Classification and Effects of Implant Surface Modification on the Bone: Human Cell–Based In Vitro Studies

Titanium and its alloys are established materials for dental implants because of their physical strength, material stability, and tissue compatibility.¹  The first generation of dental implants wa machined with a smooth surface and was used successfully clinically.²  However, the healing period for these implants may be as long as 6–9 months before they are osseointegrated enough to be loaded. Osseointegration is dependent on the direct interaction between the bone cells and the titanium surface. The design of a new and efficient implant material requires an understanding of the adhesion of osteoblasts at the bone-implant interface. Thus, the search continues for an ideal implant surface modification that can osseointegrate to bone faster and with a stronger bone-to-implant interface.

This review will present (1) the various implant surfaces and how each is achieved, (2) the parameters to characterize the implant surface morphology and how they are evaluated, and (3) the human cell–based studies to determine the strength and quality of the bone-implant contact resulting from each surface modification.

The adhesion and differentiation of osteoblastic cells are influenced by the surface properties of the dental implant.^3,4  Surface properties include chemical composition, surface energy, roughness, and topography. In dental implants, surface roughness are often modified to modulate bone apposition.^5,6  Surface roughness can be described as macro-, micro-, and nanometer-sized texture. Macro- and micrometer roughness facilitates mechanical anchorage to bone.⁵  Nanometer roughness affects the adsorption of proteins and the adhesion of osteoblastic cells. It can modulate the rate of ossteointegration.^7,8  A variety of surface treatments (Table 1) can be used to produce the desired surface topography. Commercially available implants vary in titanium composition and surface modifications (Table 2), having an understanding of these differences can help clinicians make an informed choice in implant selection for their patients.

The machined implant is turned, milled, or polished. It is minimally rough, with a surface area roughness (S_a) value of 0.3–1.0 μm.⁹  The surface morphology is determined by the manufacturing tools used, the implant material, the lubricant, and the speed at which it is machined.

The sandblasted implant is grit blasted by small particles (alumina or titanium oxide), which creates craters and ridges on impact. The surface morphology is determined by the following particle characteristics: its material, size (25, 75, and 250 μm), shape, density, and speed at which it is propelled.¹⁰ 

Laser-etched implants uses lasers as a micromachining tool to produce selective modification and generation of complex microstructures at micrometer and nanometer level. Advantages of the laser technique include the absence of chemicals and the convenience of being able to incorporate it into the routine manufacturing.

Compaction of nanoparticles on the implant surface conserves the chemistry of the underlying surface while changing or modifying the chemistry and structure of the outer surface layer.¹¹ 

Porous tantalum trabecular metal has recently been designed and developed to enhance secondary stability via bone ingrowth, and it is incorporated on the surface of the titanium dental implants.¹²  It is not a regular coating or surface treatment. The titanium alloy and the porous tantalum trabecular metal of the implant are prepared separately. The porous vitreous carbon scaffold acts as a second layer on the titanium implants. Tantalum is deposited onto the vitreous carbon scaffolds using chemical vapor deposition or infiltration and then laser welded onto the titanium alloy core.^13,14  The porous layer, with a structure similar to trabecular bone, is used to improve the bonding between the osseous tissue and the dental implants through osseointegration.^12,15,16 

The acid-etched implant is pitted by acid (HCl/H₂SO₄ mixture or 2% HF/10% HNO₃) removal of grains and grain boundaries, which are more sensitive to etching. This selective removal of the implant surface is minimally rough (S_a value of 0.3–1.0 μm).¹⁷  The surface morphology is determined by the implant material, the microstructure of the surface, the type of acid, and the soaking duration. To further improve the implant surface, the implants are blasted by particles before acid etching. This subsequent etching removes embedded blasted particles and gives a dual surface roughness (S_a value of 1–2 μm).¹⁸ 

Alkaline oxidation can be achieved by soaking the implant in high alkaline solutions followed by heat treatment. There are several examples of such methods (eg, soaking in 4–5 M sodium hydroxide solution and heat treatment at 600°C for 24 hours or soaking in boiling alkali solution of 0.2 M sodium hydroxide and heat treatment at 1400°C for 5 hours). The alkaline treatment can be preceded by acid etching to increase the porosity of the titanium surface.¹⁹ 

The anodized implant undergoes anodic oxidation, an electrochemical process in an electrolyte that results in a microstructure surface with micrometer-sized open pores. This process involves passing a current through the implant as the anode with phosphoric acid as the electrolyte to form the surface oxide. The surface morphology could be modified by varying the anode potential, the electrolyte composition, the temperature, and the current.²⁰  Ions such as phosphorus,²¹  calcium,²²  and magnesium²³  can also be integrated into the implant surface via modification of the electrolyte composition.

Peroxidation of the implant surface produces a titania gel layer through treatment with a perioxide-based chemical agent. Chemical treatment of implant surfaces with hydrogen peroxide results in chemical dissolution and oxidation of the titanium surface. When titanium surfaces react with hydrogen peroxide, titania gel layers are formed. The thickness of titania layer formed can be controlled by adjusting the treatment time, and it has been demonstrated that when immersed in simulated body fluid, thicker layers of titania gel are more favorable for the deposition of apatite.²⁴ 

This is a fluoride-modified nanostructure implant surface. The surface modification involved blasting with titanium oxide (TiO₂) and treating with dilute hydrofluoric acid. The optical interferometer microscopy data of the fluoride-modified implant surface showed a mean surface area roughness of 1.24–1.26 μm.²⁵ 

Vacuum treatment of the implant surface can be achieved by glow-discharge deposition of coating material from a solid target or by reactions in the gas phase. It can also be achieved by exposing the titanium surface to a glow-discharge of energetic ions that specifically modify the surface properties by bombardment.

Another method in vacuum treating implants is the ion implantation method; this involves the bombardment of high energy ions, which penetrate the surface of the implant.²⁶  It can be controlled by varying the concentration and the energy of the ions and can increase the corrosion resistance by forming a titanium-nitrogen (Ti-N) surface.²⁷  This method can also be used to develop antimicrobial surfaces on the implant via deposition of fluoride and silver (Ag) ions on implant surface without toxicity.²⁶ 

Plasma sprayed calcium phosphate-coated implants have improved bioactivity. Hydroxyapatite (HA) is a form of calcium phosphate coating. Plasma spraying with HA can increase the surface area, as well as increase the average surface roughness (R_a value 5.0 ± 1.0 μm).²⁸ 

The bioactive coating of titanium implants involves precipitation of calcium phosphate apatite crystals on the titanium surface. The deposition of calcium phosphate onto titanium surfaces can be achieved by using a titanium cathode and a platinum anode to generate a current producing a brushite coating, which is hydrothermally processed to apatite on the implant surface.^29,30  It can also be achieved by immersing in stimulated body fluids and producing a heterogeneous nucleation and growth of bone-like calcium phosphate crystals on the implant surface.^31,32  Calcium phosphate apatite formation on titanium metals in stimulated body fluids can be further enhanced by heat treatment after exposure to strongly acidic or alkaline solutions.^33,34 

This involves the coating of titanium implant surface with synthetic arginylglycylaspartic acid (RGD) peptides that contain binding sites for integrin receptors.³⁵ 

Antibiotics such as cephalothin, carbenicillin, amoxicillin, cefamandol, tobramycin, gentamicin, and vancomycin can bind to calcium-based coatings of implants, as well as be released from it. This antibiotic-releasing coating also retains its antimicrobial properties.³⁶ 

The implant surface can be coated with osteogenesis-stimulating agents to accelerate angiogenesis and bone formation around implants. These growth factors coating the implant can be bone morphogenetic proteins (BMPs), transforming growth factor β1 (TGF-β1), vascular endothelial growth factors (VEGFs), platelet-derived growth factors (PDGFs), or insulin-like growth factors (IGFs). BMPs can be directly incorporated into the implant surface,³⁷  or they can be incorporated via the use of a plasmid containing the BMP-encoding gene.³⁸ 

The implant surface can also be coated with bone remodeling-associated agents like bisphosphonates. Bisphosphonates have a great chemical affinity for calcium phosphate molecules and thus can be incorporated via the biomimetic coating procedure. Bisphosphonates can also be coupled with RGD peptides and chemically absorbed on titanium to produce synergistic osteogenic effects.³⁹ 

Wennerberg et al⁴⁰  published guidelines for surface topography measurement. To obtain reliable quantification, methods for topographic measurement should have both excellent vertical and lateral resolution, as well as a reasonable analysis area. A scanning electron microscope (SEM) can qualitatively evaluate surface structure in both the micro- and nano-range. Scanning electron microscopy can also obtain quantitative topographic data via stereo imaging and image analysis.

Different spectroscopic techniques used to evaluate the surface elemental composition include Auger electron spectroscopy (AES), energy dispersive X-ray spectroscopy (EDS), X-ray photoelectron spectroscopy (XPS), and secondary ion mass spectroscopy (SIMS). It is important to note that the surface composition of the implant differ from that of the corresponding bulk material.

Techniques enabling phase identification of implant surface crystal structure include Raman spectroscopy, high-resolution transmission electron microscopy (HRTEM), electron diffraction, X-ray diffraction, and electron backscattering diffraction.

Surface energy is commonly measured by the contact angle between the implant surface and liquids of differing surface tension. The surface tension is then calculated from a Zisman plot of the measured contact angle. A more simple analysis can also be done using the water contact angle as a representative parameter of the surface tension.

Different surface modification techniques have been mainly used to improve the surface roughness and hydrophilicity. Some modified surface compositions could also contain bioactive substances. Implant morphology such as grooves, ridges, and tool marks can influence the interaction between the bone and the implant. The implant morphology can also increase the overall surface area available for osseointegration. Rougher surfaces can stimulate attachment, differentiation, and proliferation of bone cells, thus increasing bone growth and mineralization. Rougher surfaces with an open structure have been shown to induce faster and more effective osseointegration. Unfortunately, this rougher surface substrate tends to accumulate bacteria.

Most techniques are applied to change the surface roughness. Surface roughening can be induced by machining, blasting, laser etching, acid/alkaline etching, anodization, and coatings. The roughness of implant surfaces treated via machining is normally less than 1 μm. The surfaces treated via blasting and plasma spraying result in the highest roughness. The size of the blasting particles determines the roughness. However, coating machined titanium or sandblasting with acid-etched titanium surfaces with polyelectrolytes does not alter the surface roughness.⁴¹  There is no consensus on the optimal implant roughness that produces the best effects on bone.

Anodized titanium substrates have numerous nanoscale surface features.⁴²  Anodization in acetic acid can produce intensely etched interlaced grooves that are distributed homogeneously on the surface. Further high density submicron scale pores can be developed with sulphuric acid anodization. When both treatments are combined, the surface created is a multilevel surface-porous anodic layers consisting of interlaced macroscopic grooves overlaid with submicrometer pores.⁴³  Furthermore, elements like calcium, oxygen, phosphorous, and sodium can be introduced onto the titanium surface during anodization.⁴⁴ 

Hydrophilicity contributes to the wettability and the surface energy of the implant surface and is affected by the surface composition. Present surface modification techniques can increase surface area and hydrophilicity.⁴⁵  Electrochemical functionalization can modify both the surface chemistry and the wettability. Electrochemical anodization of the grit-blasted and acid-etched titanium produces the most hydrophilic material compared with polished titanium and grit-blasted and acid-etched titanium without electrochemical anodization.⁴⁶  The changes of surface compositions show positive effects by incorporating hydroxide ions⁴⁷ , fluoride ions,⁴⁸  calcium ion,^49,50  phosphate ions,⁵¹  magnesium ions,⁵²  and bioactive substances.⁵³  Generally, a combination of surface modifications are used to create a titanium surface with a combination of surface properties. However, there is a lack of optimized quantitative data to support any single or combined implant surface modification.

The adhesion, differentiation, and proliferation of human osteoblastic cells on the titanium surface are crucial for successful osseointegration. This osteoblastic cell interaction with the titanium surface can be modulated by the chemical composition, the surface energy, the roughness, and the topography of the respective surfaces (Table 3).

Osteoblastic cells adhere in several different phases.⁵⁴  In the attachment phase, the titanium surface links to the osteoblast via physiochemical forces like ionic forces and van der Waals forces. In the adhesion phase, the osteoblast binds further via different biomolecules like integrins, cell surface receptors, extracellular matrix protein, cell membrane proteins, and cytoskeleton proteins.^4,54  Integrins transmit a signal from the extracellular matrix to regulate osteoblast adhesion, motility, shape, growth, and differentiation.^54,55  Integrins form focal contacts when extracellular matrix-adsorbed biomolecules interact between the bone cells and the titanium surface.^4,55 

The following titanium surface modifications can promote stronger cell adhesion. There is greater cell adhesion for aluminium oxide (Al₂O₃)-blasted and Al₂O₃-blasted and hydrofluoric acid–etched surfaces compared with polished titanium.^56,57  Niobium vapor deposition on titanium alloy promoted stronger osteoblast attachment and spreading compared with tin vapor deposition, chromium vapor deposition, and controls.⁵⁸  Niobium oxide has a significant effect favoring cell adhesion and differentiation.⁵⁹  Coating of titanium implants with titanium carbide^60,61  or with recombinant human dentin matrix protein 1⁶²  increases the proliferation, adhesion, and differentiation of the osteoblast. In addition, electrical deposition of calcium phosphate can also improve adhesion and proliferation.⁴⁴  Similarly, calcium ion–implanted titanium can enhance expression of the cell proliferation–associated nuclear Ki-67 antigen and increase the number of mitotic cells.⁶³  Osteoblast viability, adhesion, and gene expression were enhanced by the addition of 3D nanostructure to the magnesium-loaded mesoporous titanium oxide coating⁵²  or by magnesium implantation of the sandblasted and acid-etched titanium surface.⁶⁴  Quercitrin-nanocoated surfaces also promote faster stem cell adhesion and mineralization than control surfaces.⁶⁵  Compared with unmodified surfaces, cell adhesion can also be significantly enhanced by grafting peptides^50,66,67  or bioactive polymer to titanium.⁶⁸  Titanium implants modified by acrylic acid surface grafting–collagen I coupling can enhance cell adhesion compared with galvanostatic anodization.⁵³  In contrast, multilevel surface-porous anodic layers with macroscopic grooves produced by a 2-step anodization treatment can promote osteoblast adhesion and growth.⁴³  Anodic spark deposition with alkali etching,⁶⁹  ionic plasma deposition, and direct nitrogen ion immersion plasma deposition⁷⁰  of polymeric and metallic coatings can also enhance osteoblast adhesion.

However, commercially available dental implant surfaces are comparable in terms of bone cell adhesion. Baldi et al⁷¹  compared 5 different implants (Tapered Internal [BioHorizons Implant Systems], Nanotite [3i Implant Innovations],⁷²  Osseotite [3i Implant Innovations],⁷³  Straumann SLActive Standard Implant [Institut Straumann],⁷³  and SwissPlus [Zimmer Dental]) and found no significant difference in cell adhesion.

The following titanium surface modifications can promote greater cell differentiation. The osteogenic differentiation was greater on rough surfaces compared with smooth surfaces and is unaffected by the surface chemistry.⁴⁵  The expression of integrins and cell differentiation was significantly modulated by titanium nanopore size. Nanotopographic surfaces can increase cell differentiation^74,75  and was more potent in 30- and 150-nm nanopore surfaces than the 300-nm nanopore surface.⁷⁶  The surface topography may influence the phenotypic expression of osteoblast-like cells.⁷⁷  Similarly, Mendonça et al⁷⁸  reported that nanoscale alumina surface promoted greater cell differentiation and osteoblastic gene expression than machined and acid-etched surfaces. The nanostructured layer can induce overexpression of many bone turnover proteins and increase the number of osteoblast surface receptors.⁶¹  Optimum nanospacing of titania–zirconia–zirconium titanate nanotubular surfaces with inner diameter of 40 ± 12 nm induces the highest density of bone cells.⁷⁹ 

In addition, Knabe et al⁸⁰  observed that HA-coated titanium has the most effect on cell differentiation compared with the other surfaces tested. On a hydroxylated titanium surface, osteoblasts were more differentiated and exhibited increased alkaline phosphatase activity.⁴⁷  The calcium phosphate–impregnated surface induced significantly higher cell differentiation levels and alkaline phosphatase activity than the controls.⁵¹  Titanium surfaces preincubated with blood also exhibited increased alkaline phosphatase and collagen type I compared with control surfaces.⁸¹  Sandblasted and laser-pitted titanium surfaces produced higher early production of alkaline phosphatase compared with control.^82,83  Lipid-functionalized solid titanium⁸⁴ and laminin-derived functional peptide coating on pure titanium⁵⁰  support osteoblast maturation via synergistic increases in total alkaline phosphatase activity. Plasma-sprayed titanium dioxide/zirconia coating can significantly increase alkaline phosphatase activity, Runx2 expression levels, and osterix compared with titanium dioxide coating.⁸⁵  Titania-HA nanocomposite coatings promote significantly higher alkaline phosphatase activity and osteocalcin than the pure hydroxyapatite- and titanium-coated substrates.⁸⁶  Osteocalcin is a late osteoblastic differentiation marker; an increase in osteocalcin would indicate that more undifferentiated cells have switched to the osteogenic lineage.⁸⁷  Osteocalcin was higher on the sandblasted and acid etched surfaces compared with controls and was significantly higher when nanostructures are superimposed via a simple oxidation process on these surfaces.⁸⁸ 

Osteogenic differentiation on rough titanium surfaces is enhanced by early mineralization of the osteogenic matrix, and can be further enhanced by the wettability of hydrophilic surface.⁸⁹  Osteoblastic differentiation and mineralization can also be enhanced by ultraviolet radiation and alendronate sodium trihydrate treatment of the titanium surfaces.⁹⁰  Gene expression for osteoblast differentiation were up-regulated in human bone cells on carbide-coated titanium compared with uncoated titanium.⁶⁰  Cell differentiation to osteoblasts can be further induced by α2β1 integrin–mediated osteoblast response to microstructured titanium.⁹¹ 

Turned titanium surface exhibited the lowest cell proliferation.⁸³  Other smooth pretreated titanium surfaces reported a high proliferation rate, but with an immature osteoblastic phenotype that is low in integrin levels and cell differentiation.⁹²  The following are some of the titanium surface modifications that can promote greater cell proliferation. The rate of cell proliferation was significantly higher on the poly-lactic-co-glycolic acid/nano-HA nanofiber–coated and the poly-lactic-co-glycolic acid/collagen/nano-HA–coated scaffolds compared with the poly-lactic-co-glycolic acid and poly-lactic-co-glycolic acid/collagen fibers for both pure titanium and the titanium alloy.⁹³  Heo et al⁹⁴  reported that cell proliferation was higher on rough 75 μm Al₂O₃-blasted surfaces compared with machined and anodized surfaces. This sandblasted surface subjected to additional plastic deformation with a punching process can further increase the cell proliferation.⁹⁵  Laser-engineered porous titanium promotes cell viability and cell proliferation, with a greater cell response at a hemispherical porosity of 20 μm.⁹⁶  Similarly, Rosales-Leal et al⁵⁶  reported improved cell proliferation in hydrofluoric acid–etched surfaces and Al₂O₃ blasted. These acid-etched and sandblasted surfaces induced higher cell proliferation compared with plasma-sprayed surfaces.⁸⁰  Cell proliferation can be further enhanced by magnesium implantation of the sandblasted and acid-etched titanium surface⁶⁴  and by zinc-implanted pure titanium.⁹⁷  The titanium surface immobilized with bone morphogenetic proteins with⁹⁸  or without heparin modification can also increase cell proliferation compared with unmodified titanium.⁹⁹  Porous titanium implanted with adipose tissue–derived mesenchymal stem cells can also enhance cell proliferation and extracellular matrix production.¹⁰⁰ 

However, variations in methods or materials used in titanium surface modifications do not always promote cell proliferation. Klinger et al¹⁰¹  observed that cell proliferation was not statistically different between machined surfaces and rough sandblasted surfaces that were acid etched with or without hydrofluoric treatment. Furthermore, hot acid etching of sandblasted surfaces resulted in decreased cell proliferation compared with the polished surface.¹⁰²  Uggeri et al¹⁰³  reported that cell proliferation is also greater in machined smooth titanium and in zirconium oxide–sandbasted and acid-etched titanium compared with aluminium oxide–sandbasted and acid-etched titanium. In addition, Kohal et al¹⁰⁴  observed that cell proliferation was retarded in anodized surfaces for the first 7 days; however, at 28 days, cell proliferation was at the same level for all surfaces.

Rausch-Fan et al¹⁰⁵  compared hydrophobic and hydrophilic surfaces and found increased cell proliferation in hydrophobic acid–etched, followed by hydrophilic acid–etched, hydrophobic coarse grit–blasted and acid-etched, and hydrophilic coarse grit–blasted and acid-etched surfaces. Hydrophilic surfaces have a higher surface energy in comparison to hydrophobic surfaces. These hydrophilic-modified titanium surfaces support homogeneous spatial osteoblast cell growth and mineral deposition compared with hydrophobic titanium surfaces.^106,107  The chemically modified hydrophilic SLActive surfaces up-regulated more osteogenic transcription genes than the hydrophobic SLA surfaces.¹⁰⁸ 

Hydrophilicity can increase the expression of alkaline phosphatase, osteoprotegerin, and osteocalcin and can significantly decrease VEGF-A and TGF-β1.⁹¹  The plasma electrolytic oxidation coating is significantly more hydrophilic and induces significantly higher collagen synthesis compared with the plasma-sprayed hydroxyapatite coatings.¹⁰⁹  In contrast to previous mono-culture studies, the hydrophobic acid–etched surface promoted both proliferation and expression of angiogenesis-associated genes in human umbilical vein cells under coculture conditions.¹¹⁰ 

The surface chemistry, the surface topography, and the surface energy of the titanium surface have a crucial effect on osteoblast and osteocyte function. However, this relationship between patterns of gene expression and adhesion/ differentiation/ proliferation of bone cells on various titanium surfaces is not clear. This interaction is the basis for successful bone-to-implant contact during osseointegration. These in vitro studies can be used to postulate how these individual cells would react to different dental implant surface characteristics in the living body. Thus, selected titanium surface characteristics that induce favorable osteoblast function in vitro can then be further evaluated via animal and human studies. This will provide further information on how these surfaces would perform in relation to the longevity of these surface-treated implants when subjected to the oral environment and to functional loading. Animal and human studies are beyond the scope of this review and will be discussed in future reviews.

Abbreviations

AES

Auger electron spectroscopy

Ag

silver

Al₂O₃

aluminium oxide

BMPs

bone morphogenetic proteins

Ca(H₂PO₄)₂

calcium dihydrogen phosphate

CaP

calcium phosphate

Cr

chromium

EDS

energy dispersive X-ray spectroscopy

HA

hydroxyapatite

HCL

hydrochloric acid

HF

hydrofluoric acid

HRTEM

high-resolution transmission electron microscopy

H₂SO₄

sulphuric acid

IGFs

insulin-like growth factors

Mg

magnesium

NaOH

sodium hydroxide

Nb

niobium

Nb₂O₅

niobium oxide

PDGFs

platelet-derived growth factors

RGD

arginylglycylaspartic acid

S_a

surface area roughness

SEM

scanning electron microscope

SIMS

secondary ion mass spectroscopy

Sn

tin

TGF-β1

transforming growth factor β1

Ti-N

titanium-nitrogen

TiO₂

titanium oxide

UV

ultraviolet

VEGFs

vascular endothelial growth factors

XPS

X-ray photoelectron spectroscopy

ZrO₂

zirconium oxide

M.T., W.X., H.E., and J.B.S. declare that they have no conflict of interest as to the content of the manuscript. S.R.J. holds common stock in the company DENTSPLY International (York, Pa), which is a company that markets various dental implant systems. Nonetheless, S.R.J. does not believe that the abovementioned disclosure presents any conflict of interest with respect to the subject matter of this review.