Successful treatment using endosseous dental implants is dependent on the formation and maintenance of secure implant-to host bone fixation. This criterion for success has been recognized since the late 1970s as a result of early reports by Branemark and co-workers on successful use of threaded, screw-shaped cpTi implants placed using a careful two-stage surgical procedure.1 A result of those early studies was the coining, by Branemark, of the term ‘osseointegration’ which described virtual direct implant-bone contact as a necessary requirement for success of dental implants.
To ensure the maintenance of this condition during implant use, this criterion has been modified to ‘functional implant osseointegration’ implying maintenance of ‘close’ bone-implant apposition during implant loading. In keeping with the basic terminology, one can refer to ‘osseointegration potential’ of different implant designs as a means of comparing their rate of bone integration. It is believed that the ‘osseointegration potential’ may be strongly influenced by implant surface design since attachment, migration and differentiation of cells of the osteogenic lineage at the implant surface are the determinants of this response. Thus, the focus of this article is on implant surfaces, the methods used for their preparation, and, to the extent that it is known, the effectiveness of different surface designs in promoting and maintaining osseointegration. Reference is made to currently available dental implant systems and experimental surface modifications intended for increasing osseointegration potential of dental implants.
The soft connective tissue-implant interface region
While osseointegration is necessary for implant success, equally important for long-term implant maintenance is the establishment of a stable soft connective tissue-implant interface at the coronal implant region. In this regard, it is generally accepted that a smooth machined and polished surface with average surface roughness (Ra) equal to 0.1 to 0.3 m is preferred.2,3 The length of this ‘smooth’ coronal region varies with different implant designs (up to 3 mm or so in length), depending on the features at the intended implant site. Since it is known that bone will not readily juxtapose to such smooth surfaces, a longer smooth coronal region sacrifices a length of implant to which osseointegration could occur. An irregular surface while favouring osseointegration, if in contact with the oral cavity is more susceptible to micro-organism and plaque attachment.
Recently, short (6 to 12 mm), wide-diameter (5 mm) threaded implants for use in regions of limited bone volume and density (i.e. posterior jaw sites) were introduced without any ‘smooth’ coronal regions. The implant made with deeper threads along the entire implant length was designed to have a bone-interfacing surface area sufficient for implant stabilization by virtue of the deeper thread design, larger implant diameter, and threads along the entire implant length. At one-year, survival rates of 91.8% were reported and a significant number of implants showed thread exposure.4 Renouard et al expressed concern over the long-term prognosis of the ‘collarless’ implants in view of the significant bone loss observed around the majority of the implants. It appears that the requirement for a ‘smooth’ coronal region, at least sufficient to accommodate early crestal bone loss due to either biomechanical, biological,5 or bacterial6 effects is necessary.
The bone-biomaterial interface
A review of articles in current dental implant journals and marketing literature describing new designs presents a major challenge for dentists in determining whether one design or another offers clinical advantages. While there is a large volume of literature suggesting that most currently available implants work reliably and with acceptable success rates (> 90% for periods of 5 years and more) for placement in sites of good quality bone and with adequate bone volume (width and height), the situation is not as clear for more difficult to treat sites characterized by limited bone height (<8 mm or so), and low bone density (type 3 and 4 bone).7 In these situations (e.g. posterior mandible, maxilla), rapid and strong fixation to bone using diminutive implants is preferred. A number of surface modifications have been proposed to achieve this. Some of these are being used with commercially available implants while laboratory testing of others is ongoing.
A brief review of the surface designs currently in use as well as some of those under investigation is presented below. Osseointegration of all currently available implants depends primarily on mechanical (or micromechanical) interlock of implant and bone. This applies equally to so-called ‘bioactive’ implants (HA-coated) as well as Ti-based implants (cpTi or Ti alloy) with ‘passive’ surface oxide layers. Provided that the bone-interfacing implant surface has features that allow significant mechanical interlock of bone, secure long-term fixation (>10 years) with success rates greater than 90%, in general, result. Thus, implants are formed with surface geometries and textures that allow such mechanical interlock. Machined threaded, grit-blasted, acid-etched, plasma-sprayed, and porous surfaces prepared by sintering represent the currently used implant designs that achieve this condition.
Threaded Implants (Machined, Grit-blasted, Acid-etched)
Machined, screw-shaped implants are examples of designs that rely primarily on macroscopic surface features (i.e. threads) for fixation (Figs. 1a & b). As shown in Figure 1b, in addition to their macroscopic features, machining lines approximately a micron or so in width and other surface irregularities (pits, gouges, zones of cold-welded Ti) form during implant machining.
The formation of these micron-sized features is related to the deformation characteristics of the metallic implant. Ti is notoriously difficult to machine to a smooth surface finish as a result of its deformation and surface characteristics (i.e. it tends to gall). Smoother machined surfaces result for grade 4 compared with grade 1 cpTi (grade 4 has higher interstitial levels and hence higher yield strength) and Ti6Al4V compared with cpTi (again due to the alloy’s higher yield strength).
While the argument has been made that these finer features may affect cell activity thereby promoting osseointegration, possibly by enhancing osteoconductivity and/or contributing to mechanical interlocking, no clinical follow-up data has been presented to support this through a comparison of success rates of cpTi versus Ti alloy or grade 1 versus grade 4 cpTi implants, for example. The reason for success of threaded implants in favorable implantation sites is due to bone forming in close apposition to the threaded surfaces during the critical post-implantation healing period during which the implants are not in function. Implant fixation and resistance to shear forces (vertical and torque) following osseointegration is due to friction at the bone-implant interface (as with any screw system). Hence, fixation strength is directly related to implant-bone interface area. Therefore, machined threaded implants of a fixed diameter must be sufficiently long to resist the forces acting during function.
In sites of limited bone height, therefore, where longer implants (i.e. > 10 mm) can not be used, threaded implants of standard diameter (~4 mm or less) do not provide the required fixation. Similarly, in regions of low bone density or limited cortical bone thickness in which sufficient purchase of the machined threaded device is not possible, standard threaded implants must rely on length and bi-cortical fixation, if possible, for stability.
Recent developments to improve the osseointegration potential of machined threaded implants to allow possible use in more demanding situations has involved texturing the implant surface in order to i) increase potential bone-interfacing area, and ii) provide additional features to effect micromechanical interlock with bone. To achieve this, machined
threaded implants have been either grit-blasted, acid-etched, or grit-blasted (sand-blasted) and acid etched. The surface of an acid-etched Ti alloy implant is shown in Figure 2. The acid etching (in this case an acid treatment using HCl and H2SO4 solutions at elevated temperatures), results in the formation of small micron-sized dimples over the entire surface. Acid etch treatments result in some metal dissolution, the extent of metal dissolution being dependent on the strength of the acid used, temperature and time of reaction. Any foreign material inadvertently adhered to the surface might be dissolved as well thereby resulting in a cleaner implant surface. Sharp asperities, ridges or other features that might have resulted from machining and that represent higher energy regions (less stable thermodynamically) would dissolve preferentially.
Grit blasting results in an irregular surface with the surface features being related to the size and hardness of the blasting medium and blasting conditions used (pressure, incidence angle of blast, distance from surface to blasting jet). Al2O3 and TiO2 particles have been used for blasting. Studies have shown that the rate of bone integration with grit-blasted Ti samples (as determined by histology (bone contact length) and torque testing) appeared greatest for surfaces modified using 75 m-sized grit (vs 25 or 250 m Al2O3 grit).8
The concept of an optimal surface roughness and geometry for greatest osseointegration potential was also suggested to explain findings reported by Buser et al.9 In that study, a sand-blast plus acid-etch surface treatment was reported to promote highest bone adaptation and rate of interface shear strength development. Ti plasma-sprayed and machined surfaces were found to be inferior. The Ti plasma-sprayed surface had greater surface roughness and yet resulted in slower rates of bone integration in the reported animal studies.
It was proposed that this observation could be due to a non-optimal surface roughness being presented by the rougher plasma-sprayed surface. The issue is by no means well understood. At present, there is no data based on human clinical follow-up studies to suggest significant differences between grit-blasted and acid-etched, or acid-etched alone over Ti plasma-sprayed implants.
Animal studies reported on the effect of surface roughening (grit blasting vs acid etching or a combination of the two processes) clearly showed that surface roughness had a significant positive influence on osseointegration potential in the animal model studies. A certain optimal degree of surface roughness appeared to result in faster rates of osseointegration. As a result, a number of implant systems with such surface modifications have been introduced and are being used currently. Results of implant survival rates, or more significantly, cumulative success rates, over prolonged periods (>5 years) are not yet available to test the hypothesis that one surface preparation yields clinically superior results. The reason for these surface modifications is to promote higher osseointegration potential of implants in order to allow their more reliable use in lower density bone (Type 3 and 4) and in sites with limited bone volume (i.e. where implant length is limited to < 10 mm, for example).
Additionally, the concept of more rapid osseointegration is preferred since this reduces the risk, in all situations, of premature implant loading (either planned or inadvertent) leading to implant movement relative to host bone thereby compromising bone integration. In addition, faster bone adaptation to and integration with implant surfaces in the cortical implant region, reduces risks of deep recesses developing in these zones that would increase the risk of implant failure due to bacterial colonization and peri-implantitis. This is the reason why some recent designs have incorporated narrow grit-blasted or acid-etched bands just inferior to the smooth coronal collar regions. Again, confirmation of the clinical effectiveness of this approach will require long-term clinical follow-up studies comparing one design with another.
Perhaps of significance in this regard is a study by Drake et al,10 which described the beneficial effect of acid treatment in increasing Ti or Ti alloy surface hydrophilicity thereby decreasing the rate of bacterial colonization of surfaces and allowing connective tissue cells to gain access to and develop matrix at the coronal implant region. While this may suggest an advantage of the chemically-etched surfaces, it should be noted that most implants are given an acid passivation treatment as a final stage of implant preparation, be they machined, grit-blasted, plasma-sprayed or sinter-treated. While Drake’s study was limited to one bacterial strain (S sanguis), it nevertheless suggests the importance of surface chemistry in addition to topography in determining implant performance.
Surface Modification with Additive Processing:
Implant surface preparation by machining, acid etching or grit blasting , and combinations of these operations, involve the removal of material from a metallic substrate. They can be described as subtractive processes. In contrast additive processes, represented by plasma spraying and sintering, add material to a substrate to develop desired surface structures. A discussion of the preparation and properties of implant surfaces modified using such additive processes follows.
Plasma-sprayed Implants (screw-shaped and press-fit designs)
Plasma-sprayed implants (Figures 3a & b) are formed by introducing powders with particles of 100 to 300 micron size or so into a hot plasma flame. The particles are fully or partially melted in the peripheral region of the hot plasma flame (the central region of the flame reaches temperatures of 15,000 to 20,000C) and are then transported at high velocity in an ion stream (usually Ar ions) and deposited as molten splats onto a relatively cool metallic implant substrate surface (either cpTi or Ti6Al4V).
The molten particles, on impacting the substrate surface, spread over it as thin ‘splatted’ deposits and rapidly solidify and in doing so ‘freeze’ onto the pre-roughened substrate surface. Thus, during solidification which occurs at very rapid cooling rates (~106C/s), the molten material mechanically interlocks with the pre-roughened substrate, and to some extent depending on the material being deposited (Ti or HA), diffuses into and reacts with the substrate. Repeated deposition of particles onto the substrate and over previously-deposited layers through rastering of the plasma flame over the substrate results in a build-up of coating to a desired thickness, (typically 30 to 50 microns).
The final coating is characterized by a very irregular outer surface, due to the rapid solidification process, with recesses and outcroppings of deposited material. Some porosity invariably results within the coating but this is limited to a maximum of 10 volume percent or so and usually, for higher quality coatings, is well below 5 percent. The small volume of pores that form are either connected with the surface recesses or are isolated within the thickness of the coating. Very irregular surfaces result during plasma spraying with dimensions of the surface features being up to 10 to 30 microns or so in cross-section and depth. These are typically much coarser than the micron-sized etch pits and grit-blasted surface features (dimensions of up to 10 microns or so) formed on implant surfaces by the subtractive processing methods described previously.
Such plasma spray-coated implants are currently made using either Ti or HA (nominal) powders. The former results in chemically bonded coating-to-substrate structures (because of metallic bonding at the coating-substrate interface) while the latter is dependent almost entirely on mechanical interlocking between the HA layer and Ti substrate for bonding since the two materials are mutually insoluble and atomic interdifussion is limited.11
A major benefit of HA plasma-sprayed coatings is the increased osteoconductivity that has been observed resulting in faste
r rates of bone formation at an implant surface. The exact reason for this enhanced osteoconductivity is not fully understood nor is it clear that clinically significant differences derive from the use of such coatings. Proposed reasons relate to i) preferential adsorption of proteins on implant surfaces that promote pre-osteoblast cell attachment (a surface composition effect), ii) in vivo dissolution of the coating releasing Ca2+ and (PO4)3- ions that promote bone formation, and, iii) the highly irregular coating surface promoting osteoblast and matrix attachment (a surface topography effect). This latter mechanism would apply equally to HA or Ti plasma-sprayed coatings.
While HA is quite stable in vivo thereby nullifying the dissolution hypothesis, the coatings formed by plasma spray deposition of HA powder is highly heterogeneous with significant portions of the coating consisting of less stable, calcium phosphate phases such as tricalcium phosphate (TCP), tetracalcium phosphate (TTCP), and amorphous calcium phosphate (ACP) as well as CaO, another soluble phase. This is a result of the high temperature exposure of the HA powders introduced into the plasma flame, changes in Ca and P composition as a result, and the formation of an amorphous (glassy) phase during the rapid solidification process. These less stable phases are more soluble in vivo than HA and, hence, dissolution of the coatings occurs, the rate of coating degradation being dependent on the degree of heterogeneity and compositional and crystallographic variation. While this may make the coating more osteoconductive, it also raises concerns over possible dissolution debris released at the implant site.
Excessive rates of degradation and coating delamination have been reported to cause undesirable chronic inflammatory reactions at implant sites thereby potentially inhibiting osseointegration and bone bonding.12 Post-plasma spray pressurized hydrothermal treatment has been used to increase the percent HA of plasma-sprayed calcium phosphate coatings.13 Animal studies showed that the resulting 95% HA coating (+5% amorphous calcium phosphate) behaved similarly to coatings of lower % HA in terms of osseointegration and bone bonding ability.14 This finding suggests that coating dissolution and Ca2+ release does not appear to be the controlling mechanism for promoting faster osseointegration with calcium phosphate coatings.
Plasma-sprayed coatings (Ti or calcium phosphate) have been shown to result in higher interface shear strengths in most animal study reports. It is noteworthy, however, that while high shear strengths have been reported, interface tensile strengths are low (see Table 1). The mechanical interlock of bone with the surface irregularities (recesses and protrusions) of plasma-sprayed coating surfaces results in resistance to shear forces acting at the interface (either vertical or torsional). However, there is little resistance to tensile forces acting across the interface (i.e. bone does not form a 3-dimensional interlocked structure at the implant surface region). This is similar to the surfaces prepared using ‘subtractive’ surface modifications but in contrast to the porous sintered structures described below.
Sintered porous-surfaced Implants and 3-D interlocking of bone
Unlike the plasma spray coating process, sintering of Ti alloy powders to form a porous-surfaced structure is achieved by a solid state diffusion process in which metal particles (powders) achieve metallurgical bonding to form an integrally-bonded surface region. There is no localized melting and re-solidification of the metal powders during this process. Through a judicious choice of sintering parameters (temperature, time, atmosphere), structures can be formed with an interconnected porous network of desired size and volume percent porosity with pores uniformly distributed throughout the structure.
The Endopore implant (Figs. 4 & 5) is characterized by such a porous surface region consisting of approximately 35 volume percent porosity and an average pore size of approximately 100 m (range ~50 to 150 m). Under the sintering conditions used to form the surface region, the interparticle and particle-substrate junctions (sinter neck regions) are substantial (sinter neck diameter ~0.4 x particle diameter). This results in a strong sintered structure (Figs. 4a & b). Due to the atomic diffusion that results in sinter neck formation, the final implant structure is an integral combination of a solid core and a 300 m deep porous surface zone. The starting powder size and the sintering conditions used result in interconnected pores that prior studies have shown to allow rapid bone ingrowth.15 Two to three particle layers form the porous surface region thereby creating the desired 3-D porous network on implants of overall cross-sectional dimensions of 3.5, 4.1, and 5.0 mm maximum diameter (Fig. 5).
Like all implant designs, initial implant stability and lack of movement relative to host bone is necessary to allow rapid osseointegration through bone ingrowth into the porous network. This condition is more readily achieved using a tapered truncated implant that is press-fitted into a prepared recipient site. The major difference between this surface design and that of the other implants described above is that the ingrowth of bone results in both very high interface shear strength and high interface tensile strength. This is shown in Table 1 which lists average values for bone-implant interface shear and tensile strengths for different implant surface designs derived from reports in the literature using various animal models. Thus, porous-surfaced dental implants are better able to resist interfacial tensile forces due, for example, to horizontal force components acting on an implant (Fig. 6). When dental implants are placed in function, occlusal loading of implants results in compressive, shear, and tensile force components acting at the bone-implant interface. Resistance to interface tensile force components results in a more uniform distribution of stresses acting in bone surrounding the implant. This is in contrast to the stress distribution developing around any of the other designs noted above that are unable to support tensile interface forces (Fig. 6).
The expected long-term benefit of the more uniform peri-implant stress field with porous-surfaced implants is more effective bone retention and maintenance of osseointegration. These considerations also suggest that porous-surfaced implants should perform better in regions of low density bone. The early results (over four years) of Endopore implants placed in posterior mandibular sites and in the maxilla (where Type 3 and 4 bone are common) have given support to this premise. Additionally, the extremely effective bone-implant interlock resulting from 3-dimensional bone ingrowth allows much shorter implants to be used reliably in highly loaded sites. Deporter et al have reported high success rates (>99%) after 2 years of function (results unchanged with minimum function periods now approaching 3 years) for short implants (mean length = 7.7 mm) placed in posterior mandibular sites.16 Similar high success rates were also reported for porous-surfaced implants placed in the posterior maxilla17 (mean length = 6.9 mm, minimum period in function approaching 1 year).
Previously, high success rates (CSR = 93.4%) were reported for short Endopore implants (mean length = 8.7 mm) placed with overdentures in anterior mandible sites of fully edentulous patients. This study represents the longest clinical follow-up for this design with implants over 10 years in function and results unchanged to those reported at 5 to 6 years.18 Deporter et al19 have reported very high success rates (approaching 100%) with Endopore implants used for restoring single teeth in the maxilla. The free-standing implants used had a mean length of 10.1 mm and have now been in function for periods approaching 3 years.
The high success rates experienced with this design for free-standing, single tooth replacement is not surprising in view of the 3-dimensional bone ingrowth
which provides excellent resistance to imposed forces including torsional and tipping forces. Of particular note is the successful implant performance despite crown: implant length ratios far greater than the 1:2 ratios that conventional wisdom dictates as a guide for reliable implant use. This again is attributable to the excellent implant-to-bone fixation achieved through bone ingrowth into the open-pored structure. The porous-surfaced design takes full advantage of mechanical interlocking. The design, therefore, allows reliable use of shorter lengths; (currently 5, 7, 9, and 12 mm lengths are available).
The press-fit implant with its tapered truncated conical shape offers other advantages. The shape ensures self-seating into prepared sites. The tapered truncated conical shape is ideal for use with osteotomes for implant site preparation in regions with limited bone volume such as is often encountered in the maxilla.20
Sintered porous surfaces have been shown to be osteoconductive.21 Using a rabbit model, we have shown that the sintered porous surface described above results in significantly faster osseointegration compared with plasma-sprayed implants.22 Finite element studies indicated that this increased osseointegration potential of sintered porous-surfaced implants appears related to the local stress state in the interface zone that favours osteogenesis and early bone formation.23
The local stress state in host bone surrounding any implant is modulated due to the presence of the implant. The consequences of the resulting perturbed stress state is modification in normal bone remodelling. As reported by Garetto et al,24 bone remodelling around threaded cpTi implants is three to nine times faster than in similar sites without implants. This is considered necessary for long-term bone maintenance and osseointegration with dental implants. Similar effects have been observed with sintered porous-surfaced implants.
Implant surface design and the establishment of a well-bonded implant-bone interface also affects other bone remodelling events. With porous-surfaced implants, limited bone loss has been observed next to the smooth coronal implant region both in animal studies25 and in human use.18 In these situations, crestal bone loss occurs in a predictable manner over a 2- to 3-year period in humans reaching a steady state once the remodelled bone approaches the junction of the smooth-to-porous surface. The effect has been related to stress shielding and bone disuse atrophy next to the smooth collar region, a consequence of forces by-passing this region of bone since no interlock occurs there.26 Recently, Deporter has suggested a possible contribution to this effect related to the establishment of a ‘biologic width’ following implant placement.
Further studies related to this are ongoing. It is noteworthy that while some crestal bone loss invariably occurs with all implant designs, reasons for so doing vary, being related to overstressing of local bone with screw-shaped designs, for example, and understressing of bone next to the smooth regions of porous-surfaced designs. Provided that the phenomenon achieves a final stable steady state bone-implant structure with limited crestal bone loss, it is acceptable.
SUMMARY
Implant surface design and its influence on implant success
Whether press-fit or threaded implants are preferred in sites characterized by substantial bone width and height of good density is debatable. An obvious advantage of threaded implants placed in such sites is that initial implant stability is more readily ensured (assuming proper site preparation). However, proper placement of press-fit designs also results in sufficient implant stability and similar success rates have been reported.
In compromised implantation sites, where short implant lengths must be used, or in low bone density sites where rapid osseointegration is particularly desirable, surface design can have a more significant effect. The use of acid etching, grit blasting, plasma spraying, and sintering processes have been studied to develop implants that will provide greater reliability in these more challenging situations. While reports of good short-term performance have been reported for implants utilizing these different surface modification approaches, the true determination of success requires confirmation through long-term clinical follow-up. The sintered porous-surfaced designs are unique in offering true 3-dimensional implant-bone interlock and an ability to resist interface tensile forces. Results of clinical use of this design to date are encouraging.
obert Pilliar is currently a Professor in the Faculty of Dentistry (Biomaterials) and a member of the Institute of Biomaterials and Biomedical Engineering (University of Toronto) with a cross-appointment to the Department of Metallurgy and Materials Science (Faculty of Applied Science and Engineering).
Oral Health welcomes this original article. Complete references available upon request.
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