Applied Surface Science 253 (2007) 7998–8002

www.elsevier.com/locate/apsusc

Combined laser/sol–gel synthesis of calcium silicate coating on

Ti–6Al–4V substrates for improved cell integration
N. Mirhosseini a,*, P.L. Crouse a, L. Li a, D. Garrod b
a
Laser Processing Research Centre, School of Mechanical, Aerospace and Civil Engineering, The University of Manchester,
Sackville Street Building, P.O. Box 88, Manchester M60 1QD, UK
b
Faculty of Life Sciences, University of Manchester, Michael Smith Building, Oxford Road, Manchester M13 9PT, UK

Available online 2 March 2007

Abstract
New studies have shown that tricalcium silicate powder is a bioactive material and can encourage bone–implant integration. This paper reports
the synthesis of Ca2SiO4 coating on Ti–6Al–4V samples by laser irradiation under submerged conditions. The results of using a 160–1500 LDL
1.5 kW diode laser (rectangular spot = 2.5 mm  3.5 mm, l = 808 and 940 nm with equal intensities) is reported. A number of experiments were
carried out varying laser parameters, such as scanning speed and laser power. Coatings are evaluated in terms of microstructure, elemental
composition (XRD), SEM and wettability. The in vitro biocompatibility of the samples is investigated by monitoring 2T3 osteoblast cell growth on
the samples.
# 2007 Elsevier B.V. All rights reserved.

Keywords: Laser surface treatment; High power diode laser; Tricalcium silicate; Bioglass; Biocompatibility; SEM; XRD; Wettability; Cell growth; 2T3 Osteoblast
cell

1. Introduction amorphous and porous structure. There is also the risk of viruses
being carried by this layer. It was shown that shortly after
Titanium is biocompatible, corrosion resistant and tissue implantation, titanium ions could be detected almost in all parts
compatible, making it an excellent choice for bio-implants. It is of organism; therefore, the oxide layer is not a real barrier [4].
ready to adsorb protein from biological fluids and creates a Mosser et al. [5] studied the instant oxide layer of a titanium
protein film. It also supports cell growth. Titanium and implant surface 5 years after implantation and they found that
specially Ti–6Al–4V among its alloys are among the most its thickness had increased from 5 nm (initial) to 200 nm
commonly used materials in both dental and orthopaedic indicating that the implant was continuously oxidizing in the
implantations [1–3]. organism and there was no protective layer to prevent further
Both pure titanium and Ti–6Al–4V are used in dental oxidization. It is also important that the oxide layer is porous
engineering as implants and dental prostheses. One big and its quality does not improve as its thickness increases.
advantage in this field is that titanium is completely neutral The above studies showed that the oxide layer appeared as a
in taste, which makes it the ideal choice for patients who show mediator between the implant and the organism instead of an
allergic or toxic reactions to other dental alloys. insulator. However, the situation would be completely different,
A pure titanium surface can never be seen in the real world as a if there is a crystalline insulation layer on the implant surface.
result of its high reactivity. It develops an oxide layer almost There are three possible methods to form a surface layer on an
immediately [2]. The native oxide layer is thin and has an implant [4]:

 material deposition (positive material transport);

* Corresponding author.
 material removal (negative material transport);
E-mail address: n.mirhosseini@postgrad.manchester.ac.uk  material transformation without the deposition or removal
(N. Mirhosseini). (transport-free material process).
0169-4332/$ – see front matter # 2007 Elsevier B.V. All rights reserved.
doi:10.1016/j.apsusc.2007.02.172
 N. Mirhosseini et al. / Applied Surface Science 253 (2007) 7998–8002 7999

Among surface coating methods on implants, plasma spray

and sol–gel are the most widely used techniques. In plasma
spraying, there is the risk of scaling and cracking as a result of
the high temperature process, which can be considered as a
disadvantage. Another possibility is the risk of scraped material
being attached into the bone–implant interface. The critical
stage in a sol–gel process is the heat treatment, since it affects
the quality, compactness and structure of the surface layer [6].
Among laser deposition techniques Nd:YAG laser has been
used to deposit hydroxyapatite (HA), which is one of the main
inorganic chemical elements of bone, on titanium by using
pulsed laser deposition [7]. Excimer laser was also reported to
grow a HA coating on Ti–6Al–4V by using laser ablation [8].
Some new glasses and glass ceramics based on SiO2, CaO
and P2O5 with varying elemental compositions were found to
be even better than HA materials in osseointegration. Joannia
et al. have deposited thin films of bioactive glass-ceramics on a
titanium substrate using a pulsed laser deposition technique.
They have reported observing calcium phosphate precipitate on
samples after immersing in simulated body fluid (SBF),
suggesting the coating to be bioactive [9].
In another study on bioactive glass, Zhao and Chang
synthesized tricalcium silicate (Ca3SiO5) powders by sol–gel
process. Soaking the powders in simulated body fluid (SBF) for
10 days showed a dense hydroxyapatite (HA) layer formed on
the surface indicating that the powder was bioactive [10].
Fig. 1. SEM pictures: (a) cross-section and (b) top view of surface coating.
In this paper, a combined laser/sol–gel technique is used to
deposite a calcium silicate coating and determining its effect on
osteoblast response. Contact angle measurement was carried out by using SBF in
a FTA188: contact angle and surface tension analyzer
2. Experimental procedure instrument (Fig. 1). SBF is a protein-free fluid having an ion
concentration nearly equal to human blood plasma (Na+
Shot-blasted Ti–6Al–4V samples (aerospace industry grade) 142.0 mM, K+ 5.0 mM, Mg2+ 1.5 mM, Ca2+ 2.5 mM, Cl
as rectangular sheets (60 mm  50 mm  10 mm) were used 148.8 mM, HCO3 4.2 mM, HPO42 1.0 mM, SO42 0.5 mM).
as the substrate. The calcium silicate powders were synthesized It was prepared by dissolving the above ionic species in
by sol–gel method by using Ca(NO3)2
4H2O and TEOS deionised water and buffering at pH 7.4 using HCl or tris base
(tetraethoxysilane) with an initial CaO/SiO2 molar ratio of 3. (H2NC(CH2OH)3) [11,12].
Nitric acid was added as catalyst. The solution was prepared by The in vitro test was carried out by osteoblast cell growth on
adding 0.5 mol TEOS in 200 ml water under continuous samples and analysing the cell behaviour and attachment. 2T3
stirring. Calculated amount of Ca(NO3)2
4H2O was added to osteoblast cells were cultured on coated and non-coated samples
the solution and the solution was stirred for 1 h. Then, the and then the cell growth was measured by using MTT assay [13].
solution was kept in an oven at 60 8C for 24 h until complete Among the existing methods to measure the cell viability,
gelation occurred [10]. the reduction of tetrazolium salts is now recognized as a safe
Experiments were carried out by using a 160–1500 LDL and accurate technique. In this test, the yellow tetrazolium salt
1.5 kW diode laser (rectangular spot = 2.5 mm  3.5 mm, (MTT: 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium
l = 808 and 940 nm with equal intensities). A thin layer of bromide) is reduced in metabolically active cells to generate
gel was applied on the substrate surface and then scanned by the insoluble purple formazan crystals, which are solubilized by the
laser beam. A number of experiments were carried out to obtain addition of a detergent. This is a colorimetric measurement
the optimum laser parameters. The best coating was made by system which determines the reduction of a tetrazolium
applying a laser power of 120 Wand a scanning speed of 1 mm/s. component (MTT) into an insoluble formazan product [13].
Surface roughness measurements were obtained by using a
laser surface profiler developed in house. The system uses the 3. Results
principal of optical triangulation to measure short distances
with a resolution of one micron without surface contact. System 3.1. SEM and XRD
handles the data by using a computer software and is capable of
measuring Ra values on vertical and horizontal lines of the SEM micrographs of the cross-section of coated samples
specified area on the sample. (Fig. 1) revealed a thin layer coated on the surface. The
8000 N. Mirhosseini et al. / Applied Surface Science 253 (2007) 7998–8002

Table 1
Surface roughness measurements
Non-coated Coated
Average Rax (mm) 2.32 11.15
Max. Rax (mm) 4.82 30.51
Min. Rax (mm) 0.94 1.28
Std. Rax 0.94 7.4

Fig. 2. XRD analysis.

thickness of the layer based on SEM pictures was estimated as

5 mm. Fig. 1a shows a top view of the coating on the surface and
Fig. 1b shows the layer cross-section. XRD analysis confirmed Fig. 4. Contact angle measurement: (a) coated surface and (b) non-coated
a crystalline structure of Ca–Si–O compound (mostly Ca2SiO4) surface.
on the surface (Fig. 2). An EDX analysis on the coated surface
also revealed Ca, Si and O as well as Ti as the substrate (Fig. 3).
A first-order evaluation of the maximum temperature reached 3.4. MTT analysis
(for numerical details see Ref. [14]), indicates a value of just
above 1300 8C, above the b-transus of Ti–6Al–4V, but below its In this study, 2T3 osteoblast cells were cultured in a 24-well
melting point and sufficient for the vitrification of the coating. plate for 24 h in normal medium. After that, cells were
This corresponds with the observed results. incubated with the MTT reagent for 2 h and finally acidified
isopropanol (isopropanol containing 0.04 M HCl), was added
3.2. Surface roughness to lyse the cells and solubilize the coloured crystals. The
absorbance of each well was then read using a plate reader at a
Equal sized areas were considered on all samples to obtain wavelength of 570 nm. As the number of viable cells is directly
the surface roughness. Three different areas were chosen on proportional to the amount of colour produced, this information
each sample to provide reliable results. Table 1 compares the can be used to generate a calibration curve of 2T3 osteoblast
average, maximum, minimum and standard deviation of Ra cell number versus absorbance. Using this curve, enables
values along the laser beam moving direction (Rax). accurate, straightforward quantification of changes in prolif-
eration. Fig. 5 shows the MTT calibration curve relating the
3.3. Contact angle absorbance and number of viable cells. An order two function
was fitted to the curve to enable finding the number of cells
Three contact angle values were taken from each sample. from an absorbance value.
The average contact angle was measured as 858 for a non- The second MTT assay was performed on coated and non-
coated sample and 62.58 for a coated sample. Fig. 4 shows two coated samples to find the effect of coating on 2T3 osteoblast
pictures taken from a coated (a) and a non-coated (b) sample cells growth. Samples were seeded with equal number of cells
and the resultant contact angle. (5  105) for 6 days and then were incubated with MTT reagent

Fig. 5. MTT calibration curve: absorbance vs. number of viable cells (on a
Fig. 3. EDX analysis. plastic substrate).
 N. Mirhosseini et al. / Applied Surface Science 253 (2007) 7998–8002 8001

Table 2 5. Conclusion
MTT assay results
Sample No. of viable cells Percent of increase in Tricalcium silicate powder has previously been recognised
per unit area (No./mm2) number of cells as a bioactive material. Combined laser/sol–gel synthesis of
Non-coated 81.2 – calcium silicate coating on Ti–6Al–4V samples by using a high
Coated 116.08 42.9 power diode laser increased the surface roughness compared to
shot-blasted Ti–6Al–4V. The process reduced the contact angle
between the surface and a drop of SBF, indicating an improved
for 2 h followed by adding acidified isopropanol. The final wettability. It also increased the 2T3 osteoblast cell growth on
solution was given a good mix to provide a homogenous colour samples surfaces by 42.9%.
and was read using a plate reader at a wavelength of 570 nm. By Cell attachment also was observed to be uniform and evenly
using the initial cell proliferation curve provided in step one, spread on coated samples. This is an important result for a
the number of viable cells on each samples were then successful implantation with reduced healing time and a lower
calculated. risk of implant rejection.
Table 2 shows the number of viable cells per unit area of
coated and non-coated samples and the percentage of increase Acknowledgements
in cell numbers compared to the untreated sample.
The authors would like to thank Dr. David Watteridge for his
continuous support during the biological tests, Mr. Tamer Ezz
4. Discussion for his technical help in the sol–gel/laser process and Mr. Turki
Al-khaldi for his help in contact angle measurement.
In this study, surface roughness measurements indicated that
coating has increased the surface roughness of the samples from
References
2.32 to 11.15 mm. SEM pictures and XRD results (Figs. 2 and
3) confirmed a fine crystalline structure containing Ca, Si and O [1] A. Goransson, E. Jansson, P. Tengvall, A. Wennerberg, Bone formation after
mostly (Ca2SiO4) on the surface which can be the reason for 4 weeks around blood-plasma-modified titanium implants with varying
surface roughening. surface topographies: an in vivo study, Biomaterials 24 (2003) 197–205.
Contact angle value was measured to indicate the surface [2] D.M. Brunette, et al., Titanium in Medicine, Springer, London, 2001.
wettability. Previous studies have shown that roughening the [3] T.M. Lee, E. Chang, C.Y. Yang, Attachment and proliferation of neonatal
rat calvarial osteoblasts on Ti6Al4V: effect of surface chemistries of the
surface improves the wettability and therefore, reduces the alloy, Biomaterials 25 (2004) 23–32.
contact angle [15]. Laser treatment has been recognised to be [4] C. Toth, G. Szabo, L. Kovacs, K. Vargha, J. Barabas, Z. Nemeth, Titanium
able to increase the surface energy of the substrate material implants with oxidized surfaces: the background and long-term results,
which is an important factor for improving the wettability. In Smart Mater. Struct. 11 (2002) 813–818.
[5] A. Mosser, C. Speisser, D. Muster, Surface physics methods for bioma-
the study of Schakenraad et al. [16], surface energy was
terials characterization, in: D. Muster (Ed.), Biomaterials—Hard Tissue
demonstrated as a governing factor in cell adhesion and Repair and Replacement, Elsevier, Amsterdam, 1992, pp. 143–152.
proliferation. [6] Melinda Mo, Successful Osseointegration and Implant Surfaces, March
The results from this study were found to be in line with 2003.
previous research. Table 1 shows that surface roughness was [7] H. Zeng, W.R. Lacefield, XPS, EDX and FTIR analysis of pulsed laser
increased by laser coating. Contact angle measurements deposited calcium phosphate bioceramic coatings: the effects of various
process parameters, Biomaterials 21 (2000) 23–30.
confirm that coating (roughening the surface) has decreased [8] L. Cleries, E. Martinez, J.M. Fernandez-Pradas, G. Sardin, J. Esteve, J.L.
the contact angle of SBF on the surface, indicating that Morenza, Mechanical properties of calcium phosphate coatings deposited
wettability and therefore, biocompatibility has been improved by laser ablation, Biomaterials 21 (2000) 967–971.
by laser coating. [9] E. Joannia, M.C. Ferro, C.C. Mardare, A.I. Mardare, J.R. Fernandes, S.C.
Previous studies revealed that higher surface roughness and de Almeida Pina, Pulsed laser deposition of SiO2–P2O5–CaO–MgO glass
coatings on titanium substrates, Mater. Res. 7 (3) (2004) 431–436.
higher surface energy can increase the number of cells adhering [10] W. Zhao, J. Chang, Sol–gel synthesis and in vitro bioactivity of tricalcium
to the substrate and resulting in enhanced cell activity [17,18]. silicate powders, Mater. Lett. 58 (2004) 2350–2353.
In addition to that, using bioactive coatings, such as HA and [11] T. Kokubo, et al., Apatite formation on ceramics, metals and polymers
bioactive glass also was found to encourage the material induced by a CaO SiO2 based glass in a simulated body fluid, in: W.
bioactivity [9,10,12]. The coating in this study was confirmed Bonfield, G.W. Hastings, K.E. Tanner (Eds.), Bioceramics, Butterworth–
Heinmann, Oxford, 1991.
to be a Ca–Si–O compound (mostly Ca2SiO4) by XRD analysis. [12] E.S. Thian, K.A. Khor, N.H. Loh, S.B. Tor, Processing of HA-coated Ti–
MTT assay proved that the coating has encouraged 2T3 6Al–4V by a ceramic slurry approach: an in vitro study, Biomaterials 22
osteoblast cell growth by 42.9% compared to a non-coated (2001) 1225–1232.
sample. This could be as a result of both surface roughening and [13] T. Mosman, Rapid colorimetric assay for cellular growth and survival:
bioactive surface composition created by laser coating application to proliferation and cytotoxic assay, J. Immunol. Methods 65
(1983) 55–63.
(Table 2). In addition to that, cell growth was observed to [14] M. Turner, P.L. Crouse, L. Li, Comparative interaction mechanisms for
have a rather uniform distribution on the coated surface which different laser systems with selected materials on titanium alloys, E-MRS
is desirable in implantation. 2006, Strasbourg, June 2006.
8002 N. Mirhosseini et al. / Applied Surface Science 253 (2007) 7998–8002

[15] R.N. Wenzel, Resistance to solid surfaces to wetting by water, Ind. Eng. [17] L. Hao, T.H. Wang, J. Lawrence, L. Li, Manipulation of the osteoblast
Chem. 28 (1936) 988–994. response to Ti–6Al–4V titanium alloy using a high power diode laser,
[16] J.M. Schakenraad, H.J. Busscher, C.R.H. Wildevuur, J. Arends, The Appl. Surf. Sci. 247 (2005) 602–606.
influence of substratum surface free energy on growth and spreading of [18] B. Feng, J. Weng, B.C. Yang, S.X. Qu, X.D. Zhang, Characterization of
human fibroblasts in the presence and absence of serum proteins, J. surface oxide films on titanium and adhesion of osteoblast, Biomaterials
Biomed. Mater. Res. 20 (1986) 773–784. 24 (2003) 4663–4670.