journal of prosthodontic research 59 (2015) 84–95

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Review

A review of improved fixation methods for dental implants. Part II: Biomechanical integrity at bone–implant interface Yo Shibata DDS, PhDa,*, Yasuhiro Tanimoto PhDb, Noriko Maruyama DDSc, Manamu Nagakura DDSb a

Department of Conservative Dentistry, Division of Biomaterials & Engineering, Showa University School of Dentistry, Tokyo, Japan b Department of Dental Biomaterials, Nihon University School of Dentistry at Matsudo, Chiba, Japan c Department of Orthodontics, Showa University School of Dentistry, Tokyo, Japan

article info

abstract

Article history:

Purpose: The purpose of this article is to review the mechanical requirements of the tissue–

Received 30 October 2014

implant interface and analyze related theories.

Received in revised form

Study selection: The osseointegration capacity of titanium implants has been investigated

1 January 2015

over the past 50 years. We considered the ultimate goal of osseointegration to which form a

Accepted 20 January 2015

desirable interfacial layer and a bone matrix with adequate biomechanical properties.

Available online 19 March 2015

Results: Occasionally, the interface comprises porous titanium and bone ingrowth that enables a functionally graded Young’s modulus, thereby allowing reduction of stress shielding. However,

Keywords:

the optimal biomechanical connection at the interface has not yet been fully clarified. There

Titanium implant

have been publications supporting several universal mechanical testing technologies in terms of

Osseointegration

bone–titanium bonding ability, although the separation of newly formed bone quality is unlikely.

Mechanical properties

Conclusions: The understanding of complex mechanical bone behavior and size-dependent

Finite element analysis

properties ranging from a nano- to a macroscopic level are essential in the biomechanical optimization of implants. The requirements of regenerated tissue at the interface include high strength, fracture toughness related to ductility, and time-dependent energy dissipation and/or elastic–plastic stress distribution. Moreover, a strong relationship between strain signals and peri-implant tissue turnover could be expected, so that ideal implant biomechanics may enable longevity via adaptive bone remodeling. # 2015 Japan Prosthodontic Society. Published by Elsevier Ireland. All rights reserved.

Contents 1. 2. 3.

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Basic mechanical and fatigue properties of titanium . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Basic mechanical properties of bone associated with hierarchical structures . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

85 85 86

* Corresponding author at: Department of Conservative Dentistry, Division of Biomaterials & Engineering, Showa University School of Dentistry, 1-5-8 Hatanodai, Shinagawa-ku, Tokyo 142-8555, Japan. Tel.: +81 3 3784 8178; fax: +81 3 3784 8179. E-mail address: [email protected] (Y. Shibata). http://dx.doi.org/10.1016/j.jpor.2015.01.003 1883-1958/# 2015 Japan Prosthodontic Society. Published by Elsevier Ireland. All rights reserved.

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journal of prosthodontic research 59 (2015) 84–95

4.

5. 6.

1.

Nanoscale mechanical testing technology . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1. Bone mechanical characterization at material level. . . . . . . . . . . . . . . . . 4.2. Mechanical properties of mineralized tissue on titanium surfaces . . . . . 4.3. Molecular structure analysis of mineralized tissue on titanium . . . . . . . 4.4. Bone mechanotransdution . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Surface morphology-induced micromechanics and adaptive bone remodeling. Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

Introduction

Orthopedic and dental titanium implants must function as rigid osseous anchors. The biomechanical integrity of implants comprises the mechanical behavior of implant materials, surface-induced bone micromechanics, and adaptive bone remodeling. A diagram of the events presumed to occur at the bone– implant interface is shown in Fig. 1 [1,2]. Once implants are placed with intimate apposition of bone at the surgical site, the immediate response at the interface involves adsorption of tissue fluid and cell binding proteins [1,2]. This critical gap between host bone and the implant surface is likely filled by newly formed bone and a non-collageneous protein-rich cement line [3–5]. Early studies indicated that osteocalcin, osteopontin and bone sialoprotein, as well as certain plasma proteins such as a2HS-glycoprotein, predominate in the cement line [3,4]. Although this protein layer may be responsible for biological bonding to the titanium surface, the inherent mechanical weakness of the protein layer needs to be further stabilized by bone–implant mechanical interlocking [6,7]. Nevertheless, there have been a range of biomechanical issues with titanium implants, such as reduced shear load-bearing properties associated with poor mechanical interlocking at the interface [8,9]. Implant surfaces have been developed by means of several engineering processes, such as grid-blasting, which involves coating the titanium substrate with sintered beads or particles to create a porous layer [10–13], so that the micro– nanoscale external surface textures may address the aforementioned biomechanical issues. Aside from the mechanical interlocking achieved by surface processing parameters, the mechanical requirements of regenerated bone at the interface also need to be addressed. As a consequence of unfavorable osseointegration processes, the implant surface often generates fibrous woven bone and is not replaced by mature lamellar bone [14–17]. Therefore, new biomaterials for bone regenerative purposes, such as titanium implants, need to feature a surface that promotes osteogenic differentiation and proper mineralization during the initial integration stage. The biomechanical integrity of titanium implants has been evaluated based on bone–titanium contact and bone volume fraction by means of histological observation [18,19] and micro-computed tomography [20–22], respectively. However, there has been a mismatch between such observations and the mechanical stability of titanium implants [23,24], allowing us to assume that the mechanical properties of regenerated bone are not entirely associated with observed bone microhistology or densitometry.

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87 87 88 88 89 89 91 92

Understanding of basic mechanical properties for titanium and intact bones are indispensable prerequisite so that scientists will begin to adjust the concept of surface modification (characterization) techniques for titanium implants, whether such concepts achieve the biomechanical integrity at bone–implant interface. Bone is a complex hierarchical tissue with different structural levels, namely cortical and trabecular bone at the macroscale, Haversian osteons and lamellae at the microscale, and hydroxyapatite crystals and collagen fibers at the nanoscale [25,26]. The macro–microscale structural variations of bone tissue compromise a precise mechanical evaluation. In this respect, nanoscale mechanical testing technologies enable a measurement of bone mechanical properties at a material level [27] so that more accurate case simulation, according to threedimensional finite element models, is possible. Moreover, an accurate bone remodeling algorithm in the peri-implant region (adaptive bone remodeling) would be obtainable as bone morphology often relates to the stress-strain response associated with the bone mechanical properties at material level [28,29].

2. Basic mechanical and fatigue properties of titanium The mechanical mismatch between host bone and metallic implants has been a longstanding concern. For instance, the elastic modulus of bone is presumed to be 10–30 GPa, while just around 100 GPa for pure titanium, although titanium and its alloys have elastic moduli less than 50% that of cobaltchrome (approximately 230 GPa) [30]. In this context, contacted bone is often inappropriately stress shielded, and hence, implants lose supportive tissue at the peri-implant region over time [31]. Implant elasticity and the long-term bone integrity associated with adaptive bone remodeling are strongly related, as this has been well-established in a range of animal model experiments and clinical trials [32–36]. The mechanical properties of titanium previously reported are summarized in Table 1. Besides its usage in implants, titanium is mainly used as a hard tissue substitute; hence, increased fracture toughness is the basic requirement. In this respect, Ti–6Al–4V alloys have been widely applied as biomedical materials, despite the fact that the toxic element in this alloy is concerning [37]. Meanwhile, b-type titanium alloys are composed of non-toxic elements with a greater strength and toughness balance than that of a + b alloys, such as Ti–6Al–4V [37]. The elastic moduli of b-type titanium alloys are between 55–85 GPa, resulting in elasticity that is much greater than pure titanium and a + b

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journal of prosthodontic research 59 (2015) 84–95

Fig. 1 – Schematic representation of events at the bone–implant interface. (a) Protein adsorption from blood and tissue fluids, (b) protein desorption, (c) surface changes and material release, (d) inflammatory and connective tissue cells approach the implant, (e) possible targeted release of matrix proteins and selected adsorption of proteins, (f) formation of lamina limitans and adhesion of osteogenic cells, (g) bone deposition on both the exposed bone and implant surfaces, (h) remodeling of newly formed bone [1,2] (license number: 3486160959007). Table 1 – Mechanical properties of titanium alloys for biomedical applications. Alloy 1. Pure Ti grade 1 2. Pure Ti grade 2 3. Pure Ti grade 3 4. Pure Ti grade 4 5. Ti–6Al–4V ELI (mill annealed) 6. Ti–6Al–4V (annealed) 7. Ti–6Al–7Nb 8. Ti–13Nb–13Zr (aged) 9. TMZF (Ti–12Mo–6Zr–2Fe) (annealed) 10. Ti–15Mo (annealed)

Tensile strength (UTS) (MPa)

Yield strength (dy)

Elongation (%)

Modulus (GPa)

Type of alloy

240 345 450 550 860–965 895–930 900–1050 973–1037 1060–1100 874

170 275 450 485 795–875 825–869 880–950 836–908 100–1060 544

24 20 18 15 10–15 6–10 8.1–15 10–16 18–22 21

102.7 102.7 103.4 104.1 101–110 110–114 114 79–84 74–85 78

a a a a a+b a+b a+b b b b

alloys. However, the values are still much higher than that of bone. As homogeneous metallic materials, the mechanical properties of titanium and its alloys, such as tensile characteristics, fracture toughness and fatigue characteristics, have been measured by means of universal mechanical testing machines, in which testing is occasionally carried out in simulated body fluids [37].

3. Basic mechanical properties of bone associated with hierarchical structures Over a long period of human evolution and adaptation, hard tissues such as bone and teeth have developed desirable mechanical properties along with their own hierarchical structures; we can learn useful lessons from such intelligent natural biomaterials. Although the main objective of this review is to address the current biomechanical issues associated with titanium implants, a quick overview of typical natural hard tissue, such as bone, can help us to appreciate some general concepts associated with the excellent design of biomedical materials. Despite the considerable mechanical mismatch between bone and titanium, bone is still functional as a supporting tissue for much harder and stiffer titanium implants, allowing us to conclude that intact bone enables remarkable adaptation associated with strain energy dissipation in the human body. Bone consists of directional structural features across several hierarchical scales, ranging from nanoscale crystals

and molecules to the macroscopic range [38,39]. The foundational unit across the scale is a two-phase arrangement of an anisometric bone mineral (hydroxyl apatite) preferentially oriented in a collagen matrix (Fig. 2) [40,41]. The combination of nanoscale apatite crystals and collagen fibers enables light but tough mechanical properties [42,43]. Based on Griffith’s theory of brittle fracture [44], the smaller the individual sample, the higher the strength is, because smaller samples contain less and smaller defects. In this respect, nanoscale bone apatite crystals enable a higher strength than a bulk sintered apatite block. However, understanding bone mechanical properties is quite complicated because of the size-dependent mechanical properties associated with a heterogeneous, anisotropic material with complex, multiscale structural variations [45]. As a consequence of the multiscale structure, mechanical properties, such as the elastic modulus of bone, have varied dramatically in accordance with differently scaled mechanical testing technologies. The elastic modulus of large tensile cortical bone has been shown to be approximately 14–20 GPa, while that attained from micro-bending tests was 5 GPa [38]. Additionally, the effects of specimen size on the observed moduli suggest that the micro-structural constitution of bone may not be constant, similar to macroscopic observation. As mentioned above, the basic building block of bone, comprising apatite crystals and collagen fibers, is extremely small. The precise mechanical evaluation at a material level by means of universal mechanical testing technology is therefore nearly impossible.

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Fig. 2 – Schematic illustration of hierarchical structure of bone (license number: 3495260148887).

Load, P

(hmax , Pmax) S

P = Pmax Original Deformed

θ

hc

hmax 0 0

hc displacement, h

Fig. 3 – Nanoindentation force displacement curve for measurement of hardness and elastic modulus.

4.

Nanoscale mechanical testing technology

Given the above concerns, nanoscale mechanical testing technologies, such as nanoindentation, may resolve specific features of bone mechanical properties through precise measurement protocols [27]. The instrumented indentation technique has become a popular method to measure the mechanical properties of different materials at the same measurement scale, so that even multiscale structures of bone or teeth can be comparable to artificial biomaterials at the material level. Indeed, nanoindentation mechanical testing has been introduced to evaluate the intrinsic mechanical properties of regenerated bone within a titanium bone chamber, so that the mechanical quality of bone on machined and roughened titanium implants can be predicted [46]. Most nanoindentation experiments use a sharp pyramidal Berkovich indenter from which the hardness and elastic modulus can be determined based on the compliance methods advocated by Oliver and Pharr [47,48]. Briefly, the indenter tip contacts the material surface with an applied loading force followed by unloading, so that a transducer can visualize a load–displacement curve (Fig. 3). Nanoindentation uses the ‘‘area function’’ based on a certain tip geometry [48]. The area function provides a method of gaining an accurate contact area (A) against the indenter tip penetration depth. The onset of the unloading slope reveals the contact stiffness of the

material’s surface. The hardness (H) and reduced (elastic) modulus (Er) can be calculated according to the equations below, where S is the contact stiffness calculated as the slope of the unloading curve at the onset of unloading, P is the applied loading force, and A is the projected area of the indenter tip as a function of the contact depth: H ¼ P=A

(1)

pffiffiffi s p Er ¼ pffiffiffiffi 2 A

(2)

4.1.

Bone mechanical characterization at material level

The nanoindentation compliance method has been developed on the basis of a presumption that the indenter is always contacting a pure elastic–plastic, time-independent material such as ceramic or metallic materials [47,48]. However, biological tissue often generates time-dependent viscoelastic behavior, similar to polymeric artificial materials [49–51]. Although time-dependent behavior is presumed to be a function of bone durability [52–54], the delayed responses compromise the precise measurement of bone mechanical properties. According to a simple viscoelastic Maxwell model represented by a viscous dashpot and elastic spring connected in series [55], deformation of viscoelastic materials may be associated with strain rates. For instance, the viscous dashpot

journal of prosthodontic research 59 (2015) 84–95

Thold Pmax

PU

Load, P

Load, P

88

S Se

PL

0 0

Time, t

0 0

Displacement, h

Fig. 4 – Viscoelastic hysteresis loop forms during unloading.

cannot instantaneously respond to high strain rates, thereby resulting in less deformation and enhanced apparent stiffness. Alternatively, the most readily observed delayed (timedependent) response from a material surface is creep behavior during constant holding indentation, as the viscous dashpot is largely in motion against such lower strain rates [27]. Consequently, a hysteresis loop forms during unloading, upon which loading by creep strain rates generally results in overestimation of contact stiffness (Fig. 4) and the projected contact area, and hence, the observed mechanical properties such as hardness and elastic moduli are varied even in nanoindentation experiments [56,57]. Despite the substantial concerns associated with time-dependent properties, a number of bone nanoindentation experiments have already been carried out. That is to say, most researchers rule out the inaccuracy of complex elastic moduli against different strain rates due to the viscoelastic effects. Separation of elastic (storage) modulus and viscoelasticity (loss modulus) through either quasi-static or dynamic mechanical analysing has been reported [45,56,58–60]; namely, these offsets would be highly appreciated with respect to measuring bone nanomechanical properties.

4.2. Mechanical properties of mineralized tissue on titanium surfaces There are still few studies determining the biomechanical properties of the mineralized layer on titanium implants, and more studies are essential. In vitro bone cell behavior on a titanium sample can be similar to that of an in vivo afibrillar mineralized layer, and cell organization is comparable to that in a mineralized matrix on a titanium surface during culture with supplemental ascorbic acid, sodium b-glycerophosphate, and dexamethasone [61]. This is therefore an important consideration as the information obtained reflects in vivo events. The biomechanical evaluation of such micro-environments is challenging. Nano-biomechanical testing technology, such as nanoindentation, enables measurements of the mechanical properties of samples of very small volume [62– 64]. There are two crucial factors associated with the evaluation of in vitro mineralized tissue on titanium surfaces. First—and this is central to standard nanoindentation analysis—is the assumption that the surface is in ideal contact with the indenter penetration depth. In reality, all biological surfaces are characterized by some degree of roughness or heterogeneity [65]. Secondly, the

viscoelastic properties of biological tissues can cause problems [57]. In elastic–plastic materials, the obtained values are always constant, regardless of the load function. However, many polymeric biomaterials and biological tissues exhibit timedependent or viscoelastic behavior (Section 4.1) [27]. The observed effect of viscoelasticity on indentation is creep, or a sinking of the indenter tip into the sample under a constant load function. In such viscoelastic materials, the hardness and moduli usually vary with load function (or loading rate) because different loadings affect the creep or sinking behavior, leading to variations in the contact area (A) and unloading stiffness (S) in Eq. (2). A depth-dependent loading/partial unloading function consists of a number of loading/unloading portions (data points) toward the applied final loading force (Fig. 5). A partial unloading technique enables ideal contact between the indenter tip and the sample surface as a function of depth, so the depth-dependent mechanical properties of heterogeneous biological structures can be normalized in a direction perpendicular to the sample surface [66–68]. An additional holding time at each peak load, followed by an unloading portion, prevents an inappropriate unloading slope because the hold enables relaxation of the initial loading force. Numerous mineralized tissues on titanium surfaces can therefore be compared using average depth-dependent nanomechanical values. The mechanical properties of mineralized tissue are associated with maturation of the extracellular mineralized bone matrix proteins produced by osteoblasts. The increased hardness and elastic modulus of mineralized tissue is only observed on titanium surfaces with the release of hydroxyl radicals [55,66,67]. In the living body, oxidation of lysine residues in immature bone matrix proteins, such as collagen molecules, is the final enzyme reaction for cross-linking between molecules [69,70]. Therefore, the reduction of oxidation enzymes generates inferior mechanical properties of in vitro mineralized tissues, even if osteogenic differentiation is greatly increased by a supplemental anabolic agent such as bone morphogenetic protein-2 [68]. Aside from the natural oxidation enzymes, surface hydroxyl radicals may enable oxidation of the lysine residue from immature collagen molecules. Surface radical reactions on titanium may play an important role in the biomechanical considerations at the interface between healing bone and titanium implants.

4.3. Molecular structure analysis of mineralized tissue on titanium Raman spectroscopy analyses and alterations in the protein structure are of particular importance in distinguishing mature or immature calcified tissues, since subtle molecular changes often cause detectable vibrational changes [71–74]. The coupling of a Raman spectrometer with an optical microscope allows reduction of the analysis scale to the micrometer scale, with better spatial resolution than that obtained using infrared spectroscopy. There have been several characteristics associated with Raman spectra of calcified

89

Force (μN)

journal of prosthodontic research 59 (2015) 84–95

0

20

40

60

80

100

Time (sec)

Elastic modulus (GPa)

25 20 15 10 5 0 0

20

40

60

80 100 120 Contact depth (nm)

140

160

180

200

Fig. 5 – A depth-dependent load/partial unloading load function and consequent effective measurement range on cortical bone.

tissues (Fig. 6). Briefly, the ‘Amide I’ region between 1600 and 1700 cm1 revealed the occurrence of three or more bands at 1623, 1637 and 1656 cm1, attributable to characteristic collagen secondary structures [72,73]. The band at 1637 cm1 is characteristic of a collagen secondary structure, and is evident as a shoulder at the ‘Amide I’ band in the Raman spectrum of collagen. The increase in the ‘Amide I’ band shoulder at 1656 cm1 was attributable to the maturity of typical a-helix collagen secondary structures. This is related to hydrogen bonding between collagen molecules and subsequent cross-linking of tissues [75]. The Raman spectra also show bone mineral bands from apatite phosphate at 960 cm1 and carbonate at 1070 cm1. The ratio of mineral bands in the ‘Amide III’ band at around 1250 cm1, attributable to disordered and ordered bone matrix proteins, indicates the mineral to matrix ratio of the calcified tissues [74]. An inorganic apatite transformation associated with collagen cross-linking plays an important role in the biological stability of the mineralized layer at the interface between bone and titanium implants, thereby increasing the longevity of the implant prosthesis.

4.4.

lining cells for the sensing of mechanical stimuli [77]. Bone is often considered to be a fluid saturated porous network [60]. The application of mechanical strains causes an imbalance of pore pressures so that bone fluid pressure generates fluid flow. Such interstitial fluid flow is the most potential way of informing bone cells with respect to mechanical loading [78–80]. A bone nanoindentation study revealed a pile-up response against high-strain indenter penetration because of reduced fluid permeability within the bone which with time would diffuse radially and to a reduction in pile-up during constant load or displacement [60]. Along with the recovery process, proteins refolding and fluid permeability or flow back into the deformed area would occur. The large fluid flow as a consequence of pile-up response was only observed in highly mineralized intact bone while the higher collagen and fluid (porosity) content of lower mineralized bone is less able to generate a pile-up response and consequent fluid motion. Continuous balanced osseous remodeling is essential for the successful maintenance of implants. Thus, mechanotransduction due to the proper peri-implant bone tissue strains (mechanical properties) is an important determinant for prolonged biomechanical stability of titanium implant.

Bone mechanotransdution

Bone remodeling is dependent on cellular processes, namely bone formation by osteoblasts and bone resorption by osteoclasts. The detection of the applied stress is done by certain sensor cells such as bone lining cells of osteoblastic origin and osteocytes[76]. The cells transduce obtained mechanical signal and modulate bone formation and resorption (see also Section 5). The osteocytes are differentiated osteoblasts. A number of researchers have focused on the role of osteocytes in the mechanotransduction process [76]. The osteocytes form noncalcified three-dimensional network connected to the bone

5. Surface morphology-induced micromechanics and adaptive bone remodeling The mechanical stability of titanium implants has been primarily achieved by an intimate apposition in the bone surgical site and subsequent bone ingrowth to the implant surface. With increased demand for implants in orthopedic and oral/maxillofacial surgery, rough titanium surfaces have proven rather effective in improving osteogenic differentiation and micro-scale biomechanics in parallel [81,82]. Meanwhile, biomechanical evaluations of dental implants using

90

1800

1600

1400

1200

1000

800

652 633

732

793

875

958

1047 1039

1175 1156

1321 1263

1350

1479 1460 1443

1541

1620 1575

Raman intensity (a.u.)

1249

journal of prosthodontic research 59 (2015) 84–95

600

Raman shifts (cm-1)

Assignment

732

DNA

793

DNA

875

Hydroxyproline

~958

Apatite phosphate

~1070

Carbonate

1249

Collagen, amide III

1263

Collagen, amide III

1350

Glycosaminoglycans

1440~

Collagen-other proteins

1575

DNA

1600~

Collagen, amide I

Raman intensity (a.u.)

Raman shift (cm-1)

Raman shift (cm-1)

Fig. 6 – Peak assignments on mineralized tissue. the ‘Amide I’ region between 1600 and 1700 cmS1 revealed the occurrence of three or more bands at 1623, 1637 and 1656 cmS1, attributable to characteristic collagen secondary structures. The band at 1637 cmS1 is characteristic of a collagen secondary structure, and is evident as a shoulder at the ‘Amide I’ band in the Raman spectrum of collagen. The increase in the ‘Amide I’ band shoulder at 1656 cmS1 was attributable to the maturity of typical a-helix collagen secondary structures. This is related to hydrogen bonding between collagen molecules and subsequent cross-linking of tissues.

finite element analysis (FEA) have been widely performed by many researchers. As is well-known, FEA is a powerful tool for analysing stress distribution in peri-implant tissues, and most of the corresponding mechanical properties used in the FEA model, such as Young’s modulus and Poisson’s ratio of titanium, cortical and cancellous bone, are often obtained from relevant literature. Accordingly, material properties used in FEA by several studies [83–90] are shown in Table 2. By means of FEA, the effect of surface roughness treatment such as polishing, sandblasting, and plasma-spraying (or porousbeading) on the distribution of stresses at the bone–implant interface in immediately-loaded mandibular implants has been investigated [91,92]. Of these, surface treatments were represented by the coefficient of friction. More recently, enhanced mechanical interlocking has been advocated by bone tissue growth through porous titanium coated surfaces [30,93,94]. Upon titanium porosity, elastic properties closer to those of bone than to solid titanium surfaces enable reduction

of stress shielding in peri-implant tissues. In FEA, poroussurfaced implants appeared to distribute stress in a more uniform pattern around the implant when compared with smooth-surfaced implants [95]. It was indicated by FEA that the low modulus implant induces a stress distribution closer to the actual physiological phenomenon [96]. In addition to these features, the pores on titanium implants increase the surface area for osseous apposition [81,82]. Accordingly, porous coating on implants is of particular interest and has been given widespread attention by researchers as a method to reduce the stiffness mismatches between bone and titanium, achieving stable longevity of titanium implants by means of functionally graded biomechanics at the interface [30]. Aside from the load-bearing properties, a previous radiography study to examine the bone remodeling of porous-coated implants in dogs found that the implant with greater porosity of the coated area stabilized more quickly than that with a less

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Table 2 – Biomechanical evaluations of dental implant–homogeneous bone models by using FEA. Material properties useda

Author (year)

Significance

Titanium implant

Cortical bone

Cancellous bone

Harirforoush et al. (2014)

E = 104.8 GPa; n = 0.34; r = 4430 kg/m 3

E = 14 GPa; n = 0.3; r = 1300 kg/m 3

E = 0.49 GPa; n = 0.3; r = 1300 kg/m 3

Pen˜arrocha et al. (2013)

E = 117 GPa; n = 0.3

E = 13 GPa; n = 0.3

E = 1.37 GPa; n = 0.3

Ferraz et al. (2012)

E = 110 GPa; n = 0.35

E = 1.4 GPa; n = 0.3

E = 1.37 GPa; n = 0.3

Ausiello et al. (2012)

E = 120 GPa; n = 0.33

E = 13.7 GPa; n = 0.30

E = 0.50 GPa; n = 0.30

Ding et al. (2009)

E = 103.4 GPa; n = 0.35

E = 13.7 GPa; n = 0.30

E = 1.37 GPa; n = 0.30

Baggi et al. (2008)

E = 114 GPa; n = 0.34

E = 13.7 GPa; n = 0.3

E = 0.50 GPa; n = 0.3

Tanimoto et al. (2006)

E = 107 GPa; n = 0.34; r = 4500 kg/m3; h = 0.0003 E = 107 GPa; n = 0.33

E = 13.7 GPa; n = 0.30; r = 2000 kg/m3; h = 0.0100 E = 13.7 GPa; n = 0.30

E = 0.69 GPa; n = 0.30; r = 1000 kg/m3; h = 0.0381 E = 1.37 GPa; n = 0.30

Kitagawa et al. (2005) a

The resonace frequency of the implant was strongly influenced by the cortical bone to cancellous bone ratio at the implant/bone interface. The implant neck design and implantabutment joint types influence peri-implant bone stresses and abutment micromovement. The implant with microthreads showed higher stress concentration for cortical bone compared with the smooth implant, and lower stress concentration for cancellous bone. The implant thread design plays a crucial role to reduce induced stresses and damage in bone. The stress and strain on the implant-bone interfaces significantly decreased with increasing the implant diameter. The implant design (i.e., diameter and length), crestal bone geometry, and placement site affect the mechanisms of load transmission. The implant with stress-absorbing element had a higher damping effect than the implant without stress-absorbing element. The stress in bone was sensitive to the thickness and elastic modulus of cortical bone.

References

[87]

[89]

[86]

[83]

[85]

[84]

[90]

[88]

E: elastic modulus; n: Poisson’s ratio; r: density; h: loss factor.

porous coated area [97]. Thus, implant osseointegration is considered to involve a combination of initial bone ingrowth onto the surface and adaptive bone remodeling against strain energy density over time [28,29]. Bone is most likely to have the ability to adapt against strain energy. Several attempts to quantify the adaptive bone remodeling process have been reported in the literature based on the widely accepted original hypothesis of Wolff’s theory [98–101]. These studies implied that bone has internal strain sensors measuring the internal loading condition and transducers releasing the signals, so that bone remodeling is activated by a cooperative interaction between osteogenic and osteoclastic cells. The internal mechanical loading such as strain energy density in bone micro-structures may be determined mathematically by FEA [28,102–104]. An example by Weinans et al. suggested resemblance to the density distribution of a proximal human femur by means of a twodimensional FEA model [28]. Given the fact that intact bone is likely to have a selfoptimization capability, the biomechanical integrity of the implant–bone interface is a critical indicator for implant stability and longevity. Rungsiyakull et al. conducted surface morphology (porosity) optimization for titanium implants in terms of interfacial biomechanics and adaptive bone remodeling in the peri-implant region using a FEA model with a multiscale remodeling algorithm [29]. The macroscopic model generates the global responses of bone, while the microscopic level quantifies the change in the apparent

bone density of the representative volume elements. The simulation of adaptive bone remodeling in the study by Rungsiyakull et al. suggests that increasing the volume fraction of the coating particles would result in increased bone density over time, as well as achieving a better initial stability than those with a lower volume fraction [29]. As shown above in Table 1, homogeneous models for bone structure have been employed extensively in FEA studies. This is mainly because linear analysis is useful for analysing the stress distribution of homogeneous materials; however, bone is hierarchical in structure and induces timedependent bone remodeling. Accordingly, bone mechanical properties should be considered on the basis of an empirical relationship between elastic modulus and bone mineral density [105–107]. In this respect, biomechanical simulation of titanium implants should be performed with caution given to the aforementioned difficulty in terms of complex bone mechanical properties related to time-dependency [108,109].

6.

Conclusion

Although the titanium implant surface enables a biological bonding capability through the cement layer comprising body fluid or proteins, the weaknesses of the protein layer need to be further stabilized by microscale mechanical interlocking. Roughened or porous implant surfaces should accommodate

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enhanced biomechanical properties due to the maximized contact area between newly formed bone and the implant surface. Understanding titanium and basic bone mechanical properties is indispensable prerequisite for the development and clinical use of implants. Bone mechanical properties are quite complicated and are still not fully understood. Thus, precise prediction of implant biomechanics has long been a challenge. Nanomechanical testing technology is a useful tool for achieving greater precision with respect to the biomechanical response, including adaptive bone remodeling in the periimplant region.

Conflict of interest The authors have no conflicts of interest to declare.

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