelli, H. Schols, J.P. Knox, M.D. Nagel, J. Biomed. Mater. Res. A 86A (2008) 597–606. [27] E. Delvat, D.M. Gordin, T. Gloriant, J.L. Duval, M.D. Nagel, J. Mech. Behav. Biomed. Mater. 1 (4) (2008) 345–351. [28] R. Ibrahim, M. Azmirruddin, M. Jabir, M. Ridhuan, M. Muhamad, M. Rafiq, N.A. Kasim, S. Muhamad, Adv. Sci. Technol. 76 (2010) 247–252. [29] M. Jayaraman, U. Meyer, M. Bühner, U. Joos, H.P. Wiesmann, Biomaterials 25 (4) (2004) 625–631. [30] P. Whiteside, E. Matykina, J.E. Gough, P. Skeldon, G.E. Thompson, J. Biomed. Mater. Res. Part A 94 (1) (2010) 38–46. [31] S.R. Sousa, M. Lamghari, P. Sampaio, P. Moradas-Ferreira, M.A. Barbosa, J. Biomed. Mater. Res. Part A 84 (2) (2008) 281–290. [32] A. Wenneberg, Int. J. Mach. Tools Manuf. 38 (5–6) (1998) 657–662. [33] M.J. Kim, M.U. Choi, C.W. Kim, Biomaterials 27 (2006) 5502–5511. [34] O. Zinger, G. Zhao, Z. Schwartz, J. Simpson, M. Wieland, D. Landolt, B. Boyan, Biomaterials 26 (2005) 1837–1847. [35] O.M. Ferri, T. Ebel, R. Bormann, Mater. Sci. Eng., A 527 (2010) 1800–1805.
Contents lists available at SciVerse ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Cytocompatibility of titanium metal injection molding with various anodic oxidation post-treatments Clémence Demangel a, Delphine Auzène a, Muriel Vayssade b, Jean-Luc Duval b,⁎, Pascale Vigneron b, Marie-Danièle Nagel b, Jean-Claude Puippe c a b c
CRITT-MDTS, ZHT du Moulin Leblanc, 3, Bd Jean Delautre 08000 Charleville-Mézières, France Université de Technologie de Compiègne, UMR 7338 Biomécanique-Bioingénierie, BP 20529 60205 Compiègne Cedex, France Steiger Galvanotechnique, Route de Pra de Plan, 18 CH-1618 Châtel-St-Denis, Switzerland
a r t i c l e
i n f o
Article history: Received 20 September 2011 Received in revised form 26 April 2012 Accepted 22 May 2012 Available online 27 May 2012 Keywords: Metal injection molding Anodic oxidation Biocompatibility Titanium Implant
a b s t r a c t Metal injection molding (MIM) is a near net shape manufacturing method that allows for the production of components of small to moderate size and complex shape. MIM is a cost-effective and flexible manufacturing technique that provides a large innovative potential over existing methods for the industry of implantable devices. Commercially pure titanium (CP-Ti) samples were machined to the same shape as a composite feedstock with titanium and polyoxymethylene, and these metals were injected, debinded and sintered to assess comparative biological properties. Moreover, we treated MIM-Ti parts with BIOCOAT®, BIODIZE® and BIOCER®, three different anodic oxidation techniques that treat titanium using acid, alkaline and anion enriched electrolytes, respectively. Cytocompatibility as well as morphological and chemical features of surfaces was comparatively assessed on each sample, and the results revealed that MIM-Ti compared to CP-Ti demonstrated a specific surface topography with a higher roughness. MIM-Ti and BIOCER® samples significantly enhanced cell proliferation, cell adhesion and cell differentiation compared to CP-Ti. Interestingly, in the anodization post-treatment established in this study, we demonstrated the ability to improve osseointegration through anionic modification treatment. The excellent biological response we observed with MIM parts using the injection molding process represents a promising manufacturing method for the future implantable devices in direct contact with bones. © 2012 Elsevier B.V. All rights reserved.
1. Introduction Titanium and its alloys are widely used in the medical devices industry, due to a unique combination of properties including low density, high strength, corrosion resistance and biocompatibility. Titanium provides a good interface with bone, and as a consequence, titanium implants offer improved recovery and rehabilitation from injury and contribute to increased patient comfort. However, titanium production is hampered by the high cost of traditional manufacturing processes, and poor workability for complex shape production. Therefore, more efficient processing techniques such as metal injection molding (MIM) have been developed [1]. This technique was inspired from the methods developed for plastic injection molding, and is particularly relevant for complex, small size, near-net shape mass manufacturing. Furthermore it is cost-effective and enhances design flexibility without damaging mechanical properties, and therefore has relevant potential for biomedical applications. The MIM technique is carried out in four steps: (a) feedstock pellets are obtained by mixing a polymer binder and a metal powder, (b)
⁎ Corresponding author. Tel.: + 33 33 44 23 44 21; fax: + 33 33 44 23 79 42. E-mail address: [email protected] (J.-L. Duval). 0928-4931/$ – see front matter © 2012 Elsevier B.V. All rights reserved. doi:10.1016/j.msec.2012.05.037
feedstock is injected into a mold with the specific shape of the components, (c) binder is removed by catalytic and thermic methods, and (d) components are sintered in a furnace for grain cohesion and densification. In the literature, the biocompatibility of the MIM technique has only been tested for stainless steel (316L) injection molding [2]. In this study, we focused on titanium, the most frequently used metal in the biomaterial field. We evaluated the impact of the titanium MIM process and anodic oxidation post-treatments on cytocompatibility, with a particular focus on surface aspects. Since the MIM products reach a chemical composition close to raw materials, it should not affect cytocompatibility. Nevertheless, other features of MIM processing could have an impact on biological response. Because it is considered as a net-shape technique, special attention was paid to aspects of the component surface properties. Indeed it has been shown that surface roughness has a direct impact on cellular response [3]. Both chemical and structural parameters play a major role. It has been established that surface energy of a coating significantly promotes cell adhesion [4]. As a consequence, recent studies aim to distinguish the influence of the surface chemistry and topography on the cellular response [5]. The objective of our work was to compare the MIM technique to the existing approach, to fully assess its bio-application potentials and target its possible use in orthopedic medical devices. To this
1920
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
aim, we performed micrometric and nanometric topographies combined with chemical investigations of the surface. Then, we used bone explant cultures to study cytocompatibility by comparing cell behavior (migration, colonization and adhesion) between the different manufacturing techniques, which are conventional machining and metal injection molding. In addition, we assessed the ability of MC-3T3-E1 pre-osteoblasts to differentiate these conditions. All along, titanium surface modification with three different types of anodic oxidation treatments was compared to determine the optimal surface for biocompatibility. 2. Materials and methods 2.1. Surface preparation and characterization Commercially pure titanium (CP-Ti) (ACNIS, Villeurbanne, France) grade 4 bar (following surgical implants standards ISO 5832‐2 and ASTM F 67) was machined. Same dimension plates (11 mm× 55 mm× 3 mm) were injected and sintered from Catamold® Ti BASF (Ludwigshafen, Germany). Feedstock pellets composed of titanium spherical particles were embedded in plastic binder, predominantly polyoxymethylene (POM). Injection molding of the samples was performed on an Arburg (Lossburg, Germany) 320C 600–100 machine at a barrel temperature of 160 to 170 °C for the nozzle and a mold temperature of 130 °C. Material in the screw consisted of a plasticizing cylinder and mold which were specifically made to resist abrasion and corrosion due to the use of feedstock. The debinding stage was done in two steps: first we used catalytic binder removal with pure nitric acid >98% under nitrogen atmosphere at 120 °C. Treatment duration depended on sample thickness: about 1 h/mm. Second, a thermal debinding was performed at 600 °C for 3 h under argon (65 mbar) in the sintering furnace in order to remove the binder entirely. Sintering was carried out at 1300 °C under argon partial pressure (65 mbar) for 200 mn. Samples were set on zirconia plates into the furnace. Catalytic and sintering equipment were obtained from Elnik Systems (Cedar Grove, USA). Then MIM-Ti as well as CP-Ti samples underwent the same preparation steps: degreasing in an alkaline detergent, blasting with zirconia at 2 bar pressure, etching with a mixture of nitric and hydrofluoric acids (BIOETCH), and passivation in nitric acid. Three different surface treatments were applied on the sintered parts: a color anodization (BIOCOAT®), an alkaline anodization (BIODIZE®) and a glow discharge anodization (BIOCER®). These surface treatments are described in the literature [6–10] and their characteristics are summarized hereafter. The final rinsing occurred in biologically controlled water (b1 mesophylic aerobic germ per mL). The samples were dried in hot air and packaged in clean room (ISO Class 7). Sterilization with gamma rays (35 kGy) was then applied. The color anodization BIOCOAT® was essentially a titanium oxide layer TiO2 with evenly distributed thickness around the coated part. The layer was very adherent to the substrate and was generated by a transformation of the titanium metallic surface into an oxide phase. For this, the titanium part was immersed in an electrolyte consisting typically of diluted sulphuric acid or phosphoric acid and polarized anodically. Typical values of the coating thickness were between 30 and 300 nm, corresponding to 2–3 nm per volt, as described in Table 8.2 of Titanium in Medicine [11]. The coating acted as interference filter and its color was directly related to its thickness. The coating BIODIZE® consisted of an alkaline anodization according to the standard AMS 2488c. A titanium oxide TiO2 was formed by the anodic transformation of the titanium surface. The oxide content was higher at the top and decreased inwards to gradually approach the composition of the substrate. The thickness of the layer showed a gradient of oxide about 3 μm. The main characteristics of the alkaline anodizing consisted of good wear resistance, prevention and reduction of fretting wear, as well as improved resistance to fatigue by 15 to 20%. The glow discharge anodization BIOCER® was a treatment
under high voltage polarization where intense sparking occurred allowing the formation of titanium oxide. The anions present in the electrolyte were incorporated into the deposit in significant quantities, particularly phosphate (~20%), calcium (~20%) and magnesium (~7%). The typical coating thickness was between 5 and 10 μm. The deposit was amorphous and its topography presented a large amount of pores around 2 μm. MIM-Ti and CP-Ti sample features relative to surface micro and nano topography, microstructure and chemical composition were assessed with scanning electron microscopy with energy dispersive X-ray spectroscopy (SEM/EDX — Carl Zeiss, Nanterre, France), a topographic optical plateform Altisurf 500 (ALTIMET, Thonon-Les-Bains, France), an atomic force microscopy (AFM — CSM Instruments, Peseux, Switzerland), optical microscope, gas analyzers, and wettability apparatus. 2.2. Bone explant culture and assays Chick embryo organotypic culture technique developed in our laboratory [12,13] was used in this study. Bone tissue samples were obtained from the tibias of 14-day-old chick embryos. They were cut in 2 mm 3 explants and cultivated in semi-solid medium. The culture medium was Dulbecco's modified Eagle's medium (Gibco Invitrogen, Cergy-Pontoise, France) supplemented with 40% fetal bovine serum (FBS), 2% L-glutamine 200 mM, 0.15% penicillin and streptomycin mixture (100 UI/mL and 100 μg/mL). This complete medium was mixed v/v with buffered agar (2% Bacto-Agar Difco, Detroit, USA in Gey's solution). Each explant was cultured in contact with 1 cm 2 (10 mm × 10 mm square) of either Ti material surfaces to be tested or Thermanox® used as controls. Explants were grown for 14 days in a 100% humidified incubator at 37 °C with 5% CO2. Forty explants were cultured for each experimental condition. Two independent experiments were carried out. All cultured explants were stained with neutral red (2%) (Sigma Aldrich, St Quentin Fallavier, France) and cell-layers developed around the explants were measured using a stereo-microscope coupled with a camera. Cell migration area was calculated from the pictures using Image J analysis software [14]. Cell density and cell adhesion were assessed after removing the cells and counting them in a Coulter® Multisizer (Beckman-Coulter, Villepinte, France). Explants were removed from their substrata and cell-layers treated with 0.25% trypsin–EDTA (Gibco, Invitrogen, CergyPontoise, France). The cells that detached from the sample surface after incubation for 5, 10, 20, 30 and 60 min were harvested and counted. The samples were then placed in (0.25%) trypsin–EDTA for 15 min to remove any residual attached cells. The cell adhesion is expressed from an index that represents the area between the curve of cell dissociation (% of detached cells vs time) and the X axis. This value generally between 2000 and 6000 is inversely related to the cell adhesion. This is illustrated through a diagram producing three areas: strong adhesion (A b 3000), moderate adhesion (3000 b A b 4500) and weak adhesion (A > 4500). A histogram presents the migration surface in mm2. 2.3. Pre-osteoblast cultures and assays The pre-osteoblasts MC-3T3 have been widely used to assess the biocompatibility of titanium [15,16]. Pre-osteoblasts MC-3T3-E1 subclones 24 and 4 (ATCC CRL2595 and CRL2593) were cultured in minimum essential medium alpha without ascorbic acid (MEMalpha w/o ascorbic acid, Gibco Invitrogen, Cergy-Pontoise, France) supplemented with 10% FBS, L-glutamine 2 mM, penicillin (100 UI/mL) and streptomycin (100 μg/mL). 1×104 cells/cm2 were seeded on Ti material samples or on tissue culture polystyrene dishes (TCPS) (Nunc®, D. Dutscher, Brumath, France) used as a reference material. After 3 days in culture, the medium was supplemented with ascorbic acid (50 μM),
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
dexamethasone (10 − 8 M) and β glycerophosphate (10 mM) and changed every 2–3 days (all products were purchased from Invitrogen). To assess alkaline phosphatase (ALP) activity, after 10 days in culture, cells were washed briefly with Phosphate Buffer Saline (PBS) and fixed in 70% ethanol for 30 min at 4 °C. Cells were incubated for 1 h at 37 °C in a filtered staining solution made of 20 mg naphthol AS-BI phosphate dissolved in 1 mL N.N.dimethylformamide, 10 mL 200 mM Tris-200 mM maleic acid pH 8.5, 8 mL 160 mM NaCl and 6 mg fast red violet LB salt (all products were from Sigma Aldrich, St Quentin Fallavier, France). Positive red cells were observed in situ with a light microscope. In addition, ALP + cells were counted after detachment with 0.25% trypsin–EDTA treatment and cytospinning on glass slides. An alizarin red‐based assay of mineralization and immunodetection of Type I Collagen by adherent cells was performed 17 day postseeding. Some cell layers were rinsed with NaCl 0.9%, fixed in ethanol 70% for 30 min and after washing stained with alizarin red S 2% (Sigma) for 10 min. Observations were made using a light microscope. Other cell layers were fixed with paraformaldehyde 4% in PBS for 10 min, washed, and permeabilized with Triton X-100 0.5%. Cells were then incubated for 1 h with BSA 1% and then with polyclonal anticollagen I antibody (Chemicon AB749, 25 μg/mL), washed and incubated with rabbit secondary Cy3-conjugated IgG antibody (Jackson Immunoresearch, 111–165–144, 15 μg/mL) and stained with DAPI (1 μg/mL). Observations were made using Zeiss LSM 710 confocal microscope. 2.4. Statistical evaluation All statistical evaluations were performed by ANOVA using GraphPad InStat software and all the reported values are expressed as the mean± standard deviation. Tukey–Kramer multiple comparisons test and Mann–Whitney non-parametric test were used respectively for bone explants and MC-3T3-E1 cell results. A value of Pb 0.05 was considered statistically significant. 3. Results 3.1. Material and surface characterization We first investigated the BASF feedstock by SEM (20 kV), which revealed that particle diameters range from 2 μm to 35 μm. The gas characterization of the feedstock performed on gas analyzers showed that carbon content was 0.01% and oxygen rate was 0.12%. After sintering, the MIM material density was controlled and considered the primary determinant for the efficiency of the whole technique. The raw material density is typically around 4.50 g/cm 3, while in our experiment the sintered parts reached 4.32 g/cm 3 (as compared with BASF data requirement which is 4.20 g/cm 3). This density corresponded to a porosity of 4%, and pores were mostly localized on the surface of the material, with a denser and less porous core. On the other hand, MIM-Ti exhibited a specific surface aspect with larger pores, and few pores smaller than 10 μm in size on SEM micrograph (Fig. 1). MIM-Ti properties were assessed according to standard ASTM F 67 for medical devices, presenting a monophase alpha microstructure with a grain size No. 5.0 (corresponding to an average diameter of 63.5 μm). The carbon content after sintering had a range between 0.09 and 0.11% when oxygen content was 0.23%. The surface chemistry was assessed by energy dispersive X-ray spectroscopy. On each sample, titanium was the main element while carbon, oxygen and nitrogen were also detected. Calcium, phosphate and magnesium were distributed homogeneously on the surface of BIODIZE®. A notable amount of zirconium and aluminum was noticed on both CP-Ti due to grit-blasting residue. This contrasted with the absence of residue contamination on the MIM processing samples.
1921
Fig. 1. SEM micrograph on MIM-Ti surface.
No difference of contact angle was detected by wettability measurements on CP-Ti and MIM-Ti. As shown in our laboratory (unpublished study), BIOCER® treatment was characterized by a high surface energy recognizable by its high wettability when immersed in deionized water (positive meniscus). During the same test, pure titanium presented a low wettability (negative meniscus). Topographical surface examinations revealed significant differences between CP-Ti and MIM-Ti samples. At a micrometer level, the MIM-Ti surface presented a higher roughness than CP-Ti (Table 1) and peak dimension for MIM-Ti was higher and more irregular than CP-Ti (Fig. 2). These results were confirmed by the comparative Rz (average maximum height of the profile) values, which showed that peak height was about 5 μm on CP-Ti, whereas the profile of MIM-Ti had an amplitude of 20 μm and could reach amplitudes up to 80 μm in more irregular areas. At a nanometer level, the roughness was also higher for MIM-Ti by Ra (roughness average)=0.35 μm and Rms (root mean square)=0.59 μm whereas for CP-Ti, Ra=0.27 μm and Rms=0.44 μm. The AFM topography of MIM-Ti appeared deeper but smoother than CP-Ti (Fig. 3). On the CP-Ti surface, some alterations are present. Indeed, as titanium alloys present a poor workability, the grinding operation should have caused tearing and plastic deformation on the surface.
3.2. Post treatment effect on surfaces We next examined whether the specific post-treatments impacted the surface morphology. Anodic oxidation BIOCOAT® treatment with acid electrolyte did not significantly change the surface shape but coloration was observed. The oxide film, about 0.2 μm was too thin to be observed by SEM on cross section. Changes were also observed in the specific surface aspects treated with BIODIZE® and BIOCER® (Fig. 4). For instance, on surfaces treated with alkaline anodic oxidation BIODIZE®, we noted the appearance of an oxide film with micro-porosities, and surface roughness was slightly increased. Surfaces that received the BIOCER® treatment, however, developed an oxide layer that was nearly 10 μm thick with numerous porosities
Table 1 Roughness parameters on the five studied surfaces (scan length: 4.8 mm — cut-off 0. 8 mm). (μm)
MIM
CP-Ti
BIOCOAT®
BIODIZE®
BIOCER®
Ra Rz
1.7 11.6
0.7 5.0
1.8 12.2
2.0 15.5
1.8 11.3
1922
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
Fig. 2. Topographic 3D view of CP-Ti (a) and MIM-Ti (b) surfaces, 1 × 1 mm.
about 2 μm diameter, and there was an increase in surface roughness of the specimen.
Cell adhesion shown in Fig. 6b was weak on Thermanox® while it was in the mid range on CP-Ti, MIM-Ti, BIODIZE® and BIOCOAT®. In contrast, a strong adhesion was observed on BIOCER®.
3.3. Bone explant behavior 3.4. MC-3T3-E1 pre-osteoblast behavior Explants (≈2 mm 3) from tibias of 14 day-old chick embryos were cultured in semi-solid medium for 14 days in contact with Ti samples or Thermanox® control. Depending on the surface properties of the substratum, cells could escape from each bone explant, spread, migrate and proliferate, then develop a cell layer around the explant. The area of each cell layer was measured and cells were removed and counted. Cell migration areas are presented in Fig. 5a. Cells migrated similarly on all the Ti sample surfaces, but migration was significantly increased on Thermanox® control (P b 0.001). The mean number of cells per explant is shown in Fig. 5b. Cell numbers were higher on Thermanox® control than on all Ti sample surfaces (P b 0.001), and all the MIM-treated samples except for BIODIZE® showed a higher number of cells per explant than CP-Ti (P b 0.001). Fig. 6a presents the cell colonization/mm 2 of material surfaces: Thermanox® (948 cells/mm²) = MIM-Ti (1035 cells/mm²) > BIOCER® (846 cells/mm²) (P b 0.001) > BIOCOAT® (746 cells/mm²) (Pb 0.05)>BIODIZE® (635 cells/mm²) (P>0.05)>CP-Ti (411 cells/mm²) (P b 0.001). We also observed that MIM process significantly increased the cell density (P b 0.001).
Cell capability to differentiate to osteoblasts was assessed through different complementary assays carried out on Ti material samples and TCPS used as a control. In each case, when ALP activity was detected in situ after 8 days in culture (results not shown), MC-3T3E1 (subclone 24) cell layers showed ALP positive cells. Fig. 7a illustrates the quantitative evaluation assessed from cytospun cells 10 days post seeding. ALP + cell percentage in decreasing order was: MIM-Ti = BIOCER®> BIODIZE® = BIOCOAT® (P b 0.01) > CP-Ti = TCPS (P b 0.001). All layers of MC-3T3-E1 (subclone 4) exhibited calcium deposition after 17 days in culture. Alizarin red staining was especially well developed and showed numerous foci of concentration on BIOCER® (Fig. 7b). Moreover, collagen I synthesis was detected from all the experimental and control samples. Fig. 7c illustrates immunodetection of collagen I on BIOCER®. 4. Discussion This study highlights the comparison of the specific surface roughness of MIM parts, to determine the optimal surface for biocompatibility
Fig. 3. Atomic force microscopy 20 μm × 20 μm on CP-Ti (a) and MIM-Ti (b) (performed by CSM Instruments).
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
1923
Fig. 4. SEM micrograph of BIODIZE® (a) and BIOCER® (b) surfaces.
of titanium parts. One factor that is known to increase the roughness of titanium components is the initial powder particle size; however, besides increasing roughness, it has been reported that higher particle size could decrease carbon contamination [17], and does not have a direct influence on mechanical properties.
Fig. 5. (a) Migration area measurements (mm²) on the different biomaterials. ***P b 0.001 compared to Thermanox®. (b) Number of cells per explant calculation on the different biomaterials. The results are the mean ± SD of two independent experiments (n = 36). ***P b 0.001 significantly different of Thermanox®, ###P b 0.001 significantly different of CP-Ti.
Furthermore, as titanium is a reactive material and binders employed are commonly composed with organic elements, particular attention should be given to the carbon and oxygen contamination that can damage the mechanical properties of the material [18]. Moreover, commercial feedstocks revealed relative high carbon and oxygen content as also stated by Nyberg et al. [19]. In recent studies alternative binders and
Fig. 6. (a) Cell density calculation on the different biomaterials. (b) Cell adhesion assessment on the different biomaterials. The results are the mean ± SD of two independent experiments (n = 36). ***P b 0.001.
1924
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
Fig. 7. (a) Quantitative analysis of ALP expression onto the different biomaterials. **P b 0.01, ***P b 0.001. (b) Calcium deposition onto BIOCER®. (c) Collagen synthesized by the cells on BIOCER® (nuclei are stained with DAPI).
powders have been investigated to avoid interstitial contamination [20–22]. While the oxygen content in material rises, mechanical stiffness is increased at the expense of the ductility. Therefore, to preserve ductility, an oxygen rate below 0.25% is advised [23]. Other parameters such as time, temperature, vacuum level or even gas atmosphere used during sintering or removal binder steps have a significant influence on the component contamination, highlighting the importance of choosing the optimal parameters to reach the desired properties of MIM-Ti materials. Therefore, we investigated the mechanical properties of MIM-Ti under a range of manufacturing and treatment modalities. Parameters used in this present study were determined previously, and ASTM F 67 requirements were fulfilled. It has been previously reported in the literature that porosity rate of MIM-Ti is commonly about 3 to 5% [24], and that expected density is between 94 and 98% [18]. Thus, in this study, the MIM green body has been fully debinded and sintered, and MIM pore size of less than 10 μm was in accordance with previous reports [23]. Hence, MIM technique seemed to be very promising but more work was needed to reach the optimal level of decreasing grain size and gas contamination. Considering biological response, two complementary approaches were used to assess the cytocompatibility of Ti material samples. Bone explant cultures were performed to compare the ability of bone cells to migrate, to colonize the surfaces and to adhere, depending on the treatment applied to the Ti [26]. We used organotypic culture for this study because it gives the possibility to test the organ (bone) that will be impacted at the implantation site, and will be in direct contact with the material. This organotypic technique is consistent with the results obtained in cell culture [25].
Our results indicate that bone cell migration and surface colonization were promoted on Thermanox® (control), which is in accordance with other previous studies [26,27]. Cell migration was roughly the same on the different Ti material samples regardless of the treatment applied, however cell colonization was improved by the MIM process (MIM-Ti, BIOCER®, BIODIZE®, BIOCOAT®)(Fig. 5b and 6a). It has been demonstrated that MIM process does not alter L929 fibroblast behavior [28] and it is well known that Ti implant surface topography influences osteoblastic proliferation, differentiation and extracellular matrix protein expression [29]. Cell adhesion was poor on Thermanox® control, much improved on MIM-Ti, BIODIZE® and BIOCOAT®, and dramatically increased on BIOCER®. From bone explant cultures, we can therefore conclude that MIM process improves the cytocompatibility of Ti, and that among the Ti material samples tested, BIOCER® appears to be ideally suited to favor bone cell adhesion. The second approach consisted of verifying the capability of preosteoblasts to support formation, secretion and mineralization of extracellular bone matrix when seeded on the different Ti surfaces, with TCPS used as a control. For that purpose, we first checked ALP positivity as an early marker of the osteoblast differentiation. A significant increase of ALP + cells was observed from MIM-Ti and BIOCER®. Moreover, the previous observations in favor of BIOCER® were reinforced by an increased mineralization by calcium deposition. Increased collagen I synthesis reinforced the observation of efficient osteoblast differentiation. Such observations are consistent with that described by Whiteside et al. [30] who reported that titanium-based coatings produced by plasma electrolytic oxidation with cratered topography enhance human primary osteoblast proliferation, as
C. Demangel et al. / Materials Science and Engineering C 32 (2012) 1919–1925
well as cell adhesion and synthesis of collagen. Osteoblast behavior might, at least in part, depend on plasma protein adsorption on TiO2 [31]; further studies will be necessary to compare protein adsorption onto the surfaces of our Ti material samples. As the chemistry and wettability of CP-Ti and MIM-Ti samples are nearly the same (except some ceramic residues embedded in raw material), the improved biological response of MIM-Ti could be explained in a large part by the increase in specific surface features, namely the higher roughness. Indeed, it has been largely reported that surface architecture has a positive impact on the biological reaction [32–34]. Several parameters are known to influence the roughness in the MIM technique including: the ratio of binder in the feedstock [35], the size distribution of the metallic powder, and the surface aspect of the mold cavities. For BIOCER® treatment, both topographic and chemical considerations are involved, which demonstrates the potential of increasing biocompatibility by adding a surface modification treatment. In conclusion, the cytocompatibility studies suggest that MIM and BIOCER® techniques improve both osteoblast adhesion and function. BIOCER® deposition is also particularly well suited to the incorporation of bioactive substances by post-processing and represents very promising strategies to improve Ti implant bone integration. 5. Conclusion The cytocompatibility study on MIM processing pure titanium has demonstrated that this manufacturing process is suitable for the implantable device industry by enhancing biological properties in terms of adhesion, proliferation and differentiation. It has also been demonstrated that MIM technique respects standard requirements for mechanical strength, chemical composition and microstructure. Some processing parameters could be adapted in the future to find the best compromise and optimization between all these properties, including the optimization of surface parameters to favor cytocompatibility. Moreover, surface modification treatments on titanium by anodic oxidation processes have demonstrated a similar achievement to sintered material; especially for BIOCER® treatment that enriches the surface with mineral elements and clearly promotes cell adhesion. Generally this study highlights how surface aspects like roughness and chemical issue could affect cell activity. As they obtained good biological response, MIM-Ti and BIOCER® treatment would have a potential benefit in the orthopedic and dental devices manufacturing field. In this way, MIM technique presents economical advantages for such small, complex products, manufactured in high volume. Finally MIM technique versatility may be an improvement over techniques for larger purposes like sintering of new titanium alloys. Acknowledgments To CSM Instruments SA, Switzerland — Dr Guillaume Berthout, for their implication in this work with the AFM measurements for nanotopography characterizations.
1925
Appendix A. Supplementary data Supplementary data to this article can be found online at http:// dx.doi.org/10.1016/j.msec.2012.05.037. References [1] Y. Itoh, H. Miura, T. Uematsu, K. Sato, J. Solid Mech. Mater. Eng. 3 (12) (2009) 1297–1305. [2] S.H. Zhu, G.H. Wang, Y.Z. Zhao, Y.M. Li, K.C. Zhou, B.Y. Huang, J. Cent. South Univ. Technol. 12 (1) (2005) 9–11. [3] L. Le Guéhennec, A. Soueidan, P. Layrolle, Y. Amouriq, Dent. Mater. 23 (2007) 844–854. [4] Y.M. Zhang, P. Bataillon-Linez, P. Huang, Y.M. Zhao, Y. Han, M. Traisnel, K.W. Xu, H.F. Hildebrand, J. Biomed. Mater. Res. A 68A (3) (2004) 383–391. [5] D.M. Dohan Ehrenfest, P.G. Coelho, B.S. Kang, Y.T. Sul, T. Albrektsson, Trends Biotechnol. 28 (4) (2010) 198–206. [6] J.C. Puippe, Ulmer Gespräch, Galvanotechnik, 95(12), Eugen G. Leuze Verlag, 2004, pp. 2908–2915. [7] N.K. Kuromoto, R.A. Simão, G.A. Soares, Mater. Charact. 58 (2) (2007) 114–121. [8] D. Capek, M.P. Gigandet, M. Masmoudi, M. Wery, O. Banakh, Surf. Coat. Technol. 202 (8) (2008) 1379–1384. [9] Y.T. Sul, C.B. Johansson, Y. Jeong, T. Albrektsson, Med. Eng. Phys. 23 (5) (2001) 329–346. [10] X. Zhu, K.H. Kim, Y. Jeong, Biomaterials 22 (16) (2001) 2199–2206. [11] D.M. Brunette, P. Tengvall, M. Textor, P. Thomsen, Titanium in Medicine, 8, Springer-Verlag, ISBN: 3-540-66936-1, 2001, pp. 233–266. [12] J.L. Duval, M. Letort, M.F. Sigot-Luizard, Biomaterials 9 (1988) 155–161. [13] E. Leclerc, J.L. Duval, I. Pezron, F. Nadaud, Mater. Sci. Eng., C 29 (2009) 861–868. [14] M.D. Abramoff, P.J. Magalhaes, S.J. Ram, Biophotonics Int. 11 (7) (2004) 36–42. [15] B.G. Keselowsky, D.M. Collard, A.J. Garcia, Proc. Natl. Acad. Sci. U. S. A. 102 (2005) 5953–5957. [16] V. Muhonen, C. Fauveaux, G. Olivera, P. Vigneron, A. Danilov, M.D. Nagel, J. Tuukkanen, J. Biomed. Mater. Res. 88A (2009) 787–796. [17] R.M. German, PIM Int. 3 (4) (2009) 21–37. [18] K.S. Weil, E. Nyberg, K. Simmons, J. Mater. Process. Technol. 176 (1–3) (2006) 205–209. [19] E. Nyberg, M. Miller, K. Simmons, K. Weil, Met. Powder Rep. 60 (10) (2005) 8–13. [20] W. Limberg, E. Aust, T. Ebel, R. Gerling, B. Oger, Met. Powder Rep. 60 (3) (2005) 22–27. [21] G. Shibo, Q. Xuanhui, H. Xinbo, Z. Ting, D. Bohua, J. Mater. Process. Technol. 173 (3) (2006) 310–314. [22] E. Nyberg, M. Miller, K. Simmons, K.S. Weil, Mater. Sci. Eng., C 25 (3) (2005) 336–342. [23] T. Ebel, PIM Int. 2 (2) (2008) 21–30. [24] Y. Itoh, H. Miura, T. Uematsu, K. Sato, M. Niinomi, T. Ozawa, Met. Powder Rep. 64 (5) (2009) 17–20. [25] J.L. Duval, R. Warocquier-Clérout, M.F. Sigot-Luizard, Cells Mater. 2 (3) (1992) 179–192. [26] C. Bussy, R. Verhoef, A. Haeger, M. Morra, J.L. Duval, P. Vigneron, A. Bensoussan, E. Velzenberger, G. Cascardo, C. Cassinelli, H. Schols, J.P. Knox, M.D. Nagel, J. Biomed. Mater. Res. A 86A (2008) 597–606. [27] E. Delvat, D.M. Gordin, T. Gloriant, J.L. Duval, M.D. Nagel, J. Mech. Behav. Biomed. Mater. 1 (4) (2008) 345–351. [28] R. Ibrahim, M. Azmirruddin, M. Jabir, M. Ridhuan, M. Muhamad, M. Rafiq, N.A. Kasim, S. Muhamad, Adv. Sci. Technol. 76 (2010) 247–252. [29] M. Jayaraman, U. Meyer, M. Bühner, U. Joos, H.P. Wiesmann, Biomaterials 25 (4) (2004) 625–631. [30] P. Whiteside, E. Matykina, J.E. Gough, P. Skeldon, G.E. Thompson, J. Biomed. Mater. Res. Part A 94 (1) (2010) 38–46. [31] S.R. Sousa, M. Lamghari, P. Sampaio, P. Moradas-Ferreira, M.A. Barbosa, J. Biomed. Mater. Res. Part A 84 (2) (2008) 281–290. [32] A. Wenneberg, Int. J. Mach. Tools Manuf. 38 (5–6) (1998) 657–662. [33] M.J. Kim, M.U. Choi, C.W. Kim, Biomaterials 27 (2006) 5502–5511. [34] O. Zinger, G. Zhao, Z. Schwartz, J. Simpson, M. Wieland, D. Landolt, B. Boyan, Biomaterials 26 (2005) 1837–1847. [35] O.M. Ferri, T. Ebel, R. Bormann, Mater. Sci. Eng., A 527 (2010) 1800–1805.