Читать книгу Fundamentals of Implant Dentistry - John III Beumer - Страница 12
ОглавлениеOsseointegration, Its Maintenance, and Recent Advances in Implant Surface Bioreactivity
Ichiro Nishimura | Takahiro Ogawa | Basil Al-Amleh | Momen Atieh Andrew Tawse-Smith | Benjamin M. Wu
After the concept of osseointegration was introduced, a high rate of treatment success was achieved in quality bone sites with sufficient volume. The original titanium implants were available in machined surfaces or titanium plasma spray surfaces. Eventually, titanium implants with microrough surface topography were introduced that accelerated the events1 associated with osseointegration and led to stiffer bone anchoring the implants.2 This chapter discusses the biologic sequence of host tissue reactions during the process of implant osseointegration and the pathologic factors that potentially can disturb the maintenance of dental implant systems after they have been placed into function. In addition, recent advances aimed at improving the bioreactivity of implant surfaces are discussed.
Protein Adsorption (Seconds to Minutes)
Upon contact with blood, the implant surface is immediately covered by the noncellular components within the blood.3 These primarily include ions, proteins, salts, lipids, glucose, and numerous metabolic byproducts at various stages of their life cycles. All of these components, especially proteins, interact within the first second and immediately act to modify the physical-chemical-biologic properties of the dental implant surface. Among proteins in blood, albumin is the most abundant, followed by fibrinogen and gamma globulin. The initial nanometer-thick layer of proteins present chemical moieties from amino acids with charged, polar, and nonpolar functional groups. These functional groups interact with the implant surface via weak secondary bonds (hydrogen bonding, van der Waals, and electrostatic interactions), and those that bind strongly will stay longer on the material’s surface.4 The early-binding proteins that bind weakly will desorb away from the surface, displaced by stronger binders. The binding force depends greatly on the implant surface chemistry, the protein composition, and the local environment, including pH, ionic concentration, and cellular activities. Over time, the strongly bound proteins can undergo unfolding, which denatures the protein and exposes additional amino acid functional groups that further stabilize the protein-implant interaction. This time- and surface- dependent microevolution of protein composition based on kinetics and stability of protein adsorption is known as the Vroman effect and is relevant for all blood-contacting biomaterials.5
Once bound to the implant surface, the adsorbed proteins interact with the local biologic molecules via receptor-ligand interactions. These include dissolved proteins (more albumin, fibrinogen, etc) and extracellular matrix proteins such as collagen, von Willebrand factors, fibronectin, coagulation factors, complement proteins, and cell fragments such as platelets that indirectly and directly promote the initial matrix-to-cell adhesion.
Regardless of the surface treatments that have been attempted by dental implant manufacturers, protein adsorption occurs on all materials regardless of hydrophilicity levels. Regardless of surface topology and surface chemistry, some of the early-binding proteins contain binding sites that either directly or indirectly for platelet adhesion receptors and trigger the next stage: hemostasis.
Hemostasis: Platelet Plug and Fibrinogenesis (Minutes)
Cells and biomolecules in blood
The protein-modified surface dictates the kinetics and thermodynamics that platelet-surface adhesion will occur. In turn, the platelet-modified surface will influence platelet-platelet adhesion, platelet activation, fibrinogenesis, and formation of the provisional fibrin matrix. Platelets carry surface receptors suitable for attachment to exposed or damaged collagen fibers while secreting internally stored bioactive factors. In blood, platelets initially rely on their high shear stress receptors to gain initial adhesion, followed by the engagement of low shear stress receptors.6 Because blood flow is slow in dental osteotomy sites, both low and high shear stress receptors on platelet surfaces can contribute to binding. The platelet- derived factors include a series of enzymes that are essential for the cascade of the coagulation process resulting in fibrin and clot formation. These activated platelets also regulate the subsequent inflammatory response and wound healing processes. The fibrin clot not only works as a temporary “plug” to prevent further bleeding until the fibrin is formed via the intrinsic and extrinsic clotting pathways—the resultant fibrin plug, with trapped platelets inside, serves as a bioactive scaffold for epithelial and mesenchymal cell migration to commence wound repair.
Besides the injured collagen fibers and tissues, biomaterials placed in the body can activate platelets at different rates. Platelets are considered to be the first cell-like structures to adhere to the implant, and they immediately start secreting bioactive factors and organizing the fibrin clot. It takes only 2 minutes to initiate the fibrin clot formation on titanium surfaces.7 Platelet adhesion and activation on different biomaterials and material surfaces have become subject to intense investigation because the resulting fibrin clot scaffold is thought to determine inflammation behavior and subsequent wound healing around the biomaterial.
Hong et al8 reported that there was much less platelet activation on the surface of stainless steel plates than on titanium plates. When used as an endosseous implant, stainless steel is surrounded by a sustained inflammatory reaction, resulting in minimal, if any, direct bone contact.9 Therefore, the ability to activate platelets and form the fibrin clot may be an important first step in osseointegration.
Effect of implant surface modifications on fibrin clot formation
Recent research and development efforts have been directed toward creating more bioactive titanium surfaces suitable for increased platelet adhesion. Moderately rough surface topography has been shown to increase platelet activation prepared by various methods: double acid etching10 (Fig 2-1), fluoride ion–modified grit blasting,11 sandblasting, and acid etching.12
Fig 2-1 Scanning electron micrographs (SEMs) of platelet-rich plasma contact (for 30 minutes) with commercially pure titanium: (a) double acid etched; (b) 320-grit abraded; (c) machined; (d) polished. The platelet aggregation and fibrin clot formation were more significant on roughened titanium surfaces. (Reprinted from Park et al10 with permission.)
Interestingly, in the field of vascular stent development, research efforts have been directed toward decreasing the adhesion of platelets and thus minimizing thrombosis formation. In fact, the micrometer to nanometer surface topography created on the titanium vascular stent13 or polymer materials14 was shown to decrease the platelet adhesion. The stark contrast in the observations regarding endosseous implants and vascular stents that both carry moderately rough titanium surface topography may suggest that not only the surface roughness but also other factors might determine the initial host response.
Complex surface topography is generally associated with increased hydrophobicity, which prevents the adhesion of platelets and cells. Acid etching used to create microtopography increases the surface precipitation of titanium dioxide, or titania (TiO2),10 whereas alkali treatment results in the formation of charged TiO2 on the titanium surface.15 These surface modifications involving TiO2 have been postulated to control platelet adhesion and activation. TiO2 is a stable and relatively bioinert material that is largely responsible for the biocompatibility of titanium implants. However, the therapeutic role of TiO2 has not been well characterized. The zeta potential or electron charge of the surface of TiO2 is influenced by pH levels and the presence of various ions such as Ca2+. Both acidic (low pH) and alkali (high pH) treatments are known to change the zeta potential of TiO2, contributing to the modulated cell and protein adhesion behavior. Recent studies suggest that the proprietary SLActive preparation (Straumann) or postfabrication ultraviolet (UV) light treatments could increase surface hydrophilicity or surface charge of titanium implants. Characterization of their effect on the platelet behavior and fibrin clot formation has just begun,16 which may present an important clue to understanding the role of surface reactivity and zeta potential on osseointegration.
It must be noted that hydroxyapatite (HA) surfaces show somewhat different platelet adhesion and activation properties as compared with titanium surfaces. The HA surface disproportionately increases complement activation in the fibrin clot11 and increases adsorption of serum proteins.17 Therefore, new surface modifications employing a hybrid of TiO2 and HA18–22 may present a unique opportunity to expand the available armamentaria for better optimization of platelet activation and fibrin clot formation relevant to osseointegration.
Platelet activation occurs at the tissue injury site and on the surface of biomaterials. However, the tissue injury site activates fibrin clot formation much more efficiently than do titanium materials.12 Experimentally, the periodontal ligament on the freshly extracted tooth induced significantly more active clot formation than other artificial materials tested.23 Therefore, there may be a gradient of fibrin clot network around the implant that is more organized and matured on the osteotomy-wounded bone surface than on the implant surface7,23 (Fig 2-2).
Fig 2-2 (a) Diagram of fibrin clot organization around an implant immediately after placement in the osteotomy site. Platelet activation is significantly more efficient on the exposed collagen from the injured tissue than on the titanium surface. As a result, a gradient of fibrin clot (arrow) is organized from the implant surface to the bone surface. (b) A cleaned extracted human tooth with remaining periodontal ligament was dipped in a fresh extraction socket for 60 seconds, and the surface was examined by SEM. A dense fibrin clot was already formed and organized (magnification: left, ×880; right, ×4,400). (Reprinted from Steinberg and Willey23 with permission.) (c) A similar experiment was performed with a titanium plate. A titanium plate was dipped in a fresh extraction socket for 60 seconds. The fibrin clot formed a different architecture. (Reprinted from Steinberg et al7 with permission.)
Fibrin Remodeling (Days to Weeks) and Bone Formation (Weeks) to Bone Remodeling (Years)
Fibrin scaffold network and macrophage infiltration
The wound-induced fibrin clot formation results in the organization of a fibrin scaffold network necessary for the succeeding tissue repair. Although the structure of fibrin networks is determined by multiple factors such as pH, clotting rate, and coagulation factor concentrations, polymerization of fibrin molecules generally occurs within the first 24 hours of wounding. The organized fibrin network is further modified by the incorporation of fibronectin molecules, which serve as the critical factor influencing bone formation in the fibrin scaffold. A recent study suggested the presence of macrophages within the fibrin clot adjacent to a dental implant within 12 to 24 hours.24 The early and transient expression of C-X-C chemokine receptor type 4 (CXCR4; a cell surface receptor of monocytes/macrophages) in this study supports the involvement of macrophages in the process of osseointegration as well as the process of clearing the tissue debris (Fig 2-3).25
Fig 2-3 A diagram of bone formation around an implant. (a) Immediately after the fibrin clot scaffold is formed, bone marrow–derived myeloid cells called myeloid-derived suppressor cells (MDSCs) migrate into the mature fibrin clot and organize the local environment for wound repair. MDSCs stimulate new vascular formation and suppress wound-induced inflammation. (b) After 24 hours of implantation, the fibrin clot scaffold is already organized on the implant surface. Immunohistologic evaluation revealed the infiltration of CD163+ macrophages (or MDSCs) stained in brown in the fibrin scaffold. (Reprinted from Omar et al25 with permission.)
Macrophages are classically described as pro-inflammatory phagocytic cells (M1 macrophages) that clear tissue debris and eliminate bacterial infection. It has been demonstrated that there are alternative differentiation pathways generating M2 macrophages that are capable of resolving inflammation and actively inducing angiogenesis for tissue repair.24 It must be noted that the study by Omar et al25 further suggested that macrophages infiltrating the fibrin scaffold around the implant were recognized by the CD163 cell surface marker. A subset of macrophages carrying CD163 are thought to express the M2 phenotype and are considered myeloid- derived suppressor cells (MDSCs). MDSCs originate in bone marrow and resolve inflammatory reactions by suppressing T-cell activities. In addition, MDSCs induce angiogenesis and secrete a set of growth factors that support rapid wound healing.26 Therefore, the presence of macrophages and MDSCs may be critical for establishing a tissue repair environment for wound healing and bone formation.
Distance and contact osteogenesis
As seen in wound healing following tooth extraction, initial bone formation occurs in the bottom of the socket, suggesting the establishment of a tissue repair environment in the mature fibrin network (Fig 2-4). Fibronectin is a large glycoprotein with active binding sites not only to fibrin but also to other extracellular matrix (ECM) molecules and integrin-expressing cells. Incorporation of fibronectin in the fibrin network has been shown to be important for supporting macrophage function. The earliest bone formation should occur in the matured fibrin network adjacent to the osteotomy-exposed alveolar bone. An experimental implant model in mice demonstrated the early sequence of bone formation within the well-organized fibrin network that was more apparent on the bone surface.27 This study further demonstrated the highly localized fibronectin molecules associated with the bone surface fibrin network. Bone tissue formation away from the implant is called distance osteogenesis,28 which involves an ordinary sequence of bone wound healing as often seen in the tooth extraction socket or in the bone marrow ablation site.
Fig 2-4 After rat molar extraction, the fibrin clot is organized at the bottom of the extraction socket (left). The bone remodeling first occurs within the fibrin clot scaffold (right). The cervical region where the initial fibrin clot formed is less organized.
During this period, the implant surface is still associated with a less organized fibrin scaffold network. However, the implant surface fibrin network is rapidly remodeled with the incorporation of fibronectin and provides the scaffold for bone formation. Distance osteogenesis may now approach in close proximity to the implant surface, while the new bone formation can occur within the now-matured fibrin network surrounding the implant. Contact osteogenesis describes this bone formation near the implant surface, which may be significantly affected by the different environment influenced by the implant material.28 The gap between regenerating bone and the implant surface may be completely filled as early as 7 days after surgery, establishing the histologic osseointegration.
Fibrin clot formation and remodeling take place rapidly at the tissue injury site, where distance osteogenesis should be initiated immediately. Slow fibrin network maturation on the implant surface may cause delayed bone formation. In other words, contact osteogenesis around the implant occurs in a sequence, and the bone-to-implant contact (BIC) is established during the last stage of bone remodeling27 (Fig 2-5). There is a small but distinct time lag between distance osteogenesis and contact osteogenesis. However, active implant surface modifications may significantly accelerate contact osteogenesis.
Fig 2-5 (a) The first bone formation occurs within the fibrin scaffold associated with the bone tissue exposed by osteotomy. Along with the delayed organization of the fibrin scaffold on the implant surface, bone formation catches up and eventually establishes BIC. (b) An experimental implant (IMP: titanium-coated [arrowheads] plastic implant) was placed in an osteotomy site of a mouse femur. Fibrin clots were organized 1 day after implant placement (top). The fibrin scaffold associated with the bone osteotomy site and cortical bone (*) appeared to be more organized than that on the implant surface. Fibronectin (green) was found in the organized fibrin scaffold close to the bone osteotomy site (middle). Two days after implant placement, the initial bone formation was detected within the organized fibrin clot containing fibronectin, while the fibrin network (*) on the implant surface appeared to be still immature. (Reprinted from Jimbo et al27 with permission.)
Characteristics of Peri-implant Bone
Peri-implant bone, which is formed in close proximity to the implant surface, plays a central role in the sustained support of the implant. Peri-implant bone is formed within the fibrin scaffold surrounding the implant and is likely to be influenced by the implant surface topography, chemistry, and charged energy. These factors may affect the unique characteristics of the bone deposited onto the surface of the implant, which could directly or indirectly contribute to the maintenance of osseointegration. This section discusses the biomechanical characteristics, the shear strength at the bone-implant interface, and the long-term stability of peri-implant bone.
Biomechanical characteristics of peri-implant bone
Ideally, the intrinsic biomechanical properties of peri- implant bone should be capable of withstanding functional forces. It has been shown that hardness and stiffness of peri-implant bone may be associated with certain implant surface modifications. Butz et al29 employed nanoindentation assays to measure the hardness and Young modulus of peri- implant bone associated with a relatively smooth machined or double acid-etched titanium implant in a rat model. The hardness of peri-implant bone associated with a relatively smooth (machined) implant was progressively increased from 2 weeks to 4 weeks after the surgical implant placement and reached the equivalent hardness of trabecular bone. The bone hardness associated with a moderately rough (double acid-etched) implant similarly underwent the progressive increase; ultimately, however, it was found to be much harder and reached the equivalent hardness of cortical bone. Recently, a similar experiment in a rabbit model revealed that the hardness of peri-implant bone almost doubled when a moderately rough (sandblasted/acid-etched) implant surface was further modified with a nano-HA coating.30
Once osseointegration is established, the intrinsic biomechanical properties of peri-implant bone should greatly contribute to the load-bearing function. It is intriguing that peri-implant bone may reach the hardness of cortical bone around implants with moderately rough and more complex surfaces. The primary mechanism determining the bone hardness and stiffness has been debated. A positive correlation between the stiffness and bone mineral density was demonstrated in bovine cortical bone31 and porcine mandibular condyles.32
Bone is a composite tissue of collagen-based fibers and crystalline HA. The bone mineral content is regulated by the organic collagen matrix, which is largely composed of type I collagen. Fragile bone is the primary phenotype of a group of genetic disorders called osteogenesis imperfecta. Patients with these disorders experience bone fractures even during normal physical activity. A number of mutations have been discovered in type I collagen genes; however, the most severe form of osteogenesis imperfecta is associated with the genetic mutations in enzymes that control collagen cross-linking, such as prolyl-3-hydroxylase (P3H)33 and cartilage-associated protein (CRTAP).34 In addition, prolyl-4-hydroxylase (P4H) is also involved in collagen cross-linking, and collectively these enzymes are critical in determining the intrinsic bone mechanical properties. In vitro biomimetic mineralization on collagen films using a polymer-induced liquid-precursor mineralization process further supports the notion that increased collagen cross-linking significantly stimulates mineralization and increased intrinsic mechanical properties.35
With the use of genetic characterization methods, the increased expression of P4H and CRTAP has been reported in the peri-implant tissue during the early stages of osseo- integration.1,36 While type I collagen gene expression is not significantly affected by the presence of implant materials, the increased presence of collagen cross-linking enzymes associated with the implant is thought to contribute to the formation of stronger peri-implant bone29,37 (Fig 2-6).
Fig 2-6 (a) Responding to a titanium implant, peri-implant bone synthesized through contact osteogenesis acquires a unique biomechanical property. (b) Hardness and stiffness of bone formed around an implant with a machined or a double acid-etched surface were measured by a nanoindentation assay. Peri-implant bone of the roughened implant was much harder and stiffer than trabecular bone, and its biomechanical properties nearly resembled that of cortical bone. Peri-implant bone deposited on the smooth, machined implant was not as hard; however, it had increased stiffness. *P < .05; **P < .01; ***P < .001. (Reprinted from Butz et al29 with permission.)
Bone-to-implant contact and interfacial shear strength
Direct bone attachment to the implant surface is the hallmark of osseointegration. Therefore, histologic assessment of osseo- integration commonly uses the percent area of BIC. Higher failure rates in the posterior maxilla have been attributed to its relatively poor trabecular structure leading to decreased BIC. Traditionally, nondecalcified histologic ground specimens have been used to determine BIC. Significant intrasample variations in BIC have been found,38 and a small but critical discrepancy has also been reported between histologic specimens and 3D images reconstructed through microcomputed tomography (microCT).39 Therefore, the data analysis of BIC may require careful interpretation.
Recently, an increasing number of studies report that BIC does not correlate with mechanical failure. When the implant push-in test and microCT-based 3D BIC were used in a rat model, the moderately rough implant (due to double acid etching) showed three times higher shear strength than the relatively smooth machined implant.40 Because the 3D BIC was not different between these tested implants, the increased interfacial shear strength was due to the increased bone bonding to the implant surface. The mechanical interlocking mechanism for roughened implants may contribute to the increased withstanding load. However, this study indicated that epoxy resin–embedded implants showed only a small increase in the withstanding load, suggesting that biologic bone bonding may play the central role. The discrepancy between the BIC measurement and the mechanical withstanding load assay suggests that while bone formation around the implant must be a prerequisite, the development of osseointegration may rely on the actual bonding between the bone and the implant surface.
For many years, the existence of a thin layer of tissue between the bone and the implant surface has been reported in electron microscopy observations. This tissue layer is generally described as comprising an electron-dense zone 20 to 50 nm thick9,41 and a 100- to 200-nm–thick zone without typical collagen fibers,42 followed by the collagen-rich bone tissue. However, considerable structural variations of this interface tissue have been pointed out, possibly due in part to sample preparation artifacts. Davies43 proposed that the electron-dense layer might be comprised of “globular accretions” that are highly mineralized. Cross sections of globular accretions may result in the reported variation in thickness of the interface tissue layer or so-called “cement line”44 (Fig 2-7). A study using a titanium-coated polystyrene cell culture plate revealed a globular accretion–like electron-dense structure abutting the titanium layer.44 The globular accretion–like interface layer was found to contain crystalline calcium phosphates similar to HA and the previously unreported thin collagen fibers. The precise molecular composition of the interface tissue has not been elucidated. However, it is postulated that molecules comprising the interface tissue between bone and the implant surface should hold the key to the mechanical withstanding force of osseointegrated implants.
Fig 2-7 (a) Diagram of the implant-bone interface. There is a thin layer of interface zone between the peri-implant bone and the implant surface, which is thought to be composed of globular accretions. The cross section of a cluster of globular accretions may be equivalent to the zone of tissue of the cement line. It has been proposed that the molecular composition of this interface structure plays a key role in the function of osseointegration. (b) A recent in vitro study revealed that the osteogenic cells precipitated more mineralized tissue on the titanium-coated polystyrene cell culture plate (bottom) than on the control polystyrene surface (top). (c) Transmission electron microscopy suggested an electron-dense zone of globular accretions (white arrowheads) on the titanium coating (arrows). The globular accretion–like structures were interposed between the titanium coating and poorly mineralized bone (*). (d) A high magnification of the square in part c demonstrated the mineral content (arrowheads) as well as thin fibrous structures. (e) A close-up of the square in part d. The mineral content showed a crystalline structure consistent with HA. (Reprinted from Saruwatari et al44 with permission.)
It has been reported that this interface zone contains proteoglycans (PGs),45 although the amount of PGs has been debated.46,47 PGs are associated with glycosaminoglycan (GAG) side chains, which provide a sticky consistency, and therefore it has been postulated that PG-GAG in the interface zone may play a role in the bonding between bone and implant. The adhesion of in vitro mineralized tissue to a titanium disk was moderately attenuated by the treatment of GAG degrading enzymes such as chondroitinase AC, chondroitinase B, and keratinase.48 Although this study suggested a functional role of PG-GAG for bone adhesion to the implant surface, the impact of chemical degradation of PG-GAG was surprisingly small. Therefore, the shear strength of osseointegrated implants to withstand occlusal load appears to involve more complex mechanisms.
The interface tissue (also known as the cement line) contains osteopontin (OPN).49 OPN is a noncollagenous ECM molecule in bone. It has an integrin-binding sequence, suggesting cell adhesion functions. In addition, because OPN has been found in high levels in mineralized tissue of bone and teeth, its postulated functions include regulation of bone remodeling. However, genetically modified mice lacking OPN were surprisingly normal, and their skeletal tissues developed without any complications.50 The cement line of OPN-deficient mice was also found to exhibit the normal structure. Recently, a reevaluation of OPN-deficient mouse bone revealed that there was a 30% decrease in bone fracture toughness, while the bone mass remained unaffected.51 The nanoindentation assay showed that the stiffness, not the hardness, was significantly decreased. Although this conclusion is highly speculative, the high OPN content in the cement line may contribute to the increase in stiffness of the mineralized interface tissue between the bone and the implant surface, which could contribute to an increase in mechanical withstanding shear strength.
The large shear strength is due to the bone insertion sites of the ligament and tendon. Characterization of this interface zone of ligament insertion to bone repeatedly found the presence of types II, IX, and X collagen52,53 that are commonly found in cartilage tissue. In particular, type X collagen is expressed by hypertrophic chondrocytes during endochondral ossification. In the growing bone, type X collagen is co- localized with PGs and appears on the longitudinal septa of hypertrophic cartilage when the bone starts to bear the body weight.54 Type X collagen forms a network of hexagonal mesh and, when embedded in a mineralized tissue, enforces its intrinsic mechanical property. Therefore, type X collagen in the developing bone and the bone insertion sites of the ligament and tendon is thought to generate the significant shear strength to resist gravity and physical activities.
Studies involving DNA microarray reported a puzzling observation: The gene expression profile of peri-implant tissues contained not only bone-related genes but also other genes that were notably of the cartilage molecules.55–58 Those cartilage-related molecules include PGs; types II, IX, X, and XI collagen; and hyaluronan and PG link protein.59 In other words, the presence of an implant during the healing following osteotomy surgery may create a mixture of bone- and cartilage-related molecules in peri-implant bone. Recently, type X collagen was identified in the interface tissue between bone and implant.55 It may be postulated that cartilage-related molecules such as PGs and type X collagen may be involved in the interface layer between implant and bone, potentially contributing to the shear strength of implant bonding to bone55 (Fig 2-8).
Fig 2-8 (a) The entire genome microarray gene expression of peri-implant tissue. A hierarchical cluster analysis revealed that there were five major gene groups, of which Cluster 2 exhibited the genes most sensitively associated with implant osseointegration. (b) Cluster 2 included cartilage-related ECM genes (arrowheads). (c) Among cartilage- related genes, type X collagen (green, arrowheads) was identified within the interface zone between the bone and the implant surface. Bone marrow mesenchymal cells (blue). (Parts a to c reprinted from Mengatto et al55 with permission.) (d) Hypothetical structure and molecular components of the bone- implant interface tissue. The cement line is composed of crystalline calcium phosphate particles (gray sunbursts) in globular accretions containing OPN (blue bars) and type X collagen (green hexagonal mesh). These molecules may increase the stiffness and shear strength of the cement line. There is a less mineralized and relatively amorphous zone resembling cartilage tissue containing thin and sparsely arranged type II collagen fibers. The cartilage-like zone may also contain PG-GAG molecules, possibly contributing to the shock-absorbing function.
Long-term stability of peri-implant bone
The osteotomy procedure used to prepare an implant placement site creates an ablation wound in the bone marrow. Intramembranous ossification occurs during the healing of bone marrow ablation60 and tooth extraction wounds,61 thus leading to the formation of woven bone trabeculae in the marrow space. The trabecular bone formed in response to ablation wounding is then subjected to intensive remodeling and largely resorbed to create fatty bone marrow (Fig 2-9). Uniquely, bone tissue formed in the vicinity of implant surfaces appears to resist this catabolic bone remodeling and thus maintains the osseointegration for an extended period.2 Trabecular bone derived from distance osteogenesis may be relatively unstable and can disappear due to physiologic bone remodeling. On the contrary, peri-implant bone derived from contact osteogenesis appears to undergo slower bone marrow remodeling and remains around the implant for the long term (see Fig 2-9).
Fig 2-9 (a) A diagram of bone marrow ablation healing around an implant. The newly formed bone around the implant is subjected to osteoclastic bone resorption, regenerating the bone marrow space. It has been noted that peri-implant bone resists bone resorption activity. (b) MicroCT- reconstructed 3D picture depicting the persistent presence of peri-implant bone, with the surrounding bone marrow having lost its trabecular structure, in an experimental animal model using rats.
The rapid formation of bone marrow trabecular bone, perhaps with the woven bone characteristics, may occur 1 to 2 weeks after implant placement and may potentially contribute to the immediate implant stability. Whether the early woven bone can support the occlusal load has not been established. While the majority of woven bone may be resorbed, the remaining bone structures continue to mature. During the transition stage from woven bone resorption to the maturation of the small but well-organized trabecular bone, there may be a vulnerable period in which the degree of implant integration may temporarily decrease. This phenomenon has been observed in an animal model (Nishimura et al, unpublished data); however, its clinical significance has not been established.
Bone resorption is facilitated by osteoclasts. Osteoclasts are formed by fusion of monocytes under a combination of chemical cues including receptor activator of nuclear factor κB (RANK) ligand, or RANKL. During the developmental stage, RANKL is secreted from osteoblasts and hypertrophic chondrocytes. However, when bone is matured, RANKL is primarily secreted from osteocytes embedded in bone, which sensitively respond to mechanical stimuli such as occlusal loading.62 As discussed previously, the mechanical property of peri-implant bone may be harder than that of surrounding trabecular bone. It is conceivable that the increased mechanical properties of peri-implant bone may reduce the amount of matrix deformation that the embedded osteocytes experience, leading to reduced RANKL secretion under the normal occlusal force. Above a critical load threshold for peri- implant bone osteocytes, however, implant overloading can stimulate the osteocytes to initiate the secretion of RANKL, resulting in osteoclast formation and bone resorption. The role of mechanical loading on bone remodeling is discussed in further detail in chapter 3.
Osteoclasts strongly adhere to bone surface and form a ringlike apparatus, referred to as the sealing zone. Osteoclasts create an acidic milieu within the sealing zone and secrete proteinases such as cathepsin K to degenerate the organic matrix of bone. As a result, bone mineral HA and collagen matrix are removed. The osteoclast adhesion to the bone surface is required for this bone resorption process. It has been reported that the adhesion of osteoclasts is influenced by the bone surface topography. When mouse osteoclasts were cultured on titanium disks with different surface roughness ranging from 1 to 4.5 µm Ra, the sealing zone formation was shown to be disturbed by microtopographic obstacles.63 There was an inverse correlation between the stability of the osteoclast ring (ie, the structural integrity and sealing zone translocation rate of osteoclasts) and the increasing microtopography.
Because the adhesion of osteoclasts appears to be less effective on a rough surface, it may be postulated that the surface topography of peri-implant bone may be rougher than that of surrounding trabecular bone. The placement of an implant appears to influence biochemical compositions of peri-implant bone. Cartilage and bone comprise the major skeletal system, and both contain ECM such as collagen. There are distinct differences in the composition of ECM molecules; ie, types I and V collagen are predominant in bone, whereas types II, IX, X, and XI collagen are in cartilage. However, recent studies indicate that peri-implant bone may be composed of a mixture of bone and cartilage ECM. In a mouse model lacking type IX collagen, one of the cartilage ECM molecules was shown to develop an age-related osteoporosis-like phenotype.64 Type IX collagen maintains the space between the adjacent collagen fibers and has been shown to exist in a small amount in bone. The lack of type IX collagen appeared to manifest as a dense bone collagen network, resulting in the smoother bone surface. Osteoclasts were found to adhere widely to this mutant bone surface. Although highly speculative, the reduced susceptibility of peri-implant bone to osteoclastic bone resorption may in part be facilitated by its different biochemical compositions, such as increased type IX collagen, and bone surface topography.
Recent Developments Associated with the Microrough Surfaces
Attempts to improve the wettability of the implant surface
In recent years, attempts have been made to improve the wettability of the implant surfaces. The surface wettability of dental implants appears to have a significant impact on the biologic cascade of events that occur at the bone-implant interface and is modulated by the chemistry and topography of the implant surface.65 It has been shown that the wettability of the implant surfaces plays an important role in the adsorption of plasma proteins66 and the differentiation and cell adhesion of mesenchymal stem cells into bone-producing cells.67
Several methods have been used in an attempt to increase the wettability of the implant surface, including the application of fluoride ions and magnesium ions to the implant surface. This is referred to as electrowetting and allows the plasma proteins to flow freely onto the implant surface and into the irregularities of the microroughened surface immediately upon insertion of the implant. Moreover, fluoride ions on the implant surface, when applied to a titanium grit-blasted surface, increase the expression of the genes associated with the differentiation of osteoprogenitor cells68–70 and promote osseointegration during the early stages of healing.71
Another approach is to package the freshly prepared titanium implants in saline. Packaging the implants in this manner reduces the rate of contamination of the surface of the implant, which maintains the surface energy and wettability. Recent studies have also shown that photofunctionalization will increase the wettability of the implant surface.
Genetically engineered implant surfaces
Since the discovery of osseointegration-specific genes, it has been an inviting idea to imbed one or more of these genes onto the surface of the implant. There are several advantages to this approach. The genes do not degrade in these environments and can be applied to the implant surfaces in suitably low doses. Moreover, these genes are associated with the normal cell cascade of cell differentiation and function. However, there are significant disadvantages. The primary disadvantages are the cost of development and the regulatory issues that must be addressed before bringing such a product to the marketplace. Unless there is a significant clinical advantage to be gained using gene-enhanced implant systems, the cost of ensuring safety and efficacy outweighs the benefits.
Nanoenhanced, biomimetic implant surfaces
The surface of bone is highly mineralized with HA crystals and exhibits a highly complex surface topography. During the process of bone remodeling and following acid etching by osteoclasts, bone possesses distinct surface chemistry and topography. Since the phenomenon of osseointegration was introduced in the late 1970s, researchers have attempted to create so-called “biomimetic” implant surfaces; in essence, implant surfaces with surface chemistry and topography that mimic that of acid-etched bone surfaces found in vivo. Furthermore, submicron-to-nanometer titanium surface features enhance cytocompatibility properties for bone-forming cells, increasing both surface energy and cell adhesion.72
The two methods developed during the last several years—crystalline deposition of HA crystals on microrough surfaces of titanium implants18 and the application of pico-to-nanometer-thin TiO2 coating on microroughened titanium surfaces73—are the most recent attempts to mimic the physical and chemical environments found in vivo (see later in the chapter). These nanoenhanced surfaces optimize platelet activation, facilitate the adsorption of plasma proteins, and accelerate differentiation of mesenchymal stem cells into functioning osteoblasts. Furthermore, gene expression of the differentiating osteoprogenitor cells is accelerated and upregulated.18,22,74 The result is that the events associated with osseointegration are substantially accelerated.
Implant surfaces enhanced with recombinant peptides
Application of recombinant osteogenic proteins (OP-1), bone morphogenetic proteins (BMP2, BMP7), and growth factors such as platelet-derived growth factor (PDGF) to the surfaces of implants has been of interest because of the potential of enhancing osteoconductive properties of the implant surface. However, the optimal means of bonding these proteins to implant surfaces have not been determined. In addition, retention and controlled release of these proteins has been difficult. Additional disadvantages are the fact that many of these proteins are costly, they frequently are not associated with the normal cellular cascade, and they may be deactivated during sterilization of the implant. In addition, higher concentrations/doses of BMP2 have triggered troublesome side effects.75
It is possible to bind these proteins to implant surfaces. Implants with HA and chitosan coatings have been used most often. The outcomes of most of these studies indicate that binding BMPs to the implant surface significantly enhances osseointegration.76,77 Other researchers75,78 have attempted to determine whether coatings of osteogenic proteins can be used effectively for vertical augmentation of deficient ridges. In an animal study, WikesjÖ et al75 showed that coating porous titanium implant surfaces with rhBMP 2 induced significant bone formation around the neck of the implants, leading to a clinically significant vertical augmentation of the alveolar ridge. In another study using an animal model, Susin et al78 achieved a similar result with rhBMP 7.
Crystalline deposition of HA crystals
Techniques have been reported whereby nano-sized crystals of HA of a specific size (20 Nm) are deposited onto the surface of titanium implants that have previously been double acid-etched (DAE).18 A highly specific distribution and spacing of these crystals can be achieved (Fig 2-10).79 These crystals are joined to the TiO2 surfaces with covalent bonds. Based on microCT, the amount of BIC associated with nanoenhanced DEA surfaces is equivalent to the BIC of bone deposited onto untreated DEA surfaces. However, the shear bonding strengths needed to separate the implant from the bone anchoring the implant is increased dramatically. The shear bonding levels were more than doubled as compared to DEA surfaces and increased by more than 7 times compared to machined surfaces. Although the effect of crystal deposition was apparent when used on machined-surface implants, the effects were considerably more significant when used in combination with DEA microroughened surfaces, indicating a potential synergistic effect between the two surface phenomena18,79 (Fig 2-11).
Fig 2-10 (a) A DAE titanium implant surface. (b) A DAE titanium implant surface following depositions of nano-sized HA crystals. (Reprinted from Moy et al79 with permission.)
Fig 2-11 The shear strength of the bone-implant interface is dramatically increased following deposition of HA crystals on the surface of a DAE titanium implant. (Reprinted from Moy et al79 with permission.)
As mentioned previously, the hardness of peri-implant bone deposited on microrough implant surfaces is increased progressively from 2 weeks to 4 weeks after surgical implant placement and eventually approaches the hardness equivalent of cortical bone. Of interest was the fact that the hardness of peri-implant bone almost doubled when the moderately rough (sandblasted/acid-etched) implant surface was further modified with nano-HA coating.30 The reasons for this finding have yet to be clearly elucidated.
Controlled nanostructuring of titanium surfaces experimentally used to optimize nano-size
A recent study described a new method to create a uniform nanonodular surface topography on titanium implants. This method utilizes a newly discovered phenomenon of titanium nanonodular self-assembly during physical vapor or sputter deposition of titanium onto specially conditioned micro-rough titanium surfaces.80 The size of nanonodules can be controlled by altering the deposition time. The newly added nanostructures must be smaller than the existing microrough configurations of the implant surface (Fig 2-12).79 Using this self-assembly method, implants with acid-etched surfaces were converted to micro/nano-hybrid surface topographies ranging from 100 to 1,000 nm in diameter (Fig 2-13).79 The nanostructure creation effectively creates geometric undercuts and increases the surface area by up to 40% compared with the acid-etched surface with microrough surface topography.
Fig 2-12 A method has been developed for fabricating micro/nano- hybrid titanium surface topography. This increases surface area, creates mechanical undercuts, and maintains the existing microrough surface topography. (Reprinted from Moy et al79 with permission.)
Fig 2-13 (a to c) Three different sizes of nanonodules of titanium created on the acid-etched microrough titanium surfaces using a nanonodular self-assembly method. (Reprinted from Moy et al79 with permission.)
The surface topography created with this method closely resembles the surface morphology of biomineralized bone matrices.81 This nanoenhanced implant surface selectively promoted function of cultured osteoblasts but not fibroblasts. Implants with microrough surface topography promote and accelerate differentiation of osteoblasts but inhibit their proliferation. Implants prepared with microrough surfaces enhanced with nano-sized nodules substantially enhanced both of these cell activities. These biologic effects were most pronounced when the nano- nodules were 300 nm in diameter. An implant biomechanical test in a rat femur model revealed that the strength of bone-titanium integration was more than three times greater for the implants with the microrough surface enhanced with nanonodules compared to the implants with unenhanced microrough (acid-etched) surfaces. These results suggest that the establishment of uniquely functionalized nano-in- micro titanium surfaces improved osteoconductivity and may provide a biomimetic micro-to-nano-scale hierarchical model to optimize the nanofeatures of dental implant surfaces and other biomaterials.
Biologic Aging and Photofunctionalization of Implants
Many implants are packaged in plastic and are then sterilized with gamma radiation. This process contaminates the implant surface with hydrocarbons and other carbon-containing impurities. Bioreactivity of the implant surface is impaired, the surface charge is changed from positive to negative, and the surface becomes less wettable. As a result, adsorption of plasma proteins, platelet activation, and recruitment and attachment of osteogenic cells are inhibited.
Biologic aging of implant surfaces
A series of recent studies reported significant changes in the osteoconductivity and other biologic capabilities of titanium implant surfaces over time. These studies have indicated that the bioreactivity of titanium implant surfaces degrade as a function of time, and this phenomenon has been referred to the biologic aging of titanium.82–86 This time-dependent degradation can be substantial, as the strength of osseointegration measured by a biomechanical implant push-in test model can be reduced by 50% for aged titanium surfaces compared to newly prepared titanium surfaces. Moreover, a BIC area higher than 90% can be obtained for new, uncontaminated titanium surfaces compared to a BIC area less than 60% for aged surfaces. The degradation of the implant surface bioreactivity appears to be primarily associated with the reduced capability of aged titanium surfaces to adsorb plasma proteins, activate platelets, and attract osteogenic cells.82–86
Surface property changes associated with the biologic aging of titanium
An analysis of surface chemistry using x-ray photoelectron microscopy has demonstrated that the percentage of carbon molecules on titanium surfaces increases with time.82,87 The percentage of carbon, which in one study was found to be 14% on the acid-etched titanium when first prepared, increases to 63% after 4 weeks of storage under ambient conditions.82 More than half of the titanium surface does not appear as titanium at the molecular level. The increase of surface carbon is due to the deposition of carbon-containing impurities from the local environment onto titanium surfaces, consisting primarily of hydrocarbons. Significantly, the capability of titanium surfaces to attract proteins and osteogenic cells has been shown to have a strong inverse correlation with the percentage of surface carbon. This data implies that the presence of surface carbon plays a crucial role in determining bioreactivity of titanium implant surfaces.82
Implants with hydrophilic surfaces are more bioreactive. However, the contamination of the implant surface described previously degrades the hydrophilicity of titanium implant surfaces and impairs its ability to adsorb plasma proteins. Recent studies have shown that titanium surfaces, immediately after preparation, regardless of the types of processing used, display a water contact angle ranging from 0 to 5 degrees. Such surfaces are considered superhydrophilic. However, the superhydrophilic nature of freshly prepared implants gradually attenuates, and after 1 week the surface becomes hydrophobic and the contact angle increases to over 40 degrees. The contact angle for 4-week-old acid-etched implant surfaces increases to 60 degrees and has been found to be as high as 90 degrees82,83,87–90 (Fig 2-14).
Fig 2-14 Time-dependent changes in hydrophilic nature of titanium surfaces. Titanium surfaces are superhydrophilic when prepared but become hydrophobic after 1 week, and the degree of hydrophobicity increases with time. (Reprinted from Moy et al79 with permission).
Photofunctionalization
In response to the degradation of the bioreactivity of titanium implant surfaces as a function of time, recent research efforts have focused on developing a means of removing these contaminants and restoring implant surfaces to their previous level of bioreactivity prior to surgical placement. Studies to date indicate that carbon contamination of titanium implant surfaces occurs regardless of the method used to create the microsurface topography. The method described in this section has been shown to be effective in decontaminating all of the titanium implants tested.
Photofunctionalization is defined as a phenomenon whereby the physicochemical properties and bioreactivity of titanium implant surfaces are restored after UV light treatment.84,90–95 Titanium surfaces that have aged (ie, older than 1 month after preparing the surface) become hydrophobic. The contact angle of a droplet of water applied to the surface is generally above 60 degrees and often closer to or above 90 degrees on most surface types. This loss of wettability is common for all surface topographies of titanium.82,87,88,90,91,93,95,96 When water droplets are applied to the surfaces of these implants, the droplet retains its hemispheric form. After treating these titanium implant surfaces with UV light, the surfaces regain their superhydrophilicity, and the contact angle is reduced to almost zero (see Fig 2-15).
Fig 2-15 (a to c) Aged untreated implant. (d to f) Implant after photofunctionalization. Photofunctionalization improves the hemophilicity of the implant surface (part a vs part d). Photofunctionalization enhances cell adsorption of plasma proteins, enhances attraction of mesenchymal stem cells, and promotes and accelerates differentiation of mesenchymal stem cells into osteoblasts and enhances cell spreading. Note the contrast between part b and part e. The result is that the BIC area is increased to almost 100% in a rat model. Note the difference between part c (untreated surface) and part f (treated surface). (Reprinted from Moy et al79 with permission.)
As mentioned previously, soon after the implants are prepared, hydrocarbons begin to accumulate on their surfaces, most likely from the atmosphere and the surrounding environments during surface preparation and storage.17,37,82,87,95–103 UV treatment reduces the carbon percentage to less than 20% depending upon the conditions.95,96,104 Titanium surfaces covered by hydrocarbons are known to be negatively charged, which makes the implant surface repellent to plasma proteins and osteoprogenitor cells, which are negatively charged.83,84,86,94 After photofunctionalization, titanium surfaces are converted to electropositive and serve to attract plasma proteins and osteoprogenitor cells.84,86,94,105
The biologic effects of photofunctionalization are shown in Fig 2-15. Following photofunctionalization of aged implants, the strength required to separate the investing bone from the implant surfaces—as measured by a biomechanical implant push-in test model—is increased by a factor of three.95 Moreover, BIC area has been shown to increase to above 90% and in some specimens to almost 100%. In contrast, the BIC area for untreated implants is generally in the range of 50% to 60%.95,96 Photofunctionalization has been shown to be effective for all surface topographies tested (anodized, dual acid etched, sandblasted/acid etched, machined surface).88,93,96,104,106,107
In summary, in vitro studies conducted to date have demonstrated that photofunctionalizing titanium implants (1) increases adsorption of plasma proteins on the implant surface, (2) facilitates attachment of and retention of osteogenic cells to the implant surface, (3) facilitates spread of osteoblasts on the implant surface, (4) increases cell proliferation, and (5) promotes and accelerates differentiation of mesenchymal stem cells once they migrate to and bind to the surface of the implant.88,93,96,104,106,107
Clinical implications
To date, most efforts to improve the osteoconductivity of osseointegrated implants have been focused on changing the micro- or nanosurface morphology and chemistry. However, the studies cited earlier have shown that the osteoconductivity of modern implant surfaces can be dramatically enhanced by ridding the implant surface of carbon contaminants. If the degree of BIC established by photofunctionalized implants can be sustained once the implants have been placed under function, these studies imply that several difficult clinical challenges can be addressed. Shorter implants may be employed than previously have been shown to be feasible with untreated surfaces, and implants placed in poor-quality bone sites may be more predictable regardless of whether initial stabilization is accomplished.
A chairside device has been developed for commercial use that emits UV light of multiple wavelengths. The implants are photofunctionalized for 15 minutes before use in patients. Several authors have reported their initial experiences.108–112 Initial reports are promising, but long-term studies are needed to clarify the degree of advantage offered by chairside photofunctionalization.
Other applications
Photofunctionalization appears to have application beyond dental implants. Photofunctionalized titanium mesh, used to house and contain bone grafts, has been shown to be more osteoconductive and able to facilitate more bone regeneration in animal models113 and humans.110 Photofunctionalized orthodontic miniscrews have been shown to develop improved bone anchorage and better resist lateral forces.114 Photofunctionalization is also being employed with increasing frequency in orthopedic surgery.
Degradation of Titanium Dental Implant Surfaces
Dental implants are most commonly made from commercially pure titanium (cpTi) grade II or grade IV. Grade IV has higher strength and lower corrosion resistance than grade II.115 The abutment and prosthetic components are made of titanium alloy (Ti6Al4V) due to its high tensile strength; however, the corrosion resistance of Ti6Al4V is lower than cpTi.116
After surgical implant placement and prosthetic restoration, the dental implant is susceptible to biochemical degradation. Recent studies on the degradation of dental implants and prosthetic components in the presence of microorganisms and the corrosive environment of the oral cavity have gained more attention in the recent literature.117–119 Chemical (dry) and electrochemical (wet) corrosion causes different forms of degradation, which can occur in the oral cavity. Several studies118–125 have shown particles derived from dental implants in peri-implant tissues. It has been suggested that these particles are released from dental implants due to nontherapeutic and therapeutic causes. The released titanium micro- and nanoparticles are cytotoxic126 and act as foreign bodies to the immune system. As a consequence, the released particles activate osteoclasts and local inflammatory cells and increase the expression of pro-inflammatory cytokines such as tumor necrosis factor α (TNF-α), which trigger an inflammatory cascade in the peri-implant tissues.119,125
Nontherapeutic causes
Nontherapeutic causes of implant surface degradation include the presence of biofilm and inflammatory responses (ie, peri-implant diseases), wear from micromovement of contacting surfaces at the implant-abutment connection,127 or the detachment of particles from titanium implant surfaces during insertion.120 The latter is related to the corrosive effects of therapeutic agents such as citric acid, bleaching solutions, and fluoride,118,119 or surgical treatment of peri-implant diseases when implantoplasty is carried out to smoothen the implant surface.123,128
The formation of an oral biofilm can decrease the pH of the oral environment and degrade the titanium surface due to the presence of corrosive metabolites such as the production of lactic acid by Streptococcus mutans in the presence of sucrose.129 Occlusal loading can cause micromovements of the contacting surfaces at the implant-abutment connections. The combined wear and corrosion from the two contacting surfaces in the presence of a biofilm is termed tribocorrosion and can lead to degradation of the titanium implant surface into the surrounding peri-implant tissues.127,130 The mechanical wear facilitates corrosion by damaging the passive protective layer of TiO2 and exposing the implant surface to corrosion. The corroded implant surface becomes more vulnerable to further mechanical wear. Tribocorrosion is also intensified by the use of rough-surfaced dental implants, which are more susceptible to biofilm accumulation compared with smooth-surfaced implants.131 Numerous studies on surface deterioration of orthopedic implants suggest that aseptic loosening of the implant is related to the production of wear debris from the prostheses.132,133
Therapeutic causes
The most important steps in the treatment of peri-implant diseases are debridement and decontamination of the implant surface. Both debridement and decontamination can be achieved by chemical and mechanical means or a combination of both.134 Such means remove the biofilm and inflammatory tissue, but at the same time they can affect the implant surface and connection.
Chemical substances commonly used for implant surface decontamination have been shown to increase the implant surface roughness and release degradation particles, which may stimulate inflammatory responses in the peri-implant tissues.118,135
One profound mechanical technique to debride the implant surface is implantoplasty, or mechanical removal of implant threads in an attempt to decrease the surface roughness and make the implant surface less retentive of biofilm. The procedure is carried out using diamond burs under copious irrigation with sterile saline and stones to polish the implant surface.136 As a result, however, particles and debris from implant surfaces are released into the surrounding tissues despite abundant irrigation to remove debris during and after implantoplasty. In addition, implantoplasty modifies the implant design, causing more stress at the damaged area of the implant surface and surrounding bone.
Titanium brushes have been introduced for mechanical debridement and decontamination of the implant surface.137,138 Although the presence of titanium-based particles following the use of titanium brushes has not been evaluated, it has been suggested that the use of titanium brushes is less damaging to the implant surface compared with diamond burs used for implantoplasty. Similarly, the use of ultrasonic scalers to remove the biofilm from the implant surface can cause surface deterioration. This is observed both with metallic and plastic ultrasonic tips, which also generate titanium particulate that can be identified in rinsing solutions.139
In a histologic investigation of 36 peri-implantitis biopsies in humans treated with implantoplasty, foreign bodies—mostly titanium particles—were formed, and these were surrounded by chronic inflammatory infiltrates.140 Other histologic studies123,139 have confirmed the presence of micro- and nanoparticles in subepithelial connective tissues following implantoplasty. The histologic analysis showed areas of chronic inflammation with lymphocytes and plasma cells, which may suggest negative effects of titanium particles in the surrounding peri-implant tissues. Titanium particles may affect the expression of RANKL and osteoprotegerin (OPG) in osteoblastic cells and activate osteoclasts and pathologic bone remodeling.126 A synergistic effect of the presence of released titanium ions with Porphyromonas gingivalis can increase the expression of cytokine ligand 2 (CCL2), RANKL, and OPG in peri-implant tissues.126 It has been suggested that titanium ions at concentrations of 13 ppm may cause epithelial cell necrosis, but the biologic mechanism is not fully understood.141,142 Pettersson et al143 suggested that titanium particles rather than titanium ions stimulate the inflammatory response, with recent data suggesting that titanium nanoparticles have genotoxic and cytotoxic potential.144–146
The cytotoxicity has been further investigated in orthopedics and showed that particle release from prosthetic joints can be detrimental to cells and body tissues, resulting in aseptic loosening of the implant.122 Nevertheless, the content of particles needed to cause cytotoxicity is still not clear. In addition, systemic accumulation of titanium micro- and nanoparticles has also been reported in the lymph nodes, lungs, liver, and spleen.125,147
Summary
Significant advances have been made during the last 20 years to increase the bioactivity of implant surfaces. The initial efforts (titanium plasma spray surfaces, HA-coated surfaces, the original acid-etched surfaces) to improve the clinical performance of titanium dental implants had significant disadvantages. In the early 1990s, several methods (double acid etching, anodic oxidation, titanium grit blasting, sandblasting/acid etching) were developed to create microrough surface topographies on titanium implants. These surfaces were shown to dramatically improve clinical performance, particularly in poor-quality bone sites. More recently, methods have been developed to improve clinical performance by increasing or maintaining the wettability of the implant surfaces and facilitating osteoconductivity by creating biomimetic nanoenhanced implant surfaces.
Titanium materials have long been considered to be bioinert. Therefore, it has been believed that the presence of a titanium implant in an osteotomy site should not influence the wound healing process. While the mechanistic elucidation is not complete, it is increasingly clear that osseointegration is not achieved only via bone formation. Recent observations and experimental evaluations indicate that there are distinct molecular and cellular behaviors that appear to be unique to peri-implant tissue. Some of these characteristics contribute to the mechanical advantage and long-term stability of osseointegrated implants. In addition, peri-implant bone may not undergo the same biologic and pathologic sequences as tooth-bearing alveolar bone.
After manufacturing and before surgical placement, implant surfaces age, during which hydrocarbons accumulate, hydrophilicity is lost, and the electropositive surface changes to electronegative and impairs the osteoconductivity of implant surfaces. Photofunctionalization is one method to recondition the implant surface by removing hydrocarbons from the implant surface, restoring hydrophilicity and improving electrostatic charge.
After surgical placement, the degradation of implant surfaces in the presence of peri-implant diseases and the release of titanium particles during and after management of peri-implantitis are still of concern. Further studies are required to evaluate the local and systemic effects of those particles and propose guidelines for prevention, management, and maintenance of peri-implant diseases.
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