surface coatings

SURFACE COATINGS

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SURFACE COATINGS

MARIO RIZZO AND

GIUSEPPE BRUNO EDITORS

Nova Science Publishers, Inc. New York

Copyright 2009 by Nova Science Publishers, Inc. All rights reserved. No part of this book may be reproduced, stored in a retrieval system or transmitted in any form or by any means: electronic, electrostatic, magnetic, tape, mechanical photocopying, recording or otherwise without the written permission of the Publisher. For permission to use material from this book please contact us: Telephone 631-231-7269; Fax 631-231-8175 Web Site: http://www.novapublishers.com

NOTICE TO THE READER The Publisher has taken reasonable care in the preparation of this book, but makes no expressed or implied warranty of any kind and assumes no responsibility for any errors or omissions. No liability is assumed for incidental or consequential damages in connection with or arising out of information contained in this book. The Publisher shall not be liable for any special, consequential, or exemplary damages resulting, in whole or in part, from the readers use of, or reliance upon, this material. Any parts of this book based on government reports are so indicated and copyright is claimed for those parts to the extent applicable to compilations of such works. Independent verification should be sought for any data, advice or recommendations contained in this book. In addition, no responsibility is assumed by the publisher for any injury and/or damage to persons or property arising from any methods, products, instructions, ideas or otherwise contained in this publication. This publication is designed to provide accurate and authoritative information with regard to the subject matter covered herein. It is sold with the clear understanding that the Publisher is not engaged in rendering legal or any other professional services. If legal or any other expert assistance is required, the services of a competent person should be sought. FROM A DECLARATION OF PARTICIPANTS JOINTLY ADOPTED BY A COMMITTEE OF THE AMERICAN BAR ASSOCIATION AND A COMMITTEE OF PUBLISHERS. 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

Published by Nova Science Publishers, Inc. New York

CONTENTS

Preface vii

Chapter 1 State of the Art Bioactive Titanium Implant Surfaces 1 Anna Gransson 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. Perrone and 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

V.F.C. Lins , J.R.T. Branco and C.C. Berndt

Contents vi

Chapter 10 Coating of Carbon Nanotubes with Insulating Thin Layers 259 Martin Pumera

Index 265

PREFACE 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

Mario Rizzo and Giuseppe viii

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

Preface ix

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

Mario Rizzo and Giuseppe x

(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 CrFe 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.

Preface xi

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.

In: Surface Coatings ISBN: 978-1-60741-193-2Editors: M. Rizzo and G. Bruno, pp. 1-44 2009 Nova Science Publishers, Inc.

Chapter 1

STATE OF THE ART BIOACTIVETITANIUM IMPLANT SURFACES

Anna Gransson Westerlund1Dept of Biomaterials, Institute of Surgical Science,

Sahlgrenska Academy at Gteborg University, SwedenDept of Orthodontics, Institute of Odontology,

Sahlgrenska Academy at Gteborg University, Sweden

Abstract

Materials that have the ability to bond to living tissue are defined as bioactive and thefirst 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 directbone bonding to hydroxyapatite (HA). The mechanism proposed was ion exchange resultingin an apatite layer requested not only by the bone cells but also because proteins that serve asgrowth factors preferentially adsorb to this layer. The bioactive properties of these materialswere based on morphological observations of the tissue coalescence by TEM and apatiteformation in vitro and in vivo. Poor mechanical properties of these materials make themunsuitable for load-bearing, clinical applications. Therefore, experiments were made to coattitanium surfaces with calcium phosphates by plasma spraying technique. The surfaces indeedshowed rapid tissue response initially, but in later stages biodegradation and delaminating ofthe thick coating was frequently observed. Additionally, the line-of sight problem made thetechnique unsuitable to use for the coating of complex shapes.

To avoid these problems, alternative techniques have been used to make commerciallypure (CP) titanium bioactive. This article reviews recent research on bioactive titaniumimplant surfaces, focusing on five specific modifications:(I) etching with fluoride containingacids, (II) alkali-heat treatment, (III) anodization and (IV) ultra-thin coatings of calciumphosphates 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 shortbackground, 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, GteborgUniversity, Box 412, SE 405 30 Gteborg, Sweden.

Anna Gransson Westerlund2

in vitro and in vivo. Clinical evaluations will be discussed briefly. Each section is followed byan appendix with a list of references of importance for the area of interest. The references arepresented as short abstracts with similar information providing a quick overview and easycomparison of the studies.

1. Introduction

In the 1960s a new system for permanent anchorage of artificial teeth was discovered whenthe Brnemark group studied bone marrow cells in bone chambers.

The concept of osseointegration was defined in 1977 in conjunction with a 10-yearfollow-up study of titanium implants for edentulous jaws [1]. The initial definition a materialin intimate contact with living bone without intervening fibrous tissue has during the yearsbeen redefined to adapt to current knowledge.

The Brnemark system was for a long time the gold standard based mainly on goodclinical records [2]. However, in parallel implant parameters were evaluated for predictinggood osseointegration and in the 1980s Albrektsson proposed six parameters as beingimportant for the implant performancematerial compatibility, implant design and surfacequality, 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 (electropolishinganodizing), deposition (plasma-spraying, sol-gel) and biochemical [simulated body fluids(SBF), proteins] methods. The different techniques will result in a surface quality withdifferent topographical, chemical, physical and mechanical properties.

Since osseointegration depends on biomechanical bonding, i.e. ingrowth of bone intosmall irregularities of the implant, the topography and especially the roughness of theimplants 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 astandardized way have been suggested by Wennerberg and Albrektsson [5].

Furthermore, based on experimental evidence from the mid 1990s a surface roughness ofabout 1.5 m Sa (average deviation in height from a mean plane) has been defined as optimalfor osseointegration [6]. This is rougher than the original, turned Brnemark implant thatdemonstrated a surface roughness of about 0.5 m.

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 havebeen indications that surface orientation may be of importance [28, 29] for implant boneintegration, however, not evaluated in a scientifically controlled manner.

Except for the concomitant change in chemical composition when changing the surfacetopography, attempts have been made to intentionally modify chemical composition to add abiochemical bonding to the biomechanical bonding.

The theoretical benefit of a chemical bond would be earlier attachment, since it ishypothesized to occur more rapidly than bony ingrowth.

Materials that have the ability to bond to living tissue are defined as bioactive and thefirst possibly bioactive material Bio-glass was described in the 1970s by Hench and co-workers [30]. Furthermore, Jarcho and co-workers were the first to present indications of apossible direct bone bonding to hydroxyapatite (HA)[31].

State of the Art Bioactive Titanium Implant Surfaces 3

The mechanism proposed was ion exchange resulting in an apatite layer requested notonly by the bone cells but also because proteins that serve as growth factors preferentiallyadsorb to this layer. The bioactive properties of these materials were based onmorphological 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 thatbioactivity or chemical bonding are difficult to prove and that the presented evidence is of anindirect nature. Poor mechanical properties of these materials make them unsuitable for load-bearing, clinical applications. Therefore, experiments were made to coat titanium surfaceswith calcium phosphates (CaP) by the plasma spraying technique. The surfaces indeedshowed rapid tissue response initially, but in later stages biodegradation and delaminating ofthe thick coating was frequently observed [32]. Additionally, the line-of-sight problem madethe technique unsuitable to use for the coating of complex shapes.

To avoid these problems, alternative techniques have been used to make commerciallypure (cp) titanium possibly bioactive; I) etching with fluoride containing acids (fluoridatedsurfaces), II) alkali-heat treatment (alkali-heat treated surfaces), III) anodic oxidation withspecific 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 witha short background, suggested mechanism of action and performance in SBF, in vitro and invivo. Clinical evaluations will only be concluded briefly. Each section is followed by anappendix with a list of references of importance for the area of interest. The references arepresented as short abstracts with similar information providing a quick overview and easycomparison 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 beenextensively studied during the last decades [33]. The idea of using fluoride-containing acidsin low concentrations for the purpose of incorporating fluoride ions on titanium implants insmall 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 greatclinical importance, when the ion is brought in contact with the enamel. Fluoride has alsospecific attraction for skeletal tissues, e.g. trabecular bone density can be increased by thepresence of fluoride ions during remodeling [35].

The proposed effects of the fluoride ion in bone are increased proliferation of bone cellsby increasing intracellular levels of the ion, increased differentiation of mesenchymal cellsinto 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 acommercially available dental implant system that has been evaluated in approximately 5-10articles 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 hasdemonstrated good results. In addition there is an orthopedic hip implant available with someclinical documentation [44].

The possible bioactivity of titanium implant surfaces is based on its ability to give rise toearly apatite formation in SBF [37], where the fluoride-modified surface demonstrates a Ca/Pratio of 2 [45]. When adding proteins to the SBF, the fluoride-modified surface demonstratean increased apatite formation and protein adhesion compared to a blasted control [46].Furthermore, in vivo studies have demonstrated increased bone response by means ofincreased bone implant contact [38, 43, 47], bone area [47] and stability [43, 48, 49] at shorterhealing times than turned and blasted surfaces [50]. The mechanisms for the faster healingtime of the implants are not fully understood. A possible explanation is that fluoride ionmodification seems to augment the thrombogenic properties of titanium [51], anotherpossibility is that fluoride modified surfaces demonstrate increased proliferation [38] anddifferentiation [38-41] of bone cells. However, results have shown decreased cell number[40], differentiation and protein production compared to blasted controls [37]. According toother studies, the amount of fluoride ions in the surface also seems to be of importance for thebone retention [52].

Appendix - Fluoridated CP Titanium Surfaces

SBFArvidsson et al -07 [45] compared four types of possibly bioactive surfaces; a blasted surfaceprepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation orhydroxyapatit coating, where the blasted surface served as control.

Surfaces were analyzed by weight, Profilometry, SEM/EDX and XPS after immersion inSBF for 1, 2, 3, 4 and 6 weeks. The results demonstrated that the Ca/P mean ratio of all thesurfaces was approximately 1.5 after 1 week except for the fluoridated specimens thatdisplayed mean ratio of approximately 2. All surfaces showed the presence of hydroxyapatiteafter 4 and 6 weeks of immersion, but a higher degree of crystallinity at 6 weeks. It wasconcluded that differences appeared at the early SBF immersion times of 1 and 2 weeksbetween controls and bioactive surface types, as well as between different bioactive surfacetypes.

Franke-Stenport et al -08 [46] compared four types of possibly bioactive surfaces; ablasted surface prepared by alkali heat treatment, anodization (Mg ions incorporated),fluoridation or hydroxyapatit coating, where the blasted surface served as control. Surfaceswere analyzed by Profilometry, SEM/EDX and XPS after immersion in SBF with 4.5 mg/mlalbumin for 3 days, 1, 2, 3 and 4 weeks.The results demonstrated that all the bioactive surfaces initiated an enhanced calciumphosphate (CaP) formation and a more rapid increase of protein content was present on thebioactive surfaces compared to the blasted control surface. It was concluded that this might bean advantage in vivo.

In VitroEriksson et al -01 [42] compared smooth (polished) and rough (HF etched) surfaces withthick (annealed 700C) 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 celladhesion, expression of adhesion receptors and the stimulated respiratory burst, respectively.PMN cells were the dominating cell on all surfaces followed by monocytes. While cells onrough surfaces demonstrated increased expression of adhesion receptors, earlier maximumrespiratory burst occurred on the smooth surfaces. It was concluded that surface topographyhad greater impact on most cellular reactions, while oxide thickness often had a dampeningeffect.

Cooper at al -06 [38] compared grit-blasted (25 and 75 m) titanium implants with andwithout fluoride ions (various fluoride concentrations). Cell attachment, proliferation andosteoblastic gene expression were measured by SEM, Tritiated thymidine incorporation andRT-PCR, respectively. There were no differences in human mesenchymal stem cell (hMSCsOsiris) attachment between the differently modified surfaces but cells on the fluoride ionmodified 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 fluorideions and grit-blasted etched surfaces (OsseoSpeed, TiOBlast, SLA-1 and SLA-2). Cellmorphology, attachment, and osteoblastic gene expression were measured by SEM, Coultercounter (electrical conduction) and RT-PCR, respectively. There were no differences inmesenchymal pre-osteoblastic cell (HEPM 1486, ATCC) attachment, while cell morphologydiffered between the differently modified surfaces. Furthermore, cells demonstrated increasedALP gene expression on the SLA-2 surface, while cells on TiOBlast and OsseoSpeeddemonstrated increased expression of Cbfa1/RUNX-2. It was concluded that implant surfaceproperties might contribute to the regulation of osteoblastic differentiation by influencing thelevel of bone-related genes and transcription factors.

Isa et al -06 [40] compared blasted titanium implants with and without fluoride ions. Cellproliferation, alkaline phosphatase specific activity and gene expression were evaluated byCoulter counter, Spectrophotometry and RT-PCR, respectively. The number of cells humanembryonic palatal mesenchymal (HEPM) were decreased on the fluoride surface compared tothe blasted control. The gene expression was similar, except for Cbfa1, a key regulator forosteogenisis that was up regulated after 1 week on the fluoridated surface.

Stanford et al -06 [41] compared blasted titanium implants with and without fluorideions. Platelet attachment and activation were evaluated by immunofluorescence technique,while human palatal mesenchymal (HEPM 1486, ATCC) morphology and gene expressionwere evaluated by SEM and RT-PCR, respectively. The number of attached platelets wasdecreased, while activation was increased on the fluoride surface compared to the blastedcontrol. The gene expression was similar for the surfaces, except for Cbfa1 and bonesialoprotein that were increased on the fluoride modified surfaces.

Thor et al -07 [51] compared hydroxyapatite, machined, grit- blasted and fluoride ionmodified grit- blasted surfaces. The trombogenic response, platelet activation, generation ofthrombin-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 thrombingeneration and that the fluoride ion modified surface augmented the thrombogenic propertiesof titanium compared to the other surfaces.


Gransson et al -08 [37] compared four types of possibly bioactive surfaces; a blastedsurface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation orhydroxyapatit 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 varioussurfaces subjected to SBF for 72 h. Cellular attachment, differentiation (osteocalcin) andprotein production (TGF-beta(1)) was evaluated after 3 h and 10 days respectively. Theresults demonstrated that the possibly bioactive surfaces gave rise to an earlier CaP formationthan the blasted surface. Subsequent bone cell attachment was correlated to neither surfaceroughness 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 VivoEllingsen et al -95 [48] compared turned titanium implants with and without fluoride ions(various fluoride concentrations NaF). The surfaces were characterized before installation andafter push out test by SEM. It was demonstrated that fluoride modified surfaces had increasedpush 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 theturned surface it occurred in the bone-implant interface.

Ellingsen et al -04 [43] compared blasted titanium implants with and without fluorideions (HF). The surfaces were characterized by Optical Profilometry. It was demonstrated thatfluoride modified surfaces had an increased amount of bone-implant contact in a rabbit modelafter 1 and 3 months compared to untreated implants. Additionally, the fluoride modifiedsurfaces demonstrated increased RTQ and shear strengths between bone and implant after 3months. It was concluded that fluoridated implants achieved greater bone integration aftershort healing time compared to blasted controls.

Cooper at al -06 [38] compared blasted surfaces with and without fluoride ions (HF). Thesurfaces were characterized by SEM. The results demonstrated improved bone formation bymeans of bone-implant contact in a rat tibia model for the fluoridated surface compared to theblasted surface after 3 weeks.

Berglundh et al -07 [50] compared implants with a grit-blasted (TiOblast) and grit-blasted fluoride modified (OsseoSpeed) surfaces. Histological analyses were made in a dogmodel after 2 and 6 weeks. It was demonstrated that the amount of new bone formed in thevoids after 2 weeks of healing was larger at fluoride-modified implants. Furthermore theamount 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 thecontrols.

Abrahamsson et al -08 [47] compared implants with a grit-blasted (TiOblast) and grit-blasted fluoride modified (OsseoSpeed) surfaces. Histological analyses were made in a dogmodel after 2 and 6 weeks. The histological analysis demonstrated a larger area ofosseointegration and degree of bone-to-implant contact within the defect at fluoride-modifiedimplants after 6 weeks of healing.


Lamolle et al -08 [52] compared fluoride ion modified titanium implants prepared invarious HF concentrations (0,1, 0,01, 0,001 vol%). The surface topography and chemistry werecharacterized 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 andhydride content in the surface as well as the surface skewness, kurtosis and core fluidretention were positively correlated to implant retention.

Monjo et al -08 compared grit-blasted and of fluoride-modified titanium implants. Theattachment to cortical bone, [49] its association with gene expression of osteoblast (runx2,osteocalcin, collagen-I and IGF-I), osteoclast (TRAP, H-ATPase and calcitonin receptor)and inflammation (TNF-a, IL-6 and IL-10) markers from peri-implant bone tissue and bonedensity were evaluated after 4 and 8 weeks by using pull-out test, real-time RTPCR andmicro -CT respectively. The results demonstrated lower LDH and TRAP mRNA activity forfluoride modified implants after 4 weeks, however no differences in pull-out force. After 8weeks pull out force, bone density and gene expression for osteocalcin-, runX2-, collagen typI were increased compared to grit-blasted surfaces.

ClinicOsseoSpeed (Astra Tech, Gothenburg, Sweden) is a commercially available dental implantsystem that has been clinically evaluated in approximately 5-10 articles since their launch in2004. The longest follow up period is 1 year [53]. The surface has mainly been used in poorbone 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 at600 degrees result in an amorphous sodium titanate surface layer [55, 56]. The possiblybioactivity of the surfaces are based on its ability to give rise to apatite formation in SBF andhas been thoroughly investigated [37, 45, 54-63] also when adding proteins [46]. The apatiteformation 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 toapatite nucleation layers. Once this layer is formed bone like apatite growth followsspontaneously.

Furthermore, studies have demonstrated an increased [64, 65] differentiation anddecreased proliferation, differentiation and protein production of bone cells compared tountreated 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 nofurther increase or difference compared to controls at later time points [67]. If the surfacewere pre-immersed in SBF, the apatite layer on the surface significantly increases the boneresponse resulting in increased failure loads [69, 70].


Increased bone response in vivo by means of enhanced bonding strength has additionallybeen demonstrated after sodium removal in hot water immersion or, as reported lately, byimmersion in HCl [71].

If the bulk is a porous titanium material, the surface has been shown to induce ectopicbone formation in vivo in dog soft tissue model [72, 73].

This surface has so far not been applied to dental implants. However, clinical trials ofseventy hip arthroplasty patients have been successfully concluded.

Appendix - Alkali-Heat Treated CP Titanium Surfaces

GeneralKim et al -97 [55] evaluated bonding strength of the apatite layer formed in SBF on alkalitreated implant surfaces with and without subsequent heat treatment (500, 600, 700, 800C)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 theapatite layer to the alkali treated titanium surfaces that were maximized after a subsequentheat treatment in 500-600C. It was concluded that bioactive titanium metal was useful asbone substitutes, even under load-bearing conditions.

Kim et al-99 [56] compared the structure of alkali-heat treated titanium surfaces (5MNaOH 60C 24h) prepared with various hydrothermal treatment (600 or 800C). 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 600C an amorphous sodium titanate layerwith a smooth graded surface was formed, while at 800C a crystalline rutile sodium titanatewith an intervening thick oxide was formed. The apatite layer prepared in 600Cdemonstrated the tightest bond to the surface.

SBFKim et al -96 [54] evaluated apatite formation in SBF (1-4w) on titanium and titanium alloysurfaces subjected to alkali (NaOH or KOH) and heat treatment (5 C/min to 400-800 C).The surfaces were characterized by SEM-EDX, TF-XRD, ICP and pH- metry. Apatite wasformed on the SBF treated titanium and titanium alloy surfaces, though, not on cobaltchromium and stainless steal surfaces.

Kim et al -00 [54] subjected alkali-heat treated (5M NaOH 60C 24h+ 600C 1h)macroporous titanium (plasma-spraying method) to SBF. The surfaces were characterized bySEM-EDX and TF-XRD. The induction period for apatite formation was 3 days, which iscomparable to bioactive glass-ceramics A/W. It was concluded that alkali-heat treatment is aneffective method for preparation, irrespective of the surface macro-texture.

Wang et al -01 [62] compared heat-, H2O2-, and NaOH treated titanium surfaces. Thesurfaces were characterized by SEM, FTIR and XRD. Dense oxide layer, titania gel andsodium titanate gel was formed on the surfaces, respectively. Some of the specimens werepre-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 wasconcluded that bioactivity of titania gel originated from the favorable structure of the gel itselfbecause 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 opensurfaces and after water immersion (decreased ion concentration). Subsequent heat treatmentdecreased the apatite forming ability of the treated surfaces, but not the untreated titaniumsurfaces.

Takadama et al -01 [58] carefully described the apatite forming process on alkali-heattreated titanium surfaces by TF-XRD, ICP, pH-metry and XPS. It was stated that Bioactivetitanium metal with a surface sodium titanate layer forms a bone-like apatite layer on itssurface in the SBF by the following process; The Na+ ions were released from the surfacesodium titanate via the exchange with H3O+ ions in the SBF to form Ti-OH groups. These Ti-OH groups induce the apatite nucleation indirectly, by forming a calcium titanate. The initialformation of the calcium titanate may be attributable to the electrostatic reaction of thenegatively 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-heattreated titanium (5M NaOH 60C 24h + 600C 1h) subjected to SBF by TEM-EDX, ICP andpH-metry. The Ca/P ratios of the apatite were 1.4, 1.62 and 1.67 after 36, 48 and 72 hours inSBF, respectively.

Uchida et al -03 [61] compared apatite forming ability of Ti-OH with different structuralarrangements in SBF after 14 days by SEM, TF-XRD and ICP. Gels with anatase and rutilestructures induced more apatite on their surfaces compared to amorphous surfaces. It wasconcluded that crystalline planar arrangement in anatase structure was superior to rutilestructure for apatite formation.

Lu et al -04 [57] subjected an alkali-heat treated titanium (10M NaOH 60C 24h + 600C1h) 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 blastedsurface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation orhydroxyapatit coating, where the blasted surface served as control.Surfaces were analyzed by weight, Profilometry, SEM/EDX and XPS after immersion in SBFfor 1, 2, 3, 4 and 6 weeks. The results demonstrated that the Ca/P mean ratio of all thesurfaces was approximately 1.5 after 1 week except for the fluoridated specimens whichdisplayed mean ratio of approximately 2. All surfaces showed the presence of hydroxyapatiteafter 4 and 6 weeks of immersion, but a higher degree of crystallinity at 6 weeks. It wasconcluded that differences appeared at the early SBF immersion times of 1 and 2 weeksbetween controls and bioactive surface types, as well as between different bioactive surfacetypes.

Franke-Stenport et al -08 [46] compared four types of possibly bioactive surfaces; ablasted surface prepared by alkali heat treatment, anodization (Mg ions incorporated),fluoridation or hydroxyapatit coating, where the blasted surface served as control. Surfaceswere analyzed by Profilometry, SEM/EDX and XPS after immersion in SBF with 4.5 mg/mlalbumin for 3 days, 1, 2, 3 and 4 weeks.The results demonstrated that all the bioactive surfaces initiated an enhanced calciumphosphate (CaP) formation and a more rapid increase of protein content was present on thebioactive surfaces compared to the blasted control surface. It was concluded that this might bean advantage in vivo.


In VitroNishio et al -00 [65] compared titanium, alkali-heat treated titanium (5M NaOH 60C 24h +600C 1h) and alkali-heat treated titanium subjected to SBF for 2 weeks. The surfaces werecharacterized 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, whilecell number was similar for the differently modified surfaces. It was concluded that apatiteformed on the surfaces favored osteoblast differentiation and that alkali-heat treatmentfavored apatite formation.

Muramatsu et al -03 [74] compared thrombus resistance of alkali-heat treated titanium(5M NaOH 60C 24h + 600C 1h), alkali-water treated titanium (distilled water 40C 48h)and alkali-heat treated titanium subjected to SBF.The surfaces were characterized by AFM, XRD and contact angle measurement. Plateletattachment and protein adsorption were evaluated and it was concluded that SBF treatedalkali-heat treated titanium behaved thrombus resistant probably because heparin waspreferentially adsorbed to its surface.

Chosa et al -04 [64] compared TCP, titanium and SBF treated (8 days) alkali-heat treatedtitanium (5M NaOH 60C 24h + 600C 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 thetreated implants accelerated middle (OPN, BSP) and late (OSC) stage differentiation, whileearly differentiation was down-regulated (ALP, COL).

Maitz et al -05 [75] compared bioactivity of titanium following sodium plasmaimmersion, ion implantation and deposition (alkali) in SBF for 7 days. The surfaces werecharacterized by AES. In a parallel experiment, cell (rat bone marrow cells) viability,proliferation and differentiation was evaluated by LDH test, Alamar blue test and ALPactivity, respectively. It was concluded that ion implantation and deposition could wellsubstitute alkali treatment.

Gransson et al -08 [37] compared four types of possibly bioactive surfaces; a blastedsurface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation orhydroxyapatit coating, where a blasted surface served as control.Surfaces were analyzed by Profilometry, SEM and XPS after immersion in SBF for 12, 24and 72 hours.Cells Primary (human mandibular osteoblast-like cells) were cultured on the various surfacessubjected to SBF for 72 h. Cellular attachment, differentiation (osteocalcin) and proteinproduction (TGF-beta(1)) was evaluated after 3 h and 10 days respectively. The resultsdemonstrated that the possibly bioactive surfaces gave rise to an earlier CaP formation thanthe blasted surface. Subsequent bone cell attachment was correlated to neither surfaceroughness 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 VivoYan et al -97 [70] compared titanium, alkali-heat treated titanium (5M NaOH 60C 24h +600C 1h) and SBF treated (4 weeks) alkali-heat treated titanium implants. Tensile testingdemonstrated that both treated surfaces showed significantly increased failure loads after 4, 8and 16 weeks in the rabbit tibia compared to the control. Furthermore, both treated surfacesdemonstrated direct bone contact with no intervening soft tissue capsule in a histologicalevaluation after 4 weeks, whereas untreated implants formed direct contact with bone only at16 weeks.

Yan et al -97 [69] compared titanium and SBF (4weeks) treated alkali-heat treated (10MNaOH 60C 24h + 600C 1h) titanium implants. The surfaces were characterized by SEM-EPMA and TF-XRD. Tensile testing demonstrated that the treated surfaces showedsignificantly increased failure loads after 6, 10 and 25 weeks in the rabbit tibia compared tothe control. Histologic examination demonstrated that the treated surfaces demonstrated moreimmediate bone contact compared to the control titanium surface at all evaluation times.

Nishiguchi et al -99 [68] compared titanium, alkali-treated titanium and alkali-heattreated titanium implants (5M NaOH 60C 24h + 600C 1h). The surfaces were characterizedby SEM. Mechanical and histomorphometrical evaluations were performed after 8 and 16weeks in the rabbit tibia. The alkali-heat treated surfaces demonstrated direct bone-implantcontact after 8 weeks, while alkali treated implants demonstrated an intervening fibrouscapsule. Additionally, the alkali-heat treated surfaces demonstrated significantly increasedfailure load after 8 and 16 weeks. It was concluded that heat treatment is essential forpreparing a bioactive surface, even though the alkali surface had previously demonstratedapatite formation in SBF, since implants with gel surfaces are unstable and difficult topreserve and install.

Nishiguchi et al.-01 [76] compared macroporous titanium (plasma-spraying method),macroporous titanium coated with AW-glass ceramic and alkali-heat treated macroporoustitanium (5M NaOH 60C 24h + 600C 1h).Mechanical and histomorphometrical evaluations were performed after 4 and 12 weeks in dogfemur. Bone-implant contact was significantly increased on alkali-heat treated implants at 4and 12 weeks. Push out test revealed increased shear strengths for the alkali-heat treatedsurfaces compared to the other surfaces after 4 weeks. It was concluded that alkali-heattreated implants provided earlier stable fixation than control implants.

Nishiguchi et al -01 [67] compared titanium and titanium alloy implants with and withoutalkali-heat treatment (5M NaOH 60C 24h + 600C 1h).Histomorphometric evaluations and push out tests were performed after 4 and 12 weeks indog femur. Alkali-heat treated implants showed direct bone-implant contact; while alkalitreated, implants demonstrated an intervening fibrous capsule. After 4 weeks, the heat-treatedsurfaces demonstrated increased push out shear strengths compared to untreated surfaces.However, after 12 weeks the untreated implants demonstrated a catch up compared to thetreated implants.

Nishiguchi et al -03 [66] compared titanium and alkali-heat treated implants (5M NaOH60C 24h + 600C 1h). Mechanical and histomorphometrical evaluations were performedafter 3, 6 and 12 weeks in the rabbit femur. Alkali-heat treated implants demonstratedincreased bone-implant contact and increased bonding strengths (pull out test) compared tountreated surfaces at all evaluation times.


Fujibayashi et al -01 [71] evaluated the effectiveness of sodium removal from alkali-heattreated titanium surfaces, where CP titanium were used as controls. The in vivo detachingfailure load was evaluated after 4, 8, 16 and 24 weeks in rabbit tibia. Thereafter, the surfaceswere evaluated by SEM. It was concluded that sodium removal accelerated bone bondingbecause of the anatase structure. However, the adhesive strengths decreased for the sodiumfree 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 andmicro-CT/3D reconstruction. The porous alkali-heat treated surfaces demonstratedosteoinductive ability after 12 months.

Takemoto et al -05 [60] compared macroporous titanium (plasma-spraying method) withand without alkali-heat treatment (5M NaOH 60C 24h + 600C 1h). The surfaces werecharacterized by micro-CT/3D reconstruction and SEM. Mechanical tests by means ofcompression strengths, four-point binding strengths and compressive fatigue strengths wereperformed of the surface. In vitro bioactivity was evaluated in SBF for 3-7 days and in vivohistomorphometric 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, whileno apatite could be detected after 7 days on the control surfaces. Bone-implant contact andbone-area in growth were significantly higher on alkali-heat treated implants at all evaluationtimes. In addition, the surface had mechanical properties sufficient for clinical use in loadbearing conditions

Takemoto et al. -06 [73] compared ectopic bone formation of alkali-heat treated poroustitanium, 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 characterizedby SEM-EDX and TF-XRD and evaluated in dog muscle after 3, 6 and 12 months. In aparallel experiment, the surfaces were immersed in SBF for 1, 3 and 7 days. The poroussodium 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 60C24h + 600C 1h).SBF and a bone explant model (immunohistochemical staining, alkaline phosphatasehistoenzymatic localization and SEM after) were used to evaluate the surfaces after 3 and 15days respectively. Results demonstrated bone-like apatite layer on the modified surface insimulated body fluids. Furthermore, that cells from frontal and parietal bones from 21-day-oldrat fetuses can migrate from the explants and subsequently differentiate to form a mineralizednodular structure. The cells expressed alkaline phosphatase, bone sialoprotein, osteocalcinand the transcription factor, Runx2.

ClinicSo far, there are no commercially alkali-heat treated dental implant systems available.


2.3. Anodized CP Titanium Surfaces

Electrochemical modification of titanium surfaces related to implant research has beenperformed since the 1970s.

The process called anodic spark discharge (ASD) was proposed by Kurze and co-workersand was further described by Ishizawa and co-workers [77-79].

Anodized titanium surfaces have been extensively evaluated in vitro [37, 72, 80-86], invivo [77, 78, 87-113]. There are some commercially available implant systems as well, withTiUnite (Nobel Biocare, Gothenburg, Sweden) so far dominating the market. AnodizedTiUnite implants have been clinically evaluated in approximately 50 articles since theirlaunch in 2001 where the longest follow up period is 5 years [130]. This implant system ishowever 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 resultingsurfaces present heterogeneous characteristics by means of surface chemistry, oxidethickness, morphology, surface roughness, pore configurations (pore size, porosity, poredensity and crystal structure) [114, 115].

In vitro studies have demonstrated various results with either increased [72, 80] ordecreased [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 vitroinflammatory response show increased cell adherence despite similar cytokine production anddifferentiation [83].

In general but with some exceptions [88, 90, 97, 100], the anodized surfaces demonstrateincreased 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 > 600nm 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 hasbeen attributed to chemistry and a potential biochemical bond. Indications of biochemicalbonding (bioactivity) has been proposed on the basis of ultrastructural analysis of interfacialfracture (scanning electron microscopy-SEM), ion movement/exchange at the interfacialtissue (X-ray microanalysis-EDS), speed and strength of implant integration to bone (removaltorque-RTQ) [101, 102, 111, 117] and increased bone implant contact (BiC) [113]. Calcium[118, 119] and magnesium [37, 45] incorporated anodized surfaces have additionallyincreased apatite formation in SBF [120] and when adding proteins to the SBF the apatiteformation and protein content increased on the possibly bioactive titanium surfaces comparedto blasted control [46]. Furthermore, an additional hot water treatment could contribute toincreased 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 acalcium reinforced possibly bioactive surface.


Appendix Anodized CP Titanium Surfaces

GeneralIshizawa et al. -95 [79] compared anodized titanium surfaces prepared with different anodicvoltage 150-400 V (50mA/cm2), electrolytes and concentrations. Spark discharge occurred at200V. The surfaces were characterized by SEM, EDX and XRD. Calcium acetatemonohydrate and -glycerophosphate turned out to be suitable electrolytes, since the resultingCa/P had a ratio equivalent to HA. HA crystals were precipitated by an additional heattreatment.

Hall and Lausmaa -00 [122] introduced an anodized surface that later resulted in thecommercially available TiUnite. The surfaces were characterized by Optical Interferometry,SEM, AES and XRD. The surface had a roughness of 1,2 m (Ra), an oxide thickness of 1-2m at the cervical part and 7-10 m at the apical part, a pore size in the range of 1-2 m. Thesurface contained 15% Ti, 55% O, 20% C, 5% P, 1% S and 1% Si. Furthermore, it wasdemonstrated 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 andalkaline electrolytes with different electrolyte concentrations, temperature (14-42C), anodicforming 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 aSpectrophotometry system. It was concluded that colors were useful for thicknessdetermination of titanium oxide and that each electrolyte presented an individual growthconstant nm/V. Furthermore, a general trend that increased electrolyte concentration andtemperature decreased anodic forming voltage, anodic forming rate and the current efficiency,while an increased current density and surface area ratio anode/cathode increased anodicforming voltage, anodic forming rate and current efficiency. The effects of electrolyteconcentration, 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 todielectric break down and spark formation (100-400V). The surfaces were characterized byProfilometry, AES, SEM, XPS, TF- XRD and Raman Spectroscopy. The resultsdemonstrated 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 oxidationusing NaF electrolyte. The surface was characterized by field-emission scanning electronmicroscope (FE-SEM) and mechanical properties of the coatings were probed bynanoindentation.Results demonstrated that increased anodization time had no effect on tube diameter or tubewall thickness. However, coating thickness increased with time up to 2 h of anodization, atwhich point an equilibrium thickness was established. Progressively higher values of elasticmodulus were obtained for thinner films.

SBFYang 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 withoutsubsequent heat treatment (600C 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 byanodic oxidation > 90V and < 90V co-joined with heat treatment. Both the anatase and rutilewas effective for apatite formation. No apatite formed on the surfaces without spark discharge(


Franke-Stenport et al. -08 [46] compared four types of possibly bioactive surfaces; ablasted surface prepared by alkali heat treatment, anodization (Mg ions incorporated),fluoridation or hydroxyapatit coating, where the blasted surface served as control. Surfaceswere analyzed by Profilometry, SEM/EDX and XPS after immersion in SBF with 4.5 mg/mlalbumin for 3 days, 1, 2, 3 and 4 weeks.The results demonstrated that all the bioactive surfaces initiated an enhanced calciumphosphate (CaP) formation and a more rapid increase of protein content was present on thebioactive surfaces compared to the blasted control surface. It was concluded that this might bean advantage in vivo.

In VitroTakabe et al.-00 [80] compared anodized-heat treated titanium surfaces (CA/-GP) andtitanium controls. The surfaces were characterized by Profilometry and Contact AngleMeasurements. Initial cell (Rat Bone Marrow Stromal Cells) attachment, morphology andcytoskeleton were evaluated after 30, 60 and 120 minutes by Coulter counter (electricalconduction), SEM and CLSM, respectively. The anodized-heat treated surfaces were rougher,more hydrophilic and demonstrated increased cell attachment after 60 and 120 minutescompared to the controls. The cells showed a flattened surface with irregular edges andextended filipodium-like processes intimately adapted to the crystals of the surface. The actinfilament was arranged parallel to the long axis of the cells and localized in the periphery.

Rodriguez et al. -03 [84] compared 3 groups of anodized titanium surfaces (CA/-GP);with and without additionally heat treatment for 2 and 4 hours. The surfaces werecharacterized by SEM, XRD, Profilometry and EPMA. Cells (human embryonic palatalmesenchymal cells- HEPM) were used to evaluate differentiation (ALP, osteocalcin) andprotein production over an 8-day period (0, 4, 8 days). Results demonstrated thatmineralization, differentiation and protein production increased after hydrothermal-treatment.

Li et al. -04 [82] compared anodized-heat treated titanium surfaces (CA/-GP) withvarious anodic forming voltage (190-600V). The surfaces were characterized by OpticalInterferometry, SEM-EDS, XTEM and XRD. Cell adhesion (MG-63 and HumanOsteosarcoma Cells/HOS) after 3 days, proliferation after 7 days and differentiation after 10days were evaluated by SEM, Hemocytometry and Spectrophtometry (ALP activity),respectively. The surface roughness, oxide thickness and concentration of Ca and P ionsincreased with increasing voltage. In addition, there was a phase change from anatase torutile. As a result the differentiation increased (>300 V), while the proliferation decreased(>190V). Preliminary results in vivo indicated increased removal torque values after 4 weeksfor the anodized surfaces (270V).

Zhu et al. -04 [81] compared anodized titanium surfaces prepared in different electrolytes(CA/-GP and H2PO4) and anodic forming voltage (140-350V). The surfaces werecharacterized by SEM, Profilometry, XPS and Contact Angle Measurements. Cellsattachment and spread (SaOS-2) after 1 and 2 hours, proliferation and differentiation after 1, 2and 4 days were evaluated by immunohistochemistry (vinculin, phalloidin), Hemocytometryand Spectrophotometry (ALP activity), respectively. Cell attachment and proliferationincreased with increasing voltage, while differentiation was similar or decreased. The cells onthe anodized surfaces demonstrated a polygonal growth and lamellipodia, reflecting highmotility, while the control demonstrated thick stress fibers and intense focal contacts.


Kim et al. -04 [124] compared turned and anodized titanium surfaces (CA/-GP, 270V).The surfaces were characterized by Optical Interferometry, SEM and XRD. Cell (MG-63)adhesion and gene expression were evaluated after 12, 24 and 48 hours by Spectrophotometry(Crystal Violet) and Microarray technique, respectively. The anodized surfaces were rougherand displayed increased attachment of MG-63 osteoblast like cells without significantlyaffecting the gene expression.

Kim et al. -06 [86] prepared titanium surfaces by anodic oxidation (CA/-GP. Thesurfaces were characterized by scanning electron microscopy, X-ray diffraction, and electronprobe microanalysis.

Osteoblast were used to evaluate the cell differentiation. Results demonstrated thatosteoblast differentiation (ALP), increased on the anodized surfaces. It was concluded that thephenotypic expression of osteoblast was enhanced by the presence of Ca phosphate andhigher roughness on anodized surfaces.

Vanzilotta et al. -06 [118] compared CaP nucleation capacity in SBF of three surfacemodifications; etching and etching followed by either anodization or heat treatment. Thesurfaces were characterized by Profilometry, SEM-EDX and AAS, XPS before and after SBFsoaking, respectively. The Ca ion concentration decreased in the SBF solution for all surfacesfrom day 1 to day 7. The heat treated and anodized surfaces demonstrated increased CaPnucleation capacity compared to the etched surfaces, while no differences were detectedbetween the anodized and heat treated surfaces.

Gransson et al. -06 [83] compared titanium surfaces prepared by a turned, blasted,anodized and anodized surface with Mg ions incorporated.

The surfaces were characterized by Optical Interferometry. The inflammarory responsewas evaluated by cellnumber (human mononuclear cells), viability (LDH),cytokineproduction (TNF-, IL-10) and differentiation were analyzed after 24h and 72 hours.

The result demonstrated that the anodized surfaces with and without Mg ionsincorporated increased cell adherence, despite the anodized Mg ion incorporated surfacehaving a smoother character, however no differences in cytokine production anddifferentiation between the surfaces.

For all surfaces the viability was good at both 24 and 72 hours and cytokine IL-10production remained over time while TNF- and cellnumber decreased.

Das et al. -07 [125] prepared titanium surfaces with; nanotubes by anodic oxidation usingdifferent electrolyte solutions, H[3)PO(4), HF and H(2)SO(4). The surface were characterizedby field-emission scanning electron microscope (FE-SEM) fitted with an energy dispersivespectroscopy (EDS), Glancing angle X-ray diffraction (GAXRD), profilometry and contactangle measurement.

Bone cells (osteoblastic precursor cell line -OPC1) were used to study cell adhesion(Vinculin, - confocal scanning laser microscopy) and proliferation (MTT assay) anddifferentiation (alkaline phosphatase) after 3, 5 and 11 days.

The surfaces were additionally immersed in simulated body fluids for 3, 7, 14, and 21Results demonstrated distinctive cell-to-cell attachment in the HF anodized surface, cellularadherence with extracellular matrix extensions in between the cells was noticed for samplesanodized with H(3)PO(4) electrolyte. The TiO(2) layer grown in H(2)SO(4) electrolyte didnot show significant cell growth on the surface, and some cell death was also noticed. Celladhesions and differentiation were more anodized surfaces.


Das et al. -08 [126] prepared titanium surfaces with; nanotubes by anodic oxidation citricacid, sodium fluoride, and sulfuric acid as electrolyte solution with and without an additionalanodic oxidation in silver nitrate solutions

The surface were characterized by field-emission scanning electron microscope (FE-SEM) fitted with an energy dispersive spectroscopy (EDS), Glancing angle X-ray diffraction(GAXRD), profilometry and contact angle measurement. Bone cells (osteoblastic precursorcell line -OPC1) were used to study cell adhesion and proliferation (MTT assay) after 5 and11 days.

The antibacterial effect was studied using Pseudomonas aeruginosa. Resultsdemonstrated that silver-treated titania nanotube surfaces provided antibacterial properties toprevent implants against postoperative infections without interference to the attachment andproliferation of bone tissue on titanium

Bose et al. -08 [127] prepared titanium surfaces with; nanotubes by anodic oxidation(citric acid, sodium fluoride, and sulfuric acid as electrolyte solution with and without anadditional anodic oxidation in silver nitrate solutions), Tricalcium phosphate (TCP) coatingsby LENS processing with different laser power and with powder having particle sizeranging from 45 to 150 m. A titanium surface served as control.

The surfaces were characterized by (FE-SEM) (EDS), (GAXRD), profilometry andcontact angle measurement.

Bone cells (OPC1) were used to study cell adhesion (Vinculin) and proliferation (MTTassay) and differentiation (alkaline phosphatase) after 3, 7 and 11 days. Additionallymicrohardness of the coating was analysed.

Results demonstrated that anodic oxidation and laser processed TCP-coated Ti surfaceshowed enhanced cell adhesion, higher proliferation and early differentiation in comparisonto the control-Ti surface.

The TCP coating hardness was significantly increased from the base metal and furtherincreased as the volume fraction of TCP increased in the coating

Das et al. -08 [128] prepared nanotubes on titanium surfaces by anodic oxidation usingcitric acid, sodium fluoride, and sulfuric acid as electrolyte solution (20 V for 4 h). Thesurface were characterized by field-emission scanning electron microscope (FE-SEM) fittedwith an energy dispersive spectroscopy (EDS), Glancing angle X-ray diffraction (GAXRD),profilometry and contact angle measurement.

Bone cells (osteoblastic precursor cell line -OPC1) were used to study cell adhesion(Vinculin, - confocal scanning laser microscopy) and proliferation (MTT assay) anddifferentiation (alkaline phosphatase) after 3, 7 and 11 days.

The surfaces were additionally immersed in simulated body fluids for 3, 7, 14, and 21days.

The anodized nanoporous sample surfaces demonstrated increased cell adhesion,proliferation and differentiation. Apatite layer formation was non-uniform on the nanotubesurface even after 21 days in SBF.

De Angelis et al. -08 [85] compared three surfaces; titanium surfaces prepared withanodic sparc oxidation in (Ca/P Ca electrolytes), alkali etched titanium and nontreatedtitanium.

Cell (SaOS-2) attachment, morphology, viability, proliferation, metabolic activity,differentiation and mineralization were analysed by SEM (6, 24, 48 h and 4 days),Immunohistochemistry (1, 2, 4, 7days) and RT-PCR (4 and 7 days). Results demonstrated the


prepared surfaces supported cell attachment, cell proliferation, and mineralization, revealingno cytotoxicity effects. The expression of differentiation markers on the anodized surfacedemonstrated that genes related to the proliferation phase (Collagen type I, Coll I; Cbfa-1)were early expressed, whereas genes related to the mineralization phase (alkalinephosphatase, osteopontin, bone sialo protein) increased with time. Furthermore,mineralization was increased on the anodized surface.

Gransson et al. -08 [37] compared four types of possibly bioactive surfaces; a blastedsurface prepared by alkali heat treatment, anodization (Mg ions incorporated), fluoridation orhydroxyapatit coating, where a blasted surface served as control.

Surfaces were analyzed by Profilometry, and XPS after immersion in SBF for 12, 24 and72 hours.

Cells Primary (human mandibular osteoblast-like cells) were cultured on the varioussurfaces subjected to SBF for 72 h. Cellular attachment, differentiation (osteocalcin) andprotein production (TGF-beta(1)) was evaluated after 3 h and 10 days respectively. Theresults demonstrated that the possibly bioactive surfaces gave rise to an earlier CaP formationthan the blasted surface. Subsequent bone cell attachment was correlated to neither surfaceroughness 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 VivoLarsson et al. -94 [96] compared machined titanium surfaces and machined electropolishedwith and without anodization (1M acetic acid 10 and 80V). The surfaces were characterizedby SEM, AES and AFM. The surfaces differed with respect to surface oxide thickness (17-200 nm) and topography, although were similar with respect to surface composition. Bone-implant contact was evaluated in cortical bone in a rabbit model after 7 and 12 weeks. Theresults demonstrated decreased bone around the smooth electropolished surfaces compared tothe machined surfaces with similar oxide thickness and anodized implants with thicker oxidesafter 7 weeks. It was concluded that a high degree of bone contact and formation wereachieved by surface modifications with respect to oxide thickness and surface roughness.Furthermore, that a reduction in surface roughness influenced the rate of early boneformation.

Ishizawa et al. -95 [77] compared anodized titanium surfaces prepared in an electrolyte(CA/-GP, 350V) with different concentrations and with and without a subsequent heattreatment (300C , 2h) in a rabbit model. The surfaces were characterized by SEM. Turnedtitanium and a solid HA surface were used as positive and negative controls, respectively. Thepush out strengths and bone apposition increased after 8 weeks on the anodized-heat treatedsurface and were equivalent to HA ceramics. Furthermore, the anodized implants withoutheat-treatment showed increased push out strengths and bone apposition compared to theturned control surfaces. It was concluded that the good hard tissue compatibility of theimplant surfaces might be attributed to the surface roughness and the possibly inhibition oftitanium ion release.

Larsson et al. -96 [95] compared electropolished (smooth) and machined (rough) surfaceswith (thick oxide) and without (thin oxide) anodization (1M acetic acid, 80V) after 1, 3 and 6weeks. The surfaces were characterized by SEM, AES and AFM. At early stages, the smoothimplants demonstrated decreased bone-implant contact compared to the machined implants


irrespective of oxide layer thickness. At later stages, the thicker oxide layer increased thebone formation around the smooth surface, but not on the rougher machined surfaces. It wasconcluded that both topography, on the submicrometer scale, and the oxide thicknessinfluenced the bone response to titanium surfaces. Furthermore, that reduction of surfaceroughness in the initial phase decreases the rate of bone formation.

Larsson et al. -97 [94] compared electropolished (smooth) and machined (rough) surfaceswith (thick oxide) and without (thin oxide) anodization (1M acetic acid, 80V) after 1 year.The surfaces were characterized by SEM, AES and AFM. It was demonstrated that there wereno significant differences between the differently prepared implant groups after 1 year. It wasconcluded that a reduction of surface roughness from Rq 30 nm to 3 nm, which in the initialphase decreases the rate of bone formation, had no influence on the amount of bone after 1year in rabbit cortical bone

Ishizawa et al. -97 [78] compared anodized (CA/-GP) machined, grit-blasted andplasma-sprayed surfaces. A plasma-sprayed titanium surface and a solid HA surface wereused as controls. The surfaces were characterized by SEM and XRD.Bone response wasevaluated after 4 weeks in a dog model.

The anodized blasted implant showed increased bone formation compared to the smoothsurface. Furthermore, the thin HA layer demonstrated quantitatively the sameosteoconduction as the solid HA surface, however, with differed qualitatively.

Fini et al. -99 [88] compared etched (HF) titanium implants and anodized titaniumimplants (CA/-GP) prepared with and without heat-treatment. The surfaces werecharacterized by Profilometry, SEM, XRD and GD-OES.

Histomorphometric analysis demonstrated increased bone contact for the etched andanodized-heat treated surfaces compared to the anodized surfaces after 4 weeks, while theanodized-heat treated surfaces showed the highest values after 8 weeks in a rat femoralmodel.

Albrektsson et al. -00 [87] compared turned and anodized (TiUnite) titanium implants.The anodized implants demonstrated increased bone-implant contact and increased RTQcompared to the turned surfaces after 6 weeks in rabbit tibia and femur.

Gottlow et al. -00 [92] compared double etched (Osseotite) and anodized (TiUnite)implants. The anodized surfaces demonstrated increased bone-implant contact and stability bymeans of RFA and RTQ measurements compared to Osseotite after 6 weeks in rabbit femurand tibia.

Gottlow et al. -00 [91] compared double etched (Osseotite) and anodized (TiUnite)implants. The anodized implants demonstrated increased stability by means of RTQ after 10weeks in dog mandible; however, there were no differences in bone-implant contactcompared to the Osseotite implants.

Sennerby et al. -00 [99] compared insertion torque and stability of double etched(Osseotite) and anodized (TiUnite) implants. The anodized surface demonstrated an increasedinsertion torque, however, no differences in stability (RFA) after 3 weeks in rabbit tibia.

Henry et al -00 [93] compared stability of anodized (TiUnite) and turned implants after10 weeks in dog mandible. The anodized implants demonstrated a significantly increasedRTQ compared to turned implants.

Rompen et al. -00 [98] compared stability of anodized (TiUnite) and turned surfaces after3 and 6 weeks in dog mandible. It was concluded that the anodized implants maintainedhigher primary stability during 6 weeks of healing compared to the turned controls.


Sul et al. -02 [108] compared anodized (acetic acid) and turned titanium implants withvarious oxide thicknesses (600-1000 nm and 17-200 nm, respectively) in rabbit tibia. Thesurfaces were characterized by Laser Scanning Profilometry, SEM, XPS, TF-XRD and AES.There were no differences in ALP and ACP activity between the surfaces with different oxidethickness. However, implants with an oxide thickness > 600 nm demonstrated increasedbone-implant contact compared to the control surfaces. The increased bone response wasascribed the oxide properties including oxide thickness, pore size distribution, porosity andcrystal structure.

Sul et al. -02 [106] compared anodized (acetic acid) and turned titanium implants withvarious oxide thickness (20-1000 nm) in rabbit tibia. The surfaces were characterized byLaser Scanning Profilometry, SEM, XPS, Raman spectroscopy, TF-XRD and AES. Implantswith an oxide thickness > 600 nm demonstrated increased RTQ values compared to thinnerlayers, though there were no significant differences in RFA between the surfaces. Theincreased bone response were ascribed the oxide properties including oxide thickness,micropore configuration and crystal structure.

Sul et al -02 [104] compared calcium io

surface coatings

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