Calcium Phosphate Coating Produced by a Sputter Deposition Process
Joo L. Ong, Yunzhi Yang, Sunho Oh, Mark Appleford, Weihui Chen, Yongeing Liu, Kyo-Han Kim, Sangwon Park, Jeol Bumgardner, Warren Haggard, C. Mauli Agrawal, David L. Carner, and Namsik Oh
Abstract The properties of implant surfaces play critical roles in inducing a biological response. In the case of dental and orthopedic implants, deposition of calcium phosphate (CaP) coatings on these implant surfaces are often employed as means of enhancing implant osseointegration with the bone. Although most implants are coated using a plasma spraying process, sputtering is currently being accepted by some implant vendors as one of the means for depositing thin CaP coatings on dental and orthopedic implants. Acceptance of the sputtering technology and recent research are indications that the sputtering process is promising and has potential for eliminating some of the problems associated with the plasma-spraying process. This chapter discusses some of the various modes of sputtering, properties of thin CaP coatings, and the biological responses to these coatings in vitro and in vivo. The limitations and strengths of the sputtering process are also addressed.
coating [4–8]. In fact, HA coatings have continuously been reported to enhance clinical success since Furlong and Osborn’s first clinical trials [9], with three current studies reporting a less than 2% failure rate during a mean follow-up study of 10 years [6, 10, 11].
The current commercial process used by many vendors to deposit CaP or HA coatings on metallic implants is reported to be plasma spraying or arc plasma spraying [12, 13]. At the present time, minimal requirements for HA coatings have been described in the U.S. Food and Drug Administration (FDA) guide-lines as well as in the International Organization for Standardization guideguide-lines [14, 15]. Although continuous improvements in plasma-spraying technology and indications of clinical success with the use of plasma-sprayed HA-coated implants have been reported, their use has initiated substantial controversy in the field of dental and orthopedic implantology. Reported problems have included variations or poor coating–metal adhesion strength, nonuniformity in coating thickness among vendors, alteration in structural and chemical properties during the coating process, and nonuniformity in coating density [3–9, 16, 17]. As shown in Fig. 7.1, failure of the plasma-sprayed HA coating was observed at 12 weeks’ implantation and 1 year’ loading in a dog’s mandible after pullout testing. Fracture was observed at the Ti–HA coating interface as well as within the coating. As a result of the poor coating–metal adhesion strength, there is controversy over the benefits and success of HA-coated implants. How-ever, problems observed with plasma-sprayed HA-coated implants are not short-comings inherent in the rationale for HA coating but, rather, in the plasma-spray technology currently used for depositing HA coatings on implant surfaces.
Aside from the inherent coating technology, there is no standard guideline for producing HA coatings on implant surfaces. Investigations have suggested that all HA coatings are not the same, and variations in coating properties have attributed to many conflicting animal and clinical observations [4–6]. Findings
Fig. 7.1 Scanning electron microscopy (SEM) photomicrographs of a fractured interface.
a Plasma-sprayed hydroxyapatite (HA) coating after 12 weeks of implantation.
b Plasma-sprayed HA coating after 1 year of loading. C, coating; B, bone; Ti, titanium substrate. Arrows indicate cracks at the coating–metal interface after pullout testing
of plasma-sprayed HA-coated implants suggesting promotion of early implant stability are not universal, and these equivocal findings likely stem from the fact that HA coatings are poorly characterized. Typical plasma-sprayed HA coat-ings have been reported to consist primarily of partially dehydrated HA, with amorphous CaP and other more soluble CaP phases such as tricalcium phos-phate (TCP) produced during the high-temperature deposition process. In addition, the crystallinity for plasma-sprayed HA coatings produced under a normal deposition process is approximately 65%, even though the plasma-spraying technology has the capability of producing coatings with crystallinity in the range of 30% to 70%. At the extreme, some implant vendors (e.g., Zimmer Dental, Carlsbad, CA, USA) have increased their HA coating crystal-linity beyond 70% on their implant surfaces.
The ratio of HA to TCP has been reported to be crucial for bone regenera-tion. It is known that CaP phases as well as coating crystallinity affected the dissolution rate of an HA coating. An increase in amorphous and TCP phases predispose the coatings to higher dissolution rates. In addition, the formation of bone apatite-like minerals is promoted in the presence of dissolved calcium and phosphorus ions localized in the pericellular environment. Similarly, the pre-sence of high Ca ion concentration is known to induce apoptosis of bone-forming cells and the formation of osteoclasts. As shown in Fig. 7.2 and by other investigators [18, 19], reduced cell viability and osteoblast differentiation were observed as a result of rapid dissolution. As an example of the effect of coating dissolution on bone responses in vivo, Fig. 7.3 shows bone formation to be highly dependent on coating crystallinity [20]. As indicated in the figure,
Fig. 7.2 Cell viability of ATTC CRL 1486 human embryonic palatal mesenchymal cells, an osteoblast precursor cell line, as measured by [3-(4, 5-dimethylthiazol-2-yl)-5-(3-carboxy-methoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium, inner salt] (MTS) assay
bone–implant interfacial strength was significantly affected by varying the HA coating crystallinity. These studies suggested a further need to investigate coat-ing crystallinity that would result in optimum dissolution properties and ulti-mately provide maximum bone–implant interfacial strength. Given the vast number of deposition processes [21–64], as shown in Table 7.1, this chapter focuses on sputtering technology.
Fig. 7.3 Push-out strength at the implant–bone interface using a rat femur model
Table 7.1 Hydroxyapatite coating produced using various deposition technologies
Technique Thickness Advantages Disadvantages Refs.
Thermal spraying
30–200 mm High deposition rates; low cost
Line of sight technique; high temperatures induce decomposition;
rapid cooling produces amorphous coatings
[21–30]
Sputter coating 0.5–3.0 mm Uniform coating thickness on flat substrates; dense coating
Line of sight technique;
expensive, time-consuming;
produces amorphous coating
[31–39]
Pulsed laser deposition
0.05–5.00 mm Coating with crystalline and amorphous;
coating with dense and porous
Line of sight technique
[40–42]
Table 7.1 (continued)
Technique Thickness Advantages Disadvantages Refs.
Dynamic mixing method
0.05–1.30 mm High adhesive strength
Line of sight technique;
expensive;
produces amorphous coating
[43]
Dip coating 0.05–0.50 mm Inexpensive; coating applied quickly;
can coat complex substrates
Requires high sintering temperatures;
thermal expansion mismatch
[44–48]
Sol-gel <1 mm Can coat complex shapes; low processing temperatures;
relatively inexpensive as coating is very thin
Some processes require controlled atmosphere processing;
expensive raw materials
[49–52]
Electrophoretic deposition
0.1–2.0 mm Uniform coating thickness; rapid deposition rates;
can coat complex substrates
Difficult to produce crack-free coatings; requires high sintering temperatures
[53–60]
Biomimetic coating
<30 mm Low processing temperatures; can form bone-like apatite; can coat complex shapes;
can incorporate bone growth stimulating factors
Time-consuming;
requires
replenishment and constant pH of simulated body fluid
[61–63]
Hot isostatic pressing
0.2–2.0 mm Produces dense coatings
Cannot coat complex substrates; high temperature required; thermal expansion mismatch; elastic property differences;
expensive;
removal/
interaction of encapsulation material
[64]