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LIBRARY OF CONGRESS CATALOGING-IN-PUBLICATION DATA
Surface coatings / editors, Mario Rizzo and Giuseppe Bruno. p. cm.
Includes index.
ISBN 978-1-61668-992-6 (E-Book)
1. Surface sealers. 2. Protective coatings. I. Rizzo, Mario, 1958- II. Bruno, Giuseppe, 1959- TA418.9.C57S88 2009
667'.9--dc22
2009003249
C
ONTENTS
Preface vii
Chapter 1 State of the Art Bioactive Titanium Implant Surfaces 1 Anna Göransson Westerlund
Chapter 2 Antimicrobial Surface Coatings in Packaging Applications 45
Jari Vartiainen
Chapter 3 Environmentally Friendly Conversion Coating Applications for
Hot Rolled Steel (HRS) Prior to Powder Coating Application 93
Bulent Tepe
Chapter 4 Precise Synthesis of Amphiphilic Polymeric Nano Architectures Utilized by Metal-Catalyzed Living Ring-Opening Metathesis Polymerization (Romp)
123
Kotohiro Nomura
Chapter 5 Atmospheric Pressure Plasma Polymerisation 153 R. Morent, N. De Geyter and C. Leys
Chapter 6 Interface Research on Films and Coatings 177 Xiaolu Pang and Kewei Gao
Chapter 7 A Study on Inorganic Metallic and Dielectric Thin Films Grown on Polymeric Substrates at Room Temperature by PVD and CVD Techniques
189
P. Mandracci, R. Gazia, P. Rivolo, D. Perroneand A. Chiodoni Chapter 8 Sonochemical Coatings of Nanoparticles
on Flat and Curved Ceramic and Polymeric Surfaces
213 A. Gedanken and N. Perkas
Chapter 9 Post-Consumer PET and Post-Consumer PET-Containing Materials for Flame Spray Coatings on Steel:
Processing, Properties and Use
237
Chapter 10 Coating of Carbon Nanotubes with Insulating Thin Layers 259
Martin Pumera
P
REFACE
This book presents current research on thin films and coatings. The mechanical properties of films and coatings, which are highly affected by their microstructure and their adhesion to substrates, are reviewed. Furthermore, electronic semiconductor devices and optical coatings, which are the main applications benefiting from thin film construction are looked at.
This book discusses antimicrobial surface coatings as promising applications of advanced active food packaging systems. Ways in which they effectively control the microbial contamination of various foodstuffs are analyzed. Research that has been done in the last decade using ultrasonic waves for coating surfaces is also examined. Finally, since coatings and films mechanical properties are highly affected by their microstructure, and their adhesion to substrates, this book includes research on interface microstructure and the important role that bond formation plays on coatings and films.
Materials that have the ability to bond to living tissue are defined as “bioactive” and the first possibly bioactive material Bio-glass was described in the 1970s by Hench and co-workers. Furthermore, Jarcho and co-workers were the first to present indications of a direct bone bonding to hydroxyapatite (HA). The mechanism proposed was ion exchange resulting in an apatite layer requested not only by the bone cells but also because proteins that serve as growth factors preferentially adsorb to this layer. The “bioactive” properties of these materials were based on morphological observations of the tissue coalescence by TEM and apatite formation in vitro and in vivo. Poor mechanical properties of these materials make them unsuitable for load-bearing, clinical applications. Therefore, experiments were made to coat titanium surfaces with calcium phosphates by plasma spraying technique. The surfaces indeed showed rapid tissue response initially, but in later stages biodegradation and delaminating of the thick coating was frequently observed. Additionally, the line-of sight problem made the technique unsuitable to use for the coating of complex shapes.
To avoid these problems, alternative techniques have been used to make commercially pure (CP) titanium “bioactive”. Chapter 1 reviews recent research on “bioactive” titanium implant surfaces, focusing on five specific modifications:(I) etching with fluoride containing acids, (II) alkali-heat treatment, (III) anodization and (IV) ultra-thin coatings of calcium phosphates in sol-gels. Another possible approach to enhance the bone response is to (V) immobilize organic bio-molecules to the surface.
These five CP titanium surface modifications will be reviewed separately with a short background, suggested mechanism of action and performance in simulated body fluids (SBF), in vitro and in vivo. Clinical evaluations will be discussed briefly. Each section is followed by
an appendix with a list of references of importance for the area of interest. The references are presented as short abstracts with similar information providing a quick overview and easy comparison of the studies.
As explained in Chapter 2, antimicrobial packaging materials are interesting and promising applications of advanced active food packaging systems. They can effectively control the microbial contamination of various solid and semisolid foodstuffs by inhibiting the growth of micro-organisms on the surface of the food, which normally comes into direct contact with the packaging material. Recently, a lot of efforts has been put on the development of antimicrobial packaging, which can considerably prolong the shelf lives of packed food products and/or decrease the need of preserving agents in foods. Some promising results have been obtained of which the surface activation and coating treatments seem to offer the most applicable solutions. Antimicrobial surface treatment can be done by several ways such as coating, printing, grafting or covalent binding. Other surface pre-activation methods such as physical, chemical or enzymatic treatments or their combinations may be necessary to produce permanently coupled antimicrobial agents. By using surface treatments the harmful effects on valuable bulk properties of packaging materials can be minimized. Also the safety aspects should be easier to fulfil as migration of substances can be kept at very low level. Antimicrobial surface treatments can be completely separated from the high-volume production lines of bulk materials. They can be done with smaller scale equipment immediately before the packaging is formed ensuring the maximum antimicrobial efficiency. Development of antimicrobial packaging materials, which can be produced at commercial scale, is a challenging and promising area, where intensive research is still needed. They can be exploited in direct contact with certain foods only and each food system must be investigated separately.
Hot rolled steel (HRS) is extensively used in a wide range of applications by many different industries such as automotive, domestic appliances, defence etc. It is common knowledge that hot rolled steel comes with oxide scale, often called mill scale, on the surface, due to the hot rolling process. Despite the disadvantage of oxide scale on HRS, it is still one of the most popular materials used in industry due to its availability, cost and ease of profiling properties. One of the most important coating applications for HRS is powder coating, which has a number of advantages over its favourability to wet coating, therefore it is widely used for HRS components in industry, prior to powder coating, to increase corrosion and blister resistance and enhance adhesion pre-treatment systems are used. Pre-treatment systems usually contain five or more stages: cleaning, rinsing, conversion coating, rinsing and passivation. Conversion coating is the most important stage in the pre-treatment process and it is usually phosphating. Phosphating offers many advantages, however it is considered as a hazardous material to human health and the environment. The phosphating process creates sludge, which results in pipe and pump blockages and sludge built up in the phosphating tank. These concerns have driven chemical companies to conduct research aimed at finding a conversion coating that meets the requirements of health and safety and is environmentally friendly. Some companies have already developed environmentally friendly conversion coating systems which are promoted as ecological material and an alternative to the phosphating process.
The main objective of Chapter 3 is to evaluate the ability of commercially available environmentally friendly pre-treatment systems as a metal pre-treatment in finishing operations, to eliminate or reduce the amount of environmentally hazardous and toxic
chemicals. This objective must be accomplished whilst maintaining equal or better product performance properties, with economic benefit or no significant economic penalty to the metal finishing companies who would like to change their pre-treatment system to an environmentally friendly pre-treatment system. The evaluation focuses on technical performance and economics while validating the laboratory tests and environmental benefits.
In order to evaluate the conversion coatings’ performance studies on: corrosion behaviour, adhesion and blister resistance, salt spray, prohesion test, Electrochemical Impedance Spectroscopy (EIS) measurement, cross hatch test, conical bend test, pull-off test, humidity test and surface morphology were performed. In this chapter the most popular environmentally friendly conversion coatings were evaluated. Environmentally friendly coatings are usually Silane and Zirconium based.
Chapter 4 summarizes recent examples for precise synthesis of amphiphilic block copolymers by adopting transition metal-catalyzed living ring-opening metathesis polymerization (ROMP). In particular, unique characteristics of the living ROMP initiated by molybdenum alkylidene complexes (so-called Schrock type catalyst), which accomplish precise control of the block segment (hydrophilic and hydrophobic) as well as exclusive introduction of functionalities at the polymer chain end, enable us to provide the synthesis of block copolymers varying different backbones by adopting the “grafting to” or the “grafting from” approach. Moreover, use of the “grafting through” approach (polymerization of macromonomers) by the repetitive ROMP technique, using the molybdenum alkylidene catalysts, offers precise control of the amphiphilic block segments.
Plasma polymerisation is a unique technique for modifying material surfaces by depositing a thin polymer film. Plasma polymerised films have received a great deal of interest due to their unique characteristics. These coated films are pinhole-free and highly cross-linked and are therefore insoluble, thermally stable, chemically inert and mechanically though. Furthermore, such films are often highly coherent and adherent to a variety of substrates including conventional polymer, glass and metal surfaces. Due to these excellent properties, plasma polymerised films can offer many practical applications in the field of mechanics, electronics and optics.
Plasma polymerisation at low pressure is already a well established technology. However, the NECESSITY of expensive vacuum systems is the biggest shortcoming of this technology in industrial applications besides the limitation to batch processes. Therefore, to overcome these disadvantages, considerable efforts are made in developing alternative techniques. Atmospheric pressure plasmas are one of the most promising methods to deposit polymer films in a more flexible, reliable, less expensive and continuous way of treatment. In the last two decades, a lot of effort has been put into the development of plasma polymerisation at elevated pressure. Chapter 5 attempts to review this research and its applications in a broad perspective.
Coatings and films mechanical properties are highly affected by their microstructure, and their adhesion to substrates, which sustains their mechanical integrity, and consequently improves their properties. Interfaces with high adhesion are also known to ensure prolonged coatings lifetime. Research on interface microstructure and bond form plays a very important role on coatings and films.
In Chapter 6, interfaces between chromium oxide coating deposited by reactive radio frequency (RF) magnetron sputtering technique, chromium interlayer and steel substrate are examined with scanning electron microscopy (SEM), high resolution electron microscopy
(HREM) and atom force microscopy (AFM) focusing on the interfacial structure properties affecting the adhesion performance and surface roughness. This examination revealed the presence of several Cr–Fe phases, which may ensure good adhesion of the interlayer to the underlying steel. Furthermore, amorphous chromium and chromium oxide layers about 100 nm thick were detected at each interface, which may have some effect on corrosion resistance and growth of columnar coating microstructure. The amorphous interfacial layer detected may give novel thought when deposited thick film but small size column grains.
The deposition of both metallic and dielectric inorganic thin films on polymeric substrates is of great interest for several industrial and research applications. The growth of metallic coatings on polymers is of raising usage in order to impart specific functionalities, such as electrical, aesthetic and chemical-resistance properties, to polymeric substrates. Some examples are the substitution of chromium electroplating processes on plastics by PVD deposition in several industrial fields and the use of aluminum or silver coatings for the fabrication of hybrid fabrics. Dielectric thin films are also commonly grown on polymeric materials for several aims, including the protection of polymeric substrates from scratch, the attribution of barrier coatings to food packaging films, the incorporation of new functionalities to artificial fabrics, and the increase of biocompatibility of some kind of polymeric dental materials and prostheses.
Unfortunately, the growth of thin films on polymeric substrates suffers of several constraints, due to the peculiar properties of polymers, such as the low heat resistance, the high elasticity and the low hardness. These limitations lead to the necessity of very low processing temperatures (often as low as room temperature) in order to avoid substrate damage, and the deposition of films of very low thicknesses, in order to reduce the interface stress. Plasma-assisted PVD and CVD techniques are suitable to satisfy these requirements, since they allow very low deposition temperatures, they are suitable for the deposition of composite materials, and provide a very good control on a wide range of process parameters.
Chapter 7 deals with an experimental study of the interaction between the surface of different polymeric substrates, such as ABS, polyester, polyamide and some dental resins, with metal and dielectric coatings, such as Cr, Al, a-SiOx, grown by RF sputtering or
PECVD. Different types of surface modifications, such as plasmaassisted surface activation and deposition of interlayers, were also applied to some of the polymeric substrates in order to study their effect on the growth process of the inorganic coatings.
Several characterization techniques were used in order to analyze the materials involved in the study. The polymeric substrates and the inorganic coatings were characterized against their surface morphology by means of high resolution mechanical profilometry, optical microscopy and field emission scanning electron microscopy (FESEM), while some of the film chemical characteristics were analyzed by Fourier transform infrared spectroscopy (FTIR). Some chemical resistance tests were also performed to investigate some properties of the polymer-dielectric multilayer structures.
Chapter 8 will review the research that has been done in the last decade using ultrasonic waves for coating surfaces. Sonochemistry is a field of research in which chemical reactions occur due to a collapse of an acoustic bubble. The review will present examples limited to coating nanoparticles on ceramic bodies and polymeric surfaces. However, the same technique works also on metallic, glass, and textile surfaces. The excellent adherence of the nanoparticles to the substrate is reflected, for example, in the lack of bleaching of the nanoparticles from the polymeric substrate when deposited by the sonochemical process.
Sonochemistry is a research field where waves in the frequency range of 20 kHz - 1 MHz are the driving force for the chemical reactions. The reaction is dependent on the development of an acoustic bubble in the solution. Extreme conditions (temperature >5000 K, pressure >1000 atm and cooling rates >1011 K/sec) are developed when this bubble collapses, thus
causing the chemical reactions to occur.
The current review will introduce to the reader what kind of surfaces serve as the substrates for the coating. It will present the variety of nanoparticles that have been anchored sonochemically to the surface, and finally it will explain the role of the ultrasonic waves in depositing nanoparticles onto solid surfaces. The review will compare the deposition of newly formed nanoparticles with that of nanoparticles purchased from a commercial source.
The first chapter of this review will introduce the reader to the field of sonochemistry. The current review is a continuation of a series of previous reviews published by our group. These reviews introduced the sonochemical technique as a new means for the fabrication of nanomaterials [1], for the use of ultrasonic waves for the doping of nanoparticles into ceramic and polymer bodies [2], and for the microspherization of proteins by a sonochemical process [3]. Other review articles on similar topics have also been published [4-6]. However, no review on using the sonochemical technique for coating surfaces was found in our literature search.
In our literature search we will scan for papers published until May 2008. We will try to avoid duplication and the review will not include examples presented in previous reviews.
As presented in Chapter 9, yet with the generation of large quantities of thermoplastics, the use of the thermal spray method is a logical and efficient means of recycling thermoplastics, thereby reducing the accumulation of polymer residues. Poly (ethylene terephthalate), PET, has excellent mechanical and chemical properties, and is a potential corrosion barrier since it presents low permeability to gases and solvents. Solutions of polymer recycling using the post-consumer PET to produce polymeric and composite coatings on steels in order to improve the tribological and chemical properties of steels are reported. Thermal sprayed and re-fused PET coatings, blend coatings of PET and the copolymer of ethylene and methacrylic acid, EMAA, and PET-based composite coatings were produced. Quenched PET blends with 80% PET and 20% EMAA and quenched PET coatings showed corrosion resistance in a salt spray chamber, small friction coefficient, and adhesion, which are necessary for the application of polymeric films as protective coatings against corrosion and wear. Peeling and swelling of the thermally sprayed PET coatings did not occur in the immersion tests in gasoline, diesel oil, and alcohol for a period of 60 days. The higher corrosion resistance in H2SO4 solution was observed for the composite PET
coatings with 0.1% of glass powder and flakes, and zinc powder.
The aim of Chapter 10 is to discuss the problematic of coatings of carbon nanotubes with thin and ultrathin layers with insulating properties.
Editors: M. Rizzo and G. Bruno, pp. 1-44 © 2009 Nova Science Publishers, Inc.
Chapter 1
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Anna Göransson Westerlund
1 Dept of Biomaterials, Institute of Surgical Science, Sahlgrenska Academy at Göteborg University, SwedenDept of Orthodontics, Institute of Odontology, Sahlgrenska Academy at Göteborg University, Sweden
Abstract
Materials that have the ability to bond to living tissue are defined as “bioactive” and the first possibly bioactive material Bio-glass was described in the 1970s by Hench and co-workers. Furthermore, Jarcho and co-workers were the first to present indications of a direct bone bonding to hydroxyapatite (HA). The mechanism proposed was ion exchange resulting in an apatite layer requested not only by the bone cells but also because proteins that serve as growth factors preferentially adsorb to this layer. The “bioactive” properties of these materials were based on morphological observations of the tissue coalescence by TEM and apatite formation in vitro and in vivo. Poor mechanical properties of these materials make them unsuitable for load-bearing, clinical applications. Therefore, experiments were made to coat titanium surfaces with calcium phosphates by plasma spraying technique. The surfaces indeed showed rapid tissue response initially, but in later stages biodegradation and delaminating of the thick coating was frequently observed. Additionally, the line-of sight problem made the technique unsuitable to use for the coating of complex shapes.
To avoid these problems, alternative techniques have been used to make commercially pure (CP) titanium “bioactive”. This article reviews recent research on “bioactive” titanium implant surfaces, focusing on five specific modifications:(I) etching with fluoride containing acids, (II) alkali-heat treatment, (III) anodization and (IV) ultra-thin coatings of calcium phosphates in sol-gels. Another possible approach to enhance the bone response is to (V) immobilize organic bio-molecules to the surface.
These five CP titanium surface modifications will be reviewed separately with a short background, suggested mechanism of action and performance in simulated body fluids (SBF),
1 E-mail address: [email protected]. Phone +46 31 786 2962 Fax +46 31 7732962 Correspondence to: Anna Westerlund PhD, Specialist Orthodontist, Department of Biomaterials, Göteborg University, Box 412, SE 405 30 Göteborg, Sweden.
in vitro and in vivo. Clinical evaluations will be discussed briefly. Each section is followed by
an appendix with a list of references of importance for the area of interest. The references are presented as short abstracts with similar information providing a quick overview and easy comparison of the studies.
1. Introduction
In the 1960s a new system for permanent anchorage of artificial teeth was discovered when the Brånemark group studied bone marrow cells in bone chambers.
The concept of “osseointegration” was defined in 1977 in conjunction with a 10-year follow-up study of titanium implants for edentulous jaws [1]. The initial definition “a material in intimate contact with living bone without intervening fibrous tissue” has during the years been redefined to adapt to current knowledge.
The Brånemark system was for a long time the gold standard based mainly on good clinical records [2]. However, in parallel implant parameters were evaluated for predicting good osseointegration and in the 1980s Albrektsson proposed six parameters as being important for the implant performance—material compatibility, implant design and surface quality, status of implant bed, surgical trauma at installation and prosthetic loading [3].
There are several methods by which the titanium surface quality can be modified [4]; physical turning, blasting), chemical (acid etching, alkali), electrochemical (electropolishing anodizing), deposition (plasma-spraying, sol-gel) and biochemical [simulated body fluids (SBF), proteins] methods. The different techniques will result in a surface quality with different topographical, chemical, physical and mechanical properties.
Since osseointegration depends on biomechanical bonding, i.e. ingrowth of bone into small irregularities of the implant, the topography and especially the roughness of the implants has been an area of interest and has been the subject of numerous research efforts.
Guidelines of how to perform and present the measurements of surface topography in a standardized way have been suggested by Wennerberg and Albrektsson [5].
Furthermore, based on experimental evidence from the mid 1990s a surface roughness of about 1.5 µm Sa (average deviation in height from a mean plane) has been defined as optimal
for osseointegration [6]. This is rougher than the original, turned Brånemark implant that demonstrated a surface roughness of about 0.5 µm.
Titanium surface roughness has also demonstrated to affect protein absorption [7], inflammatory cell [8-13] and bone cell [14-27] responses in vitro. Furthermore, there have been indications that surface orientation may be of importance [28, 29] for implant bone integration, however, not evaluated in a scientifically controlled manner.
Except for the concomitant change in chemical composition when changing the surface topography, attempts have been made to intentionally modify chemical composition to add a biochemical bonding to the biomechanical bonding.
The theoretical benefit of a chemical bond would be earlier attachment, since it is hypothesized to occur more rapidly than bony ingrowth.
Materials that have the ability to bond to living tissue are defined as “bioactive” and the first possibly bioactive material Bio-glass was described in the 1970s by Hench and co-workers [30]. Furthermore, Jarcho and co-co-workers were the first to present indications of a possible direct bone bonding to hydroxyapatite (HA)[31].
The mechanism proposed was ion exchange resulting in an apatite layer requested not only by the bone cells but also because proteins that serve as growth factors preferentially adsorb to this layer. The “bioactive” properties of these materials were based on morphological observations of the tissue coalescence by transmission electron microscopy (TEM), apatite formation in SBF in vitro and in vivo. However, it must be pointed out that bioactivity or chemical bonding are difficult to prove and that the presented evidence is of an indirect nature. Poor mechanical properties of these materials make them unsuitable for load-bearing, clinical applications. Therefore, experiments were made to coat titanium surfaces with calcium phosphates (CaP) by the plasma spraying technique. The surfaces indeed showed rapid tissue response initially, but in later stages biodegradation and delaminating of the thick coating was frequently observed [32]. Additionally, the line-of-sight problem made the technique unsuitable to use for the coating of complex shapes.
To avoid these problems, alternative techniques have been used to make commercially pure (cp) titanium possibly bioactive; I) etching with fluoride containing acids (fluoridated surfaces), II) alkali-heat treatment (alkali-heat treated surfaces), III) anodic oxidation with specific ions (anodized surfaces) and IV) sol-gel processing in calcium phosphate solutions (nano HA surfaces). Another possible approach to enhance the bone response is to V) immobilize organic bio-molecules to the surface (protein covalent immobilized surfaces). These five CP titanium surface modifications will be reviewed in the following sections with a short background, suggested mechanism of action and performance in SBF, in vitro and in vivo. Clinical evaluations will only be concluded briefly. Each section is followed by an appendix with a list of references of importance for the area of interest. The references are presented as short abstracts with similar information providing a quick overview and easy comparison of the studies.
2. State of the Art CP Titanium Implant Surfaces
2.1. Fluoridated CP Titanium Surfaces
Etching of titanium surfaces with different acids to modify surface roughness has been extensively studied during the last decades [33]. The idea of using fluoride-containing acids in low concentrations for the purpose of incorporating fluoride ions on titanium implants in small amounts was presented by Ellingsen and co-workers [34].
The action of the fluoride ion has mostly been evaluated in the area of caries research, where the beneficial effect because of its high attraction for calcium and phosphate is of great clinical importance, when the ion is brought in contact with the enamel. Fluoride has also specific attraction for skeletal tissues, e.g. trabecular bone density can be increased by the presence of fluoride ions during remodeling [35].
The proposed effects of the fluoride ion in bone are increased proliferation of bone cells by increasing intracellular levels of the ion, increased differentiation of mesenchymal cells into bone cells and stimulation of endogenous growth factor production [36].
Fluoridated titanium implant surfaces have been studied both in SBF [37], in vitro [37-42], in vivo [34, 38, 43] and clinically. OsseoSpeed™ (Astra Tech, Gothenburg, Sweden) is a commercially available dental implant system that has been evaluated in approximately 5-10 articles since the launch in 2004. The longest follow up period is 1 year [55]. The surface has
mainly been used in poor bone and in early loading situations where it in general has demonstrated good results. In addition there is an orthopedic hip implant available with some clinical documentation [44].
The possible bioactivity of titanium implant surfaces is based on its ability to give rise to early apatite formation in SBF [37], where the fluoride-modified surface demonstrates a Ca/P ratio of 2 [45]. When adding proteins to the SBF, the fluoride-modified surface demonstrate an increased apatite formation and protein adhesion compared to a blasted control [46]. Furthermore, in vivo studies have demonstrated increased bone response by means of increased bone implant contact [38, 43, 47], bone area [47] and stability [43, 48, 49] at shorter healing times than turned and blasted surfaces [50]. The mechanisms for the faster healing time of the implants are not fully understood. A possible explanation is that fluoride ion modification seems to augment the thrombogenic properties of titanium [51], another possibility is that fluoride modified surfaces demonstrate increased proliferation [38] and differentiation [38-41] of bone cells. However, results have shown decreased cell number [40], differentiation and protein production compared to blasted controls [37]. According to other studies, the amount of fluoride ions in the surface also seems to be of importance for the bone retention [52].
Appendix - Fluoridated CP Titanium Surfaces
SBF
Arvidsson et al -07 [45] compared four types of possibly bioactive surfaces; a blasted surface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation or hydroxyapatit coating, where the blasted surface served as control.
Surfaces were analyzed by weight, Profilometry, SEM/EDX and XPS after immersion in SBF for 1, 2, 3, 4 and 6 weeks. The results demonstrated that the Ca/P mean ratio of all the surfaces was approximately 1.5 after 1 week except for the fluoridated specimens that displayed mean ratio of approximately 2. All surfaces showed the presence of hydroxyapatite after 4 and 6 weeks of immersion, but a higher degree of crystallinity at 6 weeks. It was concluded that differences appeared at the early SBF immersion times of 1 and 2 weeks between controls and bioactive surface types, as well as between different bioactive surface types.
Franke-Stenport et al -08 [46] compared four types of possibly bioactive surfaces; a blasted surface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation or hydroxyapatit coating, where the blasted surface served as control. Surfaces were analyzed by Profilometry, SEM/EDX and XPS after immersion in SBF with 4.5 mg/ml albumin for 3 days, 1, 2, 3 and 4 weeks.
The results demonstrated that all the bioactive surfaces initiated an enhanced calcium phosphate (CaP) formation and a more rapid increase of protein content was present on the bioactive surfaces compared to the blasted control surface. It was concluded that this might be an advantage in vivo.
In Vitro
Eriksson et al -01 [42] compared smooth (polished) and rough (HF etched) surfaces with thick (annealed 700ºC) and thin (HNO3) oxide. The surfaces were characterized by SEM,
Optical Profilometry and AES. After exposure to whole blood for 8 minutes to 32 hours, immunofluorescence and chemiluminescence techniques were used for evaluation of cell adhesion, expression of adhesion receptors and the stimulated respiratory burst, respectively. PMN cells were the dominating cell on all surfaces followed by monocytes. While cells on rough surfaces demonstrated increased expression of adhesion receptors, earlier maximum respiratory burst occurred on the smooth surfaces. It was concluded that surface topography had greater impact on most cellular reactions, while oxide thickness often had a dampening effect.
Cooper at al -06 [38] compared grit-blasted (25 and 75 µm) titanium implants with and without fluoride ions (various fluoride concentrations). Cell attachment, proliferation and osteoblastic gene expression were measured by SEM, Tritiated thymidine incorporation and RT-PCR, respectively. There were no differences in human mesenchymal stem cell (hMSCs Osiris) attachment between the differently modified surfaces but cells on the fluoride ion modified implants demonstrated an increased proliferation and differentiation (BSP, BMP-2) compared to grit-blasted implants.
Masaki et al -05 [39] compared grit-blasted titanium implants with and without fluoride ions and grit-blasted etched surfaces (OsseoSpeed, TiOBlast, SLA-1 and SLA-2). Cell morphology, attachment, and osteoblastic gene expression were measured by SEM, Coulter counter (electrical conduction) and RT-PCR, respectively. There were no differences in mesenchymal pre-osteoblastic cell (HEPM 1486, ATCC) attachment, while cell morphology differed between the differently modified surfaces. Furthermore, cells demonstrated increased ALP gene expression on the SLA-2 surface, while cells on TiOBlast and OsseoSpeed demonstrated increased expression of Cbfa1/RUNX-2. It was concluded that implant surface properties might contribute to the regulation of osteoblastic differentiation by influencing the level of bone-related genes and transcription factors.
Isa et al -06 [40] compared blasted titanium implants with and without fluoride ions. Cell proliferation, alkaline phosphatase specific activity and gene expression were evaluated by Coulter counter, Spectrophotometry and RT-PCR, respectively. The number of cells human embryonic palatal mesenchymal (HEPM) were decreased on the fluoride surface compared to the blasted control. The gene expression was similar, except for Cbfa1, a key regulator for osteogenisis that was up regulated after 1 week on the fluoridated surface.
Stanford et al -06 [41] compared blasted titanium implants with and without fluoride ions. Platelet attachment and activation were evaluated by immunofluorescence technique, while human palatal mesenchymal (HEPM 1486, ATCC) morphology and gene expression were evaluated by SEM and RT-PCR, respectively. The number of attached platelets was decreased, while activation was increased on the fluoride surface compared to the blasted control. The gene expression was similar for the surfaces, except for Cbfa1 and bone sialoprotein that were increased on the fluoride modified surfaces.
Thor et al -07 [51] compared hydroxyapatite, machined, grit- blasted and fluoride ion modified grit- blasted surfaces. The trombogenic response, platelet activation, generation of thrombin-antithrombin complex where evaluated in a slide chamber model with blood, platelet-rich and platelet poor plasma after 60 min.
The results demonstrated that whole blood was necessary for sufficient thrombin generation and that the fluoride ion modified surface augmented the thrombogenic properties of titanium compared to the other surfaces.
Göransson et al -08 [37] compared four types of possibly bioactive surfaces; a blasted surface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation or hydroxyapatit coating, where a blasted surface served as control.
Surfaces were analyzed by Profilometry, SEM and XPS after immersion in SBF for 12, 24 and 72 hours.
Cells Primary (human mandibular osteoblast-like cells) were cultured on the various surfaces subjected to SBF for 72 h. Cellular attachment, differentiation (osteocalcin) and protein production (TGF-beta(1)) was evaluated after 3 h and 10 days respectively. The results demonstrated that the possibly bioactive surfaces gave rise to an earlier CaP formation than the blasted surface. Subsequent bone cell attachment was correlated to neither surface roughness nor the amount of formed CaP. In contrast, osteocalcin and TGF-beta(1) production were largely correlated to the amount of CaP formed on the surfaces.
In Vivo
Ellingsen et al -95 [48] compared turned titanium implants with and without fluoride ions (various fluoride concentrations NaF). The surfaces were characterized before installation and after push out test by SEM. It was demonstrated that fluoride modified surfaces had increased push out values in rabbit ulna after 4 and 8 weeks compared to untreated implant surfaces. Furthermore, on the fluoride modified surfaces fractures occurred in bone, while for the turned surface it occurred in the bone-implant interface.
Ellingsen et al -04 [43] compared blasted titanium implants with and without fluoride ions (HF). The surfaces were characterized by Optical Profilometry. It was demonstrated that fluoride modified surfaces had an increased amount of bone-implant contact in a rabbit model after 1 and 3 months compared to untreated implants. Additionally, the fluoride modified surfaces demonstrated increased RTQ and shear strengths between bone and implant after 3 months. It was concluded that fluoridated implants achieved greater bone integration after short healing time compared to blasted controls.
Cooper at al -06 [38] compared blasted surfaces with and without fluoride ions (HF). The surfaces were characterized by SEM. The results demonstrated improved bone formation by means of bone-implant contact in a rat tibia model for the fluoridated surface compared to the blasted surface after 3 weeks.
Berglundh et al -07 [50] compared implants with a blasted (TiOblast) and grit-blasted fluoride modified (OsseoSpeed) surfaces. Histological analyses were made in a dog model after 2 and 6 weeks. It was demonstrated that the amount of new bone formed in the voids after 2 weeks of healing was larger at fluoride-modified implants. Furthermore the amount of bone-to-implant contact that had been established after 2 weeks in the macro-threaded portion of the implant was significantly larger at the test implants than at the controls.
Abrahamsson et al -08 [47] compared implants with a blasted (TiOblast) and grit-blasted fluoride modified (OsseoSpeed) surfaces. Histological analyses were made in a dog model after 2 and 6 weeks. The histological analysis demonstrated a larger area of osseointegration and degree of bone-to-implant contact within the defect at fluoride-modified implants after 6 weeks of healing.
Lamolle et al -08 [52] compared fluoride ion modified titanium implants prepared in various HF concentrations (0,1, 0,01, 0,001 vol%). The surface topography and chemistry were characterized by AFM, SEM, and tof-SIMS respectively.
Bone response was evaluated in a rabbit model by using a pull out test method after 4 weeks. The group of 0,01% HF demonstrated the highest retention in bone. Furthermore, fluoride and hydride content in the surface as well as the surface skewness, kurtosis and core fluid retention were positively correlated to implant retention.
Monjo et al -08 compared grit-blasted and of fluoride-modified titanium implants. The attachment to cortical bone, [49] its association with gene expression of osteoblast (runx2, osteocalcin, collagen-I and IGF-I), osteoclast (TRAP, Hþ-ATPase and calcitonin receptor) and inflammation (TNF-a, IL-6 and IL-10) markers from peri-implant bone tissue and bone density were evaluated after 4 and 8 weeks by using pull-out test, real-time RT–PCR and micro -CT respectively. The results demonstrated lower LDH and TRAP mRNA activity for fluoride modified implants after 4 weeks, however no differences in pull-out force. After 8 weeks pull out force, bone density and gene expression for osteocalcin-, runX2-, collagen typ I were increased compared to grit-blasted surfaces.
Clinic
OsseoSpeed™ (Astra Tech, Gothenburg, Sweden) is a commercially available dental implant system that has been clinically evaluated in approximately 5-10 articles since their launch in 2004. The longest follow up period is 1 year [53]. The surface has mainly been used in poor bone and in early loading situations where it in general has demonstrated good results.
2.2. Alkali-Heat Treated CP Titanium Surfaces
The Kokubo group introduced the alkali-heat treated surface in the middle of the 1990s[54]. NaOH treatment results in a sodium titanate hydrogel, and the subsequent heat treatment at 600 degrees result in an amorphous sodium titanate surface layer [55, 56]. The possibly bioactivity of the surfaces are based on its ability to give rise to apatite formation in SBF and has been thoroughly investigated [37, 45, 54-63] also when adding proteins [46]. The apatite formation process on the surfaces has been carefully described [58, 59] and is attributed to Ti-OH groups exchanging sodium ions from the material and hydronium ions from the solution. Thereafter, adsorption of calcium ions from the fluid takes place to form calcium titanate. This calcium titanate surface then causes adsorption of phosphate as well as calcium ions to apatite nucleation layers. Once this layer is formed bone like apatite growth follows spontaneously.
Furthermore, studies have demonstrated an increased [64, 65] differentiation and decreased proliferation, differentiation and protein production of bone cells compared to untreated controls in vitro [37, 63].
In vivo studies have shown increased bone response by means of bone-implant contact, detachment load and tensile failure load compared to untreated surfaces [66-70]. However, the bonding strength seems to be time dependent with an initial high bonding strength and no further increase or difference compared to controls at later time points [67]. If the surface were pre-immersed in SBF, the apatite layer on the surface significantly increases the bone response resulting in increased failure loads [69, 70].
Increased bone response in vivo by means of enhanced bonding strength has additionally been demonstrated after sodium removal in hot water immersion or, as reported lately, by immersion in HCl [71].
If the bulk is a porous titanium material, the surface has been shown to induce ectopic bone formation in vivo in dog soft tissue model [72, 73].
This surface has so far not been applied to dental implants. However, clinical trials of seventy hip arthroplasty patients have been successfully concluded.
Appendix - Alkali-Heat Treated CP Titanium Surfaces
General
Kim et al -97 [55] evaluated bonding strength of the apatite layer formed in SBF on alkali treated implant surfaces with and without subsequent heat treatment (500, 600, 700, 800ºC) and compared it to bonding strengths of apatite formed on Bioglass 45S5-type glass, glass-ceramic AW and dense sintered HA. The results showed the highest bonding strengths of the apatite layer to the alkali treated titanium surfaces that were maximized after a subsequent heat treatment in 500-600ºC. It was concluded that bioactive titanium metal was useful as bone substitutes, even under load-bearing conditions.
Kim et al-99 [56] compared the structure of alkali-heat treated titanium surfaces (5M NaOH 60ºC 24h) prepared with various hydrothermal treatment (600 or 800ºC). Furthermore, the bonding strengths of the apatite layer formed on the various surfaces after soaking in SBF. The surfaces were characterized by SEM, AES,
Raman spectroscopy, TF-XRD, XPS and ICP. At 600ºC an amorphous sodium titanate layer with a smooth graded surface was formed, while at 800ºC a crystalline rutile sodium titanate with an intervening thick oxide was formed. The apatite layer prepared in 600ºC demonstrated the tightest bond to the surface.
SBF
Kim et al -96 [54] evaluated apatite formation in SBF (1-4w) on titanium and titanium alloy surfaces subjected to alkali (NaOH or KOH) and heat treatment (5º C/min to 400-800º C). The surfaces were characterized by SEM-EDX, TF-XRD, ICP and pH- metry. Apatite was formed on the SBF treated titanium and titanium alloy surfaces, though, not on cobalt chromium and stainless steal surfaces.
Kim et al -00 [54] subjected alkali-heat treated (5M NaOH 60ºC 24h+ 600ºC 1h) macroporous titanium (plasma-spraying method) to SBF. The surfaces were characterized by SEM-EDX and TF-XRD. The induction period for apatite formation was 3 days, which is comparable to bioactive glass-ceramics A/W. It was concluded that alkali-heat treatment is an effective method for preparation, irrespective of the surface macro-texture.
Wang et al -01 [62] compared heat-, H2O2-, and NaOH treated titanium surfaces. The
surfaces were characterized by SEM, FTIR and XRD. Dense oxide layer, titania gel and sodium titanate gel was formed on the surfaces, respectively. Some of the specimens were pre-immersed in distilled water up to 5 days before SBF. The discs were arranged with (contact surface) and without (open surface) contact with the bottom of the container. It was concluded that bioactivity of titania gel originated from the favorable structure of the gel itself because it formed apatite on open surface and after water immersion, while the sodium
titanate was dependent of ion release and therefore was unable to produce apatite on open surfaces and after water immersion (decreased ion concentration). Subsequent heat treatment decreased the apatite forming ability of the treated surfaces, but not the untreated titanium surfaces.
Takadama et al -01 [58] carefully described the apatite forming process on alkali-heat treated titanium surfaces by TF-XRD, ICP, pH-metry and XPS. It was stated that ”Bioactive titanium metal with a surface sodium titanate layer forms a bone-like apatite layer on its surface in the SBF by the following process; The Na+ ions were released from the surface sodium titanate via the exchange with H3O+ ions in the SBF to form OH groups. These
Ti-OH groups induce the apatite nucleation indirectly, by forming a calcium titanate. The initial formation of the calcium titanate may be attributable to the electrostatic reaction of the negatively charged Ti-OH groups and the positively charged calcium ions in the SBF.
Takadama et al -01 [59] further described the structure of apatite formation on alkali-heat treated titanium (5M NaOH 60ºC 24h + 600ºC 1h) subjected to SBF by TEM-EDX, ICP and pH-metry. The Ca/P ratios of the apatite were 1.4, 1.62 and 1.67 after 36, 48 and 72 hours in SBF, respectively.
Uchida et al -03 [61] compared apatite forming ability of Ti-OH with different structural arrangements in SBF after 14 days by SEM, TF-XRD and ICP. Gels with anatase and rutile structures induced more apatite on their surfaces compared to amorphous surfaces. It was concluded that crystalline planar arrangement in anatase structure was superior to rutile structure for apatite formation.
Lu et al -04 [57] subjected an alkali-heat treated titanium (10M NaOH 60ºC 24h + 600ºC 1h) surface to SBF for 1 month. The apatite formed was characterized by Profilometry, SEM, TEM-EDS and TF-XRD. The study showed that octacalcium phosphate (OCP), not apatite, was formed on the surface after immersion in SBF.
Arvidsson et al -07 [45] compared four types of possibly bioactive surfaces; a blasted surface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation or hydroxyapatit coating, where the blasted surface served as control.
Surfaces were analyzed by weight, Profilometry, SEM/EDX and XPS after immersion in SBF for 1, 2, 3, 4 and 6 weeks. The results demonstrated that the Ca/P mean ratio of all the surfaces was approximately 1.5 after 1 week except for the fluoridated specimens which displayed mean ratio of approximately 2. All surfaces showed the presence of hydroxyapatite after 4 and 6 weeks of immersion, but a higher degree of crystallinity at 6 weeks. It was concluded that differences appeared at the early SBF immersion times of 1 and 2 weeks between controls and bioactive surface types, as well as between different bioactive surface types.
Franke-Stenport et al -08 [46] compared four types of possibly bioactive surfaces; a blasted surface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation or hydroxyapatit coating, where the blasted surface served as control. Surfaces were analyzed by Profilometry, SEM/EDX and XPS after immersion in SBF with 4.5 mg/ml albumin for 3 days, 1, 2, 3 and 4 weeks.
The results demonstrated that all the bioactive surfaces initiated an enhanced calcium phosphate (CaP) formation and a more rapid increase of protein content was present on the bioactive surfaces compared to the blasted control surface. It was concluded that this might be an advantage in vivo.
In Vitro
Nishio et al -00 [65] compared titanium, alkali-heat treated titanium (5M NaOH 60ºC 24h + 600ºC 1h) and alkali-heat treated titanium subjected to SBF for 2 weeks. The surfaces were characterized by SEM, TF-XRD and XPS.
Cell number (Primary rat bone marrow cells), differentiation and gene expression (OC, OP, ON COL) were evaluated by DNA content, ALP activity and Northern blot, respectively. Results demonstrated that cell differentiation increased on the apatite prepared surfaces, while cell number was similar for the differently modified surfaces. It was concluded that apatite formed on the surfaces favored osteoblast differentiation and that alkali-heat treatment favored apatite formation.
Muramatsu et al -03 [74] compared thrombus resistance of alkali-heat treated titanium (5M NaOH 60ºC 24h + 600ºC 1h), alkali-water treated titanium (distilled water 40ºC 48h) and alkali-heat treated titanium subjected to SBF.
The surfaces were characterized by AFM, XRD and contact angle measurement. Platelet attachment and protein adsorption were evaluated and it was concluded that SBF treated alkali-heat treated titanium behaved thrombus resistant probably because heparin was preferentially adsorbed to its surface.
Chosa et al -04 [64] compared TCP, titanium and SBF treated (8 days) alkali-heat treated titanium (5M NaOH 60ºC 24h + 600ºC 1h). The surfaces were characterized by SEM, TF-XRD, FTIR and XPS.
Cell (Human osteoblast SaOS-2) differentiation-related gene expression (ALP, COL, OPN, BSP, OSC) was evaluated by RT-PCR after 1, 2, 3 and 4 weeks. The results indicated that the treated implants accelerated middle (OPN, BSP) and late (OSC) stage differentiation, while early differentiation was down-regulated (ALP, COL).
Maitz et al -05 [75] compared bioactivity of titanium following sodium plasma immersion, ion implantation and deposition (alkali) in SBF for 7 days. The surfaces were characterized by AES. In a parallel experiment, cell (rat bone marrow cells) viability, proliferation and differentiation was evaluated by LDH test, Alamar blue test and ALP activity, respectively. It was concluded that ion implantation and deposition could well substitute alkali treatment.
Göransson et al -08 [37] compared four types of possibly bioactive surfaces; a blasted surface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation or hydroxyapatit coating, where a blasted surface served as control.
Surfaces were analyzed by Profilometry, SEM and XPS after immersion in SBF for 12, 24 and 72 hours.
Cells Primary (human mandibular osteoblast-like cells) were cultured on the various surfaces subjected to SBF for 72 h. Cellular attachment, differentiation (osteocalcin) and protein production (TGF-beta(1)) was evaluated after 3 h and 10 days respectively. The results demonstrated that the possibly bioactive surfaces gave rise to an earlier CaP formation than the blasted surface. Subsequent bone cell attachment was correlated to neither surface roughness nor the amount of formed CaP. In contrast, osteocalcin and TGF-beta(1) production were largely correlated to the amount of CaP formed on the surfaces.
In Vivo
Yan et al -97 [70] compared titanium, alkali-heat treated titanium (5M NaOH 60ºC 24h + 600ºC 1h) and SBF treated (4 weeks) alkali-heat treated titanium implants. Tensile testing demonstrated that both treated surfaces showed significantly increased failure loads after 4, 8 and 16 weeks in the rabbit tibia compared to the control. Furthermore, both treated surfaces demonstrated direct bone contact with no intervening soft tissue capsule in a histological evaluation after 4 weeks, whereas untreated implants formed direct contact with bone only at 16 weeks.
Yan et al -97 [69] compared titanium and SBF (4weeks) treated alkali-heat treated (10M NaOH 60ºC 24h + 600ºC 1h) titanium implants. The surfaces were characterized by SEM-EPMA and TF-XRD. Tensile testing demonstrated that the treated surfaces showed significantly increased failure loads after 6, 10 and 25 weeks in the rabbit tibia compared to the control. Histologic examination demonstrated that the treated surfaces demonstrated more immediate bone contact compared to the control titanium surface at all evaluation times.
Nishiguchi et al -99 [68] compared titanium, alkali-treated titanium and alkali-heat treated titanium implants (5M NaOH 60ºC 24h + 600ºC 1h). The surfaces were characterized by SEM. Mechanical and histomorphometrical evaluations were performed after 8 and 16 weeks in the rabbit tibia. The alkali-heat treated surfaces demonstrated direct bone-implant contact after 8 weeks, while alkali treated implants demonstrated an intervening fibrous capsule. Additionally, the alkali-heat treated surfaces demonstrated significantly increased failure load after 8 and 16 weeks. It was concluded that heat treatment is essential for preparing a bioactive surface, even though the alkali surface had previously demonstrated apatite formation in SBF, since implants with gel surfaces are unstable and difficult to preserve and install.
Nishiguchi et al.-01 [76] compared macroporous titanium (plasma-spraying method), macroporous titanium coated with AW-glass ceramic and alkali-heat treated macroporous titanium (5M NaOH 60ºC 24h + 600ºC 1h).
Mechanical and histomorphometrical evaluations were performed after 4 and 12 weeks in dog femur. Bone-implant contact was significantly increased on alkali-heat treated implants at 4 and 12 weeks. Push out test revealed increased shear strengths for the alkali-heat treated surfaces compared to the other surfaces after 4 weeks. It was concluded that alkali-heat treated implants provided earlier stable fixation than control implants.
Nishiguchi et al -01 [67] compared titanium and titanium alloy implants with and without alkali-heat treatment (5M NaOH 60ºC 24h + 600ºC 1h).
Histomorphometric evaluations and push out tests were performed after 4 and 12 weeks in dog femur. Alkali-heat treated implants showed direct bone-implant contact; while alkali treated, implants demonstrated an intervening fibrous capsule. After 4 weeks, the heat-treated surfaces demonstrated increased push out shear strengths compared to untreated surfaces. However, after 12 weeks the untreated implants demonstrated a catch up compared to the treated implants.
Nishiguchi et al -03 [66] compared titanium and alkali-heat treated implants (5M NaOH 60ºC 24h + 600ºC 1h). Mechanical and histomorphometrical evaluations were performed after 3, 6 and 12 weeks in the rabbit femur. Alkali-heat treated implants demonstrated increased bone-implant contact and increased bonding strengths (pull out test) compared to untreated surfaces at all evaluation times.
Fujibayashi et al -01 [71] evaluated the effectiveness of sodium removal from alkali-heat treated titanium surfaces, where CP titanium were used as controls. The in vivo detaching failure load was evaluated after 4, 8, 16 and 24 weeks in rabbit tibia. Thereafter, the surfaces were evaluated by SEM. It was concluded that sodium removal accelerated bone bonding because of the anatase structure. However, the adhesive strengths decreased for the sodium free surfaces.
Fujibayashi et al. -04 [72] compared ectopic bone formation of porous (plasma-spraying) and mesh titanium surfaces with and without alkali-heat treatment (sodium removed). Evaluations were performed in dog muscle after 3 and 12 months. In a parallel experiment, the surfaces were immersed in SBF for 7 days. The surfaces were evaluated by SEM and micro-CT/3D reconstruction. The porous alkali-heat treated surfaces demonstrated osteoinductive ability after 12 months.
Takemoto et al -05 [60] compared macroporous titanium (plasma-spraying method) with and without alkali-heat treatment (5M NaOH 60ºC 24h + 600ºC 1h). The surfaces were characterized by micro-CT/3D reconstruction and SEM. Mechanical tests by means of compression strengths, four-point binding strengths and compressive fatigue strengths were performed of the surface. In vitro bioactivity was evaluated in SBF for 3-7 days and in vivo histomorphometric evaluation was performed after 2, 4, 8 and 16 weeks in rabbit femur. Apatite formation in vitro was apparent after 3 days on the alkali-heat treated surfaces, while no apatite could be detected after 7 days on the control surfaces. Bone-implant contact and bone-area in growth were significantly higher on alkali-heat treated implants at all evaluation times. In addition, the surface had mechanical properties sufficient for clinical use in load bearing conditions
Takemoto et al. -06 [73] compared ectopic bone formation of alkali-heat treated porous titanium, alkali-heat treated (sodium removed by hot water) porous, and alkali-heat-treated (sodium removed by HCl and hot water) titanium surfaces. The surfaces were characterized by SEM-EDX and TF-XRD and evaluated in dog muscle after 3, 6 and 12 months. In a parallel experiment, the surfaces were immersed in SBF for 1, 3 and 7 days. The porous sodium free alkali-heat treated surfaces demonstrated osteo inductive ability after 3 months, while apatite formation could be seen on all surfaces after 1 day.
Isaac et al -08 [63] compared titanium and alkali-heat treated implants (5M NaOH 60ºC 24h + 600ºC 1h).
SBF and a bone explant model (immunohistochemical staining, alkaline phosphatase histoenzymatic localization and SEM after) were used to evaluate the surfaces after 3 and 15 days respectively. Results demonstrated bone-like apatite layer on the modified surface in simulated body fluids. Furthermore, that cells from frontal and parietal bones from 21-day-old rat fetuses can migrate from the explants and subsequently differentiate to form a mineralized nodular structure. The cells expressed alkaline phosphatase, bone sialoprotein, osteocalcin and the transcription factor, Runx2.
Clinic
2.3. Anodized CP Titanium Surfaces
Electrochemical modification of titanium surfaces related to implant research has been performed since the 1970s.
The process called anodic spark discharge (ASD) was proposed by Kurze and co-workers and was further described by Ishizawa and co-workers [77-79].
Anodized titanium surfaces have been extensively evaluated in vitro [37, 72, 80-86], in vivo [77, 78, 87-113]. There are some commercially available implant systems as well, with TiUnite™ (Nobel Biocare, Gothenburg, Sweden) so far dominating the market. Anodized TiUnite™ implants have been clinically evaluated in approximately 50 articles since their launch in 2001 where the longest follow up period is 5 years [130]. This implant system is however not claimed to be bioactive, instead the good results is explained by the topography.
Since the oxide properties can be controlled by anodic forming voltage, current density, electrolytes, electrolyte concentrations and temperature, agitation speed etc., the resulting surfaces present heterogeneous characteristics by means of surface chemistry, oxide thickness, morphology, surface roughness, pore configurations (pore size, porosity, pore density and crystal structure) [114, 115].
In vitro studies have demonstrated various results with either increased [72, 80] or decreased [37, 81, 82, 85] bone cell attachment, increased [82, 85, 86] or decreased [37, 81] differentiation and decreased protein production [37] compared to control surfaces. In vitro inflammatory response show increased cell adherence despite similar cytokine production and differentiation [83].
In general but with some exceptions [88, 90, 97, 100], the anodized surfaces demonstrate increased bone response compared to control titanium surfaces in vivo [87, 93, 98, 105, 106, 108, 112, 116]. This is attributed to the changes of topography, but also the oxide thickness, pore configurations and crystal structure of the oxide layer, where an oxide thickness of > 600 nm has demonstrated to be favorable [105, 106, 108]. When incorporating certain ions i.e. calcium [104] and magnesium [101-103, 109, 111, 117], the increased bone response has been attributed to chemistry and a potential biochemical bond. Indications of biochemical bonding (bioactivity) has been proposed on the basis of ultrastructural analysis of interfacial fracture (scanning electron microscopy-SEM), ion movement/exchange at the interfacial tissue (X-ray microanalysis-EDS), speed and strength of implant integration to bone (removal torque-RTQ) [101, 102, 111, 117] and increased bone implant contact (BiC) [113]. Calcium [118, 119] and magnesium [37, 45] incorporated anodized surfaces have additionally increased apatite formation in SBF [120] and when adding proteins to the SBF the apatite formation and protein content increased on the possibly bioactive titanium surfaces compared to blasted control [46]. Furthermore, an additional hot water treatment could contribute to increased apatite formation, enhanced bonding strengths between apatite layer and metal [121], increased differentiation and protein production in vitro [84].
Recently Biolin AB (Gothenburg, Sweden) launched OsPol™, an implant system with a calcium reinforced possibly bioactive surface.
Appendix – Anodized CP Titanium Surfaces
General
Ishizawa et al. -95 [79] compared anodized titanium surfaces prepared with different anodic voltage 150-400 V (50mA/cm2), electrolytes and concentrations. Spark discharge occurred at
200V. The surfaces were characterized by SEM, EDX and XRD. Calcium acetate monohydrate and ß-glycerophosphate turned out to be suitable electrolytes, since the resulting Ca/P had a ratio equivalent to HA. HA crystals were precipitated by an additional heat treatment.
Hall and Lausmaa -00 [122] introduced an anodized surface that later resulted in the commercially available TiUnite. The surfaces were characterized by Optical Interferometry, SEM, AES and XRD. The surface had a roughness of 1,2 µm (Ra), an oxide thickness of 1-2 µm at the cervical part and 7-10 µm at the apical part, a pore size in the range of 1-2 µm. The surface contained 15% Ti, 55% O, 20% C, 5% P, 1% S and 1% Si. Furthermore, it was demonstrated that the oxide layer strongly adhered to the underlying metal.
Sul et al. -01 [114] compared the oxide growth behavior on titanium surfaces in acid and alkaline electrolytes with different electrolyte concentrations, temperature (14-42°C), anodic forming voltage (20-130V), current forming density (5-40 mA/cm2), and agitation speed (250-800 rpm). The formed oxide surfaces were thoroughly characterized by AES and a Spectrophotometry system. It was concluded that colors were useful for thickness determination of titanium oxide and that each electrolyte presented an individual growth constant nm/V. Furthermore, a general trend that increased electrolyte concentration and temperature decreased anodic forming voltage, anodic forming rate and the current efficiency, while an increased current density and surface area ratio anode/cathode increased anodic forming voltage, anodic forming rate and current efficiency. The effects of electrolyte concentration, temperature and agitation speed were explained by the electrical double layer.
Sul et al. -02 [115] prepared anodic oxides by galvanostatic mode in acetic acid up to dielectric break down and spark formation (100-400V). The surfaces were characterized by Profilometry, AES, SEM, XPS, TF- XRD and Raman Spectroscopy. The results demonstrated a well characterized surface regarding surface roughness, oxide thickness, pore-size and distribution, chemical composition and crystal structure.
Crawford et al. -07 [123] prepared titanium surfaces with nanotubes by anodic oxidation using NaF electrolyte. The surface was characterized by field-emission scanning electron microscope (FE-SEM) and mechanical properties of the coatings were probed by nanoindentation.
Results demonstrated that increased anodization time had no effect on tube diameter or tube wall thickness. However, coating thickness increased with time up to 2 h of anodization, at which point an equilibrium thickness was established. Progressively higher values of elastic modulus were obtained for thinner films.
SBF
Yang et al. -04 [119] compared anodized titanium surfaces prepared in an electrolyte (H2SO4)
with different concentrations (0,5-3M) anodic forming voltage (90-180V), with and without subsequent heat treatment (600ºC 1h). The surfaces were characterized with SEM and TF-XRD. A simulated body fluid was used to evaluate the CaP nucleation capacity of the
surfaces after 3 and 6 days. Apatite forming ability could be attained at 3 and 6 days by anodic oxidation > 90V and < 90V co-joined with heat treatment. Both the anatase and rutile was effective for apatite formation. No apatite formed on the surfaces without spark discharge (<90V) and heat treatment indicating that a certain thickness of the titanium oxide was required for apatite formation.
Vanzilotta et al. -06 [118] compared CaP nucleation capacity in SBF of three surface modifications; etching and etching followed by either anodization or heat treatment. The surfaces were characterized by Profilometry, SEM-EDX and AAS, XPS before and after SBF soaking, respectively. The Ca ion concentration decreased in the SBF solution for all surfaces from day 1 to day 7. The heat treated and anodized surfaces demonstrated increased CaP nucleation capacity compared to the etched surfaces, while no differences were detected between the anodized and heat treated surfaces.
Arvidsson et al. -07 [45] compared four types of possibly bioactive surfaces; a blasted surface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation or hydroxyapatit coating, where the blasted surface served as control.
Surfaces were analyzed by weight, Profilometry, SEM/EDX and XPS after immersion in SBF for 1, 2, 3, 4 and 6 weeks. The results demonstrated that the Ca/P mean ratio of all the surfaces was approximately 1.5 after 1 week except for the fluoridated specimens which displayed mean ratio of approximately 2. All surfaces showed the presence of hydroxyapatite after 4 and 6 weeks of immersion, but a higher degree of crystallinity at 6 weeks. It was concluded that differences appeared at the early SBF immersion times of 1 and 2 weeks between controls and bioactive surface types, as well as between different bioactive surface types.
Cui X et al. -08 [120] prepared titanium surfaces by anodic oxidation in four different electrolytes: sulfuric acid, acetic acid, phosphoric acid and sodium sulfate solutions with different voltages (1min at room temperature). The surfaces were immersed in SBF for 1, 3 and 7 days and analyzed in TF-XRD and; FE-SEM. Results demonstrated that the anodic films consisted of rutile and/or anatase phases with porous structures on titanium metal after anodizing in H(2)SO(4) and Na(2)SO(4) electrolytes, while amorphous titania films were produced after anodizing in CH(3)COOH and H(3)PO(4) electrolytes. Moreover, titanium metal with the anatase and/or rutile crystal structure films showed excellent apatite-forming ability and produced a compact apatite layer covering all the surface of titanium after soaking in SBF for 7d, but titanium metal with amorphous titania layers was not able to induce apatite formation.
Cui X et al. -08 [121] prepared titanium surfaces by anodic oxidation in acetic acid followed by hot water and heat treatments to transform titania layers from an amorphous structure into a crystalline structure.
The apatite-forming ability of titania layers in simulated body fluid was investigated by XRD and SEM. Results indicated that hot water and/or heat treatment transformed the crystal structure of titania layers from an amorphous structure into anatase, or a mixture of anatase and rutile. It was suggested that abundance of Ti-OH groups formed by hot water treatment could contribute to apatite formation on the surface of titanium metals, and subsequent heat treatment would enhance the bond strength between the apatite layers and the titanium substrates. It was concluded that bioactive titanium metals could be prepared via anodic oxidation and subsequent hot water and heat treatment made them suitable for applications under load-bearing conditions.
