Laser Surface Treatment of Bio-implant Materials

Laser Surface Treatmentof Bio-Implant Materials

Laser Surface Treatment of Bio-Implant Materials L. Hao and J. Lawrence© 2005 John Wiley & Sons, Ltd ISBN: 0-470-01687-6

Laser Surface Treatmentof Bio-Implant Materials

Liang HaoLoughborough University, UK

Jonathan LawrenceNanyang Technological University, Singapore

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Library of Congress Cataloging-in-Publication Data

Hao, Liang.Laser surface treatment of bio-implant materials/Liang Hao, Jonathan Lawrence.

p. cm.Includes bibliographical references and index.ISBN-13 978-0-470-01687-9 (cloth : alk. paper)ISBN-10 0-470-01687-6 (cloth : alk. paper)

1. Biomedical materials–Effect of lasers on. 2. Biomedical materials–BiocompatibilityI. Lawrence, J. (Jonathan), 1970-. II. Title.[DNLM: 1. Lasers–therapeutic use. 2. Laser Surgery–methods.3. Biocompatible Materials. WB 117 H252L 2005]

R857.M3H37 20056100.2804–dc22 2005019354

British Library Cataloguing in Publication DataA catalogue record for this book is available from the British Library

ISBN-13 978-0-470-01687-9ISBN-10 0-470-01687-6

Typeset in 11/13pt Palatino by Thomson Press (India) Limited, New Delhi, India.Printed and bound in Great Britain by Antong Rowe Ltd, Chipperham, WiltshireThis book is printed on acid-free paper responsibly manufactured from sustainable forestry in which atleast two trees are planted for each one used for paper production.

To Yan-jun and My ParentsFor the world they bring to me.

To Louise and Ethan,Always there; beyond compare.

Contents

Acknowledgements xiii

Introduction xv

Bio-Implants and Surface Modification of Biomaterials xvWettability in Biomaterials Science and Modification Techniques xviLasers and Their Application for Modification of the Biomaterials xvii

1 Bioactivity and Biointegration of Orthopaedicand Dental Implants 11.1 Introduction 1

1.1.1 Biocompatibility 21.1.2 Host Response to Biomaterials 21.1.3 In vitro Models of Biological Response to Implants 2

1.2 Bioactivity of Bone Implants 31.2.1 The Mechanism of Apatite Formation 31.2.2 Functional Group 4

1.3 Biointegration of Orthopaedic and Dental Implants 51.3.1 Osseointegration 51.3.2 Bone Cell Adhesion [44] 51.3.3 Osteoblast–Material Interactions 6

1.4 Controlling the Bone–Implant Interface 61.4.1 Physicochemical Methods 71.4.2 Biochemical Methods [9] 7

2 Surface Modification of Biomaterials 112.1 Introduction 11

2.1.1 Orthopaedic and Dental Implants 112.1.2 Surface Properties of Biomaterials 12

2.1.3 Surface Analysis of Biomaterials 122.2 Ceramic Implants [65] 13

2.2.1 Nearly Bioinert Ceramics [64, 69] 132.2.2 Alumina 142.2.3 Zirconia Ceramics 14

2.3 Metallic Implants 152.3.1 Mechanical Properties 162.3.2 Corrosion 16

2.4 Surface Modification of Biomaterials 172.4.1 Introduction 172.4.2 Radiation Grafting and Photografting [76] 172.4.3 Plasma Surface Modification of Biomaterials 182.4.4 Ion Beam Processing 182.4.5 Other Methods [65] 19

2.5 Laser Surface Modification of Biomaterials 192.5.1 Introduction 192.5.2 Laser Patterning and Microfabrication 202.5.3 Pulsed Laser Deposition (PLD) of Biocompatible

Ceramics 202.5.4 Matrix-Assisted Pulsed Laser Evaporation and

MAPLE Direct Write 212.5.5 Other Laser Surface Treatments 21

3 Wettability in Biomaterials Science and ModificationTechniques 233.1 Introduction 233.2 Wettability, Adhesion and Bonding: Theoretical Background 23

3.2.1 The Wetting Process 233.2.2 Contact Angle and Work of Adhesion 243.2.3 Surface Energy and the Dispersive/Polar

Characteristics 253.2.4 Physical Bonding 283.2.5 Mechanical Bonding 283.2.6 Chemical Bonding 28

3.3 Wettability in Biomaterial Science 293.3.1 Biomaterial Interfaces [110] 293.3.2 Tensiometry 293.3.3 Interfacial Biophysics 303.3.4 Thermodynamic Concepts in Biomaterials Science 31

3.4 Current Methods of Wettability Modification 333.4.1 Chemical Reactions 333.4.2 Plasma Surface Modification 33

viii Contents

3.4.3 Ion Beam Processing 343.4.4 Radiation Grafting 343.4.5 UV and Ozone 343.4.6 Corona Discharge 353.4.7 Electrowetting 35

3.5 Laser Wettability Characteristics Modification 353.5.1 Laser Surface Modification of Ceramic Materials

for Improved Wettability 353.5.2 Laser Surface Modification of Metallic Materials

for Improved Wettability 36

4 CO2 Laser Modification of the Wettability Characteristicsof Magnesia Partially Stabilised Zirconia 374.1 Introduction 374.2 Experimental Procedures 38

4.2.1 Material Specifications 384.2.2 CO2 Laser Experimental Arrangement 394.2.3 Morphological, Chemical and Phase Analysis Procedures 394.2.4 Wettability Characteristics Analysis Procedure 40

4.3 The Effects of CO2 Laser Radiation on WettabilityCharacteristics 414.3.1 Contact Angle 414.3.2 The Effect of Surface Oxygen Content 424.3.3 The Effect of Surface Roughness 434.3.4 The Effects of Solidified Microstructures and Surface

Melting on Wettability Characteristics 454.4 Surface Energy and Its Component Parts 474.5 Identification of the Predominant Mechanisms Active

in Determining Wettability Characteristics 524.6 The Role Played by Microstructures in Terms of Crystal

Size and Phase in Effecting Surface Energy Changes 564.6.1 The Role of Crystal Size on Surface Energy 564.6.2 The Role of Phase Change on Surface Energy 61

4.7 Investigation of Wettability and Work Adhesion UsingPhysiological Liquids 61

4.8 Summary 63

5 In vitro Biocompatibility Evaluation of CO2 Laser TreatedMagnesia Partially Stabilised Zirconia 655.1 Introduction 655.2 Sample Preparation 675.3 Bone-Like Apatite Formation 67

Contents ix

5.3.1 Experimental Procedures 685.3.2 Spectral Analysis and Hydroxyl Group 695.3.3 The Correlation between OH Groups and Wettability

Characteristics 715.3.4 The Effects of CO2 Laser Treatment on the MgO–PSZ in

Simulated Body Fluids 735.4 Protein Adsorption 75

5.4.1 Experimental Procedures 765.4.2 Albumin and Fibronectin Adsorption on CO2 Laser

Treated MgO–PSZ 775.5 Osteoblast Cell Response 80

5.5.1 Experimental Procedures 815.5.2 Osteoblast Cell Response on the CO2 Laser Treated

MgO–PSZ 835.5.3 The Effect of CO2 Laser Treatment on the Osteoblast

Cell Response 895.6 Predictions for Implantation in an in vivo Clinical Situation 955.7 Summary 98

6 The Effects of CO2 Laser Radiation on the WettabilityCharacteristics of a Titanium Alloy 996.1 Introduction 996.2 Experimental Procedures 101

6.2.1 Material Specifications and Preparation 1016.2.2 CO2 Laser Surface Treatment 1026.2.3 Morphological, Chemical and Phase Analysis

Procedures 1026.2.4 Wettability Characteristics Analysis Procedure 102

6.3 The Effects of CO2 Laser Radiation on WettabilityCharacteristics 1036.3.1 Contact Angle 1036.3.2 Morphological Analysis and Its Effect on Wettability

Characteristics 1036.3.3 Phase and Chemical Analysis and Its Effects

on Wettability Characteristics 1056.4 Surface Energy and Its Component Analysis 1096.5 Identification of the Predominant Mechanisms Active

in Determining Wettability Characteristics 1116.6 Investigation of Wettability and Work Adhesion Using

Physiological Liquids 1146.7 Summary 116

x Contents

7 In vitro Biocompatibility Evaluation of CO2 LaserTreated Titanium Alloy 1177.1 Introduction 1177.2 Sample Preparation 1197.3 Bone-Like Apatite Formation on Titanium Alloys 120

7.3.1 Experimental Procedures 1217.3.2 The Effects of CO2 Laser Treatment on the Ti–6Al–4V

in Simulated Body Fluid 1217.4 Protein Adsorption 123

7.4.1 Experimental Procedures 1237.4.2 Albumin and Fibronectin Adsorption on CO2 Laser

Treated Titanium Alloy 1247.5 Osteoblast Cell Adhesion 127

7.5.1 Experimental Procedure 1277.5.2 Osteoblast Cell Response on CO2 Laser Treated

Titanium Alloy 1287.5.3 The Effect of CO2 Laser Treatment on the Osteoblast

Cell Response 1317.6 Predictions for Implantation in an in vivo Clinical

Situation 1357.7 Summary 138

8 Enquiry into Possible Generic Effects of the CO2 LaserTreatment on Bone Implant Biomaterials 1418.1 Introduction 1418.2 Ascertaining the Generic Effects of CO2 Laser Treatment

on Bioinert Ceramics 1428.2.1 Experimental Procedures 1438.2.2 Modification of the Surfaces Properties and Wettability

Characteristics of a Y–PSZ Bioinert Ceramic 1448.2.3 Identification of the Predominant Mechanism Active in

the Wettability Characteristics Modification of a Y–PSZBioinert Ceramic 149

8.2.4 Generic Effects of CO2 Laser Treatment on theWettability Characteristics of Bioinert Ceramics 150

8.2.5 CO2 Laser Induced Effects on the Cell Response on aY–PSZ Bioinert Ceramic 153

8.2.6 Generic Effects of CO2 Laser Treatment on theCell Response on Bioinert Ceramics 157

8.3 Ascertaining the Generic Effects of CO2 Laser Treatmenton Metal Implants 1578.3.1 Experimental Procedures 158

Contents xi

8.3.2 Modification of Surfaces Properties and WettabilityCharacteristics of a 316 LS Stainless Steel 159

8.3.3 Identification of the Predominant Mechanism Activein the Wettability Characteristics Modificationof a 316 LS Stainless Steel 166

8.3.4 Generic Effects of CO2 Laser Treatment on theWettability Characteristics of Biometals 168

8.3.5 CO2 Laser Induced Effects on Protein Adsorptionand the Cell Response on a 316 LS Stainless Steel 170

8.3.6 Generic Effects of CO2 Laser Treatment on ProteinAdsorption and the Cell Response on Biometals 175

8.4 Summary 176

Conclusions 179

References 185

Index 209

xii Contents

Acknowledgements

First, we gratefully appreciate the understanding and encouragement of ourfamilies who live on opposite sides of the world in China and England.

We are indebted to all the technicians in the Materials Laboratory atNanyang Technological University for their advice and assistance on opticalmicroscopy, XRD, SEM, EDX, contact angle analysis, XPS analysis andsample preparation. Many thanks also to the Doctoral Research studentsin the Materials Laboratory for sharing their knowledge on material treat-ment and analysis.

We acknowledge the tremendous contribution made by the first-rate workof the 2003–2004 Final Year Project students: Y.F. Phua, T.L. Tan, M.W. Kooand T.H. Wang.

On numerous occasions we obtained superb instruction and assistancefrom Mr Ma Dong Rui on the subject of osteoblast cell culture, for which weare extremely grateful.

Introduction

Bio-Implants and Surface Modification of Biomaterials

There is archaeological evidence that lost teeth were replaced by hand-carved ivory or wood ‘implants’ as long ago as ancient Egypt. In the early1950s Swedish orthopaedic surgeon Per Ingvar Branemark began studyingthe healing process of titanium anchoring screws, which proved to be aseminal point for modern dental and orthopaedic implants [1]. His workshowed that fusion between bone and the titanium implant could take place,a phenomenon he called ‘osseointegration’. Orthopaedic implants to treatjoint degradation due primarily to osteoarthritis, osteoporosis or injury arenow commonplace. These include hip (around 325 000 US implants in 2001)and shoulder, wrist and knee (around 300 000 US implants in 2001).Biomaterial applications make use of all classes of materials, metals, cera-mics, polymers and composite. These are divided roughly into three usertypes [2]: (a) inert or relatively inert with minimal host response; (b)bioactive, which actually stimulates bonding to the surrounding tissue;and (c) biodegradable, which resorb in the body over a period of time.

Events leading to integration of an implant into bone, which in turndetermine the performance of the device, take place largely at the tissue–implant interface. The main requirements for a biomaterial to functionproperly in an osseous site include good biocompatibility favouring boneapposition, adequate mechanical properties and the ability to ensure skeletalfunctions [3–5]. Bioactivity and biointegration are the two essential aspectsof these interactions. Bioactivity and the maintenance of skeletal functionsare usually attributed to the ability to induce an apatite layer on a material’ssurface in physiological conditions [6–8]. The close apposition between boneand an implant surface, or osseointegration, presented as the ability topromote bone cells anchorage, attachment, spreading, growth and differ-entiation [9, 10], is another key factor for successful implantation of abiomaterial for dental and orthopaedic applications [3–5].

Biointegration is the ideal outcome expected of an artificial implant. Thisimplies that the phenomena that occur at the interface between the implantand host tissues do not induce any deleterious effects such as chronicinflammatory response or formation of unusual tissues. It is, therefore, ofparamount importance to design biomaterials used in implants with the bestsurface properties. Meanwhile, these biomaterials must possess bulk proper-ties that meet other requirements, especially mechanical properties, in orderto function properly in a bioenvironment. As it is quite difficult to designbiomaterials fulfilling both needs, a common approach is to fabricatebiomaterials with adequate bulk properties followed by a special treatmentto enhance the surface properties. Hence surface modification of biomater-ials is becoming an increasingly popular method to improve device multi-functionality, tribological and mechanical properties, as well as bio-compatibility of artificial devices while obviating the needs for largeexpenses and a long time to develop brand new materials [11]. Materialscan be surface modified by using biological or physicochemical methods.

Wettability in Biomaterials Science and Modification Techniques

The wetting of a surface by a liquid and the ultimate extent of spreading ofthat liquid are very important aspects of practical surface chemistry. Evenwith all the new information of the last 20 years, however, there still remainsa great deal to learn about the mechanisms of movement of a liquid across asurface and the factors that govern such movement [12]. Biomaterialscientists have long sought a single, material-related parameter that effec-tively measures biocompatibility and might serve as a practical designguide. The theories of surface energy and wetting for such parameterspresent an attractive means to do this as surface properties are importantdeterminants of a biomaterial function [12]. The ability to control the surfacewettability of solid substrates is important in many situations. Varioussurface processes are used for modifying the surfaces of materials depend-ing on the actual material and the application. A number of laser-basedtechniques for altering the wettability characteristics of engineering materi-als have been investigated [13].

Surface sensitivity is of critical importance in biomaterials surface sciencebecause only the uppermost layers are in direct physicochemical contactwith the biological environment; consequently only the upper few molecularlayers determine biocompatibility. Thus the interfacial chemistry of concernto biomaterial scientists is determined by material composition within theupper nanometre or so. Tensiometry encompasses a broad range of related‘wetting’ techniques that measure surface energy. These include the obser-vation of contact angles, which is perhaps the most familiar and widely

xvi Introduction

applied method. Tensiometric methods have singular potential in biomater-ials surface science based on the criteria of surface sensitivity, kind ofanalytical information obtained and relevance of that information to biome-dical problems. First, with respect to surface sensitivity, wetting measure-ments are sensitive only to the upper 0.5 nm or so of a surface [14, 15] andare therefore among the most surface-sensitive techniques available. Second,tensiometry directly measures the fundamental energy at an interface thatdrives important processes such as adsorption and adhesion. This kind ofinformation must be particularly pertinent to biomaterial problems becauseof the overwhelming importance of protein adsorption and cell/tissueadhesion. Third, wetting measurements can be made using proteinaceoussaline solutions that are particularly relevant to biomedical applications.Special high-vacuum preparation techniques that might introduce experi-mental artefacts are not required.

Various methods are used to improve the surface wettability of materialsand their adhesion to other materials. Hydrophilicity is a characteristic ofmaterials exhibiting an affinity for water. Hydrophilic literally means ‘water-loving’ and such materials readily adsorb water. Hydrophobic describesmaterials possessing a characteristic that has the opposite response to waterinteraction compared to hydrophilic materials. Smaller water contact anglescorrespond to more hydrophilic surfaces and higher surface free energies.

At present, the processes available to engineers for the modification of amaterial’s wettability characteristics are invariably complex and conse-quently somewhat difficult to control. Lasers, on the other hand, can offerthe user not only an exceedingly high degree of process controllability butalso a great deal of process flexibility. There is a growing amount ofpublished work that testifies to the potential of lasers for altering the surfaceproperties of materials in order to improve their wettability characteristics.Laser radiation was found to effect significant changes in the wettabilitycharacteristics of materials. Lawrence and Li [13] have amply demonstratedthe practicability of employing different types of lasers to effect changes inthe wettability characteristics of composites and ceramics, metals andplastics for improved adhesion and bonding.

Lasers and Their Application for Modification of the Biomaterials

The laser has some unique properties for surface treatment. The electro-magnetic radiation of a laser beam is absorbed within the first atomic layersfor opaque materials, such as metals, and there are no associated hot gas jets,eddy currents or even radiation spillage outside the optically defined beamarea. In fact, the applied energy can be placed precisely on the surface onlywhere it is needed. Thus it is a true surface heater and a unique tool for

Lasers and Their Application for Modification of the Biomaterials xvii

surface engineering. Common advantages of laser surfacing compared toalternatives are [16]:

� chemical cleanliness;� controlled thermal penetration and, therefore, distortion;� controlled thermal profile and, therefore, shape and location of the heat

affected region;� less after-machining, if any, is required;� remote noncontact processing is usually possible; and� relatively easy to automate.

Surface treatment is a subject of considerable interest at present as it seemsto offer the chance to save strategic materials or to allow improvedcomponents with idealised surfaces and bulk properties. At present, thelasers are being used in the following surface modifications of the biomater-ials:

� Laser patterning and microfabrication� Pulsed laser deposition (PLD) of biocompatible ceramics� Matrix-assisted pulsed laser evaporation (MAPLE) and MAPLE direct

write (MDW)� Laser surface treatment for improving corrosion� Laser grafting� Laser treatment of plasma sprayed hydroxyapatite coatings

However, little work has been carried out to investigate employing lasers tomodify the surface properties of biomaterials in order to improve theirbiocompatibility. Having said that, it is recognised within the currentlypublished work that laser irradiation of material surfaces can affect changesin the cell adhesion on biomaterials. Lately, several publications haveinvestigated the modification of biocompatibility of a biomaterial’s surfacefollowing laser irradiation. A CO2 pulsed laser was used to graft a polymer[17] and a rubber [18]. The results showed a marked reduction of the plateletadhesion and aggregation for the modified polymer surface and cell attach-ment, with a greater degree of spreading and flattening on the unmodifiedrubber surface. L929 fibroblast cells attached and proliferated extensively onthe CO2 and KrF laser treated films [19] in comparison with unmodified PET(polyethylene teraphthalate), with surface morphology and wettability beingfound to affect cell adhesion and spreading. However, so far, no work hasinvestigated the use of laser surface modification of bioinert ceramics andbiograde metals for improved biocompatibility.

With a view to improve the biocompatibility of the dental and orthopaedicimplant, a CO2 laser was used to modify the surface properties of the widelyused bioinert ceramics and biograde metals. By varying the CO2 laser powerdensities, the work investigated how the CO2 laser affected the surface

xviii Introduction

properties of magnesia–partially stabilised zirconia (MgO–PSZ), a bioinertceramic, such as morphology, surface roughness, surface oxygen contentand rapidly solidified microstructure. It particularly analysed the change inthe wettability characteristics and basic mechanisms governing this mod-ification [20–22], since wettability is widely recognised as an importantdeterminant of cell adhesion and biomaterial’s function. Then, the bioactiv-ity evaluation of the MgO–PSZ was conducted to find whether the bone-likeapatite could form on the MgO–PSZ following CO2 laser irradiation in thestimulated body fluid and what were the functional groups for apatitenucleation [23]. Furthermore, it investigated how the CO2 laser modifiedsurface properties influence the protein adsorption [24, 25] and osteoblastcell adhesion [26] that would manipulate the biointegration between theimplant and tissue. Moreover, the effects of the CO2 laser on the modifica-tion of the wettability characteristics of a titanium alloy (Ti–6Al–4V) andthereof the adhesion with the simulated physiological liquids were analysed.Further, it observed the apatite formation on the CO2 laser treated Ti–6Al–4V alloy after soaking in simulated body fluid and the significant differenceof albumin and fibronectin adsorption between the untreated sampleand CO2 laser treated samples. More importantly, osteoblast cell adhesionand proliferation performed on the CO2 laser treated Ti–6Al–4V wascompared with untreated titanium alloy. With the aim of establishing thelaser as the innovative technique for the surface modification of the bioma-terials, the generic effects of CO2 laser irradiation on the biocompatibility ofa yttria–partially stabilised zirconia (Y–PSZ) and a biograde stainless steelare investigated. Similar effects of the CO2 laser treatment on the wettabilitycharacteristics, protein adsorption and osteoblast cell were observed on theY–PSZ and stainless steel. The study therefore proved the generic and greatpotential application of the laser surface process of widely used bioinertceramic and biograde metals.

Lasers and Their Application for Modification of the Biomaterials xix

1

Bioactivity and Biointegrationof Orthopaedic and DentalImplants

1.1 Introduction

Events leading to integration of an implant into bone, which in turndetermine the performance of the device, take place largely at the tissue–implant interface. The main requirements for a biomaterial to functionproperly in an osseous site include good biocompatibility favouring boneapposition, adequate mechanical properties and the ability to assure skeletalfunctions [3–5]. Bioactivity and biointegration are the two essential aspectsof these interactions. Bioactivity and the maintenance of skeletal functionsare usually attributed to the ability to induce an apatite layer on a material’ssurface in physiological conditions [6–8]. The close apposition between boneand an implant surface, or osseointegration, presented as the ability topromote bone cells anchorage, attachment, spreading, growth and differ-entiation [9, 10], is another key factor for successful implantation of abiomaterial for dental and orthopaedic applications [3–5]. Recent yearshave seen considerable interest in the systematic investigation of therelationship between surface chemistry [27, 28] and morphology [29] andbiological interfacial reactions. Ultimately, such studies may lead to anenhanced understanding of the surface chemical cues that guide the combi-nation of apatite induction ability with attachment, growth and differentia-tion of bone cells and to the design of improved synthetic materials fordental and orthopaedic applications.


1.1.1 Biocompatibility

A biocompatible material has been defined as a material that does not inducean acute or chronic inflammatory response and does not prevent a properdifferentiation of implant surrounding tissues [30]. It is recognised that someadverse tissue reaction around the implanted biomaterial is inevitable,owing to surgical trauma during insertion. This definition implies thatbiocompatibility depends on the purpose of the implant. Terms such as‘bioinert’ or ‘bioactive’, rather than ‘biocompatibility’, more accuratelydescribe the features required of an ideal biomaterial or device. Theseterms better describe the action or nonaction required from surroundingtissue and are thus related to the choice of materials and material character-istics (hydrophilic/hydrophobic).

1.1.2 Host Response to Biomaterials

The biocompatibility of the implant material is closely related to thereactions between the surface of the biomaterial and the inflammatoryhost response [31]. The implantation response in bone differs in someways from that taking place in soft tissue. There is an inflammatory and areparative response which occur one on the other. The reparative responsestarts 2–3 days after the implantation. The stem cells of bone develop intoosteoblasts, which form a layer near the implant together with fibroblasts.Fibroblasts, osteoblasts and capillaries penetrate into the blood clot, repla-cing it, and fill the space between the implant and bone [32]. After theformation of a collagen-rich extracellular matrix (ECM), mineralizationfollows. Normally, there are vesicles in the ECM and some of them includecalcification focuses. The presence of vesicles with biomaterial in the earlyperiod is a sign of good primary acceptance. When the membranes of thesevesicles rupture, the erupted apatite crystals unite and form calcifyingstructures [33]. Early trabecles grow and continue to mineralise, and someof them reach the implant surface. In an optimal situation, the material iscovered by bone tissue and not by fibrous capsule. The healing of bone tissuecontinues like fracture healing. Remodeling bone tissue begins after twoweeks and continues for the lifetime. When the material is biocompatible,there is an abundance of the ECM and osteoblasts. This is confirmed by theclose attachment and fast proliferation of these cells [34].

1.1.3 In vitro Models of Biological Response to Implants

In vitro models have been widely used to investigate biological responses tobiomaterials, and the materials adopted for implantology have been noexception. Given their relative ease of use and lack of expense, in vitromodels would appear to investigate the biological response to implant

2 Bioactivity and Biointegration

materials [35]. Cell culture methods have been used to evaluate the biolo-gical compatibility of materials for more than two decades [36]. Bone cellculture models are increasingly employed to study bone–biomaterial inter-actions. Most of the cultures have utilized osteoblastic cells (reviewed byCooper et al. [37]), with only a few using osteoclastic cells [38, 39]. In vitromodels have the potential to help elucidate events at the bone–implantinterface (reviewed by Davies [40]), by providing morphological, biochem-ical and molecular information regarding osteoblastic development andsynthesis of the matrix at the interface with various biomaterials.

1.2 Bioactivity of Bone Implants

For an artificial material to bond to living bone, it is essential that thematerial has the ability to form a biologically active, bone-like, apatite layeron its surface in the human body. Under normal conditions, the body fluid isalready supersaturated with respect to apatite, and once apatite on its nucleiform on the surface of a material, they can spontaneously grow by consum-ing the calcium and phosphate ions from the body fluid. The nucleation ofapatite on the surface of a material is induced by the functional groups on itssurface. Naturally, the development of bioactive materials that have improvedand ultimately the bone-like mechanical properties is desirable [41].

1.2.1 The Mechanism of Apatite Formation

The biological activity of most orthopaedic and dental biomaterials is relatedto their ability to promote the formation of a neoformed layer of carbonateapatite crystals analogous to bone mineral. This layer also associates specificbone proteins and is the starting point of bone reconstruction [42]. Ortho-paedic biomaterials may be classified as ‘passive’ or ‘active’ with regard totheir propensity simply to allow the nucleation and growth of carbonateapatite crystals from body fluids (hydroapatite, titanium oxide, neutralhydrogels, collagen) or to supply ions to build and develop this layer(bioglasses, alkaline hydroxides, Ca–P compounds, calcium carbonate).Bioactive ceramics have a common characteristic at the interface with boneafter integration. In fact, bioactive ceramics, including bioglass and hydro-xyapatite (HA) reveal a layer of apatite at the interface, mediating integra-tion with bone. Histological examination in vivo shows that this apatite layeris formed on the ceramic surface early in the implantation period andthereafter the bone matrix integrates into the apatite. Detailed characterisa-tion indicated that this apatite layer consists of nanocrystals of carbonate-ion-containing apatite that has a defective structure and low crystallinity.These features are, in fact, very similar to those of the mineral phase in bone;

Bioactivity of Bone Implants 3

hence bone-producing cells (osteoblasts) can preferentially proliferate on theapatite and differentiate to form an extracellular matrix composed ofbiological apatite and collagen. As a result, the surrounding bone comesinto direct contact with the surface apatite layer. When this process occurs, achemical bond is formed between the bone mineral and the surface apatite todecrease the interfacial energy between them. It can be concluded that anessential requirement for an artificial material to bond to living bone is theformation of a layer of biologically active bone-like apatite on its surface inthe body.

1.2.2 Functional Group

Bioactivity can be induced on surfaces of nonbioactive materials either bythe formation of the functional groups that are able to induce apatiteformation or by forming thin ceramic phases that have the potential toform the functional groups on exposure to a body environment [41]. Thecatalytic effect of the Si–OH groups and Ti–OH groups for apatite nucleationhas been proven from the observation that silica and titania gels producedby the sol-gel method form apatite on their surfaces in simulated body fluids(SBF), and these functional groups are abundant on their surfaces. Zirconia,niobium oxide and tantalum oxide gels have also been shown to form apatiteon their surface in SBF, as shown in Figure 1.1. This indicates that Zr–OH,Nb–OH and Ta–PH groups are effective for apatite nucleation. Otherassessments using self-assembled monolayers (SAM) in SBF have indicatedthat COOH and PO4H2 groups are also effective for apatite nucleation [41].

Figure 1.1 Scanning electron microscopy (SEM) photographs of the surfaces of silica(A), titania (B), zirconia (C), niobium oxide (D) and tantalum oxide (E) gels aftersoaking in an SBF for 14 d [41]


These groups have specific structures revealing negatively charge, andinduce apatite formation via formations of an amorphous calcium com-pound, e.g. calcium silicate, calcium titanate and amorphous calciumphosphate. Moreover, the efficacy of apatite nucleation of the above func-tional groups is determined, not by their composition alone but in acomplicated fashion that is dependent on their concentration and structuralarrangement.

1.3 Biointegration of Orthopaedic and Dental Implants

1.3.1 Osseointegration

Integration of the implant into the bone is a property of paramountimportance in the proper functioning of the implant; consequently extensivestudies have been carried out using different techniques to improveosseointegration. Osseointegration was defined by Branemark [43] as: ‘Adirect structural and functional connection between living bone and thesurface of a load-carrying implant.’ The integration of a biomaterial to boneinvolves essentially two processes: interlocking with bone tissue and che-mical interactions with bone constituents. It is essential for the efficacy oforthopaedic or dental implants to establish a mechanically solid interfacewith complete fusion between the material’s surface and the bone tissuewith no fibrous tissue interface.

1.3.2 Bone Cell Adhesion [44]

The term ‘adhesion’ in the biomaterial domain covers different phenomena:the attachment phase, which occurs rapidly and involves short-term eventslike physicochemical linkages between cells and materials involving ionicforces, van der Waals forces, etc., and the adhesion phase, occurring in thelonger term and involving various biological molecules: extracellular matrixproteins, cell membrane proteins and cytoskeleton proteins which interacttogether to induce signal transduction, promoting the action of transcriptionfactors and consequently regulating gene expression. It is widely acknowl-edged that a major determinant of the bone–biomaterial interfacial responseis the initial attachment, spreading and growth of osteoblasts on the implantsurface and that improvements in these processes may lead to faster andlong-term stability [3, 45]. It has been possible to demonstrate cellularattachment to implant surfaces and there is good evidence for new boneformation being initiated at a distance from but towards the surface of animplant rather than on the surface itself [3, 45]. The biocompatibility ofbiomaterials is very closely related to cell behaviour on contact with themand particularly to cell adhesion to their surface.

Biointegration of Orthopaedic and Dental Implants 5

1.3.3 Osteoblast–Material Interactions

Osteoblast–material interaction depends on the surface aspects of materials,which may be described according to their topography, chemistry or surfaceenergy. These surface characteristics determine how biological moleculeswill adsorb to the surface and more particularly determine the orientation ofadsorbed molecules [46]. They also determine the cell behaviour on contact.

It is known that osteoblast cells initially respond in a differential manner tothe material surface. The comparison of the behaviour of different cell typeson materials shows that they react differently according to surface roughness[28, 47]. Scanning electron microscopic examination of bone cells on materi-als with various surface roughness generally demonstrated that cell spread-ing and continuous cell layer formation were better on smooth surfacescompared to rough surfaces [48–50]; however, higher levels of cellularattachment have been found on rough surfaces of titanium with irregularmorphologies [51–53] in vitro. Similarly, recent studies have shown thatalkaline phosphatase (ALP) specific activity is enhanced on rough Ti andTi–6Al–4V [54, 55]. Cells grown on rougher surfaces exhibited increasedproduction of collagen [52, 54] and transforming growth factor b [54]. Thesedifferences of cell response could be attributed to either the microstructure,crystalline or chemistry, as different methods were used to obtain differentroughnesses. A contact guidance phenomenon has also been described onosteoblastic cells. On smooth surfaces, bone cells were randomly orientedalthough they were aligned parallel to the direction of the grooves in anend-to-end fashion in 5 mm deep grooves [47].

Osteoblast attachment is also affected by the chemistry of the biomaterial’ssurface, but a conclusive picture has not yet emerged. Early in vitrocytocompatibility studies focused on the morphological aspect, growthcapacity and the state of differentiation of cells on materials with variouschemical compositions. The diversity of cell responses to the differentmaterials tested highlighted the capacity of cells to distinguish the effectsof subtle changes in substratum surface chemistry. For primary bovineosteoblasts, the wettability of the surface has been shown to be one of theimportant factors [56]. On the other hand, the functional groups present onthe surface have been demonstrated to be of even greater importance thanthe wettability for the adhesion of MC3T3-E1 osteoblasts [57].

1.4 Controlling the Bone–Implant Interface

Different approaches are being used in an effort to obtain the desired bone–implant interface. As Kasemo and Lausmaa [58], among others, havedescribed, biological tissues interact mainly with the outermost atomic


layers of an implant; consequently, much effort is being devoted to methodsof modifying surfaces of existing biomaterials to achieve desired biologicalresponses. The approaches can be classified as physicochemical and bio-chemical methods.

1.4.1 Physicochemical Methods

Wettability characteristics are among the physicochemical characteristicsthat have been altered with the aim of improving the bone–implant interface[9]. Polymer surfaces were modified by glow discharge to study the effect ofsurface treatment on cell adhesion using polyethylene, polytetrafluoroethy-lene, poly(ethylene terephthalate), polystyrene and polypropylene films. Thesurface wettability of all the films decreased with respect to the length ofplasma treatment. For each of the polymers, a different dependence of celladhesion on the length of plasma treatment was observed, but, in each case,the optimal water contact angle for cell adhesion was approximately 70� [59].

Alterations in surface morphology and roughness have been used toinfluence cell and tissue responses to implants [9]. In addition to providingmechanical interlocking, surfaces with grooves can induce ‘contact gui-dance’, whereby the direction of cell movement is affected by the morphol-ogy of the substrate. For osteoblast-like cloned mouse cells (MC3T3-E1)cultured on titanium plates roughened by wire-type electric dischargemachining or plasma coating, proliferation and ALP activity were enhancedon the roughened surfaces [60].

Considering the role of electrostatic interactions in many biological events,charged surfaces have been proposed as being conducive to tissue integra-tion. Calcium phosphate coatings have been extensively investigatedbecause of their chemical similarity to bone mineral [9]. The osteoblast-likecells cultured on RKKP- and AP40-bioglass layer coated zirconia showed ahigher proliferation rate, leading to confluent cultures with higher celldensity and a generally better expression of osteoblast ALP activity incomparison with zirconia substrate [61]. Each approach, however, hasdrawbacks. Contradictory results with charged materials in bone havebeen reported: indeed both positively and negatively charged surfaceswere observed to promote bone formation. Although short-term clinicalresults have been encouraging, dissolution of coatings as well as crackingand their separation from metallic substrates remain concerns [9].

1.4.2 Biochemical Methods [9]

Biochemical methods of surface modification offer an alternative or adjunctto physicochemical and morphological methods. Biochemical surface mod-ification endeavours to utilise current understanding of the biology andbiochemistry of cellular function and differentiation. Much has been learned

Controlling the Bone–Implant Interface 7

about the mechanisms by which cells adhere to substrates, and majoradvances have been made in understanding the role of biomolecules inregulating differentiation and remodelling of cells and tissues, respectively.The goal of biochemical surface modification is to immobilise proteins,enzymes or peptides on biomaterials for the purpose of inducing specificcell and tissue responses or, in other words, to control the tissue–implantinterface with molecules delivered directly to the interface. One approach tocontrolling cell–biomaterial interactions utilises cell adhesion molecules.Since identification of the arginine–glycine–aspartic acid (RGD) sequenceas mediating attachment of cells to several plasma and extracellular matrixproteins, including fibronectin, vitronectin, type I collagen, osteopontin andbone sialoprotein, researchers have been depositing RGD-containing pep-tides on biomaterials to promote cell attachment. A second approach tobiochemical surface modification uses biomolecules having demonstratedosteotropic effects. A large amount of information has been obtained aboutbiomolecules involved in bone development and fracture healing. Bydelivering one or more of these molecules, which normally play essentialroles in osteogenesis, directly to the tissue–implant interface, bone formationmay be promoted.

To control exposure and concentration, retention and/or release of bio-molecules from implant surfaces can be altered using different methods,including adsorption, covalent immobilisation and release from coatings(Figure 1.2). The simplest way to deliver biomolecules to the tissue–implantinterface is by dipping the device in a solution of protein before inserting it.Studies using simple adsorption indicate that delivery of TGF-b to thetissue–implant interface can improve bone formation in the periprosthetic

Figure 1.2 Schematic illustration of methods for controlling retention and/or releaseof biomolecules at the tissue–implant interface [9]


gap and can enhance bone ingrowth into porous coatings [62]. Using asimilar approach, ALP adsorbed on titanium implants enhanced peripros-thetic bone formation [63]. One drawback with the adsorption method,however, is that it provides little control over the delivery, includingrelease/retention and orientation, of molecules. Proteins are initiallyretained on the surface by weak physisorption forces; then, depending onthe implant microenvironment, which varies between anatomical sites andbetween patients, they desorb from the surface in an uncontrolled manner toinitiate desired responses. Considering the necessity of specific receptor–ligand interactions for activity of many relevant biomolecules, appropriatepresentation of protein may also be needed. Although positive responseshave been observed using this simple approach, there is no indication thatthey are optimal for clinical applications.

Bonding biomolecules to implants is an alternate way of delivering themto the tissue–implant interface, albeit protein will not be released. Thisapproach is more complicated than adsorption, because of the chemistryinvolved, but the activity of molecules immobilised on plastics has beenshown to equal or exceed that of soluble protein [64]. For orthopaedic anddental applications, metal surfaces possess a relative paucity of functionalgroups needed for immobilising molecules. However, the passivating oxidefilm on these materials does have surface hydroxyl groups that providelocations for bonding using silane chemistry. This approach has been used toimmobilise peptides, enzymes and adhesive proteins on different biomater-ials, including Co–Cr–Mo, Ti–6Al–4V, Ti and NiTi [9].

Controlling the Bone–Implant Interface 9

2

Surface Modificationof Biomaterials

2.1 Introduction

Biomaterials have been studied for many years and have been defined byRatner [65] as being nonviable materials used in a medical device andintended to interact with a biological system. There is a big demand forbiomaterials to assist or replace organ functions and to improve patients’quality of life. Biomaterial applications make use of all classes of materials,metals, ceramics, polymers and composite. These are divided roughly intothree user types [2]: (a) inert or relatively inert with minimal host response;(b) bioactive, which actually stimulates bonding to the surrounding tissue;and (c) biodegradable, which resorb in the body over a period of time.

The biological responses to biomaterials and devices are largely controlledby their surface chemistry and structure. That is to say, the surfacecharacteristics play a role in the functioning of biomaterial. The rationalefor the surface modification of biomaterials is straightforward. Eitherbiological or physicochemical methods are often employed to modify thematerial surface. Various physicochemical methods will be introduced inthis chapter; among them, laser surface treatment has proved to have goodpotential for it is a unique feature.

2.1.1 Orthopaedic and Dental Implants

There is archaeological evidence that lost teeth were replaced by hand-carved ivory or wood ‘implants’ as long ago as ancient Egypt. In the early1950s Swedish orthopaedic surgeon Per Ingvar Branemark began studying


the healing process of titanium anchoring screws, which proved to be aseminal point for modern dental and orthopaedic implants [1]. His workshowed that fusion between bone and the titanium implant could take place,a phenomenon he called ‘osseointegration’.

Orthopaedic implants to treat joint degradation due primarily to osteoar-thritis, osteoporosis or injury are now commonplace. These include hip(around 325 000 US implants in 2001) and shoulder, wrist and knee (around300 000 US implants in 2001). These types of implant have undergonecontinual advancement of their metal and plastic components in order toimprove wear resistance and fixation in the insertion site; nevertheless,research continues for more wear-resistant joint interface materials, formaterials with improved mechanical properties, and to improve long-termreliability, especially as it relates to fixation. The objective, of course, is todevelop implants that can serve patients for an indefinite period.

2.1.2 Surface Properties of Biomaterials

In the case of medical implants the importance of surface science is quiteobvious [66, 67]. It has been hypothesised that tissue–biomaterial interac-tions are governed by surface properties and that the important interactionsoccur within around 1 nm of the biomaterial surface [58]. Natural biologicalstructures appear to be able to interact selectively with relevant biomoleculeswhile resisting nonspecific interactions. Furthermore, biological interfacesare highly dynamic. As pointed out by Blawas and Reichert [68], simplytrapping cells at a particular point on a surface is not enough. Cells must firstbe encouraged to ‘differentiate’ (i.e. change their behaviour to perform asrequired) and once their function is complete their activity must be ‘turnedoff’ again.

2.1.3 Surface Analysis of Biomaterials

Scanning electron microscopy (SEM) has been the traditional method forstudying the microscopic ‘surface structure’ of biomaterials. While allowingvisualisation and chemical analysis of the specimen, SEM is not a surface-specific technique in the strictest sense. Elemental microanalysis (EDAX)may also be carried out using SEM. This relies on the detection of X-rays ofcharacteristic energy, which are emitted on interaction with the electronbeam. With the recent, rapid growth of methods for preparing spatially well-defined materials, the focus of biomedical surface science is now on highspatial resolution surface chemical state analysis. The driving forces fordeveloping biomedical surface chemical state imaging techniques areaddressed below. X-ray photoelectron spectroscopy (XPS), time-of-flightsecondary ion mass spectrometry (ToF SIMS), scanning probe microscopy

12 Surface Modification of Biomaterials

(SPM) and near-edge X-ray absorption fine structure (NEXAFS) each has itsown strengths and weaknesses with respect to generating surface chemicalstate information at high spatial resolution, but together they provide apowerful set of complementary techniques. For example, XPS and ToF SIMScan be used to improve the level of chemical state information obtainablewith SPM, while SPM can be used to improve the spatial resolutionobtainable with XPS and ToF SIMS.

2.2 Ceramic Implants [65]

It is essential to recognise that no one material is suitable for all biomaterialapplications. As a class of biomaterials, ceramics, glasses and glass–ceramicsare generally used to repair or replace skeletal hard connective tissues. Theirsuccess depends upon achieving a stable attachment to connective tissue.The mechanism of tissue attachment is directly related to the type of tissueresponse at the implant–tissue interface (see Table 2.1).

2.2.1 Nearly Bioinert Ceramics [65, 69]

Bioceramics are compatible because they are composed of ions commonlyfound in the physiological environment (calcium, potassium, magnesium,sodium, etc.) and of ions showing limited toxicity to body tissue (zirconiumand titanium). Two nearly inert ceramics most used in surgical implants are

Table 2.1 Types of bioceramic–tissue attachment and their classification [65]

Type of attachment Example

1. Dense, nonporous, nearly inert ceramicsattach by bone growth into surfaceirregularities by cementing the device intothe tissues or by press-fitting into a defect(termed ‘morphological fixation’)

Al2O3 (single crystal and polycrystalline)ZrO2 (partially stabilised zirconia)

2. For porous inert implants, bone ingrowthoccurs that mechanically attaches the boneto the material (termed ‘biological fixation’)

Al2O3 (polycrystalline)Hydroxyapatite-coated porous metals

3. Dense, nonporous surface-reactive ceramics,glasses and glass–ceramics attach directlyby chemical bonding with the bone(termed ‘bioactive fixation’)

Bioactive glassBioactive glass–ceramicsHydroxyapatite

4. Dense, nonporous (or porous) resorbableceramics are designed to be slowlyreplaced by bone

Calcium sulfate (plaster of Paris)Tricalcium phosphateCalcium–phosphate salts

Ceramic Implants [65] 13

alumina and zirconia. The characteristics of bioinert ceramics for biomedicalapplication are shown in Table 2.2. High-strength ceramics used forimplants are very inert in the body and exhibit minimal ion release. Inertbioceramics undergo little or no chemical change during long-term exposureto the physiological environment.

2.2.2 Alumina

High-density, high-purity (>99.5%) alumina is used in load-bearing hipprostheses and dental implants because of its excellent corrosion resistance,good biocompatibility, high wear resistance and high strength [69, 71].Although some dental implants are single-crystal sapphires most Al2O3

devices are very fine-grained polycrystalline a-Al2O3 produced by pressingand sintering at T ¼ 1600–1700 �C. Alumina has been used in orthopaedicsurgery for nearly 20 years. Its use has been motivated largely by twofactors: its excellent biocompatibility and very thin capsule formation, whichpermits cementless fixation of prostheses and its exceptionally low coeffi-cients of friction and wear rates.

2.2.3 Zirconia Ceramics

Zirconia is also exceptionally inert in the physiological environment andzirconia ceramics have an advantage over alumina ceramics of higherfracture toughness and higher flexural strength and lower Young’s modulus[69]. Partially stabilised zirconia (PSZ) is a ceramic that has found wideusage in medical and dental surgery. During a heating process, zirconia willundergo a phase transformation process. The change in volume associatedwith this transformation makes the usage of pure zirconia in many applica-tions impossible. Addition of some oxides, such as calcia (CaO), magnesia(MgO) and yttria (Y2O3), into the zirconia structure in a certain degree

Table 2.2 Properties of bioinert ceramics [70]

Property Units Alumina MgO–PSZ Y–PSZ (TZP)

Chemical composition 99.9 %MgO Al2O3 þ ZrO2 þ MgO ZrO2 þ Y2O3

(8–10 mol %) (3 mol %)Density g/cm3 �3.97 5.74–6 >6Bending strength MPa >500 450–700 900–1200Compression strength MPa 4100 2000 2000Young modulus GPa 380 200 210Fracture toughness KIC MPa/m 4 7–15 7–10Hardness HV 0.1 2200 1200 1200


results in a solid solution, which is a cubic form and has no phasetransformation during heating and cooling. This solid solution material istermed ‘stabilised zirconia’. Within medicine it is commonly used tofabricate hip ball joints, knee, thumb, etc., while in dentistry it is used tomanufacture dental implants, dental posts, brackets and inlays. Some zirconiaimplants for medical and dental applications are shown in Figure 2.1.

The published results of in vitro wear tests demonstrated that zirconia hasa superior wear resistance. Saikko [72] showed no wear of zirconia femoralheads on his hip simulator wear test against the 10.9 mm ultra highmolecular weight polyethylene (UHMWPE) cup, and Oka et al. [73] demon-strated the high wear resistance of zirconia against UHMWPE and thesuperiority of zirconia ceramics even over alumina ceramics in terms oflow wear and low friction.

2.3 Metallic Implants

Metals have been the primary materials in the past for damaged humanbones due to their superior mechanical properties [74], albeit dangerous ionsthat are released in vivo from these alloys. Although pure metals aresometimes used, alloys (metals containing two or more elements) frequentlyprovide improvement in material properties, such as strength and corrosionresistance. Three material groups dominate biomedical metals: 316 L stain-less steel, cobalt–chromium–molybdenum alloy, and pure titanium andtitanium alloys. The main considerations in selected metals and alloys forbiomedical applications are biocompatibility, appropriate mechanical prop-erties, corrosion resistance, and reasonable cost.

Figure 2.1 Medical-grade zirconia used as (a) femoral balls, (b) thumb and (c)dental implant [69]

Metallic Implants 15

2.3.1 Mechanical Properties

The mechanical properties of materials are of great importance whendesigning load-bearing orthopaedic and dental implants. Some mechanicalproperties of metallic biomaterials are listed in Table 2.3.

With a few exceptions, the high tensile and fatigue strength of metals,compared with ceramics and polymers, make them the material of choice forimplants that carry mechanical loads. The elastic moduli of the metals list inTable 2.3 are at least seven times greater than that of natural bone. Thismismatch of mechanical properties can cause ‘stress shielding’, a conditioncharacterised by bone resorption (loss of bone) in the vicinity of implants.The key problems associated with the use of these metallic femoral stems arethus the release of dangerous particles from wear debris, the detrimentaleffect on the bone remodelling process due to stress shielding and alsoloosening of the implant tissue interface. It has been shown that the degreeof stress shielding is directly related to the difference in stiffness of bone andimplant material [75]. Titanium alloys are favourable materials for ortho-paedic implants due to their good mechanical properties. Titanium, how-ever, does not bond directly to bone, resulting in loosening of the implant.Undesirable movements at the implant–tissue interface results in failurecracks of the implant.

2.3.2 Corrosion

The physiological environment is typically modelled as a 37 �C aqueoussolution, at pH 7.3, with dissolved gases (such as oxygen), electrolytes, cellsand proteins. Immersion of metals in this environment can lead to corrosion,which is deterioration and removal of the metal by chemical reactions.During the electrochemical process of corrosion, the metallic biomaterial canrelease ions, which may reduce the biocompatibility of materials andjeopardise the fate of implants. For example, the type and concentration ofreleased corrosion products can alter the functions of cells in the vicinity ofimplants as well as of cells as remote locations after transport of the

Table 2.3 Selected properties of metallic biomaterials

Young’s Yield Tensile Fatiguemodulus, strength, strength, limit,

Material E (GPa) sy (MPa) sUTS (MPa) send (MPa)

Stainless steel 190 221–1213 586–1351 241–820Cobalt–chromium 210–253 448–1606 655–1896 207–950

(Co–Cr) alloysTitanium (Ti) 110 485 760 300Ti–6Al–4V 116 896–1034 965–1103 620Cortical bone 15–30 30–70 70–150


corrosion by-products to distant sites inside the body. These circumstancesbecome stronger possibilities in the bodies of sick and elderly patients,who are the largest group of recipients of prostheses. Titanium and its alloys,as well as cobalt–chromium alloys, have more favourable corrosion resis-tance for long-term implant applications such as joint and dental prostheses.

2.4 Surface Modification of Biomaterials

2.4.1 Introduction

Biointegration is the ideal outcome expected of an artificial implant. Thisimplies that the phenomena that occur at the interface between the implantand host tissues do not induce any deleterious effects such as chronicinflammatory response or formation of unusual tissues. It is, therefore, ofparamount importance to design biomaterials used in implants with the bestsurface properties. Meanwhile, these biomaterials must process bulk proper-ties that meet other requirements, especially mechanical properties, in orderto function properly in a bioenvironment. As it is quite difficult to designbiomaterials fulfilling both needs, a common approach is to fabricatebiomaterials with adequate bulk properties followed by a special treatmentto enhance the surface properties. Hence surface modification of biomaterialsis becoming an increasingly popular method to improve device multi-functionality, tribological and mechanical properties, as well as biocompat-ibility of artificial devices while obviating the need for large expenses and along time to develop brand new materials [11]. Materials can be surface-modified by using biological or physicochemical methods. A few of the morewidely used physicochemical methods are briefly described here.

2.4.2 Radiation Grafting and Photografting [76]

Radiation is widely used in biomaterials science for surface modification,sterilization and to improve bulk properties. The use of gamma, ultraviolet(UV) and electron beam radiation has enabled the biomaterial scientist toperform bulk and surface modifications that improve the biological responseof materials and, subsequently, the performance of many medical devices.Radiation grafting has proven to be a simple technique that enables controlplacement of bioactive molecules on a polymer surface. Radiation-inducedcross-linking has allowed the tailoring of the composition and properties ofhydrogels to meet numerous biomedical applications. In addition, photo-cross-linking can serve to enhance the tribological properties of load-bearingcomponents of the total artificial knee and hip. UV radiation appears to havethe potential to facilitate in situ curing of adhesives, in situ production andmodification of devices and the generation of smart biomedical devices suchas biochips.

Surface Modification of Biomaterials 17

2.4.3 Plasma Surface Modification of Biomaterials

In the plasma surface modification process, glow discharge plasma iscreated by evacuating a vessel, usually quartz because of its inertness, andthen refilling it with a low-pressure gas. The gas is then energised usingtechniques such as radiofrequency energy, microwaves and alternatingcurrent or direct current. The energetic species in gas plasma include ions,electrons, radicals, metastables and photons in the short-wave ultraviolet(UV) range. These energy transfers are dissipated within the solid by avariety of chemical and physical processes, to result in surface modification.Plasma-based techniques combining the advantages of conventional plasmaand ion beam technologies are effective methods for medical implants withcomplex shapes [77]. In particular, modification of the surface energetics ofthe materials can improve the adhesion strength, surface and coatingproperties, and biocompatibility, to name just a few [78]. However, theapparatus used to produce plasma depositions can be expensive [65] and thechemistry produced on a surface can be ill-defined.

2.4.4 Ion Beam Processing

Biomaterial modification by ion beam processing is becoming popular forimproving medical device function, biocompatibility and as a new mutationbreeding method. Ion beam base processes, such as ion implantation and ionbeam assisted deposition (IBAD), can provide beneficial surface layers withdesirable properties without detrimentally affecting the bulk properties. Theion beam method injects accelerated ions with energies ranging from 101 to106 eV (1 eV ¼ 1.6 � 10�19 J) into the surface zone of a material in order toalter its surface properties. Important potential applications for biomaterialsinclude modification of hardness (wear), lubricity, toughness, corrosion,conductivity and bioreaction [79]. The primary advantage of ion implanta-tion is selective surface modification without detrimentally affecting bulkproperties. The drawbacks of this process are the high cost and the relativelyshallow depth of modification.

Ion beam assisted deposition (IBAD) is a vacuum deposition process thatcombines physical vapour deposition (PVD) with ion beam bombardment.The major feature of IBAD is bombardment with a certain energy (rangingfrom several hundred to several thousand eV) ion beam during the deposi-tion of coating. IBAD is used in hydroxyapatite coating preparation, dia-mond-like carbon (DLC) film, C–N film and other coatings. Another ionbeam process is ion beam texturing (IBT). IBT has the ability to createdesirable microfeatures and macrofeatures on the biomaterials to meet therequirement of biocompatibility in vivo.


2.4.5 Other Methods [65]

Silanization

Silane reactions can be used to modify hydroxylated or amine-rich surfaces.Since glass, silicon, germanium, alumina and quartz surfaces, as well asmany metal oxide surfaces, are all rich in hydroxyl groups, silanes areparticularly useful for modifying these materials.

Langmuir–Blodgett Deposition

The Langmuir–Blodgett (LB) deposition method covers a surface with ahighly ordered layer. Each of the molecules that assemble into this layercontains a polar head group and a nonpolar region.

Self-Assembled Monolayers

Self-assembled monolayers (SAMs) are surface coating films that sponta-neously form highly ordered structures (two-dimensional crystals) on spe-cific substrates.

Surface-Modifying Additives

Certain components can be added in low concentrations to a material duringfabrication and will spontaneously rise to and dominate the surface.

Conversion Coating

Conversion coatings modify the surface of a metal into a dense oxide-richlayer that imparts corrosion protection, enhanced adhesivity and sometimeslubricity to the metal.

Parylene Coating

Parylene (para-xylylene) coatings occupy a unique niche in the surfacemodification literature because of their frequent application and the goodquality of the thin-film coatings formed.

2.5 Laser Surface Modification of Biomaterials

2.5.1 Introduction

Lasers can rapidly and specifically induce surface changes in organic andinorganic materials. The advantages of using lasers for such modification arethe precise control of the frequency of the light, the wide range of frequen-cies available, the high energy density, the ability to focus and raster the

Laser Surface Modification of Biomaterials 19

light, the possibilities for using both heat and specific excitation to effectchange and the ability to pulse the source and control reaction time.Treatments are pulsed (100 nanoseconds to picoseconds pulse times) andcontinuous wave (CW), with interaction times often less than 1 microsecond.Laser-induced surface alterations include annealing, etching, deposition andpolymerisation.

2.5.2 Laser Patterning and Microfabrication

Laser patterning is based on the possibility of focusing an intense laser beamat certain spots on a surface, where the high beam intensity causes evapora-tion of the material. By this approach, pits can be produced down to �1 mm,in the size range of interest to match cell sizes. By controlled motion of thebeam (either by using clever optics or by sample motion), predesignedpatterns can be made. With a kinoform, it is possible to ‘laser-machine’multiple pits in a surface at once [80]. A new method combines micro-photolithographical techniques with laser excimer beam technology to createsurfaces with well-defined three-dimensional microdomains in order todelineate critical microscopic surface features governing material–cell inter-action [81]. Most laser-based patterning techniques use UV photoablation tomicromachine biological substrates in order to generate mesoscopic pat-terns, and arrays of viable cells are required to fabricate next-generationtissue-based sensing devices to build three-dimensional cellular structuresfor advanced tissue engineering and to separate selectively and culturedifferentially microorganisms for a variety of basic and applied researchapplications [82]. Recently, laser etching (or laser ablation) and microlitho-graphy have been adopted to achieve micrometer dimensions with highprecision in order to develop a large number of miniaturised systems for theanalysis of biological tissues. Excimer laser etching was used to microtexturea biocompatible substratum for high-contrast microscope cell analysis[83].

2.5.3 Pulsed Laser Deposition (PLD) of Biocompatible Ceramics

Pulsed laser deposition (PLD) is a new technique for the deposition of thinfilms of biocompatible ceramics [84]. Pulsed laser deposition is especiallywell suited to the deposition of bone-like ceramics (e.g. hydroxyapatite (HA)and calcium phosphates) on to metal, ceramic, semiconductor or polymersubstrates for potential application in medical implants, prosthetic devicesand biocompatible probes or sensors. The degree of control over filmcharacteristics offered by PLD exceeds that of other known depositiontechniques presently applied to production of thin films of biocompatibleceramics. It is anticipated that PLD will develop into the technique of choice


for the manufacture of implant or prosthetic devices comprising biocompa-tible films on structurally robust substrates. Pulsed laser ablation is a newmethod for deposition of thin layers of HA on to biomaterial surfaces.Differences in cell spreading were apparent which were correlated with thefluence used to deposit the HA [85]. Pulsed laser ablation and deposition ofbioactive glass [86, 87] have been performed and the plume and filmcompositions have been characterised. All the elements present in the targethave been found in the gaseous phase.

2.5.4 Matrix-Assisted Pulsed Laser Evaporationand MAPLE Direct Write

Two techniques, matrix-assisted pulsed-laser evaporation (MAPLE) andMAPLE direct write (MDW) were developed to deposit biomaterial thinfilms [88]. MAPLE involves dissolving or suspending the biomaterial in avolatile solvent, freezing the mixture to create a solid target and using a lowfluence pulsed laser to evaporate the target for deposition of the soluteinside a vacuum system. Using simple shadow masks, i.e. lines, dots andarrays, pattern features with length scales as small as 20 mm can be depositedusing multiple materials on different types of substrates. MAPLE utilises alow fluence pulsed UV laser and a frozen target consisting of a dilutemixture of the material to be deposited and a high vapour-pressure solvent.MDW uses pulsed laser radiation to directly transfer material from a ribbonto a substrate. Patterns with a spatial resolution of approximately 10 mm canbe written directly. Biomaterials ranging from polyethylene glycol to eukar-yotic cells (Chinese hamster ovaries) were deposited with no measurabledamage to their structures or genotype. Deposits of immobilized horse-radish peroxides (an enzyme) in the form of a polymer composite with aprotective coating, i.e. (polyurethane) retained their enzymatic functions. Adopamine electrochemical sensor was fabricated by MDW using a naturaltissues/graphite composite. The novelty of the MDW process is that theinteraction of the incident laser pulse with the coating on the ribbon cantransfer the micrometre-size powder, nanopowders and especially thechemical precursors to form a densely packed composite on the receivingsubstrate [89].

2.5.5 Other Laser Surface Treatments

Laser Surface Treatment for Improving Corrosion

It was found that the corrosion behaviour of NiTi samples was improved byexcimer laser surface melting [90]. Laser treatment improvement resistanceis explained by a combination of the homogenisation of the surface bymelting, the hardening due to N incorporation and the thickening of the

Laser Surface Modification of Biomaterials 21

oxide layer. Moreover, excimer laser surface treatment in air showed aremarkable improvement in pitting corrosion resistance for 316LS biogradestainless steel. The results show that excimer laser surface melting caneffectively eliminate carbides and second phases alike, while also servingthe function of homogenising the microstructure. N2 induced into the laser-treated surface could promote new precipitates and as a result lowered thecorrosion resistance of 316LS stainless steel [91] and Ti–6Al–4V alloy [92].

Laser Grafting

For the purpose of improved surface hydrophilicity and biocompatibility ofethylene–propylene rubber, 2-hydroxyethyl methacrylate (HEMA) andN-vinyl pyrrolidone (NVP) have been grafted on to the surface of thispolymer using a CO2 pulsed laser at different fluence (output power J/cm2)as the excitation source [17]. Moreover, ethylene–propylene rubber (EPR)based vulcanizates have been surface-grafted with acrylamide (AAm) andHEMA using a CO2 pulsed laser as the excitation source [18]. Surface hydro-philicity (measured by the water drop contact angle) increased for the graftedsamples and comparative results indicate that the adhesion of macrophagesto EPR samples modified with AAm and HEMA, with no respiratory burstand cellular damage, is significantly lower than their adhesion on unmodi-fied surfaces, which show an activated state of the attached macrophages.

Laser Treatment of Plasma Sprayed Hydroxyapatite Coatings [93]

The three requirements generally expected of biomaterials coating are:crystallinity, porosity and adhesion. Crystallinity is essential because amor-phous coatings are more resorbable. The study found that laser treatment ofplasma-sprayed coatings led to a wide range of microstructures. Theporosity of the coatings was reduced significantly. Nd:YAG (neodymium-doped yttrium aluminium gainet) laser (pulsed) treatment significantlychanges the characteristics of the plasma-sprayed coating microstructurein several ways. It ranges from a flat and smooth surface profile containingfine grains to an irregular surface comprising re-melted particles, sphericalpores and tracks. Laser treatment of plasma-sprayed HA coatings basicallygenerates a molten layer that rapidly solidifies.


3

Wettability in BiomaterialsScience and ModificationTechniques

3.1 Introduction

The wetting of a surface by a liquid and the ultimate extent of spreading thatliquid are very important aspects of practical surface chemistry. Even withall the new information of the last 20 years, however, there still remains agreat deal to learn about the mechanisms of movement of a liquid across asurface and the factors that govern such movement [12]. Biomaterialscientists have long sought a single, material-related parameter that effec-tively measures biocompatibility and might serve as a practical designguide. The theories of surface energy and wetting for such parameterspresent an attractive means to do this as surface properties are importantdeterminants of a biomaterial function [12]. The ability to control the surfacewettability of solid substrates is important in many situations. Varioussurface processes are used for modifying the surfaces of materials depend-ing on the actual material and the application. A number of laser-basedtechniques for altering the wettability characteristics of engineering materi-als have been investigated [13].

3.2 Wettability, Adhesion and Bonding: Theoretical Background

3.2.1 The Wetting Process

The term wetting in its most general sense is used to denote the displace-ment of air from a liquid or solid surface by water or any aqueous or molten


solution [94]. Wetting is fundamentally a thermodynamic process and thechanges in free energy that may occur determine whether or not wetting willhappen, at what rate it will proceed and how far it will progress against theexternal forces.

3.2.2 Contact Angle and Work of Adhesion

When a drop of liquid is in free space it is drawn into a spherical shape bythe tensile forces of its surface tension, which results from the attractive andrepulsive forces that exist between the molecules of the liquid. When such adrop of liquid is brought into contact with a flat solid surface, the final shapetaken by the drop, and thus whether it will wet the surface or not, dependsupon the relative magnitudes of the molecular forces that exist within theliquid (cohesive) and between the liquid and the solid (adhesive) [95]. Theindex of this effect is the contact angle, y, which the liquid subtends withthe solid. In practice, for wetting to occur the contact angle should be lessthan 90�. If the contact angle is greater than 90� then the liquid does not wetthe solid surface and no adhesion takes place [95]. Figure 3.1 shows aschematic view of a liquid droplet on a solid surface.

The contact angle is related to the solid and liquid surface energies, gsv andglv, and the solid–liquid interfacial energy, gsl, through the principal ofvirtual work expressed by Young’s equation:

gsv ¼ glv cos yþ gsl ð3:1Þ

If an equilibrium for the droplet of liquid melt shown in Figure 3.1 isestablished, then the relation of y to gsv, glv and gsl is described by therearranged Young’s equation:

cos y ¼ gsv � gsl

glv

ð3:2Þ

Clearly, to achieve wetting gsv should be large, while gsl and glv should besmall. Hence liquids of a lower surface tension will always spread over asolid surface of higher surface tension in order to reduce the total free energyof the system [96, 97]. Whether the drop of liquid spreads across the solid

γsv

γsl

γlv

θ

Figure 3.1 Schematic of the wetting of a solid medium by a liquid melt [95]

24 Wettability in Biomaterials Science

surface to wet the surface and provide a coating or remains as a finite dropwith an equilibrium angle is dependent upon the spreading coefficient S. Forspreading to occur spontaneously

S ¼ glvðcos y� 1Þ > 0 ð3:3Þ

The adhesion intensity of a liquid to a solid surface is known as the workof adhesion Wad and is given by the Young–Dupre equation

Wad ¼ glvð1 þ cos yÞ ð3:4Þ

Based on the nature of the attractive forces existing across the liquid–solidinterface, wetting can be classified into the two broad categories of physicalwetting and chemical wetting. In physical wetting the attractive energyrequired to wet a surface is provided by the reversible physical forces, suchas the van der Waals and dispersion forces. In chemical wetting adhesion isachieved as a result of reactions occurring between the mating surfaces,giving rise to chemical bonds [98].

3.2.3 Surface Energy and the Dispersive/Polar Characteristics

The intermolecular attraction that is responsible for surface energy, g, resultsfrom a variety of intermolecular forces whose contribution to the totalsurface energy is additive [99]. The majority of these forces are functionsof the particular chemical nature of a certain material, and as such the totalsurface energy comprises gp (polar or nondispersive interaction) and gd

(dispersive component, since van der Waals forces are present in all systemsregardless of their chemical nature); therefore, the surface energy of anysystem can be described by [100]

g ¼ gd þ gp ð3:5Þ

Similarly, Wad can be expressed as the sum of the different intermolecularforces that act at the interface [101]:

Wad ¼ Wdad þ W

pad ¼ 2 gd

svgdlv

� �12 þ 2 gp

svgplv

� �12 ð3:6Þ

If a liquid that has both dispersive and polar forces is in contact with asolid surface where the surface energy is due to dispersion forces only, thenthe relationship between the contact angle and the surface energies of theliquid and solid are given by [101, 102]

cos y ¼2 gd

svgdlv

� �12

glv

� 1 ð3:7Þ

Wettability, Adhesion and Bonding 25

However, by equating Equation (3.7) with Equation (3.4), the contact anglefor solid–liquid systems where both dispersion forces and polar forces arepresent can be related to the surface energies of the respective liquid andsolid by

cos y ¼2 gd

svgdlv

� �12 þ 2 gp

svgplv

� �12

glv

� 1 ð3:8Þ

Therefore, from Equation (3.8), one can estimate the dispersive componentof a solid substrate surface energy, gd

sv, by plotting the graph of cos y against(gd

lv)12=glv. This is shown in Figure 3.2 for a theoretical liquid system on any

solid substrate.Thus, according to Fowkes [100], the value of gd

sv is estimated by thegradient (¼ 2 [(gd

sv) ]12) of the line (----) that connects the origin (cos y ¼ �1)

with the intercept point (cos y against (gdlv)

12=glv) of the straight line (—)

correlating the data point with the abscissa at ‘cos y ¼ 1 In contrast’, it is notpossible to determine the value of the polar component of a solid substratesurface energy, gp

sv, directly from cos y against (gdlv)

12=glv. This is because the

intercept of the straight line (cos y against (gdlv)1=2=glv) is at 2(gp

svgplv)

12=glv, and

thus only refers to individual control liquids and not the control liquidsystem as a whole. However, it has been established that the entire amountof the surface energies due to dispersion forces either of the solids or theliquids are active in the wettability performance [100, 103]. As such, it ispossible to calculate the dispersive component of the work of adhesion, Wd

ad,by using only the relevant part of Equation (3.6). Thus

Wdad ¼ 2 gd

svgdlv

� �12 ð3:9Þ

−1

−0.5

0

0.5

1

0 0.04 0.08 0.12 0.16

Cos

θ

Figure 3.2 Plot of cos y against ðgdlvÞ

12=glv for a theoretical liquid system on any solid

substrate


If one plots a graph of Wad against Wdad for the solid substrate, then for each

particular liquid in a given system in contact with the solid surface, Wad,which was determined from Equation (3.4), can often be correlated withWd

ad, which was determined from Equation (3.9), by the straight linerelationship

Wad ¼ aWdad þ b ð3:10Þ

Therefore, for a solid substrate the constants a and b can be deducedrespectively by calculating the gradient of the best-fit straight line and byextrapolating the best-fit straight line to find the intercept point on theaxis. Also, if one plots a graph of (gp

lv) against (gdlv), then for the liquids in a

given liquids system, (gplv) can often be correlated with (gd

lv) by the straightline relationship

gplv

� �12 ¼ c gd

lv

� �12 þ d ð3:11Þ

Again, for a solid substrate the constants c and d can be deduced respectivelyby calculating the gradient of the best-fit straight line and extrapolating thebest-fit straight line to find the intercept point on the axis. By introducingEquation (3.10) into Equation (3.6) and rearranging, then

Wpad ¼ ða � 1ÞWd

ad þ b ð3:12Þ

or, alternatively,

gpsv

� �12 gp

lv

� �12 ¼ a � 1ð Þ gd

sv

� �12 gp

sv

� �12 þ b

2ð3:13Þ

By introducing Equation (3.11) into Equation (3.13) and differentiatingwith respect to (gd

lv)12, considering that (gd

sv)12 and (gp

lv)12 are constant, then the

following can be derived:

gpsv

� �12 ¼

gdsv

� �12 a � 1ð Þc

ð1:14Þ

Since gdsv for the solid substrate can be determined previously directly from

the plot of cos y against (gdlv)

12=glv, then it is possible to calculate gd

sv for thesolid substrate Equation (3.14) directly. By employing this approach it ispossible to determine, from y measurements and the control liquid surfaceenergy properties, the changes in the wettability characteristics of thematerials.

Wettability, Adhesion and Bonding 27

3.2.4 Physical Bonding

Physical bonding is essentially the effect that occurs when two perfectly flatsurfaces are brought together to atomic interaction distances, resulting inlocal atomic rearrangement and consequently adhesion. A typical exampleof physical bonding is that of van der Waals bonding. The energy differencebetween the specific surface energy of one material and that of the other isthe work of adhesion. The work of adhesion can yield a theoretical breakingstress similar to the strength of either of the materials used. Physical bondingprovides a useful guideline to the selection of materials that will bond welltogether.

3.2.5 Mechanical Bonding

Mechanical bonding basically refers to the interlocking microstructure ofrough surfaces to provide tensile strength and, in the case of shear, frictionalstrengthening. During the bonding process, the liquid or melt can flow withvarying degrees of ease into cavities and asperities; a ductile metal or glassmelt can conform to a rough solid substrate surface, or a vapour can depositin surface asperities. The solid substrate surface may be roughed by meansof acid or base chemical attack, grinding, grit or sand blasting, or lasertreatment to enhance mechanical bonding. The effectiveness of these differ-ent surface-roughening techniques is entirely dependent upon their opti-mum application as well as on the specific methods and materials beingused. In addition, chemical interaction between materials that are beingbonded can lead to mechanical bonding. Further, the increase in the surfacearea of a mechanically roughened surface can affect an increase in the levelof physical bonding.

3.2.6 Chemical Bonding

Considerable research into the various chemical mechanisms that can bepresent during the bonding process is in process. Although most of theresearch is qualitative or semi-quantitative in nature, it is providing a usefulbackground of chemical data that is contributing to a basic understanding ofthe principles of chemical bonding. A chemical bond is formed at aninterface when a balance of bond energies and a continuous electronicstructure are present across the interface for any two dissimilar phases.This structure occurs when a thermodynamically stable chemical equili-brium exists at the interface and is essentially achieved by chemical reactionsat the interface. Generally, equilibrium compositions (which can be deter-mined if an equilibrium phase diagram of the two phases being bonded isavailable) at the interface are attained at the reaction temperature veryrapidly.


3.3 Wettability in Biomaterial Science

From a historical perspective, Baier’s proposal that critical surface energycan be directly linked to biocompatibility is perhaps the most penetratingconcept among the few generalities offered to explain rules of biocompat-ibility. This theory, in its most general form, recognises that surface energymust control the way biologic fluids interact with materials and that thisinteraction, in turn, must primarily influence tissue and cell reactions. Asexamples, Baier pioneered the use of Zisman’s critical surface tension as anindicator of blood compatibility [104, 105] and bioadhesion [106, 107].Neumann et al. employed their ‘equation-of-state’ approach to calculateinterfacial tensions from contact-angle measurements that, in turn, wereused to predict cell adhesion [108] and thromboresistance [109]. Whereasconcepts such as these have served as useful general guidelines or ‘rules ofthumb’ for biomaterials design, each has fallen far short of being the desiredquantitative predictor of biocompatibility, particularly when applied toproteinaceous environments.

3.3.1 Biomaterial Interfaces [110]

Surface sensitivity is of critical importance in biomaterials surface sciencebecause only the uppermost layers are in direct physicochemical contactwith the biological environment; consequently, only the upper few mole-cular layers determine biocompatibility. Chemical events such as acid–basereactions, hydrogen bonding and ion exchange occur over atomic bond-length distances. Longer-range hydrophobic forces can extend up to about10 nm and are responsible for nonspecific adsorption, adhesion and surface-induced water structure within this zone. Thus the interaction of a materialwith the biological environment occurs at or through the narrow regiontermed the interface. Therefore, the interfacial chemistry of concern tobiomaterial scientists is determined by the material’s composition withinthe uppermost few nanometres.

3.3.2 Tensiometry

Tensiometry encompasses a broad range of related ‘wetting’ techniquesthat measure surface energy. These include the observation of contactangles, which is perhaps the most familiar and widely applied method.Tensiometric methods have singular potential in biomaterials surface sciencebased on the criteria of surface sensitivity, kind of analytical informationobtained and relevance of that information to biomedical problems. First,with respect to surface sensitivity, wetting measurements are sensitive onlyto the upper 0.5 nm or so of a surface [14, 15] and are therefore among themost surface-sensitive techniques available. Second, tensiometry directly

Wettability in Biomaterial Science 29

measures the fundamental energy at an interface that drives importantprocesses such as adsorption and adhesion. This kind of information mustbe particularly pertinent to biomaterial problems because of the overwhelm-ing importance of protein adsorption and cell/tissue adhesion. Third,wetting measurements can be made using proteinaceous saline solutionsthat are particularly relevant to biomedical applications. Special high-vacuum preparation techniques that might introduce experimental artefactsare not required.

3.3.3 Interfacial Biophysics

The physicochemical nature of biomaterial interfaces was considered, lead-ing to the conclusion that interfacial energy is a primary determinant ofbiocompatibility. A colloid science theory that quantifies interactions atsmall distances and has biomaterial applications is the so-called DerjaguimLandau Verwey Overbek (DLVO) theory [111, 112], shown in the Figure 3.3.DLVO illustrates the relationship between particle (cell) distance from thesurface and repulsive (electrostatic) and attractive (namely van der Waals)interaction energies. The basis of this theory is that attractive van der Waalspotentials and repulsive electrostatic forces are additive. Formulation ofthese interaction potentials can be quite detailed for each case, but thequalitative predictive aspects for a macroscopic substrate are quite straight-forward. From a physicochemical point of view, the kinetics of adhesion canbe described as long-range interactions [100] and short-range interactions(acid–base, hydrogen bonds) [113]. Others describe these forces as dispersiveand polar [114].

10

50 100

−10

G

GE

G

GVDW

D(Å)

Figure 3.3 Interaction energies between a particle (cell) approaching a solidsurface [65]


The total interaction energy (G, solid line) is composed of attractive vander Waals forces GVDWð Þ and repulsive electrostatic forces (GE). A secondaryminimum can be observed at approximately 100 A

�and a primary minimum

at a distance <5 A�. An energy barrier can be observed at approximately 40 A

�

from the surface:

G ¼ GE þ GVDW ð3:15Þ

where G is the sum of energy forces involved in Equation (3.15). GVDW is�Ha=(6D) (which represents the van der Waals interaction between aparticle with a radium a and a distance D to a flat substratum), H is theHamaker constant and GE (:) Zl � Z2 is the electrostatic interaction between aparticle and a solid plate (in which Z1 and Z2 are the zeta potentials of theparticle and the solid substratum).

Long-range interaction forces probably result in bringing a particle intothe secondary minimum at approximately 100 A

�from the surface, but only a

little energy is needed to remove the particle from the surface again. Short-range interactions between a particle and a solid can only take place atdistances <20 A

�.

3.3.4 Thermodynamic Concepts in Biomaterials Science

An important conceptual tool in modelling of biological response is surfacethermodynamics. A principal utility of thermodynamics is the ability tohandle multicomponent systems in a phenomenological manner. Surfacehydrophobicity, or wettability, is an important determinant of cell adhesion.It is related to surface free energy and is typically evaluated by the watercontact angle. Smaller water contact angles correspond to more hydrophilicsurfaces and higher surface free energies. In general, more hydro-philic substrates support cell adhesion and spreading to a greater extentthan hydrophobic materials, which have low surface free energies. Substra-tum surface free energy is related to cell spreading, as illustrated inFigure 3.4.

Poor spreading on hydrophobic substrata and good spreading on hydro-philic substrata can be observed in both the absence and presence of pre-adsorbed serum proteins. Specifically, work with wettability gradient sur-faces has shown that protein adsorption occurs to a greater extent onhydrophobic substrata than on hydrophilic substrata, that exchange of apre-adsorbed protein by another protein occurs more readily on hydrophilicsubstrata than on hydrophobic substrates, that adsorption-induced confor-mational changes are greater on hydrophobic substrates than on hydrophilicsubstrates and that cell adhesion reaches a maximum on moderatelyhydrophilic substrates that have contact angles around 60�.

Wettability in Biomaterial Science 31

Thermodynamically, the process of adhesion and spreading of cells from aliquid suspension on to a solid substrate can be described by [114]

�Fadh ¼ gcs � gcl � gsl ð3:16Þ

in which �Fadh is the interfacial free energy of adhesion, gcs is the cell–solidinterfacial free energy, gcl is the cell–liquid interfacial free energy and gsl isthe solid–liquid interfacial free energy. If �Fadh < 0, adhesion and spreadingare energetically favourable, while if �Fadh > 0, adhesion and spreading areunfavourable.

Figure 3.5 illustrates the relationship between �Fadh and the substratumsurface free energy (or wettability). It should be noted that hydrophobicsubstrata (gs < 40 erg=cm2) do not promote adhesion of fibroblasts.

1.0

0.5

00 10 50 100

Rel

ativ

ely

Cel

lSp

read

ing

sv (mJ/m2)γ

(Dotted line and solid line represents cell spreading in the absence of proteins and presence of proteins respectively)

Figure 3.4 Cell spreading as a function of substratum surface free energy [115]

∆Fadh

+10

−10

50 100

s[erg.cm−2]

FA

VO

RA

BLE

FO

R A

DH

ES

ION

UN

FA

VO

RA

BLE

FO

R A

DH

ES

ION

γ

Figure 3.5 Interfacial energy of adhesion �Fadhð Þ as a function of the substratumsurface energy gsvð Þ [114]


Extremely hydrophilic substrata do not promote adhesion either (e.g. high-energy methacrylates or hydrogels).

3.4 Current Methods of Wettability Modification

Various methods are used to improve the surface wettability of materialsand their adhesion to other materials. Hydrophilicity is a characteristic ofmaterials exhibiting an affinity for water. Hydrophilic literally means ‘water-loving’ and such materials readily adsorb water. Hydrophobic describesmaterials possessing characteristics that have the opposite response to waterinteraction compared to hydrophilic materials. Smaller water contact anglescorrespond to more hydrophilic surfaces and higher surface free energies.

3.4.1 Chemical Reactions

A chemical reaction can be used to leave the functional groups and changethe hydrophobic or hydrophilic behaviour of a material’s surface. Chemicalreactions were used to attach various hydrophilic functional groups and onehydrophobic group to cross-linked polystyrene resins. Those with hydro-philic groups had much better wettability by water [116]. Technicallyestablished treatments to modify the surface of silicone produced by plasmadischarge and following hydrogel coating yield to a less hydrophobicsurface [117, 118]. The wettability measurements of Si (111) surfaces treatedin aqueous HF and H2SiF6 solutions indicated that the HF-etched surface ishydrophobic y � 70�ð Þ, while the H2SiF6-treated surface is, if anything,hydrophilic y � 55�ð Þ.

3.4.2 Plasma Surface Modification

Plasma treatment can produce the extremes of both highly wettable andhighly nonwettable (or hydrophobic) surfaces. Plasma surface modificationcan be used to tailor surface energies. Hydrophilic and hydrophobic surfacescan be created on polymers through interaction with gas plasma. Usingoxygen to create hydroxyl functionality will increase the wettability of thesurface. This has been used to enhance the performance of a catheter by thecreation of a wettable surface on the polymer tubing. Most plasma cleaningoperations will result in a surface that is more wettable (hydrophilic) thanthe starting surface. The increase of the surface wettability is mainly due tothe grafting on to the NH3 plasma-treated polypropylene (PP) films ofnitrogen and oxygen polar groups [119]. The treatment of compositematerials [120] by means of a cold plasma allows both the increase of theirsurface wetting properties and the improvement of their mechanicalstrength in terms of adhesion between fibres and matrix.

Current Methods of Wettability Modification 33

3.4.3 Ion Beam Processing

Ion beam based processes, such as ion implantation and ion beam assisteddeposition (IBAD) can provide beneficial surface layers with desirableproperties without detrimentally affecting the bulk properties. Wettabilityof organic materials was changed by ion implantation [118]. Doping effectscan be obtained on electrical conductivity and wettability, induced by high-dose metal implantation. To date, ion implantation has been successfullyapplied in metal and polymer biomaterials. For the polymer substrate, aspecific functional chemical group is produced at the surface of the polymerto improve the surface wettability, anticalcific behaviour and biocompat-ibility of biomaterials. A new surface modification technique, ion assistedreaction (IAR) [121, 122], has been developed for improving wettability ofmaterials and enhancing adhesion to other materials. The contact angles ofwater drops with modified polymers were reduced more by Arþ ionirradiation with a flowing oxygen gas environment than without flowingoxygen gas.

3.4.4 Radiation Grafting

Radiation grafting can be used, for instance, for introducing polar groups inthe bulk or on the surface of nonpolar polymers, for increasing or reducingthe wettability of a polymer, for imparting a better compatibility of apolymer to a specific coating and the like. Radiation grafting and relatedmethods have been widely used for surface modification of biomaterials,and comprehensive review articles are available [123]. The radiation breakschemical bonds in the material to be grafted, forming free radicals, per-oxides, or other reactive species. These reactive surface groups are thenexposed to a monomer. The monomer reacts with the free radicals at thesurface and propagates at a free radical chain reaction, incorporating othermonomers into a surface-grafted polymer.

3.4.5 UV and Ozone

An ultraviolet (UV)–ozone oxidation process [124] is shown to be aneffective adhesion pre-treatment for polyethylene (PE) and polyetherether-ketone (PEEK). The data obtained indicate that the treatment gives con-siderable oxidation and improved wettability for PE and PEEK surfacetypes. It produces changes in surface oxygen chemistry and free energywhich improves surface polarity and wetting. Exposing PP to ozone in thepresence of UV light [125] is a simple and effective way of modifying itssurface to improve its wettability and adhesion. Atomic force microscopy(AFM) showed a dramatic change in the morphology and a clear increase inthe adhesion force resulting from the modification of a PP film by UV/ozoneexposure.


3.4.6 Corona Discharge

The corona discharge is a commonly used procedure to modify surfaces ofmaterials such as metals, glass or plastics. It is generally accepted that thistreatment modifies the substrate polymer surface properties like wettabilityand adhesion. Lee et al. [126] developed a method for preparing a wettabilitygradient on polymer surfaces using corona discharge treatment. The gra-dient was produced by treating the polymer sheets with the corona from aknife-type electrode whose power was changed gradually along the samplelength. The polymer surfaces oxidised gradually with increasing coronapower and the wettability gradient was created on the sample surfaces.

3.4.7 Electrowetting

The wettability of poly(ethylene terephthalate) (PET) films by water andaqueous solutions is increased by applying a voltage between the water anda rear electrode placed under the polymer film [127]. This electrowettingeffect can decrease contact angles by more than 30� under applied voltagesof 200 Veff. The electrowetting effect can help the solutions to wet the solidsurface.

3.5 Laser Wettability Characteristics Modification

At present, the processes available to engineers for the modification of amaterial’s wettability characteristics are invariably complex and conse-quently somewhat difficult to control. Lasers, on the other hand, can offerthe user not only an exceedingly high degree of process controllability butalso a great deal of process flexibility. There is a growing amount ofpublished work that testifies to the potential of lasers for altering the surfaceproperties of materials in order to improve their wettability characteristics.Laser radiation was found to effect significant changes in the wettabilitycharacteristics of materials.

3.5.1 Laser Surface Modification of Ceramic Materialsfor Improved Wettability

At present, very little published work exists regarding the effects of laserradiation on the wettability characteristics of ceramics materials. Indeed, thepublished work is predominantly concerned with the use of excimer laserradiation. Kappel [128] has shown that the texturing of ceramics (with anexcimer laser of 248 nm) can improve the adhesion strength by up to 20 %.Such an improvement is said to be due to the formation of raised micro-scopic protrusions over the surface. The wettability characteristics of the

Laser Wettability Characteristics Modification 35

selected ceramic materials, ceramic tile (SiO2/Al2O3-based), clay quarry tile(SiO2/Al2O3/Fe2O3-based) and Al2O3 and SiO2–TiO2 (crystalline), wereimproved after high-power diode laser (HPDL) treatment [129–131]. Thechanges in surface roughness, surface oxygen content and surface energyresulted in the enhancement of the wettability characteristics. Recently,Lawrence [132] conducted work on ceramics and metals to isolate each ofthese mechanisms, which permitted the magnitude of their influence to bequalitatively determined. Also, for ordinary Portland cement (OPC), surfaceenergy, by way of microstructural changes, was seen to influence a change inthe wettability characteristics, while surface roughness was found to play aminor role in inducing changes in the wettability characteristics.

3.5.2 Laser Surface Modification of Metallic Materialsfor Improved Wettability

It is recognised within the currently published work that laser irradiation ofmaterial surfaces can affect their wettability characteristics. Previously Heitzet al. [133], Henari and Blau [134] and Olfert et al. [135] have found thatexcimer laser treatment of metals results in improved coating adhesion. Theimprovements in adhesion were attributed to the fact that the excimer lasertreatment resulted in a smoother surface and as such enhanced the action ofwetting. It was demonstrated by using contact angles that CO2 lasertreatment was sufficient to produce a fully wettable surface [136]. Self-fluxing Fe–Cr–Ni–B–Si alloy powders with various Ni contents were laserclad on medium carbon steel substrates [137]. The wettability increases withincreasing Ni content in the cladding alloy. Good wettability of the claddingalloy on the substrate has a beneficial effect on crack prevention in thecladding layer, because it reduces the formation of pores in the claddinglayer, such pores usually being where stress concentrates and crackinginitiates. Lawrence and Li [138] compared the interaction of CO2, Nd:YAGand the HPDL; excimer laser radiation with the surface of the mild steelstudied was found to effect changes in the wettability characteristics of thematerial. It was observed that interaction of the mild steel with Nd:YAG andHPDL radiation brought about an improvement in the wettability character-istics of the steel. In contrast, interaction of the mild steel with CO2 andexcimer laser radiation resulted in a depreciation of the wettabilitycharacteristics of the steel.


4

CO2 Laser Modification of theWettability Characteristics ofMagnesia Partially StabilisedZirconia

In this chapter the modification of the wettability characteristics of an MgO–PSZbioceramic following CO2 laser irradiation is investigated. To study the wett-ability characteristics, contact angles between a set of control test liquids and thesurface of the untreated and CO2 laser treated MgO–PSZ were measured. Theinvestigation revealed that CO2 laser treatment occasioned a marked increase inthe wettability of the MgO–PSZ. The influential factors active in determining thewettability characteristics were analysed and the primary mechanism wasdeduced.

4.1 Introduction

During recent decades a vast number of materials have been tested aspotential biomaterials. Ceramic implants have aroused great interest becauseof their excellent compatibility with the physiological environment [71, 139].Partially stabilised zirconia (PSZ) has found wide usage in medical anddental surgery. Within orthopaedics, PSZ is commonly used as femoralhead, artificial knee, bone screws and plates, while in dentistry it is used tomanufacture dental implants, dental posts, crowns, brackets and inlays.Magnesia–partially stabilised zirconia (MgO–PSZ) was the very first zirconiaimplant approved by the US Food and Drug Administration (FDA).


However, although bioinert ceramics, including MgO–PSZ, have long beenappreciated for their biocompatibility, they have often clinically failed due tolack of direct bonding with bone, i.e. insufficient osseointegration [140]. Thewettability and interfacial interactions in a bioceramic–body–liquid systemwere investigated and low adherence was observed at the interface betweenthe ZrO2 and the body liquids [141]. For an implant to be successful, closeapposition of bone to the surface of the implant (osseointegration) isessential. From a historical perspective, Baier’s proposal that critical surfaceenergy can be directly linked to biocompatibility is perhaps the mostpenetrating concept among the few generalities offered to explain rules ofbiocompatibility [104–107]. This theory, in its most general form, recognisesthat surface energy must control the way biologic fluids interact withmaterials and that this interaction, in turn, must primarily influence tissueand cell reactions [12].

Since the applied energy of lasers can be placed precisely on a surface,lasers provide the contemporary scientist and engineer with a controllableand flexible tool for surface engineering. Kappel [128] has shown thatthe texturing of ceramic (with an excimer laser of 248 nm) can improvethe adhesion strength by up to 20 %. Such an improvement was said to bedue to the formation of raised microscopic protrusions over the surface.Lawrence and Li have demonstrated the practicability of employing differ-ent types of lasers to effect changes in the wettability characteristics ofceramics [130, 131, 142]. Yet, despite a growing amount of work conductedwith engineering materials, no work has been conducted to assess laserprocessing as a means for enhancing the wettability characteristics of abioinert ceramic.

4.2 Experimental Procedures

4.2.1 Material Specifications

The material investigated was a 4 % MgO–PSZ obtained in sheet form withdimensions of 50 � 50 � 2:15 mm3 (Goodfellow, Ltd). For experimentalpurposes, the sheet was cut into 10 blocks each of 50 � 12 � 2:15 mm3 witha cutting machine (Miniton; Struers, GmbH) using a diamond-rimmedcutting blade and used as received prior to CO2 laser treatment. The10 blocks were then divided into two groups of five samples, with thegroups being: untreated and CO2 laser treated. The main physical propertiesof the MgO–PSZ were a density of 5.74 g/cm3, a specific heat at 25 �C of 400–500 J/K kg and thermal conductivity at 20 �C of 1.5–2.5 W/mK. The mainmechanical properties were a compressive strength of 1500–2000 MPa, atensile modulus of 200 GPa and a Vickers hardness of 1200 kgf/mm2.

38 CO2 Laser Modification of the Wettability Characteristics

4.2.2 CO2 Laser Experimental Arrangement

A 3 kW CO2 laser emitting with a wavelength of 10.6mm was used in thisstudy. The laser produced a transverse electromagnetic multimode (TEM01)beam and was operated in the continuous wave (CW) mode. Figure 4.1illustrates the full arrangement with a five-axis workstation (TLC105;Trumpf, Ltd). It can be seen that a series of optical units was used to deliverthe CO2 laser beam to the workpiece through the laser head, which waspositioned by means of two linear axes (y and z axes) and two rotary axes(b and c axes). The defocused CO2 laser beam was traversed a single timeacross the surface of the MgO–PSZ samples placed on the stage using thex axis. The fumes produced were removed with an extraction system, whileO2 process gas with 2 bar pressure was used to shield the laser optics andassist the surface treatment.

4.2.3 Morphological, Chemical and Phase Analysis Procedures

The surface roughness of the samples was measured by a surface profil-ometer (Surface Tester SV-600; Mitutoyo, Inc.). The surface and cross-sectioncharacteristics of the untreated and CO2 laser treated MgO–PSZ sampleswere examined without etching using optical microscopy and scanningelectron microscopy (SEM)(JSM 5600LV; JEOL, Ltd). The samples weresectioned with a cutting machine (Miniton; Struers, GmbH) using adiamond-rimmed cutting blade in order to analyse the cross-section. Thesectioned samples were polished with 180, 400, 800 and 1000 grit SiCabrasive papers and then fine-polished using cloths and diamond pastesdown to 3mm. On account of the nonconductive nature of the MgO–PSZ, itwas necessary to coat the samples with Au in order to conduct the SEMexamination. In order to determine any changes that may have occurred inthe chemical make-up of the MgO–PSZ following CO2 laser treatment, thesamples were examined using energy dispersive X-ray (EDX) analysis to

Figure 4.1 Schematic diagram of the set-up for the CO2 laser treatment experiments

Experimental Procedures 39

obtain the chemical constituents and X-ray photoemission spectroscopy(XPS) to ascertain the elemental content. XPS (AXIS Ultra; Kratos, Inc.)analysis was performed on the sample surface using a spectrometer withmonochromatic Al Ka (1486.71 eV) X-ray radiation (15 kV and 10 mA) and ahemispherical electron energy analyser. The changes in the phase of theMgO–PSZ were detected by means of X-ray diffraction (XRD) (PW 1830;Philips Ltd) analysis with a Cu Ka source working at 30 kV and 20 mA acrossa range of 20–80 2-theta.

4.2.4 Wettability Characteristics Analysis Procedure

To investigate the effects of CO2 laser irradiation on the wetting and surfaceenergy characteristics of the MgO–PSZ, a set of sessile drop control experi-ments were carried out using glycerol, formamide, etheneglycol, polyglycolE-200 and polyglycol 15-200 with known values of glv, gd

lv and gplv [143],

detailed in Table 4.1. The contact angles, y, of the test liquids on theuntreated and CO2 laser treated MgO–PSZ were determined in atmosphericconditions at 25 �C using a sessile drop measuring machine (FTA

�125; First

Ten A�ngstroms, Inc.).

In order to estimate the influence of contaminant layers on the measure-ment results, the specimens of the untreated MgO–PSZ were cleaned withacetone in an ultrasonic bath for 2 hours, rinsed with distilled water severaltimes and dried in a vacuum oven at 90 �C for 12 hours. All of the fivecontrol test liquids were used to measure y on the cleaned samples. It wasobserved that the value of y on the cleaned samples was lower than on theas-received (not cleaned) samples by 1.5, 1.2, 1.0, 0.9 and 0.8� for glycerol,formamide, etheneglycol, polyglycol E-200 and polyglycol 15-200, respec-tively. It appears that any contaminants on the surface of the MgO–PSZ haveonly a slight influence on y. Because the contaminant is a minor factor activein the wettability characterisation, it is reasonable to preclude any cleaningpre-treatment for the practical application of CO2 laser treatment. Therefore,to explore the potential of CO2 laser treatment as an industrial andeconomical processing technique for altering the wettability characteristicsof MgO–PSZ, the work was conducted in a normal atmospheric environ-ment without pre-cleaning.

Table 4.1 Total surface energy (glv), dispersive (gdlv) and polar

(gplv) components for the selected test liquids [143]

Liquid glv (mJ/m2) gdlv (mJ/m2) gp

lv (mJ/m2)

Glycerol 64.0 34 30Formamide 58.3 32.3 26.0Etheneglycol 48.3 29.3 19.0Polyglycol E-200 43.5 28.2 15.3Polyglycol 15-200 36.6 26.0 10.6


Each measurement of y lasted for 3 minutes with profile photographs ofthe sessile drop being obtained every minute and a mean value beingsubsequently determined. After the test liquid drops for each liquid attachedand rested on the MgO–PSZ surface, the drops consistently reached anequilibrium state in around 6 seconds. Thereafter they remained motionlessand the magnitude of the y changed little with time. On average only �0.5�

deviation of the y for each test liquid was observed during the 3 minutes ofmeasuring when photographs were taken every minutes, indicating that theshape of the drop was stable in its equilibrated state. The difference betweenthe y value on the left-hand side and the right-hand side of the sessile drop isvery small, which generates a �0.36� deviation. In turn, the average totaldeviation for the y measurement was �0.86�.

4.3 The Effects of CO2 Laser Radiation on Wettability Characteristics

4.3.1 Contact Angle

When a drop of liquid is brought into contact with a flat solid surface, thefinal shape taken by the drop, and thus whether or not it will wet the surface,depends upon the relative magnitudes of the molecular forces that existwithin the liquid (cohesive) and between the liquid and the solid (adhesive)[95]. The index of this effect is y, the angle at which the liquid subtends withthe solid. In practice, for wetting to occur y should be less than 90�. If y isgreater than 90� then the liquid does not wet the solid surface and noadhesion takes place [95].

Optical micrographs of sessile drops of glycerol placed on an MgO–PSZsample before and after CO2 laser irradiation are shown in Figure 4.2. They values and other characteristics of the sessile drops are also provided.

Figure 4.2 Contact angels for glycerol on (a) the untreated MgO–PSZ and (b) CO2

laser treated MgO–PSZ (power density of 1.6 kW/cm2 and traverse speed of2000 mm/min)

The Effects of CO2 Laser Radiation 41

All the values of y obtained with the five control test liquids and the MgO–PSZ when untreated and CO2 laser treated at various power densities aregiven in Table 4.2. As Table 4.2 shows, with all the wetting control testliquids used the MgO–PSZ consistently experienced a significant reduc-tion in y as a result of interaction with the CO2 laser beam. Since the effectsof any contamination on the surface of the MgO–PSZ has been shown tohave a negligible effect on the value of y, then the sharp reduction in y isclearly the product of the CO2 laser treatment instead of pollutionelimination.

4.3.2 The Effect of Surface Oxygen Content

The observed increase in the wetting performance of the MgO–PSZ wouldhave certainly been influenced by the increase in the surface oxygen contentof the MgO–PSZ (as a result of CO2 laser treatment), as this is known toincrease the likelihood of wetting [13, 144–147]. Wetting is governed by thefirst atomic layers of the surface of a material. Therefore, in order todetermine accurately the element content of oxygen on the surface of theMgO–PSZ, it was necessary to examine the surface using XPS.

As can be seen from Figure 4.3, augmentation of the surface oxygencontent of the MgO–PSZ after interaction with the CO2 laser beam wasobserved. The values obtained show that the surface oxygen contentincreased from an initial value of 41.6 to 64.3 at %, while the y for glyceroldecreased from an initial value of 79� to 40�. It can be seen from Figure 4.3that an overall increase of 22.7 at % in the amount of surface oxygen on theCO2 laser treated MgO–PSZ occurred due to the oxidisation of the MgO–PSZsurface during melting and re-solidification. When the surface oxygencontent increased to a value of 64.3 at %, the y decreased to its minimum

Table 4.2 Mean values of contact angles formed between the untreated and CO2

laser treated MgO–PSZ for various power densities (traverse speed of 2000 mm/min)and the selected wetting control test liquids at 25 �C

Contact angle, y (deg)—————————————————————————————————————————————

CO2 laser treated (kW/cm2)——————————————————————————————

Test liquid Untreated 0.5 0.9 1.6 1.9 2.5

Glycerol 79 76 62 40 50 54Formamide 73 71 57 36 44 50Etheneglycol 61 60 48 29 38 41Polyglycol E-200 53 51 40 26 33 36Polyglycolycol 15-200 35 33 28 19 22 27


value. This indicates that oxygen enrichment of the CO2 laser treated MgO–PSZ surface was active in promoting wetting and adhesion. Such a finding issimilar to that of Song and Netravali [146, 147], who observed that surfaceoxygen content increased after laser treatment and, in turn, effected areduction in the y. However, when the surface oxygen content increasesbeyond 64.3 at %, y increased despite the further increase in surface oxygencontent. This suggests that other mechanisms are active and more dominant,thereby causing y to increase.

4.3.3 The Effect of Surface Roughness

By varying the CO2 laser operating parameters it was possible to obtain anarrow range of surface roughness values. The values of y for glycerol incontact with the CO2 laser treated MgO–PSZ samples were obtained atvarious points across this narrow range of surface roughness values. AfterCO2 laser treatment, the roughness of the MgO–PSZ surface increasedwith power density. This is likely to be the result of turbulent convection inthe melt pool caused by the TEM01 mode of the CO2 laser beam. Indeed,the power density distribution of the CO2 laser appears to be multimode. Itis evident that the CO2 laser beam does not display a maximum peak in themiddle of the beam; rather the peak is around the midpoint of the beam(see Figure 4.4). Based on the convective currents caused by a Gaussianbeam profile [148, 149] and the intensity distribution of a TEM01 beamprofile, a schematic diagram of the convection currents within the CO2

Figure 4.3 Relationship between the y (glycerol) for the untreated and CO2 lasertreated MgO–PSZ and surface oxygen content


laser treated track on the MgO–PSZ was generated and is given inFigure 4.4.

When a laser beam has a sufficient power density, convective currents willbe generated by the surface tension gradients resulting from the temperaturegradient at the free surface. The strong turbulent convection in the centre ofthe TEM01 laser beam will be very complex and will change with the powerdensity of the CO2 laser beam and the diameter of the beam.

A model similar to that for heterogeneous solid surfaces can be developedin order to account for surface irregularities, being given by Wenzel’sequation [150]

ra gsv � gslð Þ ¼ glvcos yW ð4:1Þ

where ra is the roughness factor defined as the ratio of the true to apparentsurface areas and yW is the contact angle for the wetting of a rough surface. Ifra is large, i.e. the surface is rough, then cos yW is large and yW will decreaseif y was originally less than 90�. Therefore, when ra increases yW decreases ify was originally less than 90�. In the instance when y is originally greaterthan 90�, the situation is vice versa.

As one can see from Figure 4.5, the surface of the untreated MgO–PSZsample was very smooth, with an average surface roughness value (Ra) ofaround 0.295 mm. At this Ra value, y was measured as 79�. As shown inFigure 4.5, the CO2 laser treatment consistently generated a surface that wasrougher, to varying degrees, and gave rise to values of y that were alwaysmuch lower than on the untreated MgO–PSZ surface. Having said that,Figure 4.5 presents some very interesting findings that give tremendousinsight into the mechanisms and effects of the CO2 laser surface treatment ofthe MgO–PSZ. First, with only a relatively small increase in the surface

Figure 4.4 Schematic diagram of the convection currents generated within the CO2

laser meltpool


roughness of less than 0.5 mm, the minimum y value of 40� occurred at asurface roughness of 0.717mm. Although this observation is in line withEquation (4.1), which states that an increase in surface roughness ought toeffect a decrease in y when it is below 90�, this increase in surface roughnessof the MgO–PSZ occasioned after CO2 laser treatment is extremely small andis unquestionably not proportional to the considerable reduction in y. Anexplanation for this observation lies in the fact that the CO2 laser treatmentsimultaneously effects changes in many other surface properties of theMgO–PSZ besides surface roughness. Second, it can be seen fromFigure 4.5 that even with significant increases in Ra, beyond 0.717mm thevalue of y, rather than decreasing, actually experienced a slight increasefrom 40 to 54�. This observation not only patently contradicts Equation (4.1)but a previous finding [151] in which higher surface roughness corre-sponded to a lower value of y on ZrO2. Again, this is further evidencethat aspects of the surface properties of the MgO–PSZ other than surfaceroughness are active.

4.3.4 The Effects of Solidified Microstructures and Surface Meltingon Wettability Characteristics

Exposure of the MgO–PSZ to CO2 laser irradiation resulted in rapid heatingof the surface, for most materials typically 103–105 K/s [152], which subse-quently led to the melting and re-solidification of the MgO–PSZ surface.

Figure 4.5 Relationship between y (glycerol) for the untreated and CO2 laser treatedMgO–PSZ and surface roughness


Indeed, after the CO2 laser treatment a distinctive microstructure wasgenerated on the MgO–PSZ surface, as can be seen in Figure 4.6. Suchmicrostructures are indicative of the incidence of rapid solidification.

The various degrees of rapid solidification generated by the different laserpower densities employed resulted in the diverse range of microstructuresgenerated on the MgO–PSZ surface, as shown in Figure 4.7. In order to

Figure 4.6 Typical SEM surface images of the MgO–PSZ (a) before and (b) after CO2

laser treatment at power density of 1.6 kW/cm2

Figure 4.7 Relationship between y (glycerol) on the untreated and CO2 laser treatedMgO–PSZ and power density and solidified microstructure


simplify the analysis, the structures were defined according to the mainstructures on the MgO–PSZ after CO2 laser treatment. It can be seen fromFigure 4.7 that the surface structures of the MgO–PSZ play a significant rolein determining y. The mere re-ordering of the crystals that occurred at lowpower densities appears to have only a slight effect on y, reducing it from79 to 76�. At around 0.9 kW/cm2, y decreases markedly from 77 to 62�, with ahexagonal structure being generated on the MgO–PSZ. With furtherincreases in power density to around 1.6 kW/cm2, y decreases from 62 to40�, with some cells beginning to form in the region due to the hightemperature peak caused by the intensity distribution of the CO2 laserbeam. One reason for this sharp reduction in y may be related to the onsetof melting at this power density, which therefore implies that melting isan essential prerequisite for a significant reduction in y. Further increasesin power density resulted initially in an increase in y from 40 to 50�, whichcorresponds to the formation of uniform cell structures, and then a furtherincrease to 54�, at which point dendrites and coral structures were observed.Indeed, work conducted by Zhang, Yue and Man [153] found that consider-able improvement in the bond strength of an Si3N4 ceramic could be realisedonly when excimer laser treatment of a structural alloy steel (SAE 4340)resulted in surface melting. Similarly, Lawrence [154] observed a sharpreduction in y at the point of melting for an Al2O3/SiO2-based oxidecompound after high-power diode laser (HPDL) treatment.

4.4 Surface Energy and Its Component Parts

It is possible to adequately estimate the dispersive component of theMgO–PSZ surface energy, gd

sv, by plotting the graph of cos y against(gd

lv)12=glv according to Equation (3.8). Thus, according to Fowkes [100], the

value of gdsv is estimated by the gradient (¼ 2(gd

sv)12) of the line that connects

the origin (cos y ¼ �1) with the intercept point of the straight line (cos yagainst (gd

lv)12=glv) correlating the data point with the abscissa at cos y ¼ 1. The

values of gdsv for the untreated and CO2 laser treated MgO–PSZ could be

calculated from Figure 4.8.Comparing the ordinate intercept points of the untreated and CO2 laser

treated MgO–PSZ–liquid systems in Figure 4.8, it can be seen clearly that forthe untreated and CO2 laser treated MgO–PSZ at lower power density of0.5 kW/cm2, the best-fit straight line intercepts the ordinate closer to theorigin. This is noteworthy since the intercept of the ordinate close to theorigin is characteristic of the dominance of dispersion forces acting onthe MgO–PSZ material–liquid interfaces of the untreated and low CO2

laser power density treated sample, resulting in poor adhesion [100, 101].On the other hand, the best-fit straight line of the samples treated at higher

Surface Energy and Its Component Parts 47

CO2 laser power densities (>0.5 kW/cm2) intercept the ordinate consider-ably high above the origin. The highest intercept point is found for thesample CO2 laser treated with 1.6 kW/cm2 power density. An interception ofthe ordinate above the origin is indicative of the action of polar forcesacross the interface, in addition to dispersion forces, and hence improvedwettability and adhesion is promoted [100, 101]. Furthermore, becausenone of the best-fit straight lines intercept below the origin, it can besaid that the development of an equilibrium film pressure of adsorbedvapour on the MgO–PSZ surface (untreated and CO2 laser treated) did notoccur [101].

It is not possible to determine the gpsv value of the MgO–PSZ directly from

Figure 4.8. This is because the intercept of the straight line (cos y against(gd

lv)12=glv) is at 2(gp

sv)(gdlv)

12=glv, and so only refers to individual control liquids

and not the control liquid system as a whole. Still, it has been establishedthat the entire amount of the surface energies due to dispersion forces eitherof the solids or the liquids are active in the wettability performance [100,103]. As such, it is possible to calculate the dispersive component of the workof adhesion, Wd

ad, by Equation (3.9).Table 4.3 shows the values of Wad calculated using Equation (3.4) and the

values of Wdad calculated using Equation (3.9) for both the untreated and CO2

laser treated MgO–PSZ with various power densities. Figures 4.9 and 4.10show the best-fit straight line plots of Wad against Wd

ad for the MgO–PSZwhen it is both untreated and CO2 laser treated. From the plots of Wad

against Wdad one can see that the experimental results reveal that for each

0.00 0.03 0.06 0.09 0.12 0.15 0.18

–0.9

–0.6

–0.3

0.0

0.3

0.6

0.9

cos

Untreated 0.5 kW/cm2

0.9 kW/cm2

1.6 kW/cm2

1.9 kW/cm2

2.5 kW/cm2

( )γ γl lvd

v1 2/

/

θ

Figure 4.8 Plot of cos y against (gdlv)

12=glv for the MgO–PSZ in contact with the wetting

control test liquids, before and after CO2 laser treatment with various power densities


particular control liquid in contact with both the untreated and CO2 lasertreated MgO–PSZ surfaces, Wad, determined from Equation (3.4), can becorrelated with Wd

ad, determined from Equation (3.9), by the straight linerelationship presented by Equation (3.10). Consequently, the values of a forvarious CO2 laser parameters shown in Table 4.3 were determined by thebest-fit straight line of Wad against Wd

ad.As Figure 4.11 shows, a linear relationship between gd

lv and gplv of the

control test liquids’ surface energies was observed which satisfiedEquation (4.2) and thereby allowed the constant, c, to be calculated. Sincegd

sv has already been determined for the untreated and CO2 laser treated

Table 4.3 Values of Wad and Wdad for the control test liquids and the determined

constant, a, from the plots of Wad against Wdad for untreated (UT) and CO2 laser treated

MgO–PSZ with various power densities

Power Work of adhesion Wadð Þ Dispersive work of adhesion Wdad

� �

density(kW/cm2) a Glyc Form Ethel P1 P2 Glyc Form Ethel P1 P2

UT 2.41 76.2 75.2 71.5 69.6 66.6 76.1 74.2 70.7 69.3 66.60.5 2.37 79.4 77.5 72.5 70.9 67.3 77.2 75.2 71.6 70.3 67.50.9 3.03 92.8 89.8 80.7 77.0 68.8 77.7 75.7 72.1 70.7 67.91.6 4.25 113.3 105.5 90.3 82.7 71.4 80.9 78.8 75.1 73.7 70.81.9 3.42 105.0 98.5 86.5 80.0 70.6 80.5 78.4 74.7 73.3 70.42.5 3.17 101.8 95.6 84.5 78.7 70.3 77.7 75.2 71.6 70.3 67.5

Note: Glyc, glycerol; Form, formamide; Ethel, Etheneglycol; P1, polyglycol e-200; P2, polyglycol 15-200.

Figure 4.9 Plot of Wad against W dad for the untreated MgO–PSZ


MgO–PSZ from Figure 4.8, then it is possible to calculate gpsv for untreated

and CO2 laser treated MgO–PSZ using Equation (3.14). The values for gsv, gdsv

and gpsv of the untreated and CO2 laser treated MgO–PSZ are given in

Figure 4.12. As one can see from Figure 4.12, CO2 laser treatment of the

Figure 4.10 Plot of Wad against W dad for the CO2 laser treated MgO–PSZ (1.6 kW/cm2

and 2000 mm/min)

0 1 2 3 4 5 6 7 80

1

2

3

4

5

6

7

8

( )1 / 2γ

lvd

()1

/ 2γ lvd

Figure 4.11 Plot of (gplv)

12 against (gd

lv)12 for the control test liquids


surface of the MgO–PSZ leads to an overall increase in the gsv, while, moreimportantly, also significantly increasing gp

sv. The increase in gsv of the MgO–PSZ was primarily attributed to the increased gp

sv value, since gdsv was almost

similar for all the samples. The increase, in particular the increase in gpsv, had

a positive effect upon the action of wetting and adhesion [103] as primarilyboth dispersion and polar forces were active to a greater extent [100, 155].The changes in the surface energy are thought to be due to the fact that theCO2 laser treatment of the MgO–PSZ results in the melting of the surface, atransition that is known to cause an increase in gp

sv [141] and hence animprovement in the wettability characteristics.

Moreover, as can be seen from Figure 4.12, gsv and gpsv changed depending

on the microstructures obtained at different CO2 laser parameters. When ahexagonal microstructure was obtained on the surface of the MgO–PSZ, amarked increase in gsv and gp

sv occurred, as is evident from Figure 4.12. Whenthe MgO–PSZ was treated with a relatively medium CO2 laser powerdensity, cell formation on the MgO–PSZ surface was induced and themaximum value of gsv and gp

sv was achieved. Figure 4.12 suggests that theonset of melting initiated the cell formation. With relatively high CO2 laserpower densities, coral and dendritic microstructures were apparent on the

Figure 4.12 Relationship between surface energy (gdsv, gp

sv and gsv) for the CO2 lasertreated MgO–PSZ and power density


surface of the MgO–PSZ. The formation of these microstructures wasaccompanied by a reduction in gsv and gp

sv from the maximum value.

4.5 Identification of the Predominant Mechanisms Activein Determining Wettability Characteristics

Regardless of the CO2 laser power density employed, noticeable differencesin surface roughness, microstructure, surface oxygen content and surfaceenergy of the MgO–PSZ were occasioned simultaneously. All these surfaceproperties will have been a factor in determining the wettability character-istics of the MgO–PSZ. First, CO2 laser treatment of the surface of the MgO–PSZ generated a rougher surface and thereby reduced y. Second, the increasein the surface oxygen content of the MgO–PSZ resulting from CO2 lasertreatment will be influential in the promotion of wetting, because an increasein surface oxygen content inherently effects a decrease in y and vice versa.Lastly, an increase in gp

sv resulting from the melting and re-solidification ofthe surface of the MgO–PSZ occurred. This naturally created a differentmicrostructure that quite possibly improved the action of wetting andadhesion. Still, the foregoing sections have revealed that each of thesefactors did not play an equal role in governing the wettability character-istics of the MgO–PSZ. It is, therefore, essential to identify the effect ofeach factor and find the predominant mechanism active in governing thewettability characteristics of the MgO–PSZ. Figure 4.13 shows the relation-ship between the wettability characteristics (denoted by cos y for glycerol)and the influential factors of surface roughness, gp

sv and surface oxygencontent.

As is evident from Figure 4.13, the rougher surface of the modified samplehas a higher value of cos y than the smooth, untreated sample. Nevertheless,the change in cos y is not proportional to the alteration in the surfaceroughness, with cos y increasing sharply up to a surface roughness of0.717 mm, then decreasing despite a considerable increase in surface rough-ness. This signifies that other mechanisms, namely the surface oxygencontent and gp

sv, may play a more predominant role in influencing thewettability characteristics of the CO2 laser treated MgO–PSZ than surfaceroughness.

It is apparent from Figure 4.13 that cos y increased with increasing surfaceoxygen content below 63.4 at %. This is in accord with the established theoryfor the relationship between the surface oxygen content of a material andtheir propensity for wetting. However, when the value of the surface oxygencontent exceeded 63.4 at %, cos y decreased, indicating that the surfaceoxygen content is not a major factor active in changing the wettability of theMgO–PSZ.


A clear relationship between the value of cos y and gpsv is observed in

Figures 4.13. Figure 4.14 reveals that an increase in gpsv will in turn cause a

rise in cos y. From this it is evident that gpsv influences the wettability

characteristics of the MgO–PSZ. Indeed, it was found by Lawrence [154]that surface energy was the most predominant factor governing the wettingcharacteristics of a SiO2/Al2O3-based ceramic following irradiation with anHPDL.

From the above discussion it is not clear whether the surface roughness,the surface energy (by way of microstructural changes) or the surfaceoxygen content alone, or a combination thereof, are the principal factorsinfluencing the observed changes in the wettability characteristics of theMgO–PSZ after CO2 laser surface treatment. Therefore, several stages offine grinding were used to isolate the various influential factors detailedabove and thus analyse and establish qualitatively the effect each one had onthe wettability characteristics of the MgO–PSZ. In the first stage, the surfacesof the untreated MgO–PSZ and CO2 laser treated MgO–PSZ (1.6 kW/cm2)with the largest change in y were ground with grinding paper (180 grit SiC)

Figure 4.13 Relationship between cos y for glycerol and the surface roughness, thesurface oxygen content, gp

sv and microstructures of the untreated and CO2 lasertreated MgO–PSZ

Identification of the Predominant Mechanisms Active 53

for 3 minutes, while still retaining the CO2 laser treated microstructure. Inthis way it was possible to investigate the effects of the surface oxygencontent, which exists within the first atomic layers of the material. In orderto evaluate the influence of the CO2 laser induced microstructure, anintermediate grinding stage using grinding paper (400 grit SiC) for 3 minuteswas used to remove the microstructure. In the final stage, both the untreatedand the CO2 laser treated samples were ground down further with grindingpaper (800 grit SiC) for 3 minutes to study the effect of surface roughness.The observed changes to surface roughness, O2 content and y (glycerol)effected by these steps are given in Table 4.4.

After the first grinding stage, a large difference in y for glycerol wasobserved between the untreated and the CO2 laser treated samples, with y

Table 4.4 The contact angle (for glycerol), surface roughness and surface oxygencontent of the untreated and CO2 laser treated MgO–PSZ following the fine grindingstages

Untreated CO2 Laser Treated———————————————————— ———————————————

Polishing steps Ra (mm) O2 (at %) y (deg) Ra (mm) O2 (at %) y (deg)

Unpolished 0.30 41.6 79.1 0.72 64.3 39.4180 grit SiC 0.22 41.5 81.9 1.90 42.0 43.7400 grit SiC 0.08 41.7 82.3 1.46 41.8 77.3800 grit SiC 0.06 41.8 82.3 1.23 41.7 77.9

Figure 4.14 Relationship between cos y for glycerol and gpsv on the untreated and

CO2 laser treated MgO–PSZ


increasing slightly from 39.4 to 43.7� for the CO2 laser treated sample, whiley for the untreated sample increased from 79 to 82� (Table 4.4). From thisobservation it is reasonable to suggest that the retained CO2 laser inducedmicrostructure in the CO2 laser treated MgO–PSZ could be the mechanismresponsible for the large different y from the untreated sample. Furthermore,the surface oxygen content of the CO2 laser treated sample was found tohave reduced from 64.3 to 42.0 at %, a level similar to that of the untreatedsample, 42.2 at %. It is, therefore, quite possible that the decreased surfaceoxygen content could be the factor influencing the general increase in y. It isinteresting to note that although the surface roughness of the CO2 lasertreated sample increased from 0.72 to 1.90mm, the value of y increased, thusimplying that surface roughness does not have as great an influence on thewetting characteristics of the MgO–PSZ as that of the surface oxygencontent.

This proposition was borne out somewhat when the samples were groundfurther. In this subsequent stage, the CO2 laser induced microstructureand heat affected zone (HAZ) were removed from the CO2 laser treatedsample. The surface oxygen content on the untreated and CO2 laser treatedsamples were practically the same as the original untreated value of 41.6at %. Significantly, y for the CO2 treated samples was 77.3�, a value close tothe original untreated value of 79.1�. Basically, the removal of the CO2 laserinduced microstructure alone appears to have brought about an increase in yto around the original level, since the surface oxygen content was almost thesame value in both ground stages. Such findings reveal unequivocally thatthe microstructure is by far the predominant mechanism governingthe wettability characteristics of the MgO–PSZ. The effect of the surfaceroughness was studied through a further step. This step generated asmoother surface by reducing the Ra from 1.72 to 1.40 mm. For the CO2

laser treated sample, this step caused y to change slightly from 77.3 to 77.9�.For the untreated sample, the final ground stage caused Ra to changeconsiderably, reducing from 0.30 to 0.08 mm. However, as seen fromTable 4.4, y only increased very marginally from 79.1 to 82.3�. Despite themagnitude change in surface roughness, y had varied only slightly in thisrange of surface roughness. Such a finding indicates that changes insurface roughness are insignificant, as reflected by the corresponding changein y. It is therefore reasonable to assume that the surface energy differencebrought about by microstructural changes is the primary influentialfactor governing changes in y and, in turn, the wettability characteristicsof the MgO–PSZ. What is more, the surface oxygen content was alsofound to influence changes in the wettability characteristics of the MgO–PSZ but to a much lesser extent, while surface roughness was shown to playa very minor role in inducing changes in the wettability characteristics of theMgO–PSZ.

Identification of the Predominant Mechanisms Active 55

4.6 The Role Played by Microstructures in Terms of Crystal Sizeand Phase in Effecting Surface Energy Changes

The surface energy has been identified as the main mechanism governingthe modification of wettability characteristics of the MgO–PSZ and is variedwith the surface microstructure. It is believed that the changes in surfaceenergy of the MgO–PSZ were attributed to the change in microstructures inthe form of crystal sizes and phase changes.

4.6.1 The Role of Crystal Size on Surface Energy

After CO2 laser radiation, the modified surface of the MgO–PSZ exhibits atypical microstructure of rapid solidification. As the XRD patterns of theuntreated and CO2 laser treated MgO–PSZ given in Figure 4.15 show, after

Figure 4.15 XRD analysis of the MgO–PSZ surface (a) before and (b) after CO2 lasertreatment (1.6 kW/cm2) (c ¼ cubic, t ¼ tetragonal, m ¼ monoclinic)


CO2 laser treatment a distinct phase change took place. The diffractionpatterns of 2-theta between 28 and 32� are of particular relevance (shown inFigure 4.15) and should be emphasised [156]. The peak at 29� belongs to theoverlap diffraction of the (111) cubic phase and (101) tetragonal phase(denoted as c(111) and t(101)). The peaks of the cubic and tetragonal phasesare all overlapped with others for the untreated and CO2 laser treatedsamples (see Figure 4.15). The peaks at 28.2 and 31.5� belong to the (111)monoclinic phase (m (111)). An increase in the peak at 29� signifies theincrease in the tetragonal phase.

After CO2 laser treatment, the peak at 29�, which shifts to 30.5�, was seento rise while the peaks at 28.2 and 31.5� disappeared (see Figure 4.15),indicating that the tetragonal phase was increasing while the monoclinicphase was decreasing in the MgO–PSZ surface. As shown in Figure 4.16, therelative intensity of the tetragonal phase on the MgO–PSZ treated with aCO2 laser power density below 1.6 kW/cm2 increased with the powerdensity and then decreased slightly as the power density increased beyond1.6 kW/cm2. In contrast, the monoclinic phase decreased as the tetragonalphase increased.

Figure 4.16 The XRD pattern of the MgO–PSZ with various power densities between 27and 32� 2-theta angle

The Role Played by Microstructures in Terms of Crystal Size and Phase 57

According to the equilibrium phase diagram for MgO–ZrO2 [157], themonoclinic phase is stable below 1240 �C and the tetragonal phase is stablebetween 1240 and 1400 �C. Above 1400 �C, tetragonal and cubic phasescoexist, with increasing temperature, tetragonal transforms to the cubicphase. From 2370�C to the melting temperature (2600 �C) the stable phaseis a cubic structure. With the increase in CO2 laser power density, the surfacetemperature of the MgO–PSZ will increase and create the phase transforma-tion on the surface. As is evident from Figure 4.16, the decrease in themonoclinic phase is obvious while the tetragonal phase increases greatly

when the power density is above 0.9 kW/cm2. Indeed, pyrometer readingsin the range 250–2000 �C showed that the surface temperature of theMgO–PSZ was above the 2000 �C when the CO2 laser power density was1.6 kW/cm2. Since the monoclinic–tetragonal transformation temperature is1240 �C, then under these conditions the monoclinic phase transformed tothe tetragonal phase and the tetragonal phase reached the highest density. Ithas been reported that the CO2 laser cladding of ZrO2 composite coatingshowed a higher tetragonal phase in the laser-clad ZrO2 ceramic layer thanthat in the original ZrO2 powder [158]. Moreover, the plasma sprayed 8 wt %Y2O3–PSZ coating generated the metastable tetragonal phase (peak at 29�) inthe as-sprayed condition after Nd:YAG laser treatment [159].

Moreover, the intensity of the peak at 30.5� assigned as the (111) plane ofthe tetragonal lattice overwhelms the others, which indicated the tendencyof crystal orientation. The crystal size in the direction perpendicular to thehkl plane, Dhkl, is expressed by the Scherrer equation [160]

Dhkl ¼Kl

b cos að4:2Þ

where l is the wavelength of the X-ray (1.54056 A�

for the Ka line of Cu in thisexperiment), a is the Bragg angle, b is the expansion of the XRD peak causedby the crystal size and K is the Scherrer constant. We take the full-width half-maximum (FWHM) of the peak at (111) in the XRD analysis as b, thecrystallite size as Dhkl and K as 0.91. The crystallite sizes in the untreated andCO2 laser treated MgO–PSZ at various power densities are listed inTable 4.5.

As shown in Table 4.5, the crystal sizes in the MgO–PSZ following CO2

laser radiation are consistently larger than the untreated sample, implyingthat the crystal grew after CO2 laser radiation. A high heat input from a laserbeam facilitates surface localized melting at a very high efficiency. Its abilityto maintain a cold substrate while melting a thin surface layer of materialresults in rapid quenching of the molten layer once the beam is removed.Thermal gradients at the liquid–solid interface layer are very steep andcause crystal growth taking place along the thermal gradient. The power


density of the CO2 laser treatment has a significant effect on the crystal size.Generally, the crystal sizes in the MgO–PSZ increase with the increasingpower density, with the largest crystal size of around 103.1 nm occurring at apower density of 1.6 kW/cm2. Furthermore, it was found that the crystal sizevaries similarly as surface energy with power density, as shown inFigure 4.17, signifying that the crystal size is correlated with the surfaceenergy of the MgO–PSZ.

According to the classical theory of nucleation and growth in solids [161],a crystallite nucleate in the form of critical embryos grows by accretingatoms from the surroundings. Assuming the embryo has a spherical shape

Table 4.5 Crystallite size calculated from the XRD of theuntreated and CO2 laser treated MgO–PSZ on the basis of theScherrer equation

Power Density(kW/cm2) FWHM a (deg) D (nm)

Untreated 4.187 15.26 34.70.5 3.489 15.25 41.50.9 2.318 15.22 65.41.6 1.396 15.20 103.11.9 2.415 15.24 60.12.5 2.268 15.18 64.0

Figure 4.17 Relationship between surface energy and crystal size of the CO2 lasertreated MgO–PSZ and power densities

The Role Played by Microstructures in Terms of Crystal Size and Phase 59

of radius, r, the free surface energy of an unstrained spherical particle, G, isexpressed as

GðrÞ ¼ 43pr3GV þ 4pr2gt ð4:3Þ

where G is the free energy of an spherical particle, GV is the free energy perunit volume of a crystal and gt is the surface energy of the tetragonal crystalas Figure 4.16 showed, the crystals of the CO2 laser treated MgO–PSZ aretetragonal. Therefore, by assuming the number of crystals in the CO2 lasertreated surface to be n in unit area, the total surface energy of MgO–PSZ, gsv

(considering that half of the surface of the crystal is covered by neighbourcrystals), may be estimated by the expression

gsv ¼ 12npr2gt ¼ 1

8npD2gt ð4:4Þ

where D is crystal size. Since the main crystals remain as tetragonalstructures in the MgO–PSZ, it is reasonable to assume that gt does notchange with the crystal size. During CO2 laser processing, the n couldremain the same as the crystal grows. If this is so, then gsv is proportional

to D2, which is in agreement with the linear relationship between g12sv and

D as shown in Figure 4.18. Therefore, the larger crystal size can be attributedto the increase in the surface energy of the MgO–PSZ. Indeed, Man et al. [153,162] found that excimer laser treatment induced a conical structure, amicroscale of the peak and valley structure, providing an extra adherentsurface area for a strong adhesion joint on Si3N4 and LT35 surfaces. The joint

Figure 4.18 Relationship between g12

SV and crystal size ðDÞ of the MgO–PSZ


strength increased with the height of the cones and the laser energy density.Kappel [128] showed that the improved adhesion strength occasioned byexcimer laser texturing of ceramics is due to the formation of raisedmicroscopic protrusions over the surface.

4.6.2 The Role of Phase Change on Surface Energy

The previous XRD analysis revealed that the (101) tetragonal phaseincreased and the (111) monoclinic phase decreased in the CO2 laserinteraction layer on the MgO–PSZ. Furthermore, the relative intensity ofthe tetragonal phase increased when power density was below 1.6 kW/cm2

and then decreased slightly as the power density increased further. TheMgO–PSZ CO2 laser treated at 1.6 kW/cm2 power density has the highestsurface energy (108.9 mJ/m2), as discussed previously, and corresponds tothe highest intensity of the tetragonal phase. When the power density is at1.9 and 2.5 kW/cm2, the surface temperature of the MgO–PSZ increasedto well above the melting temperature, thereby generating more cubic phaseand reducing the tetragonal phase. Associated with this phenomenon,the surface energy begins to decrease from 108.9 to 80.7 mJ/m2 and thenfinally to 74.9 mJ/m2. These phase changes could be represented by thedifferent microstructures on the CO2 laser treated MgO–PSZ with variouspower densities. It is noticeable that tetragonal intensity varies, in a similarmanner to the surface energy, with the CO2 laser power density. Thisindicates that the tetragonal intensity is closely correlated with the surfaceenergy of the MgO–PSZ. In fact, it has been found that at T ¼ 0 K, the surfaceenergy of the (101) tetragonal phase is 45 % higher than that of the (111)monoclinic phase and 95 % higher than that of the (111) cubic phase [163].Based on this it is believed that the phase change (an increase in thetetragonal phase and a decrease in the monoclinic phase) resulted in thehigher surface energy of the MgO–PSZ following the CO2 laser irradiation.

4.7 Investigation of Wettability and Work Adhesion UsingPhysiological Liquids

In order to simulate the biological environment, the physiological fluids andsimulated physiological liquids used for the wetting experiments werehuman blood, human blood plasma, simulated body fluid (SBF) and SBF þBSA (bovine serum albumin). The SBF was prepared by dissolving reagent-grade chemicals, NaCl, NaHCO3, KCl, K2HPO4 3H2O, MgCl2 6H2O, CaCl2and Na2SO4 in ion-exchanged and distilled water, buffered at pH 7.25at 37 �C with tris(hydroxymethyl) aminomethane ([CH2OH]3CNH2) and1 m hydrochloric acid (HCl). The SBF has an inorganic ion concentrationclose to that found in human blood plasma, as shown later in Table 5.1.

Investigation of Wettability and Work Adhesion 61

SBF þ BSA was prepared using BSA (Sigma A-9306, lot 106H0300) dissolvedin SBF at pH 7.25 with a concentration of 4 mg/ml, to be used for proteinstudies to simulate human albumin because it is very similar to the sequenceof amino acid units.

The values of y formed between the selected and simulated physiologicaltest liquids and untreated and CO2 laser treated MgO–PSZ at power densityof 1.6 kW/cm2 alloy are shown in Table 4.6. It clearly reveals that the yvalues of all body fluids on the CO2 laser treated MgO–PSZ are lower thanthe untreated specimens, indicating the wettability characteristics of thematerial with the body fluids improved obviously after CO2 laser treatment.

Further, according to Equation (3.4), the decrease in y resulted in anincrease in the Wad of the MgO–PSZ towards the physiological and simu-lated physiological liquids. Using the referenced glv value of human blood(47.5 mJ/m2), human blood plasma (50.5 mJ/m2) [141], SBF (72.5 mJ/m2) andSBFþBSA (54.0 mJ/m2) [164], the work adhesion, Wad, of the Ti–6Al–4Valloy towards these body fluids were determined through Equation (3.4), asshown in Figure 4.19. A discernable increase in Wad of body fluids can be

Table 4.6 Mean values of contact angles formed between the selected andsimulated physiological test liquids and the untreated and CO2 laser treated MgO–PSZ

Contact angle, y (deg)—————————————————————————————

MgO–PSZ Human blood Human blood plasma SBF SBFþBSA

Untreated 54.7 57.9 78.8 67.9CO2 laser (1.6 Kw/cm2) 35.8 39.2 56.2 48.5

Figure 4.19 Work adhesion of body fluids for untreated, mechanically roughenedand CO2 laser treated MgO-PSZ


seen on the MgO–PSZ following CO2 laser treatment. Moreover, Wad

increased as the CO2 laser power density increased.Since biomaterials first contact a proteinaceous liquid phase, almost

aqueous in nature, leading to surface reorganization of proteins followedby cell attachment on biomaterials, wettability characteristics, by controllingthe interaction with physiological fluids, would primarily influence cellbehaviour on biomaterials. Wetting of the solid surface is a predictiveindex of cytocompatibility [165]. Moreover, the improvements of Wad

towards these fluids would imply better suitability of titanium in a biolo-gical environment.

4.8 Summary

The results presented in this chapter are a clear indication that CO2 lasersurface treatment of the MgO–PSZ brought about a reduction in the yformed between the MgO–PSZ and the control test liquids, indicating thatthe wettability characteristics of the material were modified. The extent ofthis wettability characteristics modification was varied by manipulation ofthe CO2 laser parameters.

Changes in the wettability characteristics of the MgO–PSZ were attributedto the following factors: (a) an increase in surface roughness; (b) incorpora-tion of oxygen at the MgO–PSZ surface resulting from CO2 laser treatment;and (c) the increase in the polar component, gp

sv, of the surface energyresulting from the melting and re-solidification of the MgO–PSZ surface. Thechanges in gp

sv resulted from the melting and solidification of the MgO–PSZsurface, with the value varying with the solidified microstructure. Thecellular microstructure obtained by the CO2 laser induced rapid solidifica-tion corresponded to the maximum value of gp

sv. Indeed, it was found thatthe surface energy of the MgO–PSZ increased as the crystal size andtetragonal phase present increased after the CO2 laser treatment.

Further analysis revealed that surface energy, by way of microstructure, wasthe primary influential factor governing changes in y and hence the wettabilitycharacteristics of the MgO–PSZ. Incorporation of oxygen at the surface wasalso shown to influence, to a lesser extent, changes in the wettabilitycharacteristics, while surface roughness was found to play a varying minorrole in inducing changes in the wettability characteristics of the MgO–PSZ.

Summary 63

5

In vitro BiocompatibilityEvaluation of CO2 LaserTreated Magnesia PartiallyStabilised Zirconia

This chapter is concerned with comparatively evaluating the biocompatibility of theCO2 laser treated MgO–PSZ. An investigation of the bioactivity of the CO2 lasermodified MgO–PSZ in simulated body fluids (SBF) was conducted. Thereafter theeffect of the CO2 laser treatment on the OH groups, the correlation between OHgroups and polar surface energy and the effect of the OH groups on the apatiteformation were studied. In addition, ellipsometry was used to investigate thealbumin and fibronectin adsorption on the untreated and CO2 laser treatedMgO–PSZ bioceramic. The relationship between the protein adsorption and surfaceproperties of the MgO–PSZ was discussed. Finally, the in vitro behaviour of humanfetal osteoblastic (hFoB) cells on the untreated and CO2 laser treated MgO–PSZ wasstudied. An evaluation of osteoblast cell adhesion and proliferation was alsoconducted to determine the effect of surface properties on the osteoblast cell adhesionand growth, elucidating the mechanisms active in the osteoblast cell response andthus deducing that the main factors are active.

5.1 Introduction

The biological activity of most orthopaedic and dental biomaterials is relatedto their ability to promote the formation of a neoformed layer of carbonateapatite crystals analogous to bone mineral. This layer also associates specificbone proteins and is the starting point in bone reconstruction [42]. The


nucleation of apatite on the surface of a material is induced by the functionalgroups on its surface [166].

The adsorption of proteins on to a biomaterial surface from the surround-ing fluid phase is rapid, with the surface properties of the biomaterial deter-mining the type, amount and conformation of the adsorbed proteins [167].The structure and composition of the adsorbed protein layer determine thetype and extent of the subsequent biological reactions, such as osseointegra-tion [168]. The adsorption of the protein layer may also be critical in terms ofproviding attachment sites for bone cells such as osteoblasts and theirprogenitors [169].

The cellular behaviour on a biomaterial is an important factor determiningthe biocompatibility. Osteoblasts are anchorage-dependent cells that mustadhere to substrate surfaces prior to undergoing subsequent cell functionssuch as proliferation, synthesis of collagen and other extracellular matrixproteins, etc. Cell adhesion is one of the initial events essential to subsequentproliferation and differentiation of cells before tissue formation. The wholeprocess of adhesion and spreading of the cell after contact with biomaterialsconsists of cell attachment, growth of filopodia, cytoplasmic webbing andflattening of the cell mass, and the ruffling of peripheral cytoplasm, whichprogress in a sequential fashion [170].

With the aim being to improve the biocompatibility (bioactivity andbiointegration) of a magnesia–partially stabilised zirconia (MgO–PSZ), CO2

laser radiation was used to generate surface properties that would promote abetter biological response. For an artificial material to bond to living bone, itis essential that the material has the ability to form a biologically active,bone-like, apatite layer on its surface in the human body. On account of thisthe bioactivities of the untreated and CO2 laser treated MgO–PSZ at differentlaser power densities were evaluated by observing the bone-like apatiteformation on their surface after soaking in simulated body fluids (SBF).Protein adsorption and osteoblast cell response were performed on theuntreated and CO2 laser treated MgO–PSZ in order to assess their propen-sity for biointegration, because protein adsorption is an almost immediateevent occurring upon implantation of metals and mediates, prior even to cellresponse and tissue–implant interactions [9]. In addition, it is widelyacknowledged that a major determinant of the bone–biomaterial interfacialresponse is the initial attachment, spreading and growth of osteoblasts onthe implant surface and that improvements in these processes may lead tofaster and more extensive implant integration and higher long-term stability[171]. Indeed, Vitale Brovarone et al. [172] investigated the in vitro behaviourof samples coated with a glass-matrix/zirconia particle composite by meansof soaking in SBF followed by scanning electron microscopy (SEM) observa-tion and X-ray diffraction (XRD) analysis. Rosengren et al. [173] studied in vitrothe adsorption of proteins from diluted human plasma on hydroxyapatite,

66 In vitro Biocompatibility Evaluation

alumina and zirconia with regard to total protein binding capacity, relativebinding capacity for specific proteins and flow-through and desorption pat-terns. Josset et al. [174] evaluated the biocompatibility of two implantablematerials, zirconia and alumina ceramics, in vitro using human osteoblast cellcultures. Furthermore, Bosetti et al. [61] used methods of soaking in SBF,protein adsorption and cell culture for an in vitro characterisation of abiomedical device.

5.2 Sample Preparation

The MgO–PSZ, with properties as described in Section 4.2.1, was subjectedto various in vitro evaluations to determine the effects of CO2 laser surfacetreatment on bioactivity. In order to perform all of the in vitro tests, theMgO–PSZ sheet was cut into 30 blocks each of 50 � 12 � 12:15 mm3 with acutting machine (Miniton; Struers, GmbH) using a diamond-rimmed cuttingblade, used as received prior to CO2 laser treatment. The 30 blocks were thendivided into two groups of 15 samples, with the groups being untreated andCO2 laser treated. The CO2 laser processing of the materials was conductedin the same manner as described in Section 4.2.2. For the in vitro apatiteformation test, the samples used were CO2 laser treated power densitiesranging from 0.6 to 2.5 kW/cm2. For the protein adsorption test, onlysamples that were CO2 laser treated with laser power densities of 0.9 and1.6 kW/cm2 were used. Samples CO2 laser treated with power densitiesranging from 0.6 to 2.5 kW/cm2 were used in the in vitro osteoblast celladhesion and proliferation test, while samples CO2 laser treated with powerdensities of 0.9 and 1.6 kW/cm2 were used in the evaluation of cell functions.Untreated samples were used as control in all of the in vitro tests.

5.3 Bone-Like Apatite Formation

Histological examinations in vivo show that an apatite layer is formed on theceramic [41] surface early in the implantation period and thereafter the bonematrix integrates into the apatite. This apatite layer consists of nanocrystalsof carbonate-ion-containing apatite that has a defective structure and lowcrystallinity. These features are, in fact, very similar to those of the mineralphase in bone and hence bone-producing cells (osteoblasts) can preferen-tially proliferate on the apatite and differentiate to form an extracellularmatrix composed of biological apatite and collagen. As a result, the sur-rounding bone comes into direct contact with the surface apatite layer. Whenthis process occurs a chemical bond is formed between the bone mineral andthe surface apatite to decrease the interfacial energy between them. It can be

Bone-Like Apatite Formation 67

concluded from these findings that an essential requirement for an artificialmaterial to bond to living bone is the formation of a layer of biologicallyactive bone-like apatite on its surface in the body [41]. There have beenconsiderable efforts for improving the bioactivity of zirconia inert biocera-mics. It has been revealed that a zirconia gel forms an apatite on its surface inSBF, indicating that the Zr–OH group is able to induce apatite nucleation[166]. The investigation of the apatite-forming ability of zirconia gels withdifferent structures shows that specific structures of the Zr–OH group intetragonal or monoclinic zirconia are effective for inducing apatite nuclea-tion [175–177]. For the purpose of the apatite formation, the chemicaltreatment has been used to produce the Zr–OH group on a zirconia/aluminacomposite subjecting the composite to H3PO4, H2SO4, HCl or NaOH aqu-eous solution treatments [178] and on zirconium metal treated with aqueousNaOH [179].

5.3.1 Experimental Procedures

FTIR Analysis

The optical adsorption spectra were measured at room temperature bymeans of a Fourier transfer infrared (FTIR) (FTS135; Bio-Rad, Inc.) spectro-meter over the 500–5000 cm�1 range at a resolution of 1 cm�1.

Soaking in Simulated Body Fluid

The samples were soaked in an acellular SBF [41], having an ion concentra-tion nearly equal to that of human blood plasma. This solution, whosecomposition is reported in Table 5.1, was prepared by dissolution of high-purity reagents in distilled water, and was buffered at 7.25 with 50 mMtrishydroxymethyl amino ethane and 45 mM hydrochloric acid.

The untreated and CO2 laser treated samples (treated with various powerdensities) were immersed in 30 mL SBF in a polyethylene bottle at 37 �C,without stirring. After 14 days they were removed from the solution, gentlywashed in distilled water and dried at room temperature. The soakedsamples were then characterised by SEM and EDX, the details of which

Table 5.1 Ionic concentration and pH of SBF in comparison with those in humanblood plasma [41]

Concentration

Ion Naþ Kþ Mg2þ Ca2þ Cl� HCO�3 HPO2�

4 SO2�4 pH

SBF 142.0 5.0 1.5 2.5 148.8 4.2 1.0 0.5 7.40Blood plasma 142.0 5.0 1.5 2.5 103.8 27.0 1.0 0.5 7.40


are given in Section 4.2.3. The samples for SEM observations were simplydried and covered by a thin gold layer to guarantee the conductivity.

5.3.2 Spectral Analysis and Hydroxyl Group

The main regions in the 750–950 cm�1 range are ascribed to ZrO2 stretchingmodes, as shown in Figure 5.1, owing to the fact that it is similar to infrared(IR) peak of the 20 mol % Al2O3-doped ZrO2 nanoparticles reported pre-viously [180] and the IR peak of the Al2O3–ZrO2 nanopowders after laserablation [181]. The vibration around 3300–3500 cm�1 in the FTIR spectra (seeFigure 5.2) could be attributed to the OH groups. Indeed, OH stretchingvibrations around this region have been observed by other workers onAl2O3–ZrO2 nanopowders after Nd:YAG laser ablation [181] and on theFe-doped crystals after laser irradiations [182].

As one can see from Figure 5.2, the absorption coefficient of the OH groupon the MgO–PSZ in this region increased after CO2 laser irradiation andvaried with the power density employed. For the untreated sample and theCO2 laser treated sample (power density of 0.6 kW/cm2), the absorptionpeaks from 3200 to 3600 cm�1 are not obvious, indicating that no OH groupsbonded on these samples. In contrast, the absorption peaks at this region canbe clearly observed on the samples following the CO2 laser irradiation withpower densities of 0.9, 1.6 and 1.9 kW/cm2, denoting that OH groups existedon these samples. The highest absorption coefficient of OH groups wasobtained on the sample that had been treated at 1.9 kW/cm2. This findingshows that the OH groups increased with the CO2 laser power density used.This relationship was also seen by Zeng, Yung and Xie [183] on the OH

1000 1500 2000 2500 40003000 3500

2.5 kW/cm2

1.9 kW/cm2

1.6 kW/cm2

0.9 kW/cm2

0.6 kW/cm2

Untreated

Wavenumber (cm−1)

Abs

orba

nce

(arb

. uni

ts)

0.6 arb. units

Figure 5.1 Infrared spectra of the untreated and CO2 laser treated MgO–PSZ (treatedwith different power densities)


groups bonded on to copper after CO2 laser treatment; nevertheless, theabsorption peaks in this region on the sample treated with a CO2 laser powerdensity of 2.5 kW/cm2 was not obvious, suggesting that the OH bond doesnot increase linearly with the CO2 laser power density. The explosiveevaporation due to the super-high temperature on the MgO–PSZ surfacetreated at this power density caused water vaporisation and disappearanceof the OH band. The phenomena of losing OH groups was also found onthe human dentine after Er:YAG (erbium-doped YAG) laser irradiation[184].

CO2 is a weak acid and is known to absorb on ZrO2 in the form of bothcarbonate and bicarbonate species [185]. Carbonate structures are formed viathe interaction of CO2 with zirconium cations in the lattice, as well as with asurface oxygen atom, whereas bicarbonate structures are formed via theinteraction of CO2 with a hydroxyl group. The peak from 2800 to 3000 cm�1

testified to the existence of carbonate structures on the MgO–PSZ surface.The change of carbonate structures has the same trend as the OH groupsdiscussed above.

The formation of the hydroxyl groups on the MgO–PSZ is due to thereactions of the zirconia with water vapour in air during CO2 laser proces-sing. The hydroxyl ion is a common impurity in insulating crystals and byinteracting with other impurities it gives rise to new complexes. The OH-stretching frequency is a very sensitive probe of the hydroxyl environment.Proper thermal treatments and isotopic substitutions allowed the stretchingmode absorption lines to be assigned to the defects in which OH� isembedded and to supply possible models for them [185]. Crystal growth

3000 3500 4000

2.5 kW/cm2

1.9 kW/cm2

1.6 kW/cm2

0.9 kW/cm2

0.6 kW/cm2

Untreated

Wavenumber (cm−1)

Abs

orba

nce

(arb

. uni

ts) 0.05 arb.units

Figure 5.2 Infrared spectra of the hydroxyl groups presents on the surface of theuntreated and CO2 laser treated MgO–PSZ (treated with different power densities)


from the melt is commonly carried out in air atmosphere as air alwayscontains a certain degree of humidity from which OH� ions are incorporatedinto the lattice [186]. CO2 laser irradiation is a thermal process. When thelaser fluence exceeds the ablation threshold, the irradiated surface experi-ences melting, followed by evaporation, whereupon the particles emit fromthe surface. At a higher fluence, the amount of particles increases and theybreak out quickly from the superheated surface to produce a high-densityvapour plume wherein a portion of the particles are ionised due to thethermal ionisation. The main reactions to occur in the melt ceramic andvapoured flume are

ZrO2 , Zr4þ þ 2O 2� ð5:1Þ

and the following oxidoreduction reactions occur at the melt–atmosphereinterface:

O 2� ! 12 O2 þ 2e� ð5:2Þ

H2O þ e� ! 12 H2 þ OH� ð5:3Þ

the whole reaction being

O 2� þ 2H2O ! 2OH� þ H2 þ 12 O2 ð5:4Þ

Finally, the OH� ions produced according to Reaction (5.4) would beincorporated with one, two and three and four surface Zr4þ respectively.According to the classification proposed by Tsyganenko and Filimonov[187], the OH groups bonded in the spectral ranges at 3770 and3680 cm�1 are typical one and three surface Zr4þ cations, respectively, intetragonal zirconia while the OH groups at 3775 and 3675 cm�1 bonds areone and more than one (possibly three) surface Zr4þ ions, respectively.

It has been speculated that surface melting was occasioned on the MgO–PSZ treated by the CO2 laser treatment with 1.6 kW/cm2 power density. Inturn the Zr4þ ion and OH� were produced and reaction between these ionsbrought about the Zr–OH group on the MgO–PSZ. The relatively highamounts of the hydroxyl groups bonded on to the modified samples with1.6 and 1.9 kW/cm2 were associated with the melting and chemical reactionon the MgO–PSZ surface.

5.3.3 The Correlation between OH Groupsand Wettability Characteristics

The values of the surface energy have been calculated for the MgO–PSZtreated by the various power densities in detail (see Chapter 4) and are


shown in Table 5.2. It is found that the absorption coefficient of theOH group (see Figure 5.2) and gp

sv on the MgO–PSZ increased after CO2

laser irradiation (see Table 5.2) and varied with the power densityemployed.

The untreated sample and the samples treated with the lower powerdensities had relatively low absorption peaks of the OH groups and lowergp

sv. On the other hand, the relatively high absorption peaks of the OHgroups and higher surface energy existed on the samples following the CO2

laser irradiation with power densities of 1.6 and 1.9 kW/cm2. Moreover,when the OH groups decreased on the sample treated at the power densityof 2.5 kW/cm2, there was a corresponding decrease in gp

sv. It is also inter-esting to notice that the melting of the MgO–PSZ was the fundamentalreason for the induction of the OH groups and improvement of the surfaceenergy. This finding implied that there was a correlation between the OHgroups and gp

sv. Indeed, Takeda et al. found that the surface OH groupsgoverned the wettability of commercial glasses [188] and adsorption proper-ties of metal oxide films [189]. A previous study [190] indicated that in thecase of cassiterite its wettability strongly depends on the acid–base interac-tions (polar component) resulting from the presence of OH groups andphysically adsorbed water on it. For the surface of the ‘dry’ cassiterite itssurface free energy practically results only from Lifshitzvan der Walls(dispersive component) intermolecular interactions. Most metal oxides arehydroxylated under normal conditions, i.e. at room temperature and whenwater or its vapour has had access to the surface. It was stated that the acid–base component of surface energy of the zirconia probably depends on thedensity of OH groups on the surface of the solids studied [191]. Indeed, theacid–base component of surface energy presented the majority of the forcesas the functions of the particular chemical nature of a certain materialcorresponding to gp

sv [192]. As such, it is believed that the CO2 laser inducedhydroxyl groups could be a major factor influencing gp

sv and, in turn, thewettability characteristics of the MgO–PSZ.

Table 5.2 Determined surface energy values for the MgO–PSZ before and afterCO2 laser treatment (treated with various power densities and traverse speed of2000 mm/min)

CO2 laser treated (kW/cm2)Surface Energy ———————————————————————————————————(mJ/m2) Untreated 0.5 0.9 1.6 1.9 2.5

gdsv 42.7 43.8 44.4 48.2 47.5 48.2gp

sv 10.1 10.4 21.9 60.7 33.2 26.7gsv 52.8 54.2 66.3 108.9 80.7 74.9


5.3.4 The Effects of CO2 Laser Treatment on the MgO–PSZin Simulated Body Fluids

As one can see from Figure 5.3, very small amounts of sediment are apparenton the surface of the untreated MgO–PSZ (Figure 5.3(a)) and the 0.6 kW/cm2 CO2 laser treated MgO–PSZ (Figure 5.3(b)). In contrast, consider-able amounts of sediment were clearly discernible on the samples thatwere CO2 laser treated with power densities of 0.9 (Figure 5.3(c)), 1.6

Figure 5.3 SEM images of the MgO–PSZ soaked in the SBF: (a) untreated, (b) 0.6 kW/cm2, (c) 0.9 kW/cm2, (d) 1.6 kW/cm2, (e) 1.9 kW/cm2 and (f) 2.5 kW/cm2


(Figure 5.3(d)), 1.9 (Figure 5.3(e)) and 2.5 kW/cm2 (Figure 5.3(f)), with thehighest amount of sediment being observed on the sample that was CO2

laser treated at 1.9 kW/cm2.The EDX analysis shows that most particles on the soaked surface are

NaCl sediments as the elements of Na and Cl, as shown in the Figure 5.4(a).Apatites were only found on the samples treated at power densities of the1.6 and 1.9 kW/cm2. As shown in Figure 5.4(a), only a few apatites formedon the sample treated at a power density of 1.6 kW/cm2, with only a smallamount of Ca element shown in the EDX analysis given in Figure 5.4(a).However, on the sample treated with a CO2 laser power density of 1.9 kW/cm2, some apatites were observed. One of them is shown in Figure 5.4(b)with the Ca:P ratio about 1.65. This Ca:P ratio exhibits the calcium phosphatetransforms into apatite [41].

The Effect of OH Groups

There was no occurrence of apatite formation on the untreated and certainCO2 laser treated samples (0.6, 0.9 and 2.5 kW/cm2) that displayed fewhydroxyl groups. On the other hand, some apatites formed on other CO2

Figure 5.4 SEM image and EDX analysis of the apatite formed on the surface of theMgO–PSZ when treated with power densities of (a) 1.6 kW/cm2 and (b) 1.9 kW/cm2


laser treated samples (1.6 and 1.9 kW/cm2) that did display any hydroxylgroups. This finding suggests that the hydroxyl group on the MgO–PSZcould be the predominant factor governing the formation of the apatites. Thehydroxyl groups on the MgO–PSZ surface certainly generate Zr–OH groups,which have been shown to be functional groups for the formation of theapatite [166]. It is suggested that Zr–OH functional groups formed on thesamples in the CO2 laser processing at certain parameters and that suchfunctional groups naturally brought about the nucleation of the apatite onthese samples in the simulated body fluid environment. The nucleation ofthe apatite could yield the apatite formation and bone-bonding ability to theMgO–PSZ modified to have Zr–OH groups on the surface.

The Effect of Wettability

It was found that there were more sediments and apatites on the surfacewith the higher surface energy than on the surface with the lower surfaceenergy. In the process of Ca–P precipitation, the variations of the Gibbsfunction ð�GÞ of the MgO–PSZ with the higher surface energy should begreater, compared to that of the MgO–PSZ surface with lower surfaceenergy. This finding, agreeing with the study by Feng et al. [193], suggestedthat the adsorption and reaction would more easily have occurred on thesurface with the higher surface energy, especially the polar component ofsurface energy, which would be beneficial to the chemical force andbonding.

5.4 Protein Adsorption

The molecules involved in cell adhesion and spreading include extracellularmatrix molecules, transmembrane receptors and intracellular cytoskeletalcomponents. Among the extracellular matrix proteins shown to mediate cellattachment to substrates, fibronectin is protein found in many extracellularmatrices and in blood plasma and serves as an attachment molecule betweenthe substrate and cell membrane of anchorage-dependent cells. It is knownthat the ligand fibronectin connects to the cell membrane via integrinreceptors. The activation of integrins triggers cytoplasmic reactions, andthereby stimulates the intracellular signalling pathway and subsequentlycellular functions such as proliferation and differentiation [169]. On the otherhand, human albumin is a nonadhesive protein for osteoblasts [194].Albumin is the major protein component of serum and dominates theadsorption of phenomena on medical implants in the first stage of contactwith body fluids. Human serum albumin or bovine serum albumin (BSA)coatings are often used as a passivating agent to prevent the adhesion of cellsand thrombus formation [195].

Protein Adsorption 75

An important fact to highlight is that different surfaces provide verydifferent opportunities for protein binding. The investigation of competitiveprotein adsorption showed that BSA in a single-component solutionadsorbed on to a hydrophobic surface two times more than that on to ahydrophilic surface [196]. The surface wettability of biomaterials affects theability of cells to reorganise pre-adsorbed fibronectin and to form their ownmatrix by secreted fibronectin. Moderate wettable and hydrophilic surfacesare ideal for better interactions with cells while hydrophobic substratainhibit early and late matrix formations. It is possible that there is a criticalvalue in the strength of fibronectin adsorption that regulates the ability ofcells to construct a fibronectin matrix [197].


Protein Adsorption

The proteins used for this study were human serum albumin and humanplasma fibronectin (Calbiochem, Inc.). Prior to the adsorption of 1 mg/mlalbumin in phosphate buffered salines (PBS), MgO–PSZ samples wererinsed with deionised water. The individual samples were transferred intoa 24-well tissue culture plate. Thereafter, 2.5 ml of prepared albuminsolution was added into each well. Adsorption proceeded for 1 hour in anincubator at 37 �C. After adsorption was complete, the samples were driedwith N2 and immediately transferred to an ellipsometer for measurement ofthe adsorbed protein layer. The above procedure was repeated with a0.2 mg/ml concentration of fibronectin in PBS.

Ellipsometric Measurement

Human plasma fibronectin was measured using an automatic ellipsometerequipped with a 633 nm helium–neon laser (L117F; Gaertner, Inc). Thethickness and refractive indices of protein films were determined using anellipsometer computer program with an accuracy of 3 A

�. Four ellips-

ometer measurements at different locations on each sample were takenand the average value was calculated.

Statistics

Statistical analysis was performed with an SPSS v.12 software package(SPSS/PC, Inc.). Data are reported as mean SD (standard deviation) at asignificance level of p < 0:05. After having verified normal distribution andhomogeneity of variances, one-way ANOVA and Scheffe’s post hoc multiplecomparison tests were done.


5.4.2 Albumin and Fibronectin Adsorption on CO2

Laser Treated MgO–PSZ

The thickness of the absorbed fibronectin layer was less on the untreatedsample than that on the CO2 laser treated sample, as shown in Figure 5.5.The statistical analysis reveals that the thickness of absorbed fibronectin onthe untreated sample was similar to that on the sample CO2 laser treated atthe power density of 0.9 kW/cm2, but significantly less than that on thesample CO2 laser treated at the power density of 1.6 kW/cm2 ðp < 0:01Þ. Onthe other hand, the thickness of the absorbed albumin layer on the untreatedMgO–PSZ was higher than that on the CO2 laser modified sample, as shownin Figure 5.5. The statistical analysis reveals that the thickness of theabsorbed albumin layer on the untreated sample was significantly higherthan on the CO2 laser treated samples ðp < 0:01Þ.

From Figure 5.5 it is apparent that the CO2 laser power density applied inthe experiments was negatively correlated to the amounts of albumin, butpositively correlated with the fibronectin. The results showed that the CO2

laser treatment promoted the adsorption of the fibronectin on the MgO–PSZand the amount of the adsorbed fibronectin was positively correlated withthe CO2 laser power density applied in the experiments, as shown in

300

400

500

600

700

Untreated CO2 laser CO2 laser

0.9 kW/cm2 1.6 kW/cm2


0.9 kW/cm2 1.6 kW/cm2

Fibronectin Albumin

Thi

ckne

ss o

f A

dsor

bed

Pro

tein

Lay

er (

Å)

* ***

Figure 5.5 Thickness of the adsorbed fibronectin and albumin layer on the untreatedand CO2 laser treated MgO–PSZ (treated with different power densities). For thefibronectin adsorption, there was a significant statistical difference in thickness be-tween the untreated MgO–PSZ and the sample CO2 laser treated at 1.6 kW/cm2, andno statistical difference between the untreated MgO–PSZ and the sample CO2 lasertreated at 0.9 kW/cm2. For the albumin adsorption, there was a significant statisticaldifference in thickness between the untreated MgO–PSZ and the samples that wereCO2 laser treated at 0.9 and 1.6 kW/cm2, and no statistical difference between theCO2 laser treated samples ð�p < 0:05Þ


Figure 5.5. However, the thickness of the adsorbed human serum albuminlayer on the untreated MgO–PSZ is higher than that on the CO2 lasermodified MgO–PSZ and was negatively correlated with the CO2 laserpower densities. This finding is perhaps not so surprising as the variousCO2 laser power densities used brought about changes in wettabilitycharacteristics and surface roughness of the MgO–PSZ: it is known thatprotein adsorption is influenced by the surface topography (roughness) [198]and the surface chemistry (wettability characteristics) [199].

The Effects of Surface Roughness

By altering the CO2 laser power density, it was possible to obtain the nar-row range of surface roughness values detailed in Figure 5.6 (see alsoSection 4.3.3). The experimental results given in Figure 5.6 reveal that theamount of fibronectin adsorption increased, while the amount of albuminadsorption decreased with the surface roughness of the MgO–PSZ. Thesetrends in absorption for the fibronectin and the albumin were verified to alarge extent by the results of the statistical analysis.

0.2 0.4 0.6 0.8 0.2 0.4 0.6 0.8300

400

500

600

700

Surface Roughness, Ra (µm)

Thi

ckne

ss o

f A

dsor

bed

Pro

tein

Lay

er (

Å)

Fibronectin Albumin

**

**

Figure 5.6 The relationship between the thickness of adsorbed fibronectin andalbumin layer and Ra of the MgO–PSZ. For fibronectin adsorption, there was asignificant statistical difference in thickness between the sample with Ra of 0.295 mmand the sample with Ra of 0.717 mm, and no statistical difference between the samplewith Ra of 0.295 mm and the sample with Ra of 0.313 mm. For albumin adsorption,there was a significant statistical difference in thickness between the sample with Ra

of 0.295 mm and the samples with Ra of 0.313 and 0.717 mm, and no statisticaldifference between the sample with Ra of 0.313 mm and the sample with Ra of0.717 mm ð�p < 0:05Þ


The relationship between the albumin adsorption and surface roughnessgiven in Figure 5.6 is consistent with the findings of other researchers,insofar as bovine serum was adsorbed preferentially on to the smoothsubstratum [200, 201], thus implying that the surface roughness of theMgO–PSZ is one of the factors active in albumin adsorption. It wasexplained, however, by MacDonald et al. [202] that by roughing the surfaceof Ti one would obtain a more hydrophilic surface. This increase in surfacehydrophilicity of the Ti, according to Serro et al. [201], will consequentlyresult in lower albumin adsorption. The relationship between the fibronectinadsorption and surface roughness given in Figure 5.6 is in agreement withprevious reports. Deligianni et al. [200] found that a roughened Ti alloysample adsorbed much more fibronectin than a smooth sample. The muchhigher affinity of rough substrata to fibronectin could be the driving force forpreferential adsorption of fibronectin; however, other researchers havereported that the amounts of immobilised fibronectin on the rough titaniumwere 50 % lower than those adsorbed on the smooth one [203]. This decreasewas noticed when the roughness was produced by polishing or sandblast-ing, followed by acid attack, which is an indication that the chemical ormechanical manufacturing process, used to achieve the surface texture,might influence the protein adsorption behaviour of a surface [203].Hence, a simple conclusion would be difficult to execute for the relationshipbetween the amplitude of surface roughness and protein adsorption. It mustbe noted that in this work the CO2 laser treatment effects change in othersurface properties besides roughness. It is most likely that the surfaceroughness plays a role in the protein adsorption, but its effects correlate toand are less than the wettability characteristics of the MgO–PSZ.

The Effects of Wettability Characteristics

The previous results are a clear indication that interaction of the CO2 laserbeam with the MgO–PSZ brought about a decrease in y, which in someinstances was considerable. This in turn naturally gave rise to improvedwettability characteristics.

As one can see from Figure 5.7, as the wettability characteristics of theMgO–PSZ increased, the adsorbed amounts of fibronectin increased, whilethe adsorbed amounts of albumin decreased. Indeed, this observation wassupported somewhat by the results of the statistical analysis. The results ofthe albumin adsorption are consistent with the previous finding that theincrease in surface hydrophilicity of Ti results in lower albumin adsorption[201], showing that the wettability characteristics of the MgO–PSZ could bethe main factor active in the albumin adsorption. The results of the adsorp-tion of fibronectin show that it increased on the hydrophilic surface. Theprevious investigation [204] on the extent of fibronectin adsorption as


compared to its biological activity on hydrophobic and hydrophilic surfacessuggested the possibility that fibronectin was adsorbed in two differentconformations when incubated with the surfaces at low concentrations, withthe more active conformation on the hydrophilic surfaces. The resultsshowed that the antiplasma fibronectin antibody appeared to bind to theconformation of fibronectin adsorbed on hydrophilic surfaces much betterthan the conformation of fibronectin adsorbed on hydrophobic surfaces[204]; therefore, the wettability characteristics of the MgO–PSZ could bethe predominant mechanism governing the fibronectin adsorption. It isnoticeable that considerable change in the gp

sv, instead of the minor differencein gd

sv, was the main mechanism governing the wettability characteristics ofthe MgO–PSZ after CO2 laser irradiation, indicating that the albumin andfibronectin adsorption on the MgO–PSZ surfaces was probably due to thepolar and chemical interactions [205].

5.5 Osteoblast Cell Response

The development of bone–implant interfaces depends on the direct interac-tions of bone matrix and osteoblasts with the biomaterial. There is a

0.0 0.2 0.4 0.6 0.8 0.0 0.2 0.4 0.6 0.8300

400

500

600

700

Wettability, cos (glycerol)θ

Thi

ckne

ss o

f A

dsor

bed

Pro

tein

Lay

er (

Å)

Fibronectin Albumin

**

**

Figure 5.7 The relationship between the thickness of adsorbed fibronectin and albuminlayer and wettability characteristics ðcos yÞ of the MgO–PSZ. For the fibronectin adsorp-tion, there was a significant statistical difference in thickness between the sample withcos y ¼ 0:19 and the sample with cos y ¼ 0:77, and no statistical difference between thesample with cos y ¼ 0:19 and the sample with cos y ¼ 0:47. For the albumin adsorption,there was a significant statistical difference in thickness between the sample withcos y ¼ 0:19 and the sample with cos y ¼ 0:77, and no statistical difference betweenthe sample with cos y ¼ 0:47 and the sample with cos y ¼ 0:77 ð�p < 0:05Þ


substantial body of literature based on the premise that improved initialattachment of osteoblasts or osteoblast precursor cells to orthopaedicimplant surfaces may lead to improved bone integration of the implantand longer-term stability [206]. Osteoblast adhesion is a prerequisite forbone–biomaterial interaction and depends on the surface aspect of materials.A cell in contact with the surface of a material will firstly attach, adhere andthen finally spread. The quality of this adhesion will influence theirmorphology and their future capacity for proliferation and differentiation.The attachment of anchorage-dependent cells such as osteoblasts to bioma-terial surfaces is a complex process involving cell attachment and spreading[207], focal adhesion formation, and extracellular matrix formation andreorganisation [208].


Osteoblast Cell and Cell Culture

The human osteoblastic cell line hFOB 1.19 was obtained from AmericanType Culture Collection (ATCC, Inc.). The hFOB cell line was established bytransfection of limb tissue obtained from a spontaneous miscarriage. Thecells have the ability to differentiate into mature osteoblasts expressing thenormal osteoblast phenotype and provide a homogeneous, rapidly prolifer-ating model system for study of normal human osteoblast cells. Moreover, itovercomes the disadvantages of earlier in vitro model systems, namely theunknown species-specific phenotype characteristics of animal osteoblastcultures and the very slow rates of proliferation, as well as the short lifetimeof primary cultures derived from normal human bone [209]. The cells werecultured in a medium containing a 1:1 mixture of Dulbecco’s modifiedEagle’s medium without phenol red and Ham’s F-12 medium with 2.5 mML-glutamine (D-MEM/F-12 medium), supplemented with 10 % fetal bovineserum (ATCC, Inc.) and 0.3 mg/ml G418 (Calbiochem, Inc.) at 37 �C in ahumidified 5 % CO2 incubator. Osteoblasts at passage numbers 2–4 wereused in this experiment.

Cell Cytotoxicity

Cytotoxicity tests consisted of the quantification of the activity of lactatedehydrogenase (LDH) in culture medium of cells in contact with thesamples. The activity of the LDH enzyme rises when cells are damaged.Thus the LDH activity induced by the untreated and selected CO2 lasertreated specimens (0.9 and 1.6 kW/cm2) in triplicate were compared to thatinduced by a toxic agent (Triton X100 0.05 % in PBS) and to that induced by aculture polystyrene plate (NUNC, Inc.). The cell culture plate was used as anegative control and a Triton toxic agent as a positive control.

Osteoblast Cell Response 81

Cell Adhesion and Morphology

The MgO–PSZ samples were placed in a 24-well tissue culture polystyreneplate (NUNC, Inc.), sterilised in 70 % alcohol and rinsed in PBS solution. Toanalyse the osteoblast cell attachment and morphology, one untreatedsample and one CO2 laser modified sample (power density of 1.6 kW/cm2)were used in the assessment of cell morphology. The specimens were seededwith the 0.5 ml cell suspension of 1 � 105 cell/mL and analysed by SEMafter 24 hours of cell culture. For a 7 day osteoblast cell adhesion analysis,the samples were rinsed in PBS, whereupon they were seeded with 0.5 mLcell suspension (4 � 105 cell/ml). After culturing, the cells were fixed in2.5 % glutaraldehyde solution for 1 hour, washed with PBS and thendehydrated in increasing concentrations of alcohol (70, 85 and 100 %).Thereafter, the osteoblast cells were dried in a critical point dryer(CPD030; BAL-TEC, GmbH). Then the samples were examined with bySEM after sputter gold coating. For the cell adhesion analysis, three imageswere taken for the each sample at different areas and a typical one waschosen for the analysis.

Cell Proliferation

Each group of specimens used for cell proliferation tests in triplicate wasmeasured after cell culturing for 14 days. Osteoblast cells were cultured onspecimen surfaces with the 0.5 ml cell suspension of 1 � 105 cell/ml in 6-wellculture plates. The cell culture medium was changed every 3 days. At everyharvest time point, cells were detached from specimen surfaces by incuba-tion with trypsin/EDTA (ethylene diamine tetraacetic acid) (0.5 g/l trypsinand 0.2 g/l EDTA) (GIBCO, Ltd) for 5 minutes at 37 �C and each specimenwas washed with PBS. Released cells were counted with a hemocytometer,and on every specimen counting was repeated three times.

Alkaline Phosphatase Assay

For the staining of human osteoblast cells an alkaline phosphatase (ALP)assay kit (Sigma Diagnostics, Inc.) was used. After rinses with PBS, thesamples with cells were fixed by immersing in a citrate buffered acetone for30 seconds and rinsed gently with deionised water for 45 seconds. Alkalinedye mixture was added and the samples were incubated at 24 �C for30 minutes protected from direct light. They were then rinsed thoroughlyin deionised water for 2 minutes and placed in Mayer’s hematoxylin solutionfor 10 minutes. Positive staining for alkaline phosphatase (red–violet) wasidentified by light microscopy and evaluated by scoring cell rating and countaccording to the characterisation method provided.


Statistics

The statistical analysis of the results was performed with an SPSS v.12software package in the same manner as discussed in Section 5.4.1.

5.5.2 Osteoblast Cell Response on CO2 Laser Treated MgO–PSZ

Cell Cytotoxicity

LDH is a kind of enzyme in the cell. When cells are damaged or broken, theLDH will leak into the culture medium. As the concentration of LDH in themedium is proportional to the numbers of dead/damaged cells, the con-centration of LDH in the medium can reflect the cytotoxicity. The LDHactivity in the culture media obtained from cells cultured on all the testedmaterials was found to be not significantly different from the negativecontrol, as shown in Figure 5.8, indicating that untreated and CO2 lasertreated MgO–PSZ were not cytotoxic.

Cell Attachment

Figure 5.9(a) shows that no osteoblast cells were observed on the untreatedMgO–PSZ after 24 hours of cell incubation, whereas a few cells attached on

0

10

20

30

40

50

60

LD

H A

ctiv

ity

PositiveControl

NegativeControl

UntreatedMgO-PSZ

CO2 LaserMgO-PSZ

0.9 kW/cm2

CO2 LaserMgO-PSZ

1.6 kW/cm2

** * *

Figure 5.8 Results provided by assessment of cell membrane damage are expressedas LDH activity (U/L) SD at 340 nm. There was a significant statistical differencebetween the positive control and MgO–PSZ samples, and no statistical differencebetween the negative control and the untreated and CO2 laser treated MgO–PSZð�p < 0:05Þ


to the CO2 laser treated MgO–PSZ (Figure 5.9(b)). It is quite clear that theosteoblast cell attachment on the MgO–PSZ was influenced by the CO2 lasertreatment, indicating that the surface properties generated by the CO2 lasertreatment were more favourable for the osteoblast cell attachment. Further-more, the cells on the CO2 laser treated samples showed filopodia andspread well (Figure 5.9(b)), denoting good cell attachment.

Moreover, it is evident from Figure 5.10 that the osteoblast cells haddifferent morphologies in different regions of the CO2 laser treated track.Figure 5.10(a) shows that the osteoblast cells at the edge underwent initialspreading and the individual cell was found to cover an area of about 30–40mm, as shown in Figure 5.10(b). Short filopodia protruded and elongatedabout 5–10 mm from the osteoblast cell (see Figure 5.10(b)). The elongationdirection of the short filopodia implies the direction of the migration process.Conversely, the osteoblast cells at the centre reached a stage where theygrew and spread to cover a region of 60–150mm (Figure 5.10(a)). One typicalosteoblast cell (see Figure 5.10(c)) had spread completely and flattened, withthe cytoplasmic spread to cover an area of about 30–50mm as well as formingfour filopodias, two of them elongated to a length of 50–60mm. Likewise,another two osteoblast cells (Figure 5.10(d)) had a flat cytoplasm with twofilopodias elongated to 50–60mm. The morphologies of these osteoblast cellsdisplay the final stage of cell attachment. In general, osteoblast cells in thecentre spread better and reached a higher stage of cell attachment than thoseat the edge of the CO2 laser treated track. Since the CO2 laser treatmentexerted a higher photochemical effect at the centre than on the edge due tointensity distribution of the CO2 laser TEM01 beam mode (see Section 4.3.3),it could be concluded that the osteoblast cell spreading and attachment isinfluenced by the level of the CO2 laser treatment.

Figure 5.9 SEM image of hFOB human osteoblast cells after 24 hours on (a) theuntreated MgO–PSZ and (b) the CO2 laser treated MgO–PSZ at power densities of1.6 kW/cm2


Cell Growth

After a 7 day incubation period, the hFOB cells grew well and formed a layeron all of the samples (see Figure 5.11). The degree of osteoblast cell adhesionand growth in terms of cell coverage area varied with the CO2 laser powerdensity. The cover density is defined by the ratio of the osteoblast celladhesion area to the whole surface area and is used as an indication of theosteoblast cell adhesion and growth.

It has been found that the CO2 laser power density used in the treatmenthad a significant influence on the cover density of the osteoblast cells (seeFigure 5.12). For instance, compared with the untreated sample, a powerdensity of 0.9 kW/cm2 generated double cover density, while the higherpower densities of 1.6, 1.9 and 2.5 kW/cm2 brought about triple cover densityon the CO2 laser treated MgO–PSZ. Generally, the osteoblast cell coveragearea was found to increase as the power density increased when the powerdensity is less than 1.9 kW/cm2.

Figure 5.10 SEM images of hFOB human osteoblast cells after 24 hours (a) atthe interface region, (b) at the circled area at the edge and (c) and (d) at thecircled areas at the centre of the CO2 laser treated MgO–PSZ with 1.6 kW/cm2 powerdensity


Figure 5.13 indicates that the number of osteoblast cells after 14 days onthe untreated MgO–PSZ is less than the CO2 laser treated MgO–PSZ.Especially, compared with the untreated sample, the osteoblast cell growssignificantly faster on the samples that were CO2 laser treated at relativelyhigh power densities of 1.6, 1.9 and 2.6 kW/cm2. Indeed, both cell adhesionand growth in 7 days investigated by SEM and cell proliferation after

Figure 5.11 SEM images of hFOB cells on (a) the untreated MgO–PSZ and CO2 lasertreated MgO–PSZ at power densities of (b) 0.5 kW/cm2, (c) 0.9 kW/cm2, (d) 1.6 kW/cm2, (e) 1.9 kW/cm2 and (f) 2.5 kW/cm2


0.0 0.5 1.0 1.5 2.0 2.50

20

40

60

80

100

CO2 Laser Power Density (kW/cm2)

Cel

l Cov

er D

ensi

ty (

%)

* * *

Figure 5.12 The relationship between the cover density of the hFOB cells and CO2

laser power density. There was a significant statistical difference between the un-treated and CO2 laser treated MgO–PSZ samples, and no statistical difference amongthe samples CO2 laser treated at 1.6, 1.9 and 2.6 kW/cm2 ð�p < 0:05Þ

0.0 0.5 1.0 1.5 2.0 2.50

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4

6

8

10

12

14

Tot

al N

o of

Cel

ls (

×105

)

CO2 Laser Power Density (kW/cm2)

∗∗

∗

Figure 5.13 Total number of osteoblast cells on the untreated and CO2 laser treatedMgO–PSZ after 14 days. There was a significant statistical difference betweenthe untreated sample and the samples that were CO2 laser treated at 1.6, 1.9 and2.6 kW/cm2, and no statistical difference among the untreated sample and thesamples CO2 laser treated at 0.6 and 0.9 kW/cm2 ð�p < 0:05Þ


14 days counted by the hematocytometer show a similar trend in the cellgrowth on the untreated and CO2 laser treated MgO–PSZ.

Alkaline Phosphatase (ALP) Activity

Optical images of the positive staining for ALP on the untreated and selectedsamples that were CO2 laser treated at 0.9 and 1.6 kW/cm2 are shown inFigure 5.14. The cell rating was determined on the basis of quantity andintensity of precipitated dye within the cytoplasm of these cells. As can beseen from Figure 5.14, the granule on the untreated sample is small in sizeand shows faint to moderate intensity of staining. On the other hand, thegranule is medium to large in size and shows strong intensity of staining onthe sample CO2 laser treated at 0.9 kW/cm2 and brilliant intensity of stainingon the sample CO2 laser treated at 1.6 kW/cm2. One granule evidencedspreading of the cell on the sample CO2 laser treated at 0.9 kW/cm2, asshown in Figure 5.14(c). The leukocyte alkaline phosphatase activity (LAPA)scores, determined by the number of cells counted and multiplying by the

Figure 5.14 Optical image of the positive staining for alkaline phosphatase on(a) the untreated and (b) the CO2 laser treated samples (c) at 0.9 kW/cm2 and(d) at 1.6 kW/cm2


value of the cell rating, are shown in Figure 5.15. The LAPA score of the CO2

laser treated sample is clearly much higher than that of the untreatedsample. Furthermore, the CO2 laser treated MgO–PSZ samples had LAPAscores that were statistically significantly higher than that of untreatedsamples, as shown in Figure 5.15.

5.5.3 The Effect of CO2 Laser Treatmenton the Osteoblast Cell Response

The results show that the CO2 laser treatment brought about the consider-able change in the response of osteoblast cells. Moreover, the osteoblast cellresponse varied on the MgO–PSZ with the power density of CO2 lasertreatment applied. As discussed in Chapter 4, the variation of the powerdensity of CO2 laser treatment resulted in the different changes in surfaceproperties. The surface topography and wettability characteristics chemistryhave been shown to be the factors influencing the osteoblast cell response inprevious studies [210, 211] and are believed to effect the osteoblast cellresponse on the CO2 laser treated MgO–PSZ.

The Effect of Topography on the Osteoblast Cell Response

The topography has been shown to be one of the factors in influencing theosteoblast cell response [212]. As demonstrated in Chapter 4, CO2 lasertreatment generated a consistently rougher surface on the MgO–PSZ whencompared with the untreated sample and Ra increased with the power

0

20

40

60

80

100

120

LA

PA

Sco

re

Untreated CO2 laser0.9 kW/cm2

CO2 laser1.6 kW/cm2

Figure 5.15 LAPA scores of the untreated and CO2 laser treated MgO–PSZ. There wasa significant statistical difference between the untreated sample and samples CO2

laser treated at 0.9 and 1.6 kW/cm2 at p < 0:05


density of CO2 laser treatment. In this way it was possible to obtain a narrowrange of surface roughness values (see Section 4.3.3). As shown inFigure 5.16, the CO2 laser treated MgO–PSZ with rougher surfaces has ahigher osteoblast cell cover density compared with the smooth untreatedsample. This is in agreement with some reports that the rougher surface oftitanium promoted more osteoblast-like cell attachment [213]. Even so, thereis no linear relationship between the osteoblast cell cover density and Ra, asshown in Figure 5.16.

Furthermore, the different microstructures and increase in the crystal sizeswere postulated to be the factors influencing the osteoblast cell coverdensity, as shown in Figure 5.17. The degrees of the cell adhesion improvedmarkedly when an obvious microstructure change happened on the MgO–PSZ. Surface microtopography has been cited as an important factor influen-cing protein–surface and cell–surface interactions [80]. A number of reasonshave been suggested for an increased differentiation of osteoblasts onmicrostructured surfaces, such as the influence of surface structure on cellshape or the fact that the surface topography creates a specific biochemicalmicroenvironment around each cell [214]. Figure 5.17 shows that the crystalsizes in all CO2 laser treated MgO–PSZ are larger than the untreated sample.The osteoblast cell cover density generally increased with the increasedcrystal size when the power density was lower than 1.9 kW/cm2, indicatingthat crystal size could possibly influence the osteoblast cell adhesion. It ismost likely that the greater nanosurface area created by the larger crystalsize may promote interactions (such as adsorption, configuration, bioactiv-ity, etc.), of select serum proteins(s), which, subsequently, enhance osteo-blast adhesion. The study of osteoblast adhesion on nanophase ceramics haselucidated the fact that a critical grain size (between 49 and 67 nm for

0.0 0.5 1.0 1.5 2.0 2.5 3.0 3.5 4.00

20

40

60

80

100

Surface Roughness, Ra (µm)

Cel

l Cov

er D

ensi

ty (

%)

Figure 5.16 The relationship between the cover density of the hFOB cells and Ra


alumina and between 32 and 56 nm for titania) played a crucial role inmediating osteoblast adhesion to nanophase ceramics by creating a greatersurface area and promoting the interaction of protein [140]. However, alinear relationship does not exist between the cell cover density and crystalsize (see Figure 5.17). Owing to this lack of a linear relationship, a simpleconclusion would be difficult to elucidate the relation between the topogra-phy and cell behaviours, because the surface roughness is only one of thefactors affecting cell behaviours. Indeed, Hallab et al. [215] demonstratedthat surface free energy was a more important surface characteristic thansurface roughness for cellular adhesion strength and proliferation. Thus it isreasonable to postulate that the surface roughness does influence the humanosteoblast cell response; however, its effect is less than that of the surfaceenergy.

The Effect of Wettability Characteristicson the Osteoblast Cell Response

As demonstrated in Chapter 4, the wettability characteristics of the MgO–PSZ improved after CO2 laser treatment. It is believed that the changes inwettability characteristics resulted in modification of the osteoblast cellresponse. The result of the one-day cell culture on the MgO–PSZ showed

Figure 5.17 The relationship between the cover density of hFOB cells, microstructureand crystal size of the MgO–PSZ (treated with various CO2 laser power densities)


that there were no osteoblast cells attached on the untreated MgO–PSZ withlow wettability characteristics, whereas some cells had already attached andspread on the CO2 laser treated MgO–PSZ (power density of 1.6 kW/cm2)with high wettability characteristics. The difference in wettability character-istics and surface energy must be the mechanism determining the differencein the osteoblast cell attachment. The CO2 laser used in the experiment is aTEM01 multimode. The level of CO2 laser beam interaction is higher in thecentre than at the edge of the CO2 laser beam; consequently the modificationlevel of surface energy would be higher at the centre than at the edge of theCO2 laser treated track. The different wettability characteristics generated byCO2 laser treatment across the track brought about the different levels ofosteoblast spreading between cells at the edge and cells at the centre(see Figure 5.10).

As can be seen from Figure 5.15, the enhanced cell functions representedby the LAPA increase with the wettability characteristics. It is believed thatthe wettability characteristics of the MgO–PSZ was the primary factorgoverning the cell response. The effects of wettability characteristics oncell functions could result from their influence on the protein adsorption andcell adhesion. The adsorption of the proteins is the net result of various typesof interactions that depend on the nature of the protein aqueous solution.The difference in cellular response of different materials suggests that thereare differences in the organization of the adsorbed protein layer. Proteinadsorption mediated cell behaviours are regarded as fundamental reactionsat the biomaterial–tissue interface [216].

The value of cos y (glycerol) was used to express the wettability character-istics of the MgO–PSZ. The higher the value of cos y, the higher is thewettability characteristics. It is evident from Figure 5.18 that the cell growth

0.0 0.2 0.4 0.6 0.8 1.00

20

40

60

80

100


Cel

l Cov

er D

ensi

ty (

%)

Figure 5.18 The relationship between the cover density of hFOB cells and thewettability characteristics of the MgO–PSZ


increased when the wettability characteristics changed from a lower value toa moderate value (cos y from 0.2 to 0.65), indicating that generation of thewettability characteristics is the main mechanism accounting for the differ-ent amounts of osteoblast cell adhesion and growth. The finding agrees withprevious studies showing the influence of wettability on the attachment andspreading of various cells [213, 217–219]. These studies showed good cellattachment and spreading on high-energy substrata and poor cell attach-ment and spreading on low-energy substrata, which accounts for theminimal energetic state of a system in equilibrium. It was noticed that thevalue of cos y ranged between 0.6 and 0.8 and did not present a greatdisparity in cell cover density. This implied that after a certain value, furtherincreases in the wettability characteristics would not improve the cellresponse. This is similar to the finding that the highest levels of cellattachment were found on a moderately hydrophilic surface using amodel surface [220].

As demonstrated in Chapter 4, the surface roughness, surface oxygencontent and surface energy are the mechanisms governing the wettabilitycharacteristics of the MgO–PSZ. The correlation between the wettabilitycharacteristics and the osteoblast cell response implies that the surfaceroughness, surface oxygen content and gp

sv have some bearing on the cellresponse owing to the fact that these surface properties influence thewettability characteristics of the MgO–PSZ. In the consideration of surfaceroughness in Section 5.5.3, it was found that surface roughness doesinfluence the response of osteoblast cells, but its effect is slight. The surfaceoxide, among the surface characteristics, is shown to be a main factorinfluencing the cell response besides surface roughness [52, 221]. In vivostudies show that a high degree of bone contact and bone formation areachieved with titanium implants which are modified with respect to oxidethickness and surface topography [222]. As shown in Chapter 4, the CO2

laser processing of the MgO–PSZ with the oxygen shield gas resulted in theincorporation of oxygen atoms on the material’s surface layer. It is postu-lated that the surface oxygen content is one of the factors influencing cellgrowth. It can be seen from Figure 5.19 that the cell cover density on thematerial is higher when there is a higher surface oxygen content, indicatingthat the increase in surface oxygen content is attributed to better cell growth.However, there is no linear relationship between the surface oxygen contentand cell cover density, implying that some other mechanism is moredominant in governing the cell response.

In Chapter 4 it was found that changes in the wettability characteristics ofthe MgO–PSZ were primarily influenced by gp

sv of the MgO–PSZ. As one cansee from Figure 5.20, the cell cover density statistically increases as gp

sv

increases from 10 to 25 mJ/m2. Hence, the higher surface energy induced bythe CO2 laser treatment resulted in better proliferation of human osteoblast


cells on the MgO–PSZ surface than on the untreated MgO–PSZ. A furtherincrease in gp

sv from 25 to 60 mJ/m2 did not bring about an increase in cellproliferation that was statistically significant. As the gp

sv did not changemarkedly after CO2 laser treatment, the results indicated that gp

sv influencedthe behaviour of the osteoblasts on MgO–PSZ surfaces more stronglycompared to gd

sv following CO2 laser treatment, which was probably attrib-uted to the fact that the composition and the culture medium all are polar,and thus cells and the MgO–PSZ should interact mainly by polar force.

40 45 50 55 60 65 70 750

20

40

60

80

100

Surface Oxygen Content (at%)

Cel

l Cov

er D

ensi

ty (

%)

Figure 5.19 The relationship between the cover density of the hFOB cells and thesurface oxygen content of the MgO–PSZ

10 20 30 40 50 600

20

40

60

80

100

Cel

l Cov

er D

ensi

ty (

%)

psvγ (mJ/m2)

Figure 5.20 The relationship between the cover density of the hFOB cells and gpsv for

the MgO–PSZ


The behaviour of osteoblastic cells at the surface of hydroxyapatite [218]and at the surface of titanium [213] demonstrated that gp

sv plays a criticalrole.

5.6 Predictions for Implantation in an in vivo Clinical Situation

A relationship between the bioactivity/biocompatibility of a material in vivoand the material’s ability to form an apatite-like layer in vitro when soaked inaqueous solutions that imitate the inorganic components of human plasmawas propounded recently by Vallet-Regi [223]. This proposition extends toceramics, for once implanted into the body, bioactive ceramics are able tobond to bone through the formation of a hydroxyapatite surface layer. Thesehydroxyapatite layers that are formed in vivo can be closely mimickedthrough in vitro testing, usually by using SBF, which contains ionic concen-trations similar to that of human blood plasma. This allows for an evaluationof the potential for biocompatibility/bioactivity of a new ceramic, or a novelprocessing technique for a ceramic can be conducted before performing anyanimal tests.

As the preceding sections of this chapter show, CO2 laser treatment couldgenerate functional groups and subsequently facilitate the formation ofbone-like apatites on the surface of the MgO–PSZ. Building on this findingby considering the view of Vallet-Regi [223], it is reasonable to assume that,if implanted, the in vivo performance of the CO2 laser treated MgO–PSZwould be acceptable. There is certainly a considerable body of evidence thatsupports this supposition.

In comprehensive studies by Aldini et al. [224, 225] and Torricelli et al.[226], yttria–stabilised tetragonal zirconia (Y-STZ), either coated with abioactive glass termed RKKP or uncoated, was evaluated in vitro usingnormal and osteopenic bone-derived osteoblasts and in vivo using healthybone in female Sprague Dawley rats. The in vitro results suggested that theRKKP-coated Y–STZ was biocompatible and enhanced proliferation, activa-tion and differentiation of normal bone-derived osteoblasts and stimulationof osseopenic bone-derived osteoblasts when compared with the uncoatedY–STZ. To assess the in vivo performance of the RKKP-coated and uncoatedY–STZ, samples of each were implanted into the distal femurs of the rats andthen removed after 30 and 60 days from surgery, whereupon they weresubjected to a histomorphometrical analysis to assess osseointegration andbone quality around the implants. The histomorphometrical analysisshowed that the RKKP coating ensured a better osseointegration rate withhigher affinity index values than the uncoated Y–STZ, even when theosseopenic rate were used. No differences were observed at the bone–biomaterial interfaces for either material. In this instance there was a direct

Predictions for Implantation in an in vitro Clinical Situation 95

correlation between the findings of the in vitro and the in vivo investigationsin terms of biocompatibility and osseointegration, but not for the case ofbone-biomaterial interfacial characteristics.

In a study conducted to evaluate the biocompatibility of MgO-doped HA(hydroxyapatite)/b-TCP (tricalcium phosphate) biphasic ceramics, Ryuet al. [227] performed in vitro and in vivo tests. The in vitro tests were carriedout with a murine fibroblast L929 cell culture of extract from a 1 wt % MgO-doped HA/b-TCP ceramic with the aim of establishing whether theceramics were biocompatible or cytotoxic. The presence or absence of anycytotoxic effects was determined qualitatively using SEM analysis. The SEManalysis following the cell culture from the 1 wt % MgO-doped HA/b-TCPceramic revealed no morphological change, no vacuolisation or cell lysisand hence the absence of any cytotoxicity. In the first of two in vivo testssmall samples of the 1 wt % MgO-doped HA/b-TCP ceramic were insertedinto the back muscles of mature rabbits and then removed after 8 weeks forXRD analysis. The XRD analysis showed that the 1 wt % MgO-doped HA/b-TCP ceramic was indeed biocompatible as the b-TCP phase had comple-tely dissolved and an apatite layer had formed on the surface by means ofcellular activity and a resultant dissolution/precipitation process. Thesecond in vivo test involved implanting the same small 1 wt % MgO-doped HA/b-TCP ceramic samples that were inserted into the back musclesof the rabbits for 8 weeks across the proximal tibial metaphysis of maturerabbits for a further 8 weeks. From a histological examination of the samplesno inflammation or foreign body reaction such as the formation of anintervening fibrous tissue was observed; rather the new bone formed andbonded directly to the implanted 1 wt % MgO-doped HA/b-TCP ceramic.Clearly, there was a direct link between the in vitro results and theobservations made from both of the in vivo tests, with the in vitro testsindicating that the 1 wt % MgO-doped HA/b-TCP ceramic was biocompa-tible, the first in vivo tests demonstrating that the 1 wt % MgO-doped HA/b-TCP ceramic could form an apatite layer and the final in vivo testsshowing that the 1 wt % MgO-doped HA/b-TCP ceramic could supportthe bonding of new bone to its surface.

Literature exists that shows that the in vitro bioactivity of a material hasbeen observed to be dependent upon the type of aqueous solution used anddoes not always correlate to the observed in vivo behaviour. For example,apatite–wollastonite (A–W) glass–ceramics, which form a hydroxyapatitesurface layer when immersed in SBF, are unable to form such layers in trisbuffer [228]. In addition, SBF has been reported to affect the rate at whichcrystalline hydroxyapatite layers develop [229]. Indeed, the work ofGil-Albarova et al. [230] studied the in vivo behaviour of an SiO2–P2O5–CaO sol-gel glass and an SiO2–P2O5–CaO–MgO glass–ceramic, both of whichare bioactive when soaked in SBF but display different rates of apatite layer


formation. The ceramics were implanted into mature and immature NewZealand rabbits and histological results after 6 and 12 weeks revealed thatboth ceramics allowed bone growth over their surface in similar quantitiesand at similar rates by means of mesenchymal cell recruitment from thesurrounding bone.

Similarly, the in vitro and in vivo behaviour of certain ceramics has beenseen to differ on account of the ceramic material itself. For instance, sinteredhydroxyapatite exhibits in vivo bioactivity although the formation kinetics ofan apatite-like layer on its surface is very slow under in vitro conditions[228]. Additionally, Li et al. [8] reported that that Al2O3 gel did not induceapatite formation when immersed in SBF for 21 days, whereas both pureSiO2 gel and gel-derived TiO2 were hydroxyapatite inducers.

Kobayashi et al. [231] developed a composite (designated ABC) consistingof Al2O3 bead powder as an inorganic filler and bisphenol-a-glycidylmethacrylate (bis-GMA) based resin as an organic matrix, which allowsdirect bone formation on its surface in vivo. Although bioactive materialssuch as bioglass or apatite and wollastonite-containing glass–ceramic havepreviously been reported to form bone-like apatite on their surfaces in vitrounder acellular conditions via simple chemical reactions, ABC did notpresent such characteristics. Indeed, no apatite formation was detected onthe surfaces of the ABC composite after soaking in SBF for 28 days in vitro.Histological examination of rat tibiae after 8 weeks revealed that the ABCcomposite bonded to bone directly via a layer of calcium, phosphorus andalumina with no interposed soft-tissue layer. Moreover, the amount of bonedirectly apposed to the ABC composite surface was seen to increase withtime. These results imply that the ABC composite has the ability to bonddirectly with bone through a calcium–phosphorus-rich layer. It was con-cluded that this layer was induced by some property of the ABC compositethat encouraged calcification or apatite formation due to the actions ofproteins and cells in vivo.

Although the aqueous solutions employed to replicate inorganic bodyfluids are able to reproduce the process of bone-like apatite formation on thesurfaces of bioactive ceramics in vitro, they are only capable of evaluating thebone-bonding capacities of bioactive ceramics by means of simple chemicalreactions. However, once a ceramic is implanted into the body, they elicitseveral responses from living tissue that cannot be simulated in vitro. Theseresponses include protein adsorption, cell attachment and adhesion, aswell as ionic exchange. Therefore, depending upon the actual ceramicitself and the aqueous solution used, a ceramic could well be biocompatiblein vivo despite appearing to be cytotoxic in vitro. However, importantly forthis work, it is clear that if a ceramic presents itself to be biocompatibleafter in vitro testing, then it is highly likely that it will perform satifactorilyin vivo.

Predictions for Implantation in an in vitro Clinical Situation 97

5.7 Summary

A CO2 laser was used to modify the surface of the MgO–PSZ for the purposeof acquiring surface properties favouring interaction at the implant–bonetissue interface. The bioactivity of the CO2 laser modified MgO–PSZ wasinvestigated in SBF. Protein adsorption and hFOB cells were used toexamine the in vitro biological response on the MgO–PSZ following CO2

laser treatment.It was demonstrated that CO2 laser treatment could improve the bioactiv-

ity of the MgO–PSZ surface by generating functional groups to facilitate theformation of bone-like apatites. The apatite formed readily on the MgO–PSZwith relatively high amounts of hydroxyl groups, which were generated byCO2 laser treatment with power densities of 1.6 and 1.9 kW/cm2. No apatitewas observed on the untreated and CO2 laser modified samples (0.6, 0.9 and2.5 kW/cm2), which exhibited few hydroxyl groups. These observationsindicate that Zr–OH groups on the MgO–PSZ surface are the functionalgroups required to facilitate apatite formation. The melting and re-solidifica-tion on the surface of the MgO–PSZ induced by CO2 laser processingprovides the Zr4þ ion and OH� ion and therefore creates the Zr–OHgroup on the surface.

In comparison with the untreated MgO–PSZ, CO2 laser treatment broughtabout a thinner adsorbed albumin layer and a thicker adsorbed fibronectinlayer on the MgO–PSZ. As the wettability characteristics of the MgO–PSZincreased, the albumin adsorption decreased while the fibronectin adsorp-tion increased, indicating that wettability is a major factor in governingprotein adsorption. Further, the correlative effect of gp

sv observed on theprotein adsorption suggested that protein adsorption on the MgO–PSZ wasprobably due to the polar and chemical interactions.

Better osteoblast cell responses were found on the CO2 laser treated MgO–PSZ when compared with the untreated sample. The change in topographyinduced by the CO2 laser treatment was identified as being one of the factorsinfluencing the hFOB cell response, but in a minor capacity only. Theimproved wettability characteristics of the MgO–PSZ due to enhancedsurface energy, especially the polar component, brought by the CO2 lasertreatment, played a significant role in the number of initial cells that attachedand spread, thereby enhancing the long-term cell adhesion and growthpotential.

There is a reasonable body of literature to support the concept that if aceramic presents itself to be biocompatible after in vitro testing then it ishighly likely that it will perform satisfactorily in vivo. This being the case,then it is reasonable to assume that the in vivo performance of the CO2 lasertreated MgO–PSZ would be acceptable due to its excellent in vitro perfor-mance demonstrated herein.


6

The Effects of CO2 LaserRadiation on the WettabilityCharacteristics of a TitaniumAlloy

This chapter describes the modification of the wettability characteristics of atitanium alloy (Ti–6Al–4V ELI) following CO2 laser irradiation. The Ti–6Al–4Valloy is often used for the fabrication of dental and orthopaedic implants. To studythe change in the wettability characteristics of the Ti–6Al–4V alloy, contact anglesbetween selected control test liquids and the surfaces of the untreated and CO2 lasertreated Ti–6Al–4V alloy were measured. The surface properties of the untreated,mechanically roughened and CO2 laser treated Ti–6Al–4V alloy were characterisedand the effect of surface roughness, surface oxygen content and surface energy on thewettability characteristics of the Ti–6Al–4V alloy were analysed. It was apparentthat CO2 laser treatment brought about significant changes in the wettabilitycharacteristics of the Ti–6Al–4V alloy. Furthermore, the predominant mechanismsactive in determining the wettability characteristics were analysed and the primarymechanism was identified.

6.1 Introduction

During the last decade, numerous materials have been used for the fabrica-tion of dental and orthopaedic implants. The materials of choice havepredominantly been metals. Stainless steel, cobalt chromium molybdenumalloy, titanium and a multitude of titanium alloys were the materials of


choice. The following criteria define an ideal bone contacting material fororthopaedic surgery: a biocompatible chemical composition to avoid adversetissue reactions, acceptable strength, a high wear resistance to minimisewear debris, excellent corrosion resistance in the physiological milieu and amodulus of elasticity similar to that of bone to minimise bone resorptionaround the device. Titanium and its alloys, including the Ti–6Al–4V alloy,are now being used as a common material for bone implants underbiomechanical loading conditions. None of these bioinert titanium basedmaterials, however, bonds to bone and subsequently their stable fixations tothe surrounding bone have long been considered as a fundamental problemin clinical uses [232].

It is generally accepted that early surface events that occur rapidly uponimplantation of a biomaterial into biological fluids determine a subsequentresponse. These involve wetting by physiological liquids, followed byadsorption of proteins and cells to the biomaterials surface [233]. Numerousresearch groups have studied the interactions of different types of culturedcells with biomaterials with different wettability characteristics to correlatethe relationship between surface wettability and blood, cell or tissuecompatibility for polymeric materials [59, 115, 234]. Furthermore, the surfacewettability has a significant influence on the friction behaviour of a tribolo-gical system and gives an indication of its biotolerance: in a first approxima-tion, the more wettable the material, the better the human body tolerates it[235]. Clearly, techniques to control the wettability characteristics of abiomaterial’s surface and thereby improve the material’s biocompatibilityare of great interest.

Due to the rapid and specific modification of organic and inorganicmaterials, laser surface processing has aroused growing interest and beenproven to be a controllable and flexible technique for modifying the surfaceproperties of materials. It is recognised within the currently publishedwork that laser irradiation of material surfaces can affect their wettabilitycharacteristics. Previously Heitz et al. [133], Henari and Blau [134] and Olfertet al. [135] had found that excimer laser treatment of metals results inimproved coating adhesion, attributed to the fact that the excimer lasertreatment resulted in a smoother surface and as such enhanced the action ofwetting. It was demonstrated that five pulses per area of CO2 laser treat-ment were sufficient to produce a fully wettable mild steel surface. Thewettability was influenced by the surface exposing time (SET) after lasertreatment [136]. Self-fluxing Fe–Cr–Ni–B–Si alloy powders with variousNi contents were laser clad on medium carbon steel substrates [137].Lawrence and Li [138] revealed that the interaction of CO2, Nd:YAG,HPDL and excimer laser radiation with the surface of a selected mild steelgave rise to changes in the wettability characteristics of the material. It wasobserved that interaction of the mild steel with Nd:YAG and HPDL

100 The Effects of CO2 Laser Radiation

radiation brought about an improvement in the wettability characteristics ofthe steel. In contrast, interaction of the mild steel with CO2 and excimer laserradiation resulted in a depreciation of the wettability characteristics of thesteel [236]. However, despite a growing amount of work conducted withmetal, no work has been conducted so far on the feasibility of the lasersurface treatment process for the modification of the wettability character-istics of biograde metals.

6.2 Experimental Procedures

6.2.1 Material Specifications and Preparation

Medical grade titanium alloys have a significantly higher strength-to-weightratio than competing stainless steels. The range of available titanium alloysenables medical specialist designers to select materials and forms closelytailored to the needs of the application. The natural selection of titanium forimplantation is determined by a combination of most favourable character-istics, including immunity to corrosion, biocompatibility, strength, lowmodulus and density and the capacity for joining with bone and othertissue (osseointegration). The mechanical and physical properties of tita-nium alloys combine to provide implants that are highly damage tolerant.Forms and material specifications of titanium and its alloy for medicalapplication are detailed in a number of international specifications.

In this study a Ti–6Al–4V ELI alloy (F136) was used. The as-received Ti–6Al–4V alloy (ground annealed) was in the form of a round bar with adiameter of 28.5 mm (Carpenter, Inc.). For experimental purposes, the roundbar was divided into 15 sections, each of 3 mm thickness, by a cuttingmachine (Miniton; Struers, GmbH) using a diamond-rimmed blade. The 15divided sections were then separated into three groups of five samples, withthe groups being: untreated, mechanically roughened and CO2 laser treated.For the mechanically roughened group, the samples were roughened byevenly abrading the entire surface of the sample with grinding paper (180grit SiC). This was achieved by applying the grinding paper to the surface ofthe sample with moderate pressure and drawing it across the surface indifferent directions eight times. The composition of the Ti–6Al–4V alloy was:88.3–90.8 wt % Ti, 5.5–6.5 wt % Al, 3.5–4.5 wt % V, <0.08 wt % C, 0.0125 wt%H, <0.25 wt % Fe, <0.05 wt % N and <0.13 wt % O. The main physicalproperties of the Ti–6Al–4V alloy are: density of 4.42 g/cm2, melting pointof 1649 � 15 �C, a specific heat at 25 �C of 560 J/kg �C and a thermalconductivity at 25 �C of 7.2 W/mK. The mechanical properties are: tensilestrength of 1000 MPa, an elastic modulus of 114 GPa and a Rockwellhardness of 36 C.

Experimental Procedures 101

6.2.2 CO2 Laser Surface Treatment

The surface treatment of the Ti–6Al–4V alloy was completed with the 3 kWCO2 laser described in Section 4.2.2. As before, the laser was run in the CWmode. The CO2 laser beam was defocused to a 7 mm spot diameter andoperated with laser powers that yielded laser power densities of 1.3 and1.6 kW/cm2. The CO2 laser beam was traversed across the surface of the Ti–6Al–4V alloy at a speed of 4800 mm/min. The fumes produced wereremoved with an extraction system while an O2 assist gas was supplied at2 bar pressure to shield the laser optics. The set-up of the CO2 laserexperiment was as described in Chapter 4 and is shown in Figure 4.1.

6.2.3 Morphological, Chemical and Phase Analysis Procedures

The morphological characteristics of the untreated, mechanically roughenedand CO2 laser treated Ti–6Al–4V alloy samples were examined by means ofsurface roughness measurements of the samples and through SEM analysis.The chemical composition of the Ti–6Al–4V alloy before and after CO2 lasertreatment was determined using SEM, EDX and XPS analysis, while thephase was observed with XRD across a range of 30 to 80 2-theta. Theequipment details and manner in which they were operated are describedin Section 4.2.3.

6.2.4 Wettability Characteristics Analysis Procedure

To examine the wetting and surface energy characteristics of the Ti–6Al–4Valloy when in the untreated, mechanically roughened and CO2 laser treatedconditions, wetting experiments were conducted. A set of sessile dropcontrol experiments was carried out using glycerol, formamide, ethenegly-col, polyglycol E-200 and polyglycol 15-200. The characteristics of the controltest liquids are given in Table 4.1. The y values for the control test liquids onthe untreated, mechanically roughened and CO2 laser treated Ti–6Al–4Valloy were determined in atmospheric conditions at 25 �C using a sessiledrop measuring machine in the same manner as described in Section 4.2.4.

In order to estimate the influence of contaminant layers on the measured yresults, the untreated Ti–6Al–4V alloy samples were cleaned in the samemanner as described in Section 4.2.4. The values of y on the cleaned sampleswere lower than on the as-received (not cleaned) samples by 0.8, 0.7, 0.5, 0.3and 0.2� for glycerol, formamide, etheneglycol, polyglycol E-200 and poly-glycol 15-200, respectively. It is therefore reasonable to conclude that thepresence of any contaminants on the surface of the Ti–6Al–4V alloy do nothave a great influence on y. This being the case, the work was conducted in anormal atmospheric environment without pre-cleaning with the samejustification as that given in Section 4.2.4.


6.3 The Effects of CO2 Laser Radiationon Wettability Characteristics

6.3.1 Contact Angle

The mean values of y for the five control test liquids measured on theuntreated, mechanically roughened and CO2 laser treated Ti–6Al–4V alloyare given in Table 6.1. As one would expect, in accord with Equation (4.1),reductions in y for all of the control test liquids were occasioned by themechanical roughening of the surface of the Ti–6Al–4V alloy. It is clear,however, that these reductions were only moderate. In contrast, CO2 lasertreatment of the surface of the Ti–6Al–4V alloy at both 1.3 and 1.6 kW/cm2

resulted in marked decreases in y with all of the control test liquids, with thegreatest reductions in y occurring when the power density was 1.6 kW/cm2.

6.3.2 Morphological Analysis and Its Effecton Wettability Characteristics

The typical surface view of the untreated Ti–6Al–4V alloy, as shown inFigure 6.1(a), exhibited regularly ordered parallel and longitudinal grooves.From Figure 6.1(b) it can be seen that the mechanically roughened Ti–6Al–4V alloy samples also presented a grooved surface. However, in contrastwith the untreated Ti–6Al–4V alloy surface, the mechanically roughenedsamples typically displayed no surface ordering, with the grooves beingorientated in various directions. The surfaces of both of the CO2 laser treatedTi–6Al–4V alloy samples retained a similar morphology to that of theuntreated sample insofar as parallel and longitudinal grooves that areregular in size and periodicity are present (see Figures 6.1(c) and (d)).However, as one can see from Figures 6.1(c) and (d), it appears that CO2

laser radiation, when incident with the Ti–6Al–4V alloy surface, resulted in

Table 6.1 Measured contact angle values for the untreated, mechanicallyroughened and CO2 laser treated Ti–6Al–4V alloy

Contact angle, y (deg)————————————————————————————

CO2 laser treated (kW/cm2)Mechanically

Control test liquid Untreated roughened 1.3 1.6

Glycerol 85.8 78.7 74.4 70.1Formamide 83.6 74.5 72.6 67.6Etheneglycol 77.3 70.9 68.4 62.9Polyglycol E-200 71.1 64.5 63.2 57.3Polyglycolycol 15-200 67.1 55.3 53.5 48.1

The Effects of CO2 Laser Radiation on Wettability Characteristics 103

partial melting and re-solidification. By comparing Figures 6.1(c) and (d) it isevident that the melting and re-solidification of the surfaces of the Ti–6Al–4V alloy occurred in varying degrees.

The extent to which melting and re-solidification differs on the surface ofthe CO2 laser treated Ti–6Al–4V alloy samples is due entirely to thedifference in the CO2 laser power density used to treat the surfaces. Inthe case of the sample treated at the CO2 laser power density of 1.6 kW/cm2,the higher power used would result in more energy being obsorbed, therebycausing more surface melting to occur. In contrast, the sample treated at theCO2 laser power density of 1.3 kW/cm2 was treated with less power and soless energy was absorbed. It is obvious that the specimen treated at the CO2

laser power density of 1.6 kW/cm2 experienced more melting and re-solidification and so the depth of the grooves was much greater (seeFigure 6.1(d)), whereas the specimen treated at the CO2 laser power densityof 1.3 kW/cm2 underwent melting and re-solidification to a lesser extent and

Figure 6.1 Typical SEM surface images of the Ti–6Al–4V alloy (a) untreated,(b) mechanically roughened and CO2 laser treated at power densities of (c)1.3 kW/cm2 and (d) 1.6 kW/cm2


thus the depth of the grooves was shallower (see Figure 6.1(c)). It is worthremarking that since the surface morphology of the two CO2 laser treatedsamples, while being different to the untreated sample, are not dramaticallydifferent from each other, it is therefore possible to assert that the energyabsorbed was neither insufficient nor excessive.

From Figure 6.2 one can see that the surface of the Ti–6Al–4V alloy in theuntreated condition was reasonably smooth, having an average surfaceroughness (given in terms of Ra) value of 0.35mm. Figure 6.2 reveals thatthe average surface roughness value of the Ti–6Al–4V alloy did not alter agreat deal as a result of CO2 laser treatment, increasing to 0.39mm whentreated with a laser power density of 1.3 kW/cm2 and to 0.42mm whentreated with a laser power density of 1.6 kW/cm2. Conversely, Figure 6.2shows that mechanical roughening of the Ti–6Al–4V alloy surface had moreof a marked effect on surface roughness. In this instance, CO2 laser treatmentbrought about an increase in the average surface roughness to 0.47mm. Inboth instances these reductions in y with corresponding increases in surfaceroughness are in accord with Equation (4.1). These values of average surfaceroughness are reflected by the images shown in Figure 6.1.

6.3.3 Phase and Chemical Analysis and Its Effectson Wettability Characteristics

The results of an XRD analysis on the Ti–6Al–4V alloy surface before andafter CO2 laser treatment are shown in Figure 6.3. The mechanicallyroughened Ti–6Al–4V alloy samples were not subjected to an XRD analysis

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0.2

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ace

Rou

ghne

ss, R

a (µ

m)

Untreated MechanicallyRoughened

CO2 laser1.3 kW/cm2

CO2 laser1.6 kW/cm2

Figure 6.2 Mean values of surface roughness for the untreated, mechanicallyroughened and CO2 laser treated Ti–6Al–4V alloy


as the roughening process would not have induced any phase change. Onthe whole, CO2 laser treatment appears not to have altered the phasespresent in the Ti–6Al–4V alloy (see Figures 6.3(b) and (c)). The results of anEDX analysis on the surface of the Ti–6Al–4V alloy are shown in Figure 6.4.Again, it appears that CO2 laser treatment did not alter the chemical

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nsi

ty

Ti(110)

Ti(101)

Ti(102)

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V(110)

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V(211)

Al(311)

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(200)

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Rut

ile

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Al(220)

Ti(201)

Al(311)

V(211)

Rut

ile

2 theta

(a)

2 theta

(b)

2 theta

(c)

Figure 6.3 XRD analysis of the Ti–6Al–4V alloy (a) untreated and CO2 laser treated atpower densities of (b) 1.3 kW/cm2 and (c) 1.6 kW/cm2


composition of the surface of the Ti–6Al–4V alloy, since Ti, Al and V, thebasic elements of Ti–6Al–4V alloy, were still present in similar proportionson the surface of the alloy before and after CO2 laser treatment.

Having said that, the XRD analysis did reveal that an oxide diffractionpeak at 39.5� was generated after CO2 laser treatment (see Figures 6.3(b) and(c)). This finding is evidence that the surface of the Ti–6Al–4V alloy wasoxidised after the CO2 laser process, with the oxide layer consisting of rutile.This oxidation of the Ti–6Al–4V alloy samples came about because thesurface was superficially melted by the CO2 laser beam during treatments.This in turn led to oxygen diffusion through the molten material andsubsequently to the oxidation of the Ti–6Al–4V alloy surface. Of particularnoteworthiness is the observation that can be made by comparing

Figure 6.4 EDX analysis of the Ti–6Al–4V alloy (a) untreated and CO2 laser treated atpower densities of (b) 1.3 kW/cm2 and (c) 1.6 kW/cm2


Figures 6.3(b) and (c), namely that the level of oxidation, as indicated by thediffraction peak at 39.5�, increases with increasing CO2 laser power density.Such an observation supports the proposition given earlier that the oxidationof the Ti–6Al–4V alloy surface following CO2 laser treatment is broughtabout by oxygen diffusion through the molten material. Hence the moremolten material there is and the longer the material remains in the moltenstate, conditions that result from an increase in CO2 laser power density, thegreater the degree of oxidation.

XPS was used to examine the surface oxygen content on the surface of theTi–6Al–4V alloy before and after CO2 laser treatment (see Figure 6.5). As isevident from Figure 6.5, the surface oxygen content on the Ti–6Al–4V alloyincreased after CO2 laser treatment. This increase in the oxygen content onthe surface of the Ti–6Al–4V alloy was naturally due to oxidisation of theCO2 laser treated surfaces, which was shown to occur from the XRD analysis(see Figures 6.3(b) and (c)). A similar observation was found in the work ofLawrence and Li [138], where the surface oxygen content of a carbon steelincreased after CO2 laser treatment. The specimen CO2 laser treated at apower density of 1.6 kW/cm2 had the highest surface oxygen content. Thisfinding further supports the proposition that with an increase in laser powercomes a greater degree of melting, thereby inducing a larger amount ofoxygen to be absorbed into the Ti–6Al–4V alloy surface.

From the results of y measurements between the control test liquids andthe Ti–6Al–4V alloy in various conditions of treatment (see Table 6.1), it is

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CO2 laser1.3 kW/cm2

CO2 laser1.6 kW/cm2

Surf

ace

Oxy

gen

Con

tent

(at

%)

Figure 6.5 Mean values of the surface oxygen content for the untreated, mechani-cally roughened and CO2 laser treated Ti–6Al–4V alloy


clear that the CO2 laser treatment increased the wettability characteristics ofthe Ti–6Al–4V alloy. This observed increase in the wetting performanceof the Ti–6Al–4V alloy would have certainly been influenced by the increasein the oxygen content of the Ti–6Al–4V alloy surface as a result of the CO2

laser treatment (see Figure 6.5), since this is known to increase the likelihoodof wetting [13, 146].

6.4 Surface Energy and Its Component Analysis

Measurements of y between the Ti–6Al–4V alloy and the control test liquids(see Table 6.1) shows that CO2 laser surface treatment improved thewettability characteristics of the material. As has already been demonstratedin Chapter 4, it is possible to estimate reasonably accurately the gd

sv of the Ti–6Al–4V alloy by plotting the graph of cos y against (gd

lv)12=glv) in accordance

with Equation (3.8). Figure 6.6 shows the best-fit plot of cos y against(gd

lv)12=glv) for the untreated, mechanically roughened and CO2 laser treated

Ti–6Al–4V alloy–experimental control liquids system.Comparing the ordinate intercept points of the untreated, mechanically

roughened and CO2 laser treated Ti–6Al–4V alloy–liquid systems, as shownin Figure 6.6, it can be seen clearly that for the untreated and mechanicallyroughened Ti–6Al–4V alloy, the best-fit straight line intercepts the ordinate

0.00 0.03 0.06 0.09 0.12 0.15 0.18 0.21 0.24−1.0

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UntreatedMechanically Roughened

CO2 laser (1.3 kW/cm2)

CO2 laser (1.6 kW/cm2)

cos θ

( ) lvdlv γγ /

2/1


12=glv for the untreated, mechanically roughened

and CO2 laser treated Ti–6Al–4V alloy in contact with the control test liquids

Surface Energy and Its Component Analysis 109

closer to the origin. This is noteworthy as the intercept of the ordinate closeto the origin is characteristic of the dominance of dispersion forces acting onthe Ti–6Al–4V alloy material–liquid interfaces of the untreated and mechani-cally roughened sample, resulting in poor adhesion [100, 101]. On the otherhand, the best-fit straight line of samples treated by the CO2 laser interceptsthe ordinate considerably higher above the origin. An interception of theordinate above the origin is indicative of the action of polar forces across theinterface, in addition to dispersion forces, and hence improved wettabilityand adhesion are promoted [100, 101]. Furthermore, because none of thebest-fit straight lines intercepts below the origin, it can be said that thedevelopment of an equilibrium film pressure of adsorbed vapour on the Ti–6Al–4V alloy surface (untreated, mechanically roughened and CO2 lasertreated) did not occur [101].

As was shown previously in Chapter 4, in order to determine the gpsv of

the Ti–6Al–4V alloy, it is necessary to calculate values of Wad calculatedusing Equation (3.4) and Wd

ad calculated using Equation (3.9). Both Wad

and Wdad are related by the straight line relationship represented by Equation

(3.10). On account of this it is possible from the best-fit straight line plots ofWad against Wd

ad to determine the constant, a, for each separate condition ofthe Ti–6Al–4V alloy: 2.38 in the untreated condition, 2.66 when mechanicallyroughened, 2.80 when CO2 laser treated at 1.3 kW/cm2 and 2.84 when CO2

laser treated at 1.6 kW/cm2. Since a linear relationship exists betweenthe dispersive and polar components of the control liquids’ surface energiesthat satisfies Equation (3.11), the constant, c, is determined to be 2.9(see Section 4.4). It is therefore possible to calculate gp

sv directly forthe untreated, mechanically roughened and CO2 laser treated Ti–6Al–4Valloy using Equation (3.14) with the determined values of the constants,a and c.

As can be seen from Figure 6.7, CO2 laser treatment of the surface of theTi–6Al–4V alloy led to an increase in the total surface energy. In particular,the CO2 laser treatment brought about a considerable increase in thevalue of gp

sv, which is known to have a positive effect upon the action ofwetting and adhesion [237]. The changes in the surface energy values of theTi–6Al–4V alloy are thought to be due to the fact that CO2 laser treatmentresults in the melting and re-solidification of the Ti–6Al–4V alloy surface, atransition that is known to effect an increase in gp

sv [103, 130, 131]. Thisexplains why, after CO2 laser treatment, the Ti–6Al–4V alloy specimens havea higher propensity for wetting. As before, owing to the higher degree ofmelting on the Ti–6Al–4V alloy surface occasioned by CO2 laser treatment ata power density of 1.6 kW/cm2, the increase in the value of gp

sv was morethan for the Ti–6Al–4V alloy sample CO2 laser treated at a power density of1.3 kW/cm2.


6.5 Identification of the Predominant Mechanisms Activein Determining Wettability Characteristics

Based on the findings of the preceding sections it is apparent that modifica-tions to the wettability characteristics of the Ti–6Al–4V alloy following CO2

laser treatment are attributable to changes in the surface roughness, surfaceO2 content and surface energy. The dependency of y to surface roughness iswell known and, moreover, surface roughness has been identified as thepredominant factor governing changes in wettability characteristics of steelafter surface treatment with various lasers [138, 236]. Additionally, thesurface oxygen content has been found to be a factor contributing to theenhancement of the wettability characteristics of a number of materials:MgO–PSZ after CO2 laser treatment (see Chapter 4) and steel followingirradiation with a HPDL [236]. Further, increases in gp

sv due to laser-inducedsurface melting and re-solidification of steel [138, 236] and various ceramicmaterials have been seen to be influential in effecting improvements in thewettability characteristics of the materials. What is more, the increase in gp

sv

of MgO–PSZ after CO2 laser surface treatment was identified as the majorfactor in determining the wettability characteristics of MgO–PSZ (seeChapter 4). Naturally, it would be advantageous to determine whether thechanges to the surface roughness, the surface oxygen content or the surface

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CO2 laser1.3 kW/cm2

CO2 laser1.6 kW/cm2

Surf

ace

Ene

rgy

(mJ/

m2 )

Dispersive ComponentPolar ComponentTotal

Figure 6.7 Measured total surface energy gsvð Þ, dispersive (gdsv) and polar (gp

sv)components for the untreated, mechanically roughened and CO2 laser treatedTi–6Al–4V alloy

Identification of the Predominant Mechanisms 111

energy, namely the increases in gpsv (which is determined by the micro-

structure), influenced the observed increase in the wettability of the Ti–6Al–4V alloy after CO2 laser surface treatment, either independently or incombination with one another. To do this, several stages of fine grindingwere used to isolate the various influential factors detailed above and thusanalyse and qualitatively establish the effect each one had on the wettabilitycharacteristics of the Ti–6Al–4V alloy.

The fact that the value of y was seen to decrease by simply roughening thesurface mechanically (see Table 6.1) suggests that the surface roughness mayplay a major part in determining the wettability characteristics of the Ti–6Al–4V alloy. Having said that, Table 6.1 shows that the value of y wasconsistently lower on the CO2 laser treated samples than on the mechani-cally roughened sample, despite the fact that the surface roughness valuewas higher on the mechanically roughened sample. This implies that notonly are other factors besides surface roughness active in promoting areduction in y, they are significant in their influence.

In the first fine grinding stage the surfaces of the untreated and CO2 lasertreated Ti–6Al–4V alloy (1.6 kW/cm2) were ground down to an Ra value ofaround 0.30 mm with grinding paper (800 grit SiC), while still retaining theCO2 laser treated microstructure. In this way it was possible to isolate andassess the effect of surface roughness. In addition the first fine grinding-down stage would allow one to investigate the effects of the surface oxygencontent as it is present only within the first atomic layers of the alloy. Inorder to evaluate the influence of the CO2 laser induced microstructure, andin turn the surface energy, a second fine grinding-down stage was under-taken using grinding paper (800 grit SiC) to remove the microstructure butretain the surface roughness value obtained after the first fine grinding stage.Following on from this, previously un-ground samples of both the untreatedand the CO2 laser treated Ti–6Al–4V alloy were ground up with grindingpaper (400 grit SiC) to an Ra value of around 0.45mm in order to confirm theeffect of surface roughness and investigate the role played by surface energy.In this first fine grinding-up stage the CO2 laser induced microstructure wasretained. A second fine grinding-up stage was then carried out to remove theCO2 laser induced microstructure and hence confirm the effect of surfaceroughness. The observed changes to surface roughness, surface oxygencontent and y for glycerol brought about after each of these fine grindingstages are given in Table 6.2.

Confirmation of the predominant role played by surface roughness isclearly evident by the results obtained after the first fine grinding-downstage. From Table 6.2 it can be seen that for both the untreated and CO2 lasertreated Ti–6Al–4V alloy samples, y increases and in the case of the CO2 lasertreated sample, the increase is considerable. However, despite the factthat the surface roughness values for the untreated and CO2 laser treated


Ti–6Al–4V alloy samples are similar after the first fine grinding-down stage,there is a discernible difference in the value of y, being 0.9� lower for the CO2

laser treated sample. Whereas the surface oxygen content of the untreatedTi–6Al–4V alloy sample remained around the original value of 23.32 at %,the surface oxygen content of the CO2 laser treated Ti–6Al–4V alloy samplereduced to a level similar to that of the untreated sample. Because the firstfine grinding-down stage does not remove the CO2 laser induced micro-structure, this difference in y can, therefore, be taken as an indication that gp

sv

is active in determining the wettability characteristics of the Ti–6Al–4V alloy.The situation after the first fine grinding-up stage presented revealing

results. As one would expect, the increase in surface roughness occasionedby the first fine grinding-up stage caused the y measured for the untreatedTi–6Al–4V alloy sample to decrease. For the CO2 laser treated Ti–6Al–4Valloy sample the first fine grinding-up stage actually caused the value of y toincrease by 0.2�, despite the increase in the surface roughness to virtually thesame value as the untreated Ti–6Al–4V alloy sample. Since the surfaceoxygen content of the untreated Ti–6Al–4V alloy sample remained aroundthe original value, while the surface oxygen content of the CO2 laser treatedTi–6Al–4V alloy sample reduced to a level similar to that of the untreatedsample, then this observation implies that the surface oxygen content isindeed active in determining y. Moreover, because the CO2 laser inducedmicrostructure was retained after the first fine grinding-up stage, then it isreasonable to assert that gp

sv is the factor responsible for the fact that y is 1.9�

degrees lower on the CO2 laser treated Ti–6Al–4V alloy sample than that ofthe untreated Ti–6Al–4V alloy sample. Thus the claim that gp

sv plays a role ininfluencing changes in the wettability characteristics of the Ti–6Al–4V alloyis substantiated.

Further confirmation of the action of gpsv with regard to the wettability

characteristics of the Ti–6Al–4V alloy can be found from an examination of

Table 6.2 The contact angle, surface roughness and surface oxygen content of theuntreated and CO2 laser treated (1.6 kW/cm2) Ti–6Al–4V alloy following the finegrinding stages

Untreated CO2 laser treated

O2 y O2 yFine polishing stages Ra (mm) (at %) (glycerol) Ra (mm) (at %) (glycerol)

Unpolished 0.35 23.32 85.8 0.39 41.80 74.4Grinding-down stage 1 0.30 23.33 87.1 0.31 23.34 86.2Grinding-down stage 2 0.30 23.31 87.1 0.30 23.33 86.9Grinding-up stage 1 0.45 23.32 76.5 0.46 23.33 74.6Grinding-up stage 2 0.46 23.32 76.4 0.46 23.33 76.4

Identification of the Predominant Mechanisms 113

the results of the second fine grinding-down stage. The values obtained for ygiven in Table 6.2 show that after the second fine grinding stage, whichremoved the CO2 laser induced microstructure, y increased slightly from86.2 to 86.9�. No change in the value of y for the untreated Ti–6Al–4V alloysample was observed after the second fine grinding-down stage, nor wasany change in the surface oxygen seen. This suggests that the increase in ywas due to the removal of the CO2 laser induced microstructure and, in turn,the reduction in gp

sv, which was presumably reduced to a value around thatof the untreated Ti–6Al–4V alloy sample. When the CO2 laser inducedmicrostructure was removed by the second grinding-up stage, the value ofy increased again to 76.4�, a value almost equal to that of the untreated Ti–6Al–4V alloy sample.

Although the first and second fine grinding stages of the grinding-downand grinding-up processes revealed that gp

sv influenced the wettabilitycharacteristics of the Ti–6Al–4V alloy, a comparative examination of thesurface roughness, surface oxygen content and y values for the untreatedand CO2 laser treated Ti–6Al–4V alloy samples shows that its influence isrelatively small. If one considers the first fine grinding-down stage, for theuntreated Ti–6Al–4V alloy sample the surface roughness was decreasedfrom 0.35 to 0.30 mm, which gave rise to an increase in y of 1.3�, with thesurface oxygen content remaining around the original value of 23.32 at %. Incontrast, the surface roughness of the CO2 laser treated Ti–6Al–4V alloysample after the first fine grinding-down stage was reduced from 0.39 to0.31mm, which in turn caused an increase in y of some 11.8�. This increase iny for the CO2 laser treated Ti–6Al–4V alloy sample is clearly dispropor-tionate to the reduction in the surface roughness. However, unlike the casewith the untreated Ti–6Al–4V alloy, the surface oxygen content of the CO2

laser treated sample was found to have reduced from 41.8 to 23.34 at %, alevel similar to that of the untreated samples. This suggests that therelatively large increase in y experienced by the CO2 laser treated Ti–6Al–4V alloy sample is correlated with the change in the surface oxygen content.As such, it is reasonable to conclude that the wettability characteristics of theTi–6Al–4V alloy are, after surface roughness, influenced predominantly bythe surface oxygen content and, to some extent, by the microstructure.

6.6 Investigation of Wettability and Work AdhesionUsing Physiological Liquids

In order to simulate the biological environment, the physiological fluids andsimulated physiological liquids used for the wetting experiments werehuman blood, human blood plasma, simulated body fluid (SBF) andSBF þ BSA (bovine serum albumin), as described in Section 4.7.


The values of y formed between the selected and simulated physiologicaltest liquids and untreated, mechanically roughened and CO2 laser treatedTi–6Al–4V alloy are shown in Table 6.3. It clearly reveals that the y values ofall body fluids on the CO2 laser treated Ti–6Al–4V alloy are lower than theuntreated and mechanically roughened specimens, indicating that the wett-ability characteristics of the material with the body fluids obviouslyimproved after CO2 laser treatment. The results detailed previously showclearly that interaction of the CO2 laser beam with the Ti–6Al–4V alloy hadresulted in the lower y formed between the physiological liquids. Sincebiomaterials first contact a proteinaceous liquid phase, almost aqueous innature, leading to surface reorganization of proteins followed by cellattachment on biomaterials, wettability characteristics, by controlling theinteraction with physiological fluids, would primarily influence cell beha-viour on biomaterials. Wetting of the solid surface is a predictive index ofcytocompatibility [165].

Further, according to Equation (3.4), the decrease in the y resulted in theincrease of the work adhesion of Ti–6Al–4V towards the physiological andsimulated physiological liquids. Using the referenced glv value of humanblood (47.5 mJ/m2), human blood plasma (50.5 mJ/m2) [141], SBF (72.5 mJ/m2) and SBFþBSA (54.0 mJ/m2) [164], the work adhesion, Wad, of the Ti–6Al–4V alloy towards these body fluids was determined through Equation(3.4), as shown in Figure 6.8. A discernable increase in Wad of body fluids canbe seen on the Ti–6Al–4V alloy following CO2 laser treatment. Moreover,Wad increased as the CO2 laser power density increased. Owing to the factthat biomaterials first contact a proteinaceous liquid phase, almost aqueousin nature, leading to surface reorganization of proteins followed by cellattachment on biomaterials, wettability characteristics, by controlling theinteraction with physiological fluids, would primarily influence cell beha-viour on biomaterials. Wetting of the solid surface is a predictive index ofcytocompatibility [165]. Moreover, the improvements of Wad towards thesefluids would imply better suitability of titanium in the biological environment.

Table 6.3 Mean values of contact angles formed between the simulated physiolo-gical test liquids and the untreated, mechanically roughened and CO2 laser treatedTi–6Al–4V alloy


Ti–6Al–4V alloy Human blood Human blood plasma SBF SBFþBSA

Untreated 58.3 63.2 82.8 60.9Mechanically roughened 56.9 61.7 80.2 59.1CO2 laser (1.3 kW/cm2) 50.6 55.4 71.0 53.5CO2 laser (1.6 kW/cm2) 48.3 53.6 68.2 51.2

Investigation of Wettability and Work Adhesion 115

6.7 Summary

The results presented in this chapter are a clear indication that CO2 lasersurface treatment of the Ti–6Al–4V alloy brought about a reduction in the yformed between the Ti–6Al–4V alloy and simulated physiological liquids,indicating that the wettability characteristics of the material were modified.Such results reveal that the CO2 laser could be a suitable tool to modify thebiograde metals for improved biocompatibility. It was found that themodification of surface roughness, surface oxygen content and surfaceenergy of the Ti–6Al–4V alloy following laser treatment were the factorsinfluencing the wettability characteristics. The predominant mechanismsactive in determining the wettability characteristics of the Ti–6Al–4V alloyfollowing CO2 laser irradiation were identified. It was found that thewettability characteristics of the Ti–6Al–4V alloy were, after surface rough-ness, influenced predominantly by the surface oxygen content and, to someextent, by the microstructure. A reduction in y contributes to enhancementin work adhesion of physiological liquids on the Ti–6Al–4V alloy followingCO2 laser treatment. The improvements in work adhesion of the Ti–6Al–4Valloy surface towards these fluids would mean better suitability of the Ti–6Al–4V alloy to be used as a biomaterial after laser treatment.

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CO2 laser1.6 kW/cm2

Wor

k A

dhes

ion,

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(mJ/

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Human bloodHuman blood plasmaSBFSBF+BSA

Figure 6.8 Work of adhesion of body fluids for the untreated, mechanically rough-ened and CO2 laser treated Ti–6Al–4V alloy


7

In vitro BiocompatibilityEvaluation of CO2 LaserTreated Titanium Alloy

This chapter is concerned with comparatively evaluating the biocompatibility of aCO2 laser treated titanium alloy (Ti–6Al–4V). An investigation of the bioactivityof the CO2 laser modified Ti–6Al–4V alloy in simulated body fluid (SBF) wasconducted and the formation of bone-like apatite was found on samples that wereCO2 laser treated at certain power densities. In addition, ellipsometry was usedto investigate the albumin and fibronectin (protein) adsorption on the untreatedand CO2 laser treated Ti–6Al–4V alloy and significant changes in proteinadsorption were observed on the samples following CO2 laser treatment, beingthe result of the CO2 laser modified surface properties. Finally, the in vitrobehaviour of the hFOB cell response was conducted to determine the effect ofsurface properties on the osteoblast cell adhesion and growth, thereby elucidatingthe mechanisms active in the osteoblast cell response and correlating them withthe enhanced wettability characteristics of the Ti–6Al–4V alloy after CO2 lasertreatment.

7.1 Introduction

Titanium and some of its alloys are now dominant biomaterials because oftheir good biocompatibility. Commercially pure titanium (cp Ti) implantsare alloplastic materials used as the foundation for replacing teeth indentistry and are also used for orthopaedics; however, this interactiondoes not involve a chemical bond with bone. The lack of ability to bond


chemically may lead to slow fixation of cp Ti dental implants and to theirgradual loosening over a long period [238].

Different approaches are being used in an effort to obtain the desiredbone–implant interface. The implant should present a surface conductive toor that will induce osseointegration [9]. It was recently demonstrated byKokubo et al. [239] that an in vitro chemical deposited bone-like apatite on cpTi could be induced by an alkali and heat treatment process followed by asimulated body fluid (SBF) soaking. This apatite layer is an essentialrequirement for artificial materials to bond to living bone [41]. Furthermore,wettability, which controls the way biological fluids interact with materials,is among the physicochemical characteristics that have been altered with theaim of improving the bone–implant interface. The cells attached to carboxy-licylic-acid-terminated hydrophilic monolayers were about two times morethan those attached to methyl-terminated hydrophobic monolayer over90 minutes [217]. Radiofrequency glow discharge has been used to increasesurface energy and to enhance cell binding [240, 241]. Alterations in surfacemorphology and roughness have been used to influence cell and tissueresponse. Classically, to improve bone tissue integration on implant sur-faces, various techniques have been used to increase the roughness of theimplant surfaces [242–244]. Many in vivo studies have compared the effi-ciency of various surface treatments in mechanically and morphologicallyimproving bone tissue integration of implants. Various results have beenobtained, depending on the roughness amplitude but also on the methodused to produce the surface roughness [242–246].

Due to the rapid and specific modification of organic and inorganicmaterials, laser surface processing has aroused growing interest and beenproven to be a controllable and flexible technique for modifying the surfaceproperties of materials. Yet little work has been carried out to investigateemploying lasers to modify the surface properties of biomaterials in order toimprove their biocompatibility. Having said that, it is recognised within thecurrently published work that laser irradiation of material surfaces can affectchanges in cell adhesion on biomaterials. Lately, several publications haveinvestigated the modification of biocompatibility of biomaterials’ surfacesfollowing laser irradiation. A CO2 pulsed laser were used to graft a polymer[17] and a rubber [18]. The results showed a marked reduction of the plateletadhesion and aggregation for the modified polymer surface and cell attach-ment with a greater degree of spreading and flattening on the unmodifiedrubber surface. Dadsetan et al. [19] found that L929 fibroblast cells attachedand proliferated extensively on CO2 and KrF laser treated PET films incomparison with unmodified PET, with surface morphology and wettabilitybeing found to affect cell adhesion and spreading. More recently, Hao andLawrence found that the laser generated surface properties on magnesiapartially stabilised zirconia (MgO–PSZ) resulted in bone apatite formation in


stimulated body fluid [23], favourable albumin [24] and fibronectin adsorp-tion [25] and better human fibroblast response [247] and human osteoblastcell adhesion [248] and functions [26]. However, no work has so farinvestigated the use of lasers to alter the biocompatibility of the biogrademetals.

With the aim of improving the biocompatibility (bioactivity and biointe-gration) of a titanium alloy (Ti–6Al–4V), a CO2 laser was used to generatethe favourable surface properties of the material for better biologicalresponse. The bioactivities of untreated and CO2 laser treated Ti–6Al–4Valloy were evaluated by observing the bone-like apatite formation on theirsurface after soaking in SBF, because for an artificial material to bond toliving bone, it is essential that the material has the ability to form abiologically active, bone-like, apatite layer on its surface in the humanbody. The biointegrations of the untreated and CO2 laser treated Ti–6Al–4V alloy were assessed by protein adsorption and osteoblast cell response.Protein adsorption is the almost immediate event that occurs upon implan-tation of metals and mediates subsequent cell response and tissue–implantinteractions [9]. In addition, it is widely acknowledged that a major deter-minant of the bone–biomaterial interfacial response is the initial attachment,spreading and growth of osteoblasts on the implant surface and thatimprovements in these processes may lead to faster and more extensiveimplant integration and higher long-term stability [171]. Indeed, the inves-tigation of apatite formation in SBF was applied by Gil et al. [238] and Wanget al. [249] to evaluate the bioactivity of titanium alloys. In vitro tests ofprotein adsorption and osteoblast cell interactions are now well establishedto assess the biocompatibility of biomaterials and are widely used toevaluate the osseointegration of titanium alloys [171, 221]. In addition, thebetter performance of apatite formation on titanium alloy in SBF in vitro wasconfirmed by the beneficial effect on the osteoconduction in vivo experimentconducted by Lu et al. [250]. Of great significance is the fact that in vitro cellculture models have the potential to help elucidate events at the bone–implant interface (reviewed by Davies [251]), by providing morphological,biochemical and molecular information regarding osteoblastic developmentand synthesis of the matrix at the interface with various biomaterials.

7.2 Sample Preparation

Full details of the properties of the Ti–6Al–4V alloy used for in vitroevaluation of the potential of CO2 laser surface treatment to enhancebioactivity are to be found in Section 6.2.1. The Ti–6Al–4V alloy wassupplied as a round bar with a diameter of 28.5 mm. To prepare theTi–6Al–4V alloy for the in vitro tests, it was sectioned to produce 30 samples

Sample Preparation 119

(discs) each of 3 mm thickness with a cutting machine (Miniton; Struers,GmbH) using a diamond-rimmed cutting blade. As in Chapter 6, theTi–6Al–4V alloy samples were used as received prior to CO2 laser treatment.The 30 discs were then divided into two groups of 15 samples, with thegroups being: untreated and CO2 laser treated. The procedure adopted forthe CO2 laser processing of the Ti–6Al–4V alloy is described in Section 6.2.2.In this instance, however, only CO2 laser power densities of 1.3 and 1.6 kW/cm2 were used, since the findings of Chapter 6 showed these two CO2 laserpower densities were the most effective. Untreated samples were used as thecontrol for the in vitro apatite formation and protein adsorption tests.Mechanically roughened samples were used as a control for the osteoblastcell culture as the aim in this test was to compare the effect of CO2 lasertreatment with the traditional method employed to improve cell adhesion(surface roughening by mechanical means). Details of the mechanical rough-ening procedure are given in Section 6.2.1.

7.3 Bone-Like Apatite Formation on Titanium Alloys

In a different model, the mechanism of apatite formation on an amorphoussodium titanate formed on titanium metal was also examined by XPS. In theSBF, the sodium titanate releases Naþ ions via exchange with the H3Oþ ionsin the fluid to form Ti–OH groups on its surface. The Ti–OH groups formedimmediately combine with Ca2þ ions in the fluid to form amorphouscalcium titanate. This calcium titanate later combines with phosphate ionsin the fluid to form amorphous calcium phosphate with a low Ca:P ratio. Thecalcium phosphate transforms into apatite, which exhibits a Ca:P ratio of1.65, and contains a small concentration of Mg and Na, similar to bonemineral [41]. To reveal the reasons why this complex process is required forapatite formation, the zeta potential of the surface of sodium titanate wasmeasured by laser electrophoresis at various SBF soaking times. It was foundthat the surface of the sodium titanate was highly negatively chargedimmediately after it was soaked in the SBF. The surface potential increasedwith increasing soaking time up to a maximum positive value. Thereafter, itdecreased with increasing soaking time, reached a negative value again andfinally converged to a constant negative value. On the basis of this finding,the complex process of apatite formation described above is well interpretedin terms of the electrostatic interaction of the functional groups with the ionsin the fluid. The Ti–OH groups formed on the surface of sodium titanateafter soaking in SBF are negatively charged and, hence, combine selectivelywith the positively charged Ca2þ ions in the fluid to form calcium titanate.As the calcium ions accumulate on the surface, the surface gradually gainsan overall positive charge. As a result, the positively charged surface


combines with negatively charged phosphate ions to form amorphouscalcium phosphate. This calcium phosphate spontaneously transforms intothe apatite, because the apatite is the stable phase in the body environ-ment [41].


Soaking in Simulated Body Fluid

The untreated and CO2 laser treated Ti–6Al–4V alloy samples were soakedin an acellular SBF [41] having an ion concentration nearly equal to that ofhuman blood plasma. This solution was prepared by dissolution of high-purity reagents in distilled water and was buffered at 7.25 with 50 mMtrishydroxymethyl amino ethane and 45 mM hydrochloric acid. Theuntreated and CO2 laser treated Ti–6Al–4V alloy samples with variouspower densities were immersed in 30 ml of SBF in a polyethylene bottle at37 �C, without stirring. After 14 days they were removed from the solution,gently washed in distilled water and dried at room temperature. The soakedsamples were then characterised by SEM and EDX. The samples for SEMobservations were simply dried and covered by a thin gold layer toguarantee the conductivity.

7.3.2 The Effects of CO2 Laser Treatment on the Ti–6Al–4Vin Simulated Body Fluid

Figure 7.1 shows the morphologies of the untreated and CO2 laser treatedTi–6Al–4V alloys after 7 days of immersion in SBF. On the untreated sample,very small precipitants could be observed and no apatite was found (seeFigure 7.1(a)). While on the CO2 laser treated Ti–6Al–4V alloys, some apatitenuclei were observed in Figures 7.1(b) and (c), indicating that laser process-ing brought about the bioactivity to the nonbioactive Ti–6Al–4V alloy.The EDX analysis revealed an average Ca:P ratio of 1.57 in these apatitenuclei.

Hydroxide ions (OH�) attach to metal cations. It has been reported that thehydrated surface of titanium contains more hydroxyl groups, thus improv-ing the bioactivity of titanium [249]. The titanium alloy oxide surface incontact with water is believed to be highly hydroxylated [211]. It is demon-strated that the laser surface processing brought about the oxidation of theTi–6Al–4V alloy as the surface oxygen content of the laser modified Ti–6Al–4V alloy is much higher than the untreated sample. Feng et al. [193] foundthat during heat treatment, oxygen and water molecules in atmospherereacted with the surfaces, resulting in an increase in the amount of OHgroups as well as thickening of the oxide films. The OH groups originatedfrom chemisorption. Since chemisorption belongs to the molecular layer,

Bone-Like Apatite Formation on Titanium Alloys 121

when active sites on titanium surfaces that could induce chemisorption ofwater were saturated, the amount of hydroxyl groups would tend to beconstant.

Oxide-covered titanium is simple in terms of composition, and thepossible species that may be released are Hþ or OH� ions and the ionsrelated to titanium. Healy and Ducheyne [252] have suggested that when inSBF titanium was subjected to passive dissolution and, within a soakingperiod of up to 400 hours, the passive dissolution was governed by thehydrolysis of the titanium oxide and the equilibration of the surface with theSBF, which resulted in the formation of OH� and Ti(OH)

ð4�nÞþn . The OH�

ions were adsorbed on the oxide surface, while the titanium hydroxide ionsentered into the SBF. In the present specimen set-up, these ions could beaccumulated inside the confined space between the two contact surfaces,where the SBF was much more stagnant. Obviously, the accumulation ofOH� ions on the surface would lead to a more negatively charged surfacethat was believed necessary for apatite nucleation [253].

Figure 7.1 SEM images of the Ti–6Al–4V alloy soaked in SBF for 7 days: (a) untreated,(b) CO2 laser treated at a power density of 1.3 kW/cm2 and (c) CO2 laser treated at apower density of 1.6 kW/cm2


7.4 Protein Adsorption

The molecules involved in cell adhesion and spreading include extracellularmatrix molecules, transmembrane receptors and intracellular cytoskeletalcomponents. Among the extracellular matrix proteins shown to mediate cellattachment to substrates, fibronectin is protein found in many extracellularmatrices and in blood plasma and serves as an attachment molecule betweenthe substrate and cell membrane of anchorage-dependent cells. It is knownthat the ligand fibronectin connects to the cell membrane via integrinreceptors. The activation of integrins triggers cytoplasmic reactions, andthereby stimulates the intracellular signalling pathway and subsequently thecellular functions such as proliferation and differentiation [169]. On the otherhand, human albumin is a nonadhesive protein for osteoblasts [194].Albumin is the major protein component of serum and dominates theadsorption of phenomena on medical implants in the first stage of contactwith body fluids. Human serum albumin or bovine serum albumin (BSA)coatings are often used as a passivating agent to prevent the adhesion of cellsand thrombus formation [195].


Protein Adsorption

The proteins used for this study were human serum albumin and humanplasma fibronectin (Calbiochem, Inc.). Prior to the adsorption of 1 mg/ml ofalbumin in phosphate buffered salines (PBS), Ti–6Al–4V samples wererinsed with deionised water. The individual samples were transferred intoa 24-well tissue culture plate. Thereafter, 2.5 ml of prepared albumin solu-tion was added into each well. Adsorption proceeded for 1 hour in anincubator at 37 �C. After adsorption was complete, the samples were driedwith N2 and immediately transferred to an ellipsometer for measurement ofthe adsorbed protein layer. The above procedure was repeated with a0.2 mg/ml concentration of fibronectin in PBS.

Ellipsometric Measurement

The thickness of the adsorbed human serum albumin and human plasmafibronectin were measured using an automatic ellipsometer equipped with a632.8 nm helium–neon laser (L2W16SF.544; Gaertner, Inc.). The thicknessand refractive indices of protein films were determined using an ellip-someter computer program with an accuracy of �3 A

�. Four ellipsometer

measurements at different locations on each sample were taken and theaverage value was calculated.


Statistics

Statistical analysis was performed with an SPSS v.12 software package(SPSS/PC, Inc.) in the same manner as described in Section 5.4.1.

7.4.2 Albumin and Fibronectin Adsorption on CO2 Laser TreatedTitanium Alloy

The thickness of the adsorbed human serum albumin and human plasmafibronectin layer, which indicate the amount of the adsorbed protein onthe untreated and CO2 laser treated Ti–6Al–4V alloy, are shown in theFigure 7.2. It was found that the thickness of the albumin layer on the un-treated Ti–6Al–4V is higher than on the CO2 laser modified sample, whereasthe thickness of the fibronectin layer is less on the untreated one than on themodified sample, as shown in Figure 7.2. The CO2 laser power densityapplied in the experiment is negatively correlated with amounts of albumin,while positively correlated with the fibronectin (see Figure 7.2). On the onehand, the CO2 surface treatment promoted the adsorption of the fibronectinand the amount of the adsorbed fibronectin on the Ti–6Al–4V was positive

300

400

500

600

700

800


CO2 laser1.6 kW/cm2


CO2 laser1.6 kW/cm2

Fibronectin Albumin

Thi

ckne

ss o

f A

dsor

bed

Pro

tein

Lay

er (

Å) *

**

*

Figure 7.2 Thickness of the adsorbed fibronectin and albumin protein layer on theuntreated and CO2 laser treated Ti–6Al–4V alloys (treated with different powerdensities). For fibronectin adsorption, there was a significant statistical difference inthickness between the untreated Ti–6Al–4V alloy and the sample CO2 laser treated at1.6 kW/cm2, and no statistical difference between the untreated Ti–6Al–4V alloy andthe CO2 laser treated sample at 1.3 kW/cm2. For albumin adsorption, there was asignificant statistical difference in thickness between the untreated Ti–6Al–4V alloy andsamples CO2 laser treated at 1.3 and 1.6 kW/cm2 and no statistical difference amongthe samples CO2 laser treated at 1.3 and 1.6 kW/cm2 (�p < 0:05)


with the CO2 laser power density; on the other hand, it inhibited albuminadsorption and the amount of the adsorbed albumin was negative with CO2

laser power density. In addition, the statistical analysis reveals that thethickness of absorbed fibronectin on the untreated sample was significantlyless than on the sample CO2 laser treated at the power density of 1.6 kW/cm2, but not significantly less than on the sample CO2 laser treated at thepower density of 1.3 kW/cm2 (p < 0:01); on the other hand, the thickness ofthe absorbed albumin layer on the untreated Ti–6Al–4V alloy was signifi-cantly higher than that on the CO2 laser modified samples, as shown inFigure 7.2. As stated in Section 5.4.2, the previous study shows that proteinadsorption is influenced by the surface topography [198] and the surfacechemistry (wettability characteristics) [199].

The Effects of Surface Roughness

As discussed previously in Section 6.3.2, the change in surface roughness ofthe Ti–6Al–4V following CO2 laser irradiation is very slight, increasing byonly 3mm. Further, there is no obvious relationship between the surfaceroughness and protein adsorption due to the fact that the CO2 laser treatedsamples have more or less surface roughness than the untreated samples,but both treated samples have a similar trend for an increase in fibronectinadsorption or a decrease in albumin adsorption. It is therefore reasonable topostulate that the surface roughness is not a factor of importance when itcomes to influencing the adsorption of these proteins.

The Effects of Wettability Characteristics

As one can see from Figure 7.3, as the wettability characteristics of theTi–6Al–4V increased, the adsorbed amounts of albumin decreased. Theresults of the albumin adsorption are consistent with the previous findingthat the increase in surface hydrophilicity of Ti results in lower albuminadsorption [201], showing that the wettability characteristics of the Ti–6Al–4V could be an active factor in albumin adsorption, much better than theconformation of fibronectin adsorbed on hydrophobic surfaces [204].

The results of the adsorption of fibronectin given in Figure 7.3 show that itincreased with the increasingly wettable characteristics (hydrophilic) of theTi–6Al–4V alloy surface, implying that wettability characteristics of the Ti–6Al–4V alloy provide the predominant mechanism governing fibronectinadsorption. In addition, the statistical analysis revealed that the fibronectinadsorption on the sample with a cos y value of 0.07 was significantly lessthan on the sample with a cos y value of 0.34 and was not significantly lesson the sample with cos y of 0.27, the albumin adsorption on the sample witha cos y value of 0.07 was significantly higher than on the samples with cos yvalues of 0.27 and 0.34 and no significant difference in albumin adsorption


was observed between the sample with a cos y value of 0.27 and the samplewith a cos y value of 0.34. Indeed, a previous study by Hao and Lawrence[25] on albumin adsorption also revealed that albumin absorption decreasedon the MgO–PSZ following CO2 laser treatment due to the increasedwettability characteristics. Moreover, a previous investigation [204] on theextent of fibronectin adsorption as compared to its biological activity onhydrophobic and hydrophilic surfaces suggested the possibility that fibro-nectin was adsorbed in two different conformations when incubated withthe surfaces at low concentrations, with the more active conformation on thehydrophilic surfaces. The results showed that the antiplasma fibronectinantibody appeared to bind to the conformation of fibronectin adsorbed onhydrophilic surfaces.

As discussed previously in Section 6.5, the considerable change in gpsv as

opposed to the minor difference in gdsv was a driver for the improvement in

the wettability characteristics of the Ti–6Al–4V alloy after CO2 laser irradia-tion. The observed significant correlation between wettability characteristics

0.0 0.1 0.2 0.3 0.4300

400

500

600

700FibronectinAlbumin

Wettability, Cos (Glycerol)θ

Thi

ckne

ss o

f A

dsor

bed

Pro

tein

Lay

er (

Å)

Figure 7.3 The relationship between the thickness of the adsorbed albumin andfibronectin layer and wettability characteristics (cos y) of the Ti–6Al–4V alloy. Forfibronectin adsorption, there was a significant statistical difference in thickness be-tween the sample with cos y of 0.07 and the sample with cos y of 0.34, and nostatistical difference between the sample with cos y of 0.27 and the sample with cos yof 0.34. For albumin adsorption, there was a significant statistical difference in thick-ness between the sample with cos y of 0.07 and the samples with cos y of 0.27 and0.34, and no statistical difference between the sample with cos y of 0.27 and thesample with cos y of 0.34 (�p < 0:05)


and protein adsorption implies that albumin and fibronectin adsorption onthe Ti–6Al–4V alloy surfaces was probably due to the polar and chemicalinteractions [205]. This finding is similar to a previous investigation thatalbumin and fibronectin adsorption is mainly influenced by the change in gp

sv

of the MgO–PSZ following CO2 laser irradiation [25]. It is therefore possibleto maintain that the change in gp

sv of the Ti–6Al–4V alloy generated by theCO2 laser irradiation contributes to the inhabitation of albumin adsorptionand enhancement of fibronectin adsorption, and in turn its potential for afavourable bone cell response on the Ti–6Al–4V alloy surface.

7.5 Osteoblast Cell Adhesion

The development of bone–implant interfaces depends on the direct interac-tions of bone matrix and osteoblasts with the biomaterial. There is asubstantial body of literature based on the premise that improved initialattachment of osteoblasts or osteoblast precursor cells to orthopaedicimplant surfaces may lead to improved bone integration of the implantand longer-term stability [206]. Osteoblast adhesion is a prerequisite forbone–biomaterial interaction and depends on the surface aspect of materials.A cell in contact with a material surface will firstly attach, adhere and thenspread. The quality of this adhesion will influence their morphology andtheir future capacity for proliferation and differentiation. The attachment ofanchorage-dependent cells such as osteoblasts to biomaterial surfaces is acomplex process involving cell attachment and spreading [207], focal adhe-sion formation, and extracellular matrix formation and reorganisation [208].

7.5.1 Experimental Procedure

Cell Culture

The human osteoblastic cell line hFOB 1.19 was obtained from the AmericanType Culture Collection (ATCC) (Manassas, Inc.). These cells were culturedin a medium containing a 1:1 mixture of Dulbecco’s Modified Eagle’smedium without phenol red and Ham’s F12 medium with 2.5 mML-glutamine (D-MEM/F-12 Medium), supplemented with 10 % fetal bovineserum (ATCC) and 0.3 mg/ml G418 (Calbiochem, Inc.) in a 37 �C, humidi-fied, 5 % CO2/95 % air incubator.

Cell Cytotoxicity

Cytotoxicity tests consisted of the quantification of the activity of lactatedehydrogenase (LDH) in a culture medium of cells in contact with thesamples. The activity of the LDH enzyme rises when cells are damaged: the

Osteoblast Cell Adhesion 127

LDH activity induced by the untreated and selected CO2 laser treatedspecimens (0.9 and 1.6 kW/cm2) in triplicate were compared to that inducedby a toxic agent (Triton X100 0.05 % in PBS) and to that induced by a culturepolystyrene plate (NUNC, Inc.). The cell culture plate was used as a negativecontrol and a Triton toxic agent as a positive control.

Cell Adhesion and Morphology

For cell adhesion analysis, osteoblasts were enzymatically lifted from poly-styrene tissue culture flasks until cell confluence using 1 ml Trypsin-EDTA(0.25 % Trypsin/0.53 mM EDTA solution) before suspension in the culturemedium. The Ti–6Al–4V alloy samples were placed in a 24-well tissueculture polystyrene plate (Falcon, BP) under a sterile environment andsterilized in 70 % ethanol for 24 hours. The samples were rinsed by PBSand then were seeded with cell suspension. To analyse the cell attachmentand morphology, the untreated and laser treated specimens were seededwith a 0.5 ml cell suspension of 1 � 105 cell/ml for 24 hours of cell culture,then dehydrated in a graded ethanol series, critical point dried with CO2 andgold coated for SEM analysis.

Cell Proliferation

Cell proliferation on each specimen was measured by MTT assay. Theosteoblast cells cultured for 7 days on each specimen were gently washedwith PBS and were measured by MTT assay using 3-(4,5-dimethyl-thiazole-2-yl)-2, 5-diphenyl tetrazolium bromide (MTT; Sigma, Inc.). The MTT solu-tion was added to each specimen and the cells were incubated for 4 hours at37 �C, before replacing the medium with dimethylsulfoxide. Absorbance ofthe solution was measured by an instrument plate reader (EL312; Bio-Tek,Inc.) at 490 nm.

Statistics

Statistical analysis was performed with an SPSS v.12 software package(SPSS/PC, Inc.) in the same manner as described in Section 5.4.1.

7.5.2 Osteoblast Cell Response on CO2 Laser Treated Titanium Alloy

Cell Cytotoxicity

The cell culture plate was used as a negative control and a Triton toxic agentas a positive control for the assessment of cell membrane damage expressedas LDH activity. LDH activity in the culture media obtained from cellscultured on all the tested materials was found to be not significantly


different from the negative control, as shown in Figure 7.4, indicating thatuntreated and CO2 laser treated Ti–6Al–4V were not cytotoxic.

Cell Attachment

Generally, cells in contact with a material’s surface will firstly attach, thenadhere and finally spread. From Figure 7.5, it is quite clear that the adhesionand spreading of osteoblast cells was influenced by the CO2 laser treatment.The cells on the untreated and mechanical roughened surface present theinitial stage of the adhesion, with individual cells covering a small surfacearea and not spreading; on the other hand, the cells on the CO2 lasertreatment show a good state of adhesion and flattening to cover moresurface area.

This means that the cells showed a better adhesion on the CO2 lasertreated sample than the untreated and mechanical roughed samples, sug-gesting that the surface properties generated by the CO2 laser treatmentwere more favourable for the osteoblast cell adhesion. In addition, thenumber of cells on the untreated Ti–6Al–4V alloy is nearly the same as onthe mechanical roughened material, but is much lower than on the CO2 lasertreated Ti–6Al–4V alloy (for both power densities). Hence, the CO2 laser

0

10

20

30

40

50

60

LD

H A

ctiv

ity

(U/L

)

+ve −ve Untreated Mech.Rough

CO2 laser1.3 kW/cm2

CO2 laser1.6 kW/cm2

**

** *

Figure 7.4 LDH activity on the positive and negative controls, untreated, mechani-cally roughened and CO2 laser treated Ti–6Al–4V alloys. There was a significantstatistical difference between the positive control and the untreated, mechanicallyroughened and CO2 laser treated samples, and no statistical difference between thenegative control and the untreated and CO2 laser treated Ti–6Al–4V alloy samples(�p < 0:05)


treatment has a significant effect on the cell adhesion and growth, and playsa more important role than when the surface is roughened mechanically.

Cell Proliferation

The MTT results in Figure 7.6 reveal the cell proliferation on the surface ofuntreated, mechanically roughened and CO2 laser treated Ti–6Al–4V alloyspecimens for 7 days. It can be seen that the CO2 laser treated samples had asignificantly higher cell proliferation compared to the untreated sample andmechanically roughened sample. A slight improvement in cell proliferationwas observed on the mechanically roughened sample compared to theuntreated sample. It is observed that the cell proliferation generally increasesas power density of the CO2 laser treatment increases; however, the

Figure 7.5 SEM images of osteoblast cells on the Ti–6Al–4V alloy when (a) untreated,(b) mechanically roughened and CO2 laser treated (c) at 1.3 kW/cm2 and (d) at1.6 kW/cm2


statistical analysis shows that there are no significant differences in cellproliferation between the samples CO2 laser treated at the power densities of1.3 and 1.6 kW/cm2. In addition, the statistical analysis reveals that cellproliferation significantly improved for the CO2 laser treated samplescompared with that for the untreated sample, but was not significantlyincreased for the mechanically roughened sample. The results apparentlyrevealed that the surface generated by CO2 laser treatment is more favour-able for cell proliferation than the untreated and mechanical roughenedsurfaces.

7.5.3 The Effect of CO2 Laser Treatment on the OsteoblastCell Response

As discussed in Chapter 6, the CO2 laser surface treatment of the Ti–6Al–4Valloy was shown to result in alteration of the surface topography, oxidechemistry, wettability characteristics as well as albumin and fibronectinprotein adsorption. The changes in these surface properties certainly influ-enced the osteoblast cell response. It is demonstrated that CO2 laser surfaceprocessing brought about not only better cell adhesion but also higher cellproliferation to the treated Ti–6Al–4V alloy compared with the untreatedand mechanically roughened samples, indicating that this surface processing

0.00

0.03

0.06

0.09

0.12

0.15

0.18

MT

T O

ptic

al D

ensi

ty


CO2 laser1.3 kW/cm2

CO2 laser1.6 kW/cm2

**

Figure 7.6 MTT optical density of osteoblast cells grown on untreated and CO2 lasertreated Ti–6Al–4V alloys after 7 days of cell culture. There was a significant statisticaldifference between the untreated sample and the samples CO2 laser treated atpower densities of 1.3 and 1.6 kW/cm2, and no statistical difference among theuntreated and the mechanically roughened samples (�p < 0:05)


technique is effective and better than the traditional mechanical roughingmethod. In this study, the osteoblast cell response increases with the powerdensity of CO2 laser treatment. Such findings may further understanding ofthe particular surface properties that are most closely related to osteoblastcell attachment and function. For this reason it is suggested that either thesurface modifications investigated here or knowledge gained from theirstudy may be used to help improve the biocompatibility and osteoinductiveproperties of titanium based implants.

The Effect of Topography on the Osteoblast Cell Response

In this study there was no distinguishable difference in the cell adhesionbetween the untreated and mechanically roughened samples. In addition,statistical analysis shows that the cell proliferation on the mechanicallyroughened sample did not significantly increase compared to the untreatedsamples. On the other hand, the sample CO2 laser treated at the powerdensity of 1.6 kW/cm2, which had the lowest surface roughness as demon-strated in Chapter 6, exhibits better cell adhesion and a significant increasein cell proliferation compared to the mechanically roughened sample, whichhad the highest surface roughness. This in turn suggests that the stimulatoryeffect of surface properties followed by CO2 laser irradiation on osteoblastcell attachment is most likely to be attributable to changes in surfaceproperties other than roughness alone. Furthermore, since the CO2 laserinduced change in surface roughness is of such small magnitude comparedto the accompanying changes in oxygen chemical and wettability character-istics, the change in surface roughness is also unlikely to explain theobserved effects of CO2 laser treatment on osteoblast cell adhesion andproliferation. Indeed, the work by Macdonald et al. [254] examining athermally modified Ti–6Al–4V alloy revealed that the surface topographywas less important for determining protein and cell attachment than surfacechemistry.

The Effect of Wettability Characteristics on the OsteoblastCell Response

From the foregoing results it is clear that the CO2 laser treatment of theTi–6Al–4V alloy enhanced the wettability characteristics of the material andthat following CO2 laser treatment the Ti–6Al–4V alloy presented improvedbioactivity in terms of osteoblast cell adhesion and proliferation. It is notpossible, however, to conclude from these findings that the increasedosteoblast cell bioactivity of the Ti–6Al–4V alloy is correlated to theimprovement of the wettability characteristics of the material. The reasonthat it is not possible to claim the existence of a correlation betweenwettability and osteoblast cell bioactivity for the CO2 laser treated Ti–6Al–4V


alloy is that the CO2 laser treatment resulted in an increase in the surfaceroughness of the alloy. Because surface roughness is known to influenceosteoblast cell adhesion and proliferation, it is possible that the surfaceroughness increase alone could be responsible for the increased osteoblastcell bioactivity. Therefore, for any correlation between wettability character-istics and osteoblast cell bioactivity for the CO2 laser treated Ti–6Al–4V alloyto be confirmed or refuted, attention needs to be paid to the effects ofincreasing only the surface roughness of the alloy. This was done byincluding the mechanically roughened Ti–6Al–4V alloy sample in theanalysis.

A comparison of the MTT values for the untreated and mechanicallyroughened Ti–6Al–4V alloy samples given in Table 7.1 confirms that surfaceroughness does indeed affect osteoblast cell adhesion. As one can see fromTable 7.1, the MTT value on the Ti–6Al–4V alloy increased markedly from0.073 to 0.080 when the surface roughness was increased by mechanicalroughening from 0.35 to 0.46 mm. In addition, y was observed to decreasefrom 85.8 to 78.7�, which is perhaps not surprising as surface roughness isknown to be a factor influential in determining y, as demonstrated byEquation (4.1). Significantly, Table 7.1 shows that the surface oxygen contentand gp

sv of the sample subjected to mechanical roughening remained vir-tually the same as the untreated sample, thus indicating that the decrease iny and the increase in surface roughness alone caused the MTT value toincrease. The MTT value for the sample CO2 laser treated at 1.3 kW/cm2 wasconsiderably larger than that of the mechanically roughened sample, 0.130compared to 0.080, despite the similar values of surface roughness for thetwo samples; however, the surface oxygen content and gp

sv of the sample CO2

laser treated at 1.3 kW/cm2 were much higher. The value of y was also foundto be smaller by 4.3�. Since the surface oxygen content and gp

sv, alongwith surface roughness, have been shown to play a role in governing thewettability characteristics of the Ti–6Al–4V alloy (see Chapter 6), it is

Table 7.1 Wettability characteristics and MTT optical density of osteoblast cells grownon the untreated, mechanically roughened and CO2 laser treated Ti–6Al–4V alloyafter 7 days of cell culture

Wettability characteristics————— —————————————————————————————y (glycerol) Roughness Oxygen

Ti–6Al–4V alloy (deg) (mm) gpsv (mJ/m2) (at %) MTT

Untreated 85.8 0.35 4.85 23.32 0.073Mechanically roughened 78.7 0.46 5.10 23.35 0.080CO2 laser treated (1.3 kW/cm2) 74.4 0.39 9.43 41.80 0.115CO2 laser treated (1.6 kW/cm2) 70.1 0.41 10.11 48.78 0.130


to be expected that the CO2 laser treated sample displays the lowest valueof y.

Owing to the extensive body of literature attesting to the efficacy of surfaceroughness in enhancing biocompatibility, it is clear that the increase insurface roughness of the Ti–6Al–4V alloy following CO2 laser treatment wasresponsible, in a large part, for the observed improvement in osteoblast cellbioactivity. The wetting of a solid surface can, nevertheless, be a predictiveindex of the biocompatibility of materials involved. As such, improvementsin the wettability characteristics of the Ti–6Al–4V alloy would most possiblyresult in better biocompatibility. Indeed, a reduction in y contributes to theenhancement in Wad of SBF and SBFþBSA on the Ti–6Al–4V alloy followingCO2 laser treatment. As both SBF and SBFþBSA have close chemicalcompositions to human body fluids, the improvements to Wad of theTi–6Al–4V alloy surface towards these fluids would mean better suitabilityof the Ti–6Al–4V alloy as a biomaterial after CO2 laser treatment. Also,Hallab et al. [215] demonstrated that surface free energy was a moreimportant surface characteristic than surface roughness for cellular adhesionstrength and proliferation. Schakenraad et al. [115] found that, despite thegreat number of parameters interfering with cellular adhesion and spread-ing, the solid surface energy is apparently a dominated factor in cellularattachment to a polymer surface and remains so, even if the solid surface hasbeen covered by a protein layer. Moreover, as discussed previously,improvement of the wettability characteristics resulted in a decrease ofalbumin adsorption and an increase of fibronectin. The phenomena wouldbe beneficial to cell adhesion since albumin is a nonadhesive protein andfibronectin is an adhesion protein. A previous study by Hao and Lawrence[247, 248] also shows that enhancement of the wettability characteristics ofMgO–PSZ after CO2 laser treatment resulted in a better response of thehuman fibroblast cell and human osteoblast cell. The surface oxygen contentof the Ti–6Al–4V alloy following CO2 laser irradiation is much higher thanthe untreated sample. The improved performance of cell adhesion andproliferation is very likely to be due to the augmentation of the surfaceoxygen content. It has been found that the biocompatibility of titaniumimplants is associated with the oxide on its surface [58]. The surface oxidelayer on titanium is biocompatible and capable of interacting with surround-ing biological fluids and cells when implanted in situ [58, 255]. This layer,composed primarily of TiO2, is found superficially on both cp Ti and Ti–6Al–4V alloy [58, 256].

Based on the results of this analysis it appears that the aspects ofwettability characteristics, surface oxygen content and gp

sv, play an importantrole in promoting cell proliferation, particularly when surface roughness issimultaneously increased. Moreover, it would be reasonable to postulatethat a correlation exists between the CO2 laser induced wettability


characteristics of the Ti–6Al–4V and osteoblast cell bioactivity. Furthersupport for this postulation comes when one considers the results given inTable 7.1 for the Ti–6Al–4V alloy sample that was CO2 laser treated at1.6 kW/cm2. In this instance the CO2 laser treatment gave rise to the highestMTT value, 0.130. At the same time, in comparison with the sample CO2

laser treated at 1.3 kW/cm2, the surface oxygen content and gpsv increased

appreciably, while the surface roughness increased slightly. These increasesin the factors known to affect wettability led in turn to a decrease in y. Sincethe surface roughness of the sample CO2 laser treated at 1.6 kW/cm2 was0.04 mm less than that of the mechanically roughened sample, it is clear thatthe better wettability characteristics of the CO2 laser treated sample wereresponsible for the improved MTT value. It is also evident that CO2 lasertreatment could be a more effective way to improve osteoblast cell adhesionthan the traditional methods currently available, especially mechanicalroughening.

7.6 Predictions for Implantation in an in vivo Clinical Situation

The orthopaedic and dental implant industry is based entirely on favourableinteraction at the bone–titanium interface (osseointegration). What is more, itis clear that osseointegration is a property of titanium implant surfaces. Thesurface modification of titanium implants is an active area of research fortwo reasons. The first reason is to increase the rate of successful implantationfrom satisfactory, as it is today, to excellent. The second is to induceacceleration of normal bone healing phenomena as this would allow earlyimmediate loading of the implant, which would have significant implica-tions in terms of decreased patient morbidity, patient physicology andhealth care costs. According to Puleo and Nanci [9], there are basicallythree different approaches to the surface modification of bone-contactingtitanium implants: physicochemical methods, morphological methods andbiochemical methods. Of the three, morphological methods are the mostwidely applied. Whichever approach is adopted, cellular behaviour, namelyadhesion, morphologic change, functional alteration, proliferation and dif-ferentiation, is greatly affected by surface properties, such as roughness,chemical composition, wettability and morphology of the oxide on thetitanium.

A rapid osseointegration is associated with improved secondary stabilityand, in turn, with a favourable prognosis for long-term success of theimplant. To this end, initial stability has to be achieved through thereduction of micromotion so as to allow early fibrin adhesion, blood vesselgrowth and eventually new bone formatin [257–259]. If micromotion cannotbe reduced then, as Branemark [1] and Brunski [260] found, a fibrous tissue

Predictions for Implantation in an in vivo Clinical Situation 135

instead of a bony interface will result at the implant surface. In order toreduce micromotion initially, and improve osseointegration thereafter, manyvariations of surface geometry have been developed [261–263], for it is wellknown that surface geometry governs the interactions of proteins and cellswith the implant surface [66, 264]. Moreover, many workers [259, 265–268]have demonstrated that increased surface roughness is associated withbetter cell adherence, higher bone–implant contact (BIC) and improvedbiomechanical interaction.

Using implants fabricated from the Ti–6Al–4V alloy, Gotz et al. [269]conducted a detailed study of the effects of various laser induced surfacefinishes on osseointegration. The Ti–6Al–4V alloy implants were laser-textured with an Nd:YAG laser treated to generate pores of various sizes,while corundum-blasted (CB) Ti–6Al–4V alloy implants were used ascontrols. An in vivo evaluation of the control and laser-textured implantswas conducted using a rabbit transcortical implantation model. Initially, thesurface between the pores was polished after laser-texturing, which wasfound to result in a BIC lower than that of the CB control implants. However,when the laser-textured surfaces were blasted with 500 to 710mm of Al2O3

grit to generate a rougher surface between the pores, the osseointegrationwas markedly improved, with the BIC being higher than that of the CBcontrol implants. This observation suggests that the biological mechanismsinvolved in the bone ingrowth on the laser-textured and surface-blasted Ti–6Al–4V alloy implants preferentially used the pores to initially anchor andimplant within the newly formed bone. The roughened surface due to theAl2O3 grit blasting came into play in the latter stages of osseointegrationwhen bone growth spread on to the implant surface.

The effect of surface topography, in particular surface roughness, onosseointegration with cp Ti was investigated by Li et al. [270]. The studytook the form of a biomechanical comparison of cp Ti cylindrical solid-screwdental implants with surfaces roughened by two procedures. The twoprocedures formed two testing groups, with the first group being subjectedto acid etching (sulfuric acid–hydrochloric acid) and termed MA, while thesecond group was sandblasted (large grit of 250–500mm) and then acidetched (sulfuric acid–hydrochloric acid) and was termed SLA. Samples ineach of the two groups exhibited similar topographical features. In each casethe acid etching generated micropits of around 2mm in diameter. For theSLA samples, however, the micropits were accommodated within themacrorough texture produced by the sandblasting, whereas for the MAsamples the micropits were superimposed on a flat machined surface. Thesediffering surface topographies naturally resulted in different values ofsurface roughness: an Ra of 1.57mm for the MA samples and an Ra of2.18mm for the SLA samples. The efficacy of each surface in terms of inter-facial shear strength was assessed using an established animal model for


implant removal torque testing by means of a split-mouth experimentaldesign. The findings revealed that the SLA surface displayed enhancedinterfacial shear strength in comparison with the MA surface, with theremoval torque value being 30 % higher. Also, the SLA surfaces achieved abetter bone anchorage than the MA surfaces owing to the fact that thesandblasting carried out before the acid etching generated a rougher surface.

The biocompatibility of titanium and its alloys as an implant material isattributed to surface oxides spontaneously formed in air and/or physio-logical liquids. The natural oxide that exists on the surface of titanium andtitanium alloys is thin, approximately 3–8 nm in thickness, and is amor-phous, as well as being stoichiometrically defective. It is known that theprotective and stable oxides on titanium surfaces are able to providefavourable osseointegration. Studies, both in vitro and in vivo, by Sul et al.[271] have revealed that alterations in the surface oxide on cp Ti implantsstrongly influence the tissue response. The nature of the surface oxides canbe manipulated by thermal oxidation and/or anodic oxidation.

In terms of osteoblast cell responses, Zhu et al. [272] studied the effects ofthe topography and composition of the surface oxides of cp Ti in vitro. Thesurface oxides of the cp Ti were modified in terms of composition andtopography using anodic oxidation in two kinds of electrolytes: the firstbeing 0.2 M H3PO4 and the second being 0.03 M calcium glycerophosphate(Ca-GP) and 0.15 M calcium acetate (CA). Depending upon the type ofanodic oxidation, phosphorus (ca.10 at %) or calcium (1–6 at %) and phos-phorus (3–6 at %) were incorporated into the surface of the cp Ti in the formof phosphate and calcium phosphate. Using the SaOS-2 human osteoblast-like cell line, the study revealed that cell attachment and cell proliferationwere enhanced by the anodic oxidation. Moreover, the cell attachmentwas seen to increase as the thickness of the oxide layers increased. Incontrast, ALP activity was not found to be influenced by the presence ofthe oxides.

Yang, Ong and Tian [273] evaluated in vivo the biological response of a cpTi surface modified by sandblasting and ion implantation of amino (NH2

þ)using an established dog model. The study found that the porous nature ofthe cp Ti surface after sandblasting encouraged osseous tissues to grow intothe pores and thereby allow the formation of a gradual calcium–phosphate(Ca–P) interface layer. In addition, it was concluded that ion implantation ofthe cp Ti surface with the amino groups induced high concentrations of Caand P precipitation and more mineralization as compared with the non-ion-implanted surfaces.

A comprehensive in vitro and in vivo investigation of the bioactivity ofuncoated and collagen-coated cp Ti was conducted by Morra et al. [274]. Inthis study collagen was covalently linked to the surface of cp Ti by thedeposition of a thin film from hydrocarbon plasma followed by acrylic acid

Predictions for Implantation in an in vivo Clinical Situation 137

grafting. Cell growth experiments conducted in vitro using osteoblast-likeSaOS-2 cells revealed that the growth rate was lower on the collagen-coatedcp Ti than on the uncoated cp Ti. No significant difference in ALP produc-tion was detected between the collagen-coated and the uncoated cp Ti.Importantly, the in vivo study using a rabbit model in which samples wereimplanted into rabbit femurs revealed a significant increase in bone growthand bone-to-implant contact in the case of the collagen-coated cp Tiimplants. In addition, an in vivo study of metal implants coated withsmall, synthetic peptides revealed the stimulation of bone formation andsupported in vitro studies showing that adhesion and spreading of osteo-blasts on RGD-modified surfaces was quite substantial [275].

In a study conducted by Wu et al. [276], the bioactivity of a new kind ofanodic oxidized titanium metal was investigated in vitro and in vivo. Afterbeing immersed in SBF solution for 7 days in vitro, apatite formed andcovered almost all the surfaces of the anodic oxidized samples. Furthermore,after an in vivo animal experiment the apatite also precipitated on theinterface of the tissue and materials after 12 weeks post-operation. Ofgreat significance was the observation that there were no fibrous capsulesformed around the materials. The materials bonded with the bone verytightly and attached to the skin very closely, which would result in theachievement of the biological sealing for the bone-anchored percutaneousimplants. These positive results might be attributed to the precipitatedapatite layer formed on the surface of the bioactive oxidized titanium.

For implants fabricated from titanium or titanium alloys, the existingliterature reveals that should a surface treatment process applied to improvethe bioactivity prove not to be cytotoxic in an in vitro analysis, then it is morethan likely that the in vivo functioning of the implant will be improved. Thetreatments in current usage change only one surface characteristic, such assurface topography, surface chemistry, wettability, etc. Even so, it is evidentfrom this section how effective these treatments are at enhancing the in vivoperformance. The results presented in this chapter show that the biocompat-ibility of the Ti–6Al–4V alloy was enhanced considerably as a result of CO2

laser treatment, which changes surface topography, surface chemistry andwettability simultaneously. Thus, since the CO2 laser treatment has beenshown in vitro to be an actual means of improving the bioactivity of theTi–6Al–4V alloy, it would appear that the CO2 laser treatment would beeffective in vivo.

7.7 Summary

This study investigated the apatite formation on the untreated and CO2 lasertreated Ti–6Al–4V alloy after soaking in simulated body fluid. The protein


adsorption and hFOB cells were used to examine the in vitro biologicalresponse on the Ti–6Al–4V alloy following CO2 laser treatment.

It has been demonstrated that the CO2 laser treatment could improve thebioactivity of the Ti–6Al–4V alloy, as some apatite nuclei were only observedon the CO2 laser modified Ti–6Al–4V alloy, while no apatite nuclei appearedon the untreated sample. It is believed that the CO2 laser oxided surfacelayer generated the hydroxide ions in the water and resulted in the nuclea-tion of apatite.

The CO2 laser treatment brought about a lower amount of the adsorbedalbumin layer and a higher amount of the adsorbed fibronectin layer on themodified Ti–6Al–4V alloy compared with the untreated sample. The albu-min adsorption decreased, while the fibronectin increased with theincreased wettability characteristics of the material, indicating that thewettability characteristic is a predominant factor governing protein adsorp-tion. Further, the effect of gp

sv on protein adsorption implied that the proteinadsorption on the Ti–6Al–4V alloy was probably due to the polar andchemical interactions.

The one-day cell adhesion test showed that cells not only adhered andspread better, but also grew faster on the CO2 laser treated sample than theuntreated and mechanically roughed sample. Further, the MTT cell prolif-eration analysis revealed that the roughed surface resulted in a slightenhancement, while CO2 laser treatment brought about the considerableincrease in cell proliferation compared with the untreated sample.

Despite the fact that surface roughness is certainly one of the factorsinfluencing cell adhesion and proliferation, the results presented hereinsuggest that the aspects of wettability characteristics, surface oxygen contentand gp

sv, play an important role in promoting cell proliferation. Indeed, it wasevident that the better wettability characteristics of the CO2 laser treatedTi–6Al–4V alloy were responsible for the improved MTT value. Thus itwould be reasonable to maintain that a correlation exists between the CO2

laser induced wettability characteristics of the Ti–6Al–4V and osteoblast cellbioactivity. It is also evident that the CO2 laser treatment could be a moreeffective way of improving osteoblast cell adhesion than the traditionalmethods currently available, especially mechanical roughening.

Summary 139

8

Enquiry into Possible GenericEffects of the CO2 LaserTreatment on Bone ImplantBiomaterials

If the previously observed improvements to the surface characteristics and thebiocompatibility of the MgO–PSZ and the Ti–6Al–4V alloy after CO2 lasertreatment are common across a whole range of bioinert ceramics and biometals,then the viability of CO2 laser surface treatment of biomaterials will be reinforced.This chapter aims to establish generic relationships in terms of enhanced surfaceproperties and biocompatibility of biomaterials using CO2 laser surface treatment.Studies of the generic effects of CO2 laser treatment on the modification ofwettability characteristics and biocompatibility of a yttria partially stabilisedzirconia (Y–PSZ) bioinert ceramic and a biograde stainless steel were conducted.The findings of these studies were correlated with one another and extended not onlyto other bioinert ceramics and biometals but also to ceramics and metals in general.In this way it was possible to establish the generic nature of the advantageous effectsof the CO2 laser surface treatment of ceramics and metals with regard to improvingwettability and biocompatibility.

8.1 Introduction

Although laser surface processing has aroused growing interest and beenproven to be a controllable and flexible technique for modifying the surfaceproperties of materials, the applications of lasers in surface processing of


biomaterials are still limited. Having said that, it is recognised within thecurrently published work that laser irradiation of material surfaces can effectchanges in the biocompatibility of biomaterials. Lately, several publicationshave investigated the modification of biocompatibility of biomaterial sur-faces following laser irradiation. A CO2 pulsed laser was used to graft apolymer [17] and a rubber [18]. The results showed a marked reduction inplatelet adhesion and aggregation for the modified polymer surface and cellattachment with a greater degree of spreading and flattening on theunmodified rubber surface. Furthermore, L929 fibroblast cells attachedand proliferated extensively on the CO2 and KrF laser treated films [19] incomparison with the unmodified PET, with surface morphology and wett-ability being found to affect cell adhesion and spreading. More recently,Hao and Lawrence found that the laser generated surface properties onmagnesia–partially stabilised zirconia (MgO–PSZ) resulted in bone apatiteformation in stimulated body fluid [23], favourable albumin [24] andfibronectin adsorption [25] as well as a better human fibroblast response[247] and human osteoblast cell adhesion [248] and functions [26]. Inaddition, Hao and Lawrence demonstrated that surface treatment using anHPDL brought about higher wettability and an improved stimulated phy-siological liquid response on a biograde 316 LS stainless steel [277] andbetter osteoblast cell adhesion and proliferation on a Ti–6Al–4V alloy [278].

With the aim being to establish the laser as a generic processing techniquefor improving the biocompatibility of biomaterials, CO2 laser surface pro-cessing was applied to a yttria–partially stabilised zirconia (Y–PSZ) biocera-mic and a biograde stainless steel.

8.2 Ascertaining the Generic Effects of CO2 Laser Treatmenton Bioinert Ceramics

The Y–PSZ was introduced in the biomaterials world several years ago [279]and is frequently used in high-load-bearing sites such as artificial knee andbone screws in an orthopaedic application and dental post-crown in a dentalapplication due to their attractive mechanical properties [139]. Y-PSZceramics, whose minimal requirements as implants for surgery are nowdescribed by the standard ISO 13356 [3], are the materials selected by almostall manufacturers that are introducing zirconia ball heads into the market(see Figure 2.1). Fini et al. [280] tested various materials in healthy andosteopenic bone, and Y–PSZ performed better when implanted in thehealthy bone of rats. More than 300 000 Y–PSZ ball heads have beenimplanted [4], and only two failures were reported [5] up to now. Thereason for the large and rapid development of the zirconia hip joint head is aresult of reduced polyethylene wear rate of the mating component and

142 Enquiry into Possible Generic Effects

improved mechanical properties when compared to alumina ceramics. Theoutstanding fracture strength and toughness of Y–PSZ ceramics, which aredouble that of surgical grade alumina, allow the manufacture of a largevariety of femoral heads with a combination of various head diameters,down to 22 mm. Also, several taper designs with a high reliability areachieved. Such a high reliability is a consequence not only of the improvedmechanical properties of Y–PSZ ceramics but also of the optimization of themanufacturing process from designing to final quality controls.

The properties of laser treated ceramic surfaces are very important indetermining the effectiveness of the treatment of such materials for variousapplications. For the ceramics used as linings in high-temperature wasteincinerators, the surface roughness and wettability influence the interactioncharacteristics of the treated ceramic surfaces with the corrosive species. Ingeneral, the effects of laser treatment on the wetting characteristics ofceramics have not been fully explored yet, with only limited researchwork being conducted [129–131]. The main factors that affect the wettingcharacteristics of a surface are its composition, the content of surface oxygen,the surface morphology, surface energy and the temperature [1, 23]. As forthe Al2O3 based oxide ceramics, the change in surface morphology mainlyaffects wettability characteristics following laser surface treatment as nosignificant changes in the composition and oxygen content caused laserinduced melting and re-solidification [281]. Interactions of the bioceramicswith biofluid and cells are influenced by their surface properties, such assurface roughness and wettability characteristics. The laser can be a highlycontrollable tool to modify the surface roughness and wettability character-istics of the ceramics and can be used to improve the performance of theMgO–PSZ bioinert ceramics. It is reasonable to expect that the laser can be aneffective surface processing technique for bioinert ceramics and can alter theinteraction between the materials and the biological response.


The material under investigation was a 5 % yttria partially stabilised zirconia(Y–PSZ) sheet with dimensions of 10 � 50 � 5 mm3 (Dynamic Ceramic, Ltd).The main mechanical and thermal properties of the Y–PSZ used in this studyare: a density of 6.05 g/cm3, a compressive strength of 2000 MPa, a Vickershardness of 1400 kgf/mm2, a tensile modulus of 205 GPa, a specific heat at25 �C of 400 J/K kg and a thermal conductivity at 20 �C of 2 W/m K. TheY–PSZ sheet was CO2 laser treated as received in the same manner as des-cribed in Section 4.2.2. To investigate the effects of the CO2 laser irradiationon the wetting and surface energy characteristics of the Y–PSZ, a set ofsessile drop control experiments was carried out under the same conditionsas described in Section 4.2.4.

Generic Effects of CO2 Laser Treatment on Bioinert Ceramics 143

The human osteoblastic cell line hFOB 1.19 was obtained from theAmerican Type Culture Collection (Manassas, Inc.). These cells werecultured using the same procedure given in Section 5.4.1. The samples andosteoblast cells were prepared and examined using SEM in the same manneras described in Section 5.4.1. The specimens were seeded with a 0.5 ml cellsuspension of 4 � 105 cell/ml and then cultured with cell culture mediumand maintained in the incubator for one week. The cell culture medium waschanged every 3 days.

8.2.2 Modification of the Surfaces Properties and WettabilityCharacteristics of a Y–PSZ Bioinert Ceramic

As one can see from Table 8.1, the Y–PSZ experienced marked reductions iny with all of the control test liquids as a result of interaction with the CO2

laser beam. The CO2 laser treatment brought about a consistently roughersurface on the Y–PSZ sample compared with the untreated sample. Evidenceof this rougher surface can be seen in Figure 8.1, as well as the fact that the Ra

value increased in a linear fashion as the power density of the lasertreatment increased. The higher the power density applied, the lower they obtained. Wenzel’s equation (4.1) [150] states that when the solid surface isrough, cos yw is large and yw decreases when yw < 90�. In this study, therougher surface generated by the CO2 laser treatment at 1.8 kW/cm2 broughtabout a reduction in y in accordance with Equation (4.1) and previous workconducted by Uelzen and Muller [282]. An additional reduction in y wasobserved when the surface of the Y–PSZ was roughened further by CO2 lasertreatment at 2.3 kW/cm2.

In a similar way, Figure 8.2 reveals that the surface oxygen content of theY–PSZ increased following CO2 laser treatment and increased in a linearfashion as the laser power density increased. This observed increase in thesurface oxygen content of the Y–PSZ would have effected better wettabilitycharacteristics, since oxidation is known to increase the likelihood of wetting[13, 144–147].

Table 8.1 Mean values of the contact angle formed between the untreated andCO2 laser treated Y–PSZ for various power densities and selected control test liquids at25 �C


Polyglycol PolyglycolY–PSZ Glycerol Formamide Etheneglycol E-200 15-200

Untreated 82.4 76.2 64.4 56.8 40.2CO2 laser (1.8 kW/cm2) 74.2 70.9 59.2 53.2 38.0CO2 laser (2.3 kW/cm2) 70.5 67.7 57.0 49.8 36.9


The untreated Y–PSZ presented only a grooved surface, as evidenced inFigure 8.3(a), which may have been generated in the manufacturing proces-sing. In contrast, a hexagonal microstructure appeared on the surface of thesample CO2 laser treated at 1.8 kW/cm2(see Figure 8.3(b)), while a cellularmicrostructure appeared on the surface of the sample CO2 laser treated at

0.0

0.1

0.2

0.3

0.4

0.5

0.6

Untreated CO2 Laser1.8 kW/cm2

CO2 Laser2.3 kW/cm2

Surf

ace

Rou

ghne

ss, R

a (µ

m)

Figure 8.1 Relationship between the surface roughness and the CO2 laser powerdensity of the Y–PSZ

0

20

40

60

80


CO2 Laser2.3 kW/cm2

Surf

ace

Oxy

gen

Con

tent

(at

%)

Figure 8.2 Relationship between the surface oxygen content and the CO2 laserpower density of the Y–PSZ


2.3 kW/cm2 (see Figure 8.3(c)). In both instances, the surfaces of the modifiedY–PSZ displayed a microstructure typical of rapid solidification after CO2

laser irradiation. With a CO2 laser power density of 1.8 kW/cm2, the capped-hexagonal formation becomes the interface (see Figure 8.3(b)). When markedsupercooling is generated with a CO2 laser power density of 2.3 kW/cm2, thecellular structure developed as shown in Figure 8.3(c). Indeed, differentre-solidified microstructures generated by laser irradiation have beenreported by a number of workers conducting research into the lasertreatment of various ceramics. Pei, Omyang and Lei [158] noted that bothequiaxed and dendritic microstructures were obtained in different regions ofthe same laser clad ZrO2 layer, concluding that the differences were relatedto different cooling rates in the various regions of the laser clad ZrO2 layer.Liu [283] obtained similar results in the laser sealing Y2O3–ZrO2 and MgO–ZrO2 ceramic coatings. Hao and Lawrence found the different microstruc-tures on the MgO–PSZ generated at the different CO2 laser power densities[21]. Such differences in microstructure type and size were ascribed to

Figure 8.3 Optical images of the morphology of (a) the untreated Y–PSZ and the CO2

laser treated Y–PSZ with laser power densities of (b) 1.8 kW/cm2 and (c) 2.3 kW/cm2


the varying degrees of constitutional supercooling, which, according toMcCallum, Kramer and Weir [284], are inherent in laser processes.

As has already been demonstrated in Section 4.4, it is possible to estimategd

sv for the Y–PSZ, by plotting the graph of cos y against (gdlv)

12=glv according to

Equation (3.8), as shown in Figure 8.4. Comparing the ordinate interceptpoints of the untreated and CO2 laser treated Y–PSZ–liquid systems inFigure 8.4, it can be seen clearly that for the untreated Y–PSZ, the best-fitstraight line intercepts the ordinate closer to the origin while the best-fitstraight line of samples treated by the laser intercepts the ordinate consider-ably higher above the origin. As with the MgO–PSZ studies described inChapter 4, this shows that polar forces are active across the Y–PSZ interfaceafter CO2 laser treatment, and hence improved wettability and adhesion ispromoted. As was shown previously in Section 4.4, in order to determine gp

sv

for the Y–PSZ, it is necessary to calculate Wdad by using Equation (3.9). Both

Wad and Wdad are related by the straight-line relationship represented by

Equation (3.10). Thus, from the best-fit straight line plots of Wad against Wdad

for the Y–PSZ when it is both untreated and CO2 laser treated, it is possibleto determine the constant a for each separate condition of the Y–PSZ.For the Y–PSZ the value of a is 2.28 (untreated), 2.85 (1.8 kW/cm2) and3.10 (2.3 kW/cm2).

Since it is already known that from Equation (3.11) c is 2.9 for the set ofcontrol test liquids, it is possible to calculate gp

sv directly for the untreated

0.00 0.02 0.04 0.06 0.08 0.10 0.12 0.14 0.16 0.18

1.0

0.8

0.6

0.4

0.2

0.0

−0.2

−0.4

−0.6

−0.8

−1.0

Cos

θ

( )1/2 /dlv lvγ γ

UntreatedCO2 Laser (1.8 kW/cm2)CO2 Laser (2.3 kW/cm2)


12=glv for the untreated and CO2 laser treated Y–PSZ

in contact with the wetting control test liquids


and CO2 laser treated Y–PSZ using Equation (3.14). The determined resultsof surface energies of the untreated and CO2 laser treated Y–PSZ (at theselected CO2 laser power densities) are given in Figure 8.5. As is evidentfrom Figure 8.5, the CO2 laser treatment increased gsv of the Y–PSZ byprimarily increasing gp

sv, as gdsv was similar for all the samples. It is important

to note that because of the long-range ionic interactions between the Y–PSZand the control test liquids, it is highly likely that the thermodynamicallydefined total solid surface energy will be higher than the sum of the gd

sv andgp

sv components of the surface energy.A clear relationship between the value of cos y and the surface energy can

be observed in Figure 8.6, which reveals that an increase in the surfaceenergy will cause a rise in the value of cos y. It is therefore deemed that thesurface energy primarily influences the wettability of the Y–PSZ. Indeed, itwas found by Lawrence [154] that the surface energy was the mostpredominant factor governing the wetting characteristics of the SiO2/Al2O3 based ceramic following irradiation by the high power diode laser.What is more, Hao and Lawrence recently found that the change in surfaceenergy, represented by the change in microstructure features [285], wasidentified as the main mechanism governing the wettability characteristics ofthe MgO–PSZ following CO2 laser irradiation [22]. Since gd

sv only changedslightly after laser treatment, as shown in Figure 8.5, an appreciable increasein the total surface energy were governed by the marked enhancement in gp

sv.The increase in gsv, in particular the increase in gp

sv, has a positive effect uponthe action of wetting and adhesion [286], since primarily both dispersion andpolar forces are active to a greater extent [100, 101]. The changes in gsv arethought to be due to the fact that CO2 laser treatment of the Y–PSZ results in

0

15

30

45

60

75

CO2 Laser

2.3 kW/cm2

CO2 Laser

1.8 kW/cm2

Untreated

γγγγsv

d

γγγγsv

p

γγγγsv

Surf

ace

Ene

rgy

(mJ/

cm2 )

Figure 8.5 Relationship between the surface energy and CO2 laser power density forthe Y–PSZ


the surface melting, a transition that is known to effect an increase in gpsv

[155], and thus an improvement in the wettability and an increase in theadhesion at the interface in contact with the control liquids.

In fact, the re-solidified microstructure generated by laser surface meltingmight be attributed to the changes in the surface energy and thereof inwettability characteristics. The total surface energy values of samples withthe hexagonal structure and cellular structure were 60 and 67.3 mJ/cm2,respectively, being higher than the untreated sample, which had a totalsurface energy value of 46.7 mJ/cm2. Indeed, work conducted by Zhang,Yue and Man [153] found that considerable improvement in the bondstrength of an Si3N4 ceramic could be realised only when excimer lasertreatment of a structural alloy steel (SAE 4340) resulted in surface melting.Similarly, Lawrence [154] observed a sharp reduction in y at the point ofmelting for an Al2O3/SiO2 based oxide compound after HPDL treatment.

8.2.3 Identification of the Predominant Mechanism Activein the Wettability Characteristics Modificationof a Y–PSZ Bioinert Ceramic

It is apparent from the above discussion that the surface roughness, thesurface energy (by way of microstructural changes) and the surface oxygencontent govern the observed changes in the wettability characteristics of theY–PSZ after CO2 laser surface treatment. To determine the extent to whicheach of these factors affect the wettability characteristics of the Y–PSZ,several stages of grinding were used to isolate the various influential factorsand thus analyse and qualitatively establish their role. The grinding

30 40 50 60 70 800.0

0.1

0.2

0.3

0.4

0.5

cosθ

sv (mJ/mm2)γ

Figure 8.6 Relationship between wettability characteristics (cos y, glycerol) and gsv forthe Y–PSZ


procedures were similar to those used for MgO–PSZ which are described inSection 4.5. The observed changes to surface roughness, surface oxygencontent and y (glycerol) caused by these grinding steps are given inTable 8.2.

After the first grinding stage, a large difference in y for glycerol wasobserved between the untreated and the CO2 laser treated samples. How-ever, the decrease in the surface oxygen content of the CO2 laser treatedsample from 59.3 to 44.1 at % resulted in only a small increase in the y,increasing from 70.5 to 73.7�, implying that the decreased surface oxygencontent could be the factor influencing the general increase in y. In thissubsequent stage, the CO2 laser induced microstructure was removed fromthe CO2 laser treated sample and y for the CO2 treated samples was 80.7�

and close to the original untreated value of 79.1�. Basically, the removal ofthe CO2 laser induced microstructure alone appears to have brought aboutan increase in y to around the original level, since the surface oxygen contentwas almost the same value in both ground stages and the surface roughnessonly changed slightly. In addition, the change in the surface roughness of theuntreated sample brought about only a very slight increase in y, revealingthat surface roughness played a very minor role in inducing changes in thewettability characteristics of the Y–PSZ.

8.2.4 Generic Effects of CO2 Laser Treatmenton the Wettability Characteristics of Bioinert Ceramics

As one can see from Figure 8.7, CO2 laser surface treatment of both theMgO–PSZ and Y–PSZ bioinert ceramics caused a general reduction in y,which was decreased further as the CO2 laser power density was increased.Such decreases in y following laser surface treatment and the attendantincrease in wettability were observed by Triantafyllidis, Li and Stott [281],found for the alumina based ceramics. These results suggest that lasersurface treatment of ceramic materials generally brings about a change inthe wettability characteristics.

Table 8.2 The contact angle (for glycerol), surface roughness and surface oxygencontent of the untreated and CO2 laser treated Y–PSZ following the fine grindingstages

Untreated CO2 laser treated (2.3 kW/cm2)—————————————— ———————————————

Polishing steps Ra (mm) O2 (at %) y (deg) Ra (mm) O2 (at %) y (deg)

Unpolished 0.35 44.1 82.4 0.49 59.3 70.5180 grit SiC 0.24 44.7 83.1 0.47 44.4 73.7400 grit SiC 0.10 45.2 83.3 0.44 44.2 80.7800 grit SiC 0.09 45.2 83.3 0.43 44.1 81.9


As was discussed previously in Sections 4.3.3 and 8.2.2, the furtherreduction in y observed for the MgO–PSZ and Y–PSZ when surface treatingwith the CO2 laser is due in part to the increase in surface roughness. This isin accord with Equation (4.1) and is therefore to be expected. As shown inTable 8.3, the surface roughness of both the MgO–PSZ and Y–PSZ increasedwith increasing CO2 laser power density and as a consequence y decreasedaccordingly. The reasons for such an observation have been discussed inSections 4.3.2 and 8.2.2. In fact, it was found by Triantafyllidis, Li and Stott[281] that an almost linear increase in surface roughness with laser powerdensity was seen for laser treated Al2O3 based ceramics and y decreasedwith increasing laser power density. It is therefore speculated that the laserirradiation brought about the rougher surface and thereby genericallycontributed to the better wettability characteristics of zirconia and alumina.Similarly, the improvement in the wettability characteristics of the Y–PSZafter CO2 laser treatment (see Section 8.2.2) will have been influenced by theincrease in surface oxygen content. This is likewise the case for the MgO–PSZ (see Section 4.3.2). Indeed, it was reported by Lawrence [287] that the

0

20

40

60

80

100

Con

tact

Ang

le,

(de

g)θ

Untreated 0.9 1.8 1.6 2.3CO2 Laser(kW/cm2)

MgO-PSZYPSZ

CO2 Laser(kW/cm2)

Figure 8.7 Contact angle (for glycerol) for the MgO–PSZ and the Y–PSZ following CO2

laser treatment

Table 8.3 The surface roughness, surface oxygen content and polar surface energyof the untreated and CO2 laser treated MgO–PSZ and Y–PSZ

Bioinert ceramic MgO–PSZ Y–PSZ

CO2 laser power Untreated 0.9 1.6 Untreated 1.8 2.3density (kW/cm2)

Surface roughness, Ra (mm) 0.29 0.33 0.72 0.35 0.40 0.49Surface oxygen content (at %) 41.5 44.0 64.3 44.1 52.2 59.3gp

sv (mJ/m2) 10.1 21.0 60.7 46.7 60.0 67.3


increase in surface oxygen content was one of the factors influencing theenhanced wettability of an alumina bioceramic following HPDL surfacetreatment. As with the studies of MgO–PSZ described in Section 4.4,Figure 8.5 reveals that polar forces are active across the Y–PSZ interfaceafter CO2 laser treatment. The observed increases in gp

sv resulted from themelting and solidification of the CO2 laser treated MgO–PSZ and Y–PSZsurfaces, with the value varying with the re-solidified microstructure. Thecellular microstructure on the MgO–PSZ and Y–PSZ resulting from the CO2

laser induced rapid solidification corresponded to the maximum value ofgp

sv. This increase in gpsv has been found to play an important role in the

wettability characteristics of the bioinert ceramics examined in this work, asdiscussed in Sections 4.4 and 8.2.2. Indeed, Bradley, Li and Stott [288] foundthat CO2 laser surface treatment generated a denser, more uniform surfacelayer on alumina based materials, with the re-solidified microstructure andensuing increase in surface energy resulting in improved wettability andbonding characteristics. Similar findings were made by Lawrence [287, 289]when using CO2, Nd:YAG and HPDLs to treat the surface of an aluminabioceramic.

From the above discussion it would seem reasonable to postulate that theeffects of laser surface treatment on the surface roughness, surface oxygencontent and gp

sv of the bioinert ceramics studied in this work, as well as theceramics studied by other workers, are generic. Furthermore, based on thispostulation it is perhaps possible to assume that modification of the wett-ability characteristics of the ceramics caused by the changes to these factorsis also generic. This assertion regarding the generic nature of the changes tothe wettability characteristics of ceramics as a result of laser surface treat-ment can be substantiated somewhat by considering the information givenin Table 8.3. Here one can see that the increases in the surface roughness,surface oxygen content and gp

sv of both the MgO–PSZ and Y–PSZ followingCO2 laser treatment appear to have altered in relation to one another, anoccurrence that could only take place if the changes were generic.

Further support for the proposition that CO2 laser surface treatment bringsabout generic modification of the wettability characteristics of the MgO–PSZand Y–PSZ can be obtained from consideration of the findings arising fromthe discussions in Sections 4.5 and 8.2.3. These findings established that thepredominant factor governing modification of the wettability characteristicsof both the MgO–PSZ and Y–PSZ was the increase in gp

sv, which was yieldedby the CO2 laser induced re-solidified microstructure. In addition, Lawrence[132, 154, 289] has identified that increases in gp

sv as the result of laserinduced re-solidified microstructures as being the predominant mechanismgoverning the wettability characteristics modification of a number of cera-mic materials. Again, the incidence of gp

sv being the primary mechanism indetermining the wettability characteristics of such a wide range of ceramics


following surface treatment with a number of different lasers could only beif the changes were actually generic.

8.2.5 CO2 Laser Induced Effects on the Cell Responseon a Y–PSZ Bioinert Ceramic

The hFOB cell adhesions improved considerably on the Y–PSZ after CO2

laser treatment. As presented in Figure 8.8, few osteoblast cells were foundon the untreated Y–PSZ and covered less than 10 % of the surface area of thesample. On the other hand, highly dense osteoblast cells appeared andcovered about 70 and 90 % of the surface area of the CO2 laser treated Y–PSZat 1.8 and 2.3 kW/cm2, respectively. In addition, the morphology of osteo-blast cells on the CO2 laser treatment showed better spreading on the CO2

laser treated sample compared with the cells on the untreated sample.Typical osteoblast cells on the CO2 laser treated sample at 1.8 kW/cm2, asobserved in Figure 8.8(c), had spread completely and flattened, with the

Figure 8.8 SEM image of hFOB cells on (a) the untreated Y–PSZ and the CO2 lasertreated Y–PSZ at laser power densities of (b) 1.8 kW/cm2, (c) 1.8 kW/cm2 at highmagnification and (d) 2.3 kW/cm2


cytoplasmic material spreading to cover a larger area and elongated filopo-dias. On the untreated Y–PSZ sample the situation was different. It isdiscernible from Figure 8.8(a) that the osteoblast cells did not exhibit goodspreading and filopodias. It is very interesting to note that two osteoblastcells adhered in one of the grooves which were generated by the polishingmethods employed by the manufacturers of the Y–PSZ. The groove inquestion had a width of about 10 mm and the two osteoblast cells spreadalong the direction of the groove and connected with each other, as shown inFigure 8.9. Such an observation implies that the surface topography couldinfluence the direction of osteoblast cell spreading and growth.

Cell coverage density, represented by the ratio of the osteoblast cellcovered area and the whole surface area, is used to express the degree ofosteoblast cell adhesion. There are higher cell cover densities on the samplesfollowing the CO2 laser treatment than on the untreated sample, as shown inFigure 8.10. In the power density range employed and in the specifiedexperimental conditions, the increase in power density caused an increase incell cover density. It is certain that the levels of power densities of the CO2

laser treatment are a significant factor in promoting the hFOB cell adhesion,implying that this technique is able to improve the response of the cells tothe Y–PSZ. Indeed, the CO2 laser treatment creates the changes in thetopography and surface energy synchronously. These changes, which pri-marily determined the wettability characteristics, could be the factorsinfluencing the response of the hFOB cells.

As discussed previously, the CO2 laser treatment generated a consistentlyrougher surface on the Y–PSZ compared with the untreated sample, with thevalue of Ra increasing with increases in power density of the CO2 lasertreatment. As shown in Figure 8.10, the CO2 laser treated Y–PSZ with arougher surface had a higher cell cover density compared with the smoothuntreated sample. This agrees with some reports demonstrating that the

Figure 8.9 SEM image of two hFOB cells adhering and spreading within one of theever-present grooves on the untreated Y–PSZ


rougher surface promoted more osteoblast-like cell attachment on titanium[213], apatite–wollastonite glass–ceramic [290] and hydroxyapatite [291].Hence, one could assume that surface roughness plays a role in influencinghuman osteoblast cell adhesion on the laser treatment of the Y–PSZ. Inaddition, the microtopography could affect cell spreading, as shown inFigure 8.9. The spreading of the cells could be defined in the groove andalong the direction of the groove for about 10 mm. A previous findingrevealed that the surface topography also caused the alignment of cellsparallel to the 10 mm grooves present on the metal surfaces [292]. Surfacemicrotopography has been cited as an important factor influencing protein–surface and cell–surface interactions [80]. Curved surfaces, pits, protrusions,cavities, etc., that have sizes and radii comparable with those of thebiological entities (proteins �1–10 nm, cells 1–100 mm) induce biologicalinteractions different from those on a flat surface [80]. It is noted thatthe degrees of the cell adhesion improved markedly when obvious micro-structural change occurred on the Y–PSZ. The 70 and 90 % cell coverdensities occurred on the laser modified surface with hexagonal and cellularmicrostructures, respectively. A number of reasons have been suggestedfor an increased differentiation of osteoblasts on the microstructuredsurface, such as the influence of surface structure on cell shape or the factthat the surface topography creates a specific biochemical microenvironmentaround each cell [214]. Microstructures of about 10–20 mm induced by the

0

25

50

75

100


CO2 Laser2.3 kW/cm2

0.00

0.15

0.30

0.45

0

20

40

60

80

γ sv

(mJ/

m2 )

Ra

(µm

)C

ell C

over

Den

sity

(%

)γ

µ

Figure 8.10 Relationship between the osteoblast cell coverage density, surfaceroughness (Ra), surface energy and CO2 laser power density for the Y–PSZ


CO2 laser treatment on the polyethylene terephthalate (PET) wereregarded as one of the factors for fibroblast cell attachment and spreading[293].

As can be seen from Figure 8.10, cell adhesion increases as the surfaceenergy of the Y–PSZ increases. There were only sparsely adhered cells on theuntreated sample with a surface energy of 46.7 mJ/cm2, and an obviousimprovement of cell adhesion was found on the CO2 laser treated Y–PSZwith a surface energy of 60 mJ/cm2, while maximum cell adhesion occurredon the Y–PSZ with a surface energy of 67.3 mJ/cm2. The work on polymersand glass revealed that cell spreading and substratum surface free energyshowed a characteristic sigmoid relationship both in the presence and in theabsence of serum proteins; good spreading only occurred when the surfaceenergy was higher than approximately 57 mJ/cm2 [115]. It was found thatthe critical parameter for osteoconduction was the initial number of well-attached osteoblastic cells to the bone substitute [218]. The difference in celladhesion was therefore attributed to the difference in wettability character-istics, which is determined by the surface energy, particularly gp

sv betweenthe untreated and CO2 laser treated Y–PSZ, since the change in wettabilitycharacteristics was primarily influenced by the surface energy of the Y–PSZ,especially gp

sv. The finding agrees with previous studies showing the influ-ence of wettability on the attachment and spreading of various cells [213,217–219]. These studies showed good cell attachment and spreading onhigh-energy substrata and poor cell attachment and spreading on low-energy substrata, which accounts for the minimal energetic state of a systemin equilibrium. The behaviour of osteoblastic cells at the surface of HA [218]and at the surface of titanium [213] demonstrated that gp plays a critical role.As shown in Figure 8.5, the dispersion components (gd) of the surface energywere similar, whereas the gp

sv values were significantly different for theuntreated and CO2 laser treated Y–PSZ at various power densities. Thereforeit is possible to say that the change in osteoblast cell adhesion was mainlyrelated to gp. These results indicate that gp influenced the behaviour ofosteoblasts on Y–PSZ surfaces more strongly when compared to gd, whichcan probably be attributed to the fact that the composition of the culturemedium is all polar; thus osteoblast cells and the Y–PSZ should interactmainly by polar force. Moreover, the osteoblast cells showed better spread-ing and flattening at the CO2 laser treated sample. It is likely that the moreflattened osteoblast cells produced more collagen than less flattened osteo-blast cells [218] on the untreated sample. One important aspect to beconsidered in this study is the kinetics of osteoblast cell events. Thedifference in morphology of osteoblast cell attachment might lead to thedifference in terms of osteoblast cell growth. It was suggested that enhancinggp

sv would promote the initial osteoblast cell attachment and spreading, andthereby could bring about a large bone-like matrix synthesis.


8.2.6 Generic Effects of CO2 Laser Treatment on the CellResponse on Bioinert Ceramics

As one can see from Figure 8.11, the CO2 laser surface treatment resulted inboth the MgO–PSZ and the Y–PSZ displaying an increase in the cell coverdensity. As discussed in Section 8.2.4, the CO2 laser surface treatment yieldsgeneric enhancement of the surface roughness and wettability characteristicsof these bioinert ceramics. Further, Chapter 5 and Section 8.2.2 revealed thatimprovements in the cell response of the MgO–PSZ and the Y–PSZ after CO2

laser surface treatment were correlated directly to the CO2 laser inducedenhancement of the wettability characteristics of these ceramics. On accountof the generic effects of the laser surface treatment on the surface roughnessand wettability for other ceramics such as alumina, by extension it wouldappear that improvements in the cell response of the bioinert ceramics afterCO2 laser surface treatment are generic.

8.3 Ascertaining the Generic Effects of CO2 LaserTreatment on Metal Implants

Currently, 316 LS stainless steel is still the most used metal for internalfixation devices thanks to a favourable combination of mechanical proper-ties, acceptable biocompatibility and cost effectiveness when compared toother metallic implant materials [294]. Still, a disadvantage seen for stainlesssteel is its tendency towards corrosion under physiological conditions,

0

20

40

60

80

100

Cel

l cov

er d

ensi

ty (%

)

Untreated 0.9 1.8 1.6 2.3CO2 Laser(kW/cm2)

MgO-PSZYPSZ

CO2 Laser(kW/cm2)

Figure 8.11 Cell cover density (hFOB osteoblast cell) on the MgO–PSZ and Y–PSZfollowing CO2 laser treatment at various laser power densities

Ascertaining the Generic Effects of CO2 Laser Treatment on Metal Implants 157

causing a release of metal ions such as those of nickel (Ni) and chromium(Cr). Stainless steels are actually not immune to crevice corrosion in thehuman body, which can increase ion release in the surrounding tissues byseveral orders of magnitude.

To date, very little published work exists pertaining to the use of lasers foraltering the surface properties of stainless steel in order to improve itswettability characteristics. Notwithstanding this, Lawrence and Li [138]compared the interaction of CO2, Nd:YAG, HPDL and excimer laser radia-tion with the surface of the mild steel and found that changes took place tothe wettability characteristics. HPDL surface treatment of a common engi-neering carbon steel (EN8) was found to effect appreciable changes to thewettability characteristics of the metal [295]. These modifications have beeninvestigated in terms of the changes in the surface roughness of the steel, thepresence of any surface melting, the polar component of the steel surfaceenergy and the relative surface oxygen content of the steel.

Surface properties of biometals, such as surface roughness and wettabilitycharacteristics, influence their interactions with biofluids and cells. As a laserwas used to modify the surface properties of a mild steel in order to improvethe wettability characteristics, it is reasonable to expect that the CO2 lasertreatment can be used to modify the surface of 316 LS stainless steel andthereby improve its biological response. It is possible that successfulapplication of the laser surface treatment to titanium alloys and stainlesssteels would provide evidence of the laser’s potential for application to otherbiometals such as pure titanium and cobalt alloys.


The as-received biodur consumet type 316 LS stainless steel (of implantquality, ground annealed and cold worked) was in the form of a round barwith a diameter of 19 mm (Carpenter, Inc.). The composition and mainproperties of the steel are detailed in Table 8.4. It was cut into disks ofapproximately 3 mm thickness with a diamond-rimmed blade cutter.

Table 8.4 Composition and selected physical and mechanical properties of 316 LSstainless steel

Composition 0.030 C, 2.00 Mn, 0.75 Si, 0.025 P, 0.010 S, 17.00/19.00 Cr�, 13.00/15.00 Ni,2.25/3.50 Mo�, 0.50 Cu, 0.10 N, balance Fe (�3.3 � Mo þ Cr 26.00)

Physical Density Melting range Specific heat Thermal conductivityproperties 7.95 g/cm3 1649 � 15 �C 502 J/kg �C 16.3 W/m K

Mechanical Tensile strength Elastic modulus Hardness Rockwellproperties 883 MPa 190 GPa 36 C


The 316 LS stainless steel was used as received prior to CO2 laser treat-ment. The surface treatment was carried out with a 3 kW CO2 laser in thesame manner as described in Section 4.2.2. A series of experiments wasconducted with a wide range of CO2 laser power densities with a 5.5 mmbeam spot diameter, while the traverse speed was set at 800 mm/min. Toprotect the CO2 laser optics and assist the surface treatment of the 316 LSstainless steel, 2 bar pressure O2 process gas was supplied off-axis.

The surface properties of untreated and CO2 laser modified stainless steelwere examined in the same manner as in Section 6.2.3. To investigate theeffects of laser irradiation on the wetting and surface energy characteristicsof the 316 LS stainless steel, a set of sessile drop control experiments werecarried out in the same way as in Section 4.2.4.

The proteins used for this study were human serum albumin and humanplasma fibronectin (Calbiochem, Inc.). The adsorption of proteins wasconducted in the same manner as described in Section 5.4.1 and wasmeasured by the ellipsometer described in Section 5.4.1. The human osteo-blastic cell line hFOB 1.19 was obtained from the American Type CultureCollection (Manassas, Inc.). These cells were cultured following the sameprocedure described in Section 5.5.1. The samples and osteoblast cells were pre-pared in the same way as explained in Section 5.5.1. The specimens wereseeded with a 0.5 ml cell suspension of 1 � 105 cell/ml and then culturedwith cell culture medium and maintained in the incubator for one week. Thecell culture medium was changed every 3 days. After incubation and gentlyrinsing with PBS, the samples were processed and investigated using SEM inthe same manner as in Section 5.5.1. Cell proliferation on each specimen wasmeasured by MTT assay in the same manner as Section 5.5.1. The statisticalanalysis of the results was performed with a SPSS v.12 software package(SPSS/PC, Inc.) in the same manner as discussed in Section 5.5.1.

8.3.2 Modification of Surfaces Properties and WettabilityCharacteristics of a 316 LS Stainless Steel

As one can see from Table 8.5, the CO2 laser treated 316 LS stainless steelexperienced a consistent reduction in y with all the control test liquids used,showing that the wettability characteristics improved. The CO2 laser treat-ment at a power density of 1.8 kW/cm2 yielded the maximum decrease in y.In addition, y on the mechanically roughened 316 LS stainless steel was seento be only slightly lower than on the untreated sample.

As evident in Figure 8.12, both mechanical roughening and CO2 lasertreatment had a marked effect on the surface roughness of the 316 LSstainless steel. For the CO2 laser treated samples, the Ra value increasedwith increasing laser power density. The CO2 laser treatment at 1.8 kW/cm2

generated a surface with an increase in the Ra value of 65 %, compared with


the untreated specimen. Its effect on surface roughness is approximate to themechanical roughening, which resulted in the roughest surface andincreased the surface roughness by 80 %. It is worth remarking that even acursory cross-referencing of the results given in Table 8.5 and Figure 8.12shows that despite the higher surface roughness occasioned on the 316 LSstainless steel by mechanical roughening, the CO2 laser treatment is moreeffective in enhancing the wettability characteristics of the 316 LS stainlesssteel.

As one can see from Figure 8.13(a), the untreated 316 LS stainless steelsample presented a surface that appeared to be smooth and withoutdepressions or holes. Some irregular marks can be seen, however, whichare ‘scratch’ marks produced on the surface by the rolling process involvedduring manufacture of the sheets. It is clear from Figures 8.13(c) and (d) thatthe surface morphology of the 316 LS stainless steel was altered markedly

Table 8.5 Mean values of contact angle formed between the untreated and CO2

laser treated 316 LS stainless steel and the control test liquids at 25 �C


Polyglycol PolyglycolStainless steel Glycerol Formamide Etheneglycol E-200 15-200

Untreated 73.8 68.1 57.4 52.3 36.4Mechanically roughened 70.2 64.9 55.6 50.2 34.7CO2 laser (1.6 kW/cm2) 69.0 63.4 53.2 48.6 33.2CO2 laser (1.8 kW/cm2) 67.9 61.7 51.3 45.5 31.5

0.0

0.1

0.2

0.3

0.4


CO2 Laser1.6 kW/mm2

CO2 Laser1.6 kW/mm2

Surf

ace

Rou

ghne

ss, R

a (µ

m)

Figure 8.12 Surface roughness values for the untreated, mechanically roughenedand CO2 laser treated 316 LS stainless steel


from the untreated state after the CO2 laser treatment, regardless ofthe laser power density used. Between the two CO2 laser treated samples,the sample treated with a laser power density of 1.6 kW/cm2 exhibited amore uneven surface, which seemed to have a rougher and groove-liketexture, than the sample treated at 1.8 kW/cm2. The surface of the mechani-cally roughened sample was also altered noticeably from the untreatedstate, displaying a generally rougher surface with deep, wide grooves (seeFigure 8.13(b)).

Figure 8.12, together with Figures 8.13(c) and (d) clearly show thatinteraction of the CO2 laser beam with the surface of the 316 LS stainlesssteel effected an increase in the surface roughness. This occurrence is due tothe melting and re-solidification of the 316 LS stainless steel surface. Becausethe melt flow within the meltpool is under the influence of Marangoni forces,a more random surface morphology will result. It is well known that, on awetting surface, the rougher the surface, the higher is the wettability ofthe surface [150]. However, a review of the results given in Table 8.5 and

Figure 8.13 Typical SEM images of (a) the untreated, (b) the mechanically rough-ened and the CO2 laser treated at power densities of (c) 1.6 kW/cm2 and (d) 1.8 kW/cm2 316 LS stainless steel


Figure 8.12 shows that the CO2 laser treatment is more effective in enhancingthe wettability characteristics of the 316 LS stainless steel, in spite of the factthat mechanical roughening generated a rougher surface. This wouldsuggest that the changes in the other surface properties associated withthe CO2 laser treatment (surface oxygen content and surface energy, whichare modified simultaneously) play a significant role in modifying thewettability characteristics of the 316 LS stainless steel.

An XPS analysis was conducted to ascertain the oxygen content on thesurface of the untreated and CO2 laser treated 316 LS stainless steel. As onecan see from Figure 8.14, mechanical roughening occasioned no discerniblechange in the surface oxygen content of the 316 LS stainless steel. In contrast,the CO2 laser treatment brought about, to varying degrees, an increase inthe surface oxygen content. In particular, the sample CO2 laser treated at apower density of 1.8 kW/cm2 experienced a significant increase in the sur-face oxygen content, from 15.9 to 27.6 mass %. Such a relatively largeincrease in the surface oxygen content implies that sufficient melting onthe surface of the 316 LS stainless steel took place, which in turn gave rise tomore oxygen being adsorbed into the material surface. The surface oxygencontent of the sample CO2 laser treated at a power density of 1.6 kW/cm2

increased marginally from 15.9 to 17.7 mass %, an indication of an insuffi-cient degree of melting of the surface of the 316 LS stainless steel to effectsubstantial oxygen absorption. Based on previous work by Lawrence and Li[20, 295], which revealed that the surface oxygen content of a carbon steelincreased after the CO2 laser treatment and contributed to higher wettabilitycharacteristics, it is reasonable to suppose that changes in the surfaceoxygen content for the CO2 laser treated samples are responsible in part

0

10

20

30


CO2 Laser1.6 kW/mm2

CO2 Laser1.8 kW/mm2

Surf

ace

Oxy

gen

Con

tent

(m

ass%

)

Figure 8.14 Surface oxygen content of the untreated, mechanically roughened andCO2 laser treated 316 LS stainless steel


for the overall increase in the wettability characteristics of the 316 LSstainless steel.

The gdsv value for the 316 LS stainless steel are estimated by plotting the

graph of cos y against (gdlv)

12=glv according to Equation (3.8), as shown in

Figure 8.15. Thus, according to Fowkes [100], the value of gdsv is estimated by

the gradient (¼ 2(gdsv)

12) of the line that connects the origin (cos y¼�1) with

the intercept point of the straight line (cos y against (gdlv)

12=glv) correlating the

data point with the abscissa at cos y¼ 1. Comparing the ordinate interceptpoints of the untreated and CO2 laser treated 316 LS stainless steel-liquidsystems in Figure 8.15, it can be seen clearly that for the untreated 316 LSstainless steel, the best-fit straight line intercepts the ordinate closer to theorigin. This is noteworthy for the intercept of the ordinate close to the originis characteristic of the dominance of dispersion forces acting on the 316 LSstainless steel–liquid interfaces of the untreated and low-power densitytreated sample, resulting in poor adhesion [100, 101]. On the other hand,the best-fit straight line of samples treated by the CO2 laser intercept theordinate considerably higher above the origin. An interception of theordinate above the origin is indicative of the action of polar forces acrossthe interface, in addition to dispersion forces, and hence improvedwettability and adhesion is promoted [100, 101]. Furthermore, becausenone of the best-fit straight lines intercept below the origin, it can be saidthat the development of an equilibrium film pressure of adsorbed vapour on

0.00 0.03 0.06 0.09 0.12 0.15 0.18 0.21−1.0

−0.8

−0.6

−0.4

−0.2

0.0

0.2

0.4

0.6

0.8

1.0

Cos

θ

( )1/2/d

lv lvγ γ

UntreatedMacanically rounded

CO2 laser (1.6 kW / cm2)

CO2 laser (1.8 kW / cm2)


12=glv for the untreated, mechanically roughened

and CO2 laser treated stainless steel in contact with the wetting test control liquids


the 316 LS stainless steel surface (untreated and CO2 laser treated) did notoccur [101].

From the best-fit straight line plots of Wad against Wdad for the 316 LS

stainless steel it is possible to determine the constant, a, for each separatecondition of the untreated, mechanically roughened and CO2 laser treated316 LS stainless steel. For the 316 LS stainless steel the value of a is2.10 (untreated), 2.29 (mechanically roughened), 2.48 (1.6 kW/cm2) and2.5 (1.8 kW/cm2). As a linear relationship satisfying Equation (3.11) existsbetween the dispersive and polar components of the surface energy of thetest control liquids, then c is 2.9 for the set of test control liquids. As wasshown previously, it is possible to calculate gp

sv directly for the untreated andCO2 laser treated 316 LS stainless steel using Equation (3.14). The calculatedresults of surface energies of the untreated and CO2 laser treated stainlesssteel (at various laser power densities) are given in Figure 8.16.

This overall change in the surface energy of the 316 LS stainless steel isperhaps one of the most important outcomes of the CO2 laser treatment. Asamply demonstrated by Lawrence and Li [13] and Hao and Lawrence[20–22], the surface energy of ceramic and metals, which results from avariety of intermolecular forces existing within the materials’ elements, canbe increased by laser treatment. This fact is borne out by Figure 8.16, wherethe surface energy of the 316 LS stainless steel changed by varying degreesafter mechanical roughing and CO2 laser irradiation. It is interesting to notethat the mechanically roughened sample experienced a slight increase insurface energy, suggesting that surface roughness influences the surfaceenergy. However, the marked changes in gp

sv were brought about by CO2

laser treatment. From Figure 8.16 it can be seen that gdsv of the sample CO2

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40

Untreated Mechanically CO2 Laser CO2 LaserRoughened 1.6 kW/mm2 1.8 kW/mm2

Surf

ace

Ene

rgy

(mJ/

m2 )

γγγγsvd

γγγγsvp

γγγγsv

Figure 8.16 Surface energy values of the untreated, mechanically roughened andCO2 laser treated 316 LS stainless steel


laser treated with a power density of 1.6 kW/cm2 is lower than the untreatedsample. This sample, however, has higher wettability characteristics than theuntreated sample, denoting that enhancement in gp

sv played a major role ingoverning the improvement in the wettability characteristics of the 316 LSstainless steel. This is in accord with the fact that an increase in gp

sv has apositive effect upon the action of wetting and adhesion [237]. Indeed, thesample CO2 laser treated with a power density of 1.8 kW/cm2 has the highestvalue of gp

sv and gdsv and its wettability characteristics are, unsurprisingly, the

highest of the sample group. Based on the work of Lawrence, Li and Spencer[129], it is believed that the melting and re-solidification of the 316 LSstainless steel surface caused by CO2 laser irradiation of the surface effectedthe observed increases in gp

sv. Furthermore, it seems highly likely that theincrease in gp

sv was influential in precipitating enhancement of the wettabilitycharacteristics of the 316 LS stainless steel.

In order to simulate the biological environment, physiological fluids andsimulated physiological liquids were used for wetting experiments. Theselected physiological and simulated physiological liquids were: humanblood, human blood plasma, simulated body fluid (SBF) and SBFþBSA(bovine serum albumin). The values of y formed between the selectedphysiological and simulated physiological liquids, and the untreated,mechanically roughened and CO2 laser treated 316 LS stainless steel aregiven in Table 8.6. Here it is evident that the values of y for all the selectedphysiological and simulated physiological liquids on the surface of the CO2

laser treated 316 LS stainless steel samples were lower than on either theuntreated or mechanically roughened specimens. This is a clear indicationthat the wettability characteristics of the 316 LS stainless steel improved withregard to the selected physiological and simulated physiological liquids.Because the wetting effect of a solid surface can be a predictive index of thebiocompatibility of the material involved, improvements in the wettabilitycharacteristics of the 316 LS stainless steel would no doubt result in betterbiocompatibility.

Table 8.6 Mean values of contact angles formed between the selected andsimulated physiological test liquids and the untreated, mechanically roughened andCO2 laser treated 316 LS stainless steel

Contact Angle, y (deg)————————————————————————————

HumanStainless steel Human blood blood plasma SBF SBFþBSA

Untreated 46.3 72.1 82.9 61.4Mechanically roughened 45.2 68.3 80.6 60.0CO2 laser (1.6 kW/cm2) 43.6 66.9 78.2 60.4CO2 laser (1.8 kW/cm2) 39.2 63.5 73.2 57.1


Any reduction in y would contribute to an enhancement in Wad of theSBF and SBFþBSA on the 316 LS stainless steel following the CO2 lasertreatment. As both SBF and SBFþBSA have close chemical compositions tohuman body fluids, the augmentation of Wad towards these fluids wouldmean better suitability of the 316 LS stainless steel for use as a biomaterialafter the CO2 laser treatment. Using the referenced glv value for human blood(47.5 mJ/m2), human blood plasma (50.5 mJ/m2) [141], SBF (72.5 mJ/m2) andSBFþBSA (54.0 mJ/m2) [164], the value of Wad for the 316 LS stainless steeltowards the selected physiological and simulated physiological liquids wasdetermined by means of Equation (3.4). As Figure 8.17 shows, the decreasein y following the CO2 laser treatment did indeed yield an increase in thevalue of Wad for the 316 LS stainless steel with respect to the physiologicaland simulated physiological fluids. Moreover, Wad can be seen to increase asthe CO2 laser power density increases. Perhaps as one would expect,Figure 8.17 reveals that the mechanically roughened sample experiencedonly a very minor increase in Wad in relation to the selected physiologicaland simulated physiological liquids.

8.3.3 Identification of the Predominant Mechanism Activein the Wettability Characteristics Modification of a316 LS Stainless Steel

It is evident from the preceding discussion that the increases in the surfaceroughness, surface oxygen content and surface energy of the 316 LS stainless

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Untreated Mechanically CO2 Laser CO2 LaserRoughened 1.6 kW/cm2 1.8 kW/cm2

Wor

k A

dhes

ion,

Wad

(m

J/m

2 )

Human bloodHuman blood plasmaSBFSBF+BSA

Figure 8.17 Work of adhesion of body fluids for the untreated, mechanically rough-ened and CO2 laser treated 316 LS stainless steel


steel, which take place concurrently during the CO2 laser treatment, areresponsible for the increased potential of the steel to wet. Indeed, althoughthe surface roughness of the CO2 laser treated 316 LS stainless steel sampleswas less than that of the mechanically roughened sample, the CO2 lasertreated samples presented a surface with a greater propensity to wet. Thisreinforces the assertion that the changes in other surface properties asso-ciated with the CO2 laser treatment (surface oxygen content and surfaceenergy) play a significant role in modifying wettability characteristics of the316 LS stainless steel. Even so, it is highly likely that surface roughness playsa major part in determining the wettability characteristics of the 316 LSstainless steel as y was seen to decrease by simply roughening the surfacemechanically (see Table 8.5).

To ascertain the extent to which each of these factors affect the wettabilitycharacteristics of the 316 LS stainless steel, several stages of grinding wereused to isolate the various influential factors and thus analyse and qualita-tively establish their role. The grinding procedures were similar to thoseused for the Ti–6Al–4V alloy described in Section 6.5. The observed changesto the surface roughness, surface oxygen content and y caused by thesegrinding steps are given in Table 8.7.

The principal role played by surface roughness is certainly affirmed by theresults obtained after the first fine grinding-down stage. From Table 4.7 itcan be seen that for both the untreated and CO2 laser treated stainless steelsamples, y increases, and in the case of the CO2 laser treated sample, theincrease is considerable. The combination of a decrease in the surface oxygencontent of the CO2 laser treated sample from 27.66 to 15.98 at % to a similaroriginal level to the untreated one and surface roughness from 0.31 to0.19mm resulted in an increase in y, increasing from 67.9 to 72.8�. On theother hand, the change in Ra for the untreated sample ground up from 0.19to 0.38mm generated a similar change in y from 73.8 to 68.4�. The fact that the

Table 8.7 The surface roughness, surface oxygen content and contact angle (forglycerol) of the untreated and CO2 laser treated 316 LS stainless steel (1.8 kW/cm2)following the fine grinding stages

Untreated CO2 laser treated

Ra O2 y (glycerol) Ra O2 y (glycerol)Fine polishing stages (mm) (at %) (deg) (mm) (at %) (deg)

Unpolished 0.19 15.96 73.8 0.31 27.66 67.9

Grinding-down stage 1 0.17 15.92 74.2 0.19 15.98 72.8Grinding-down stage 2 0.15 15.91 74.4 0.17 15.96 74.0

Grinding-up stage 1 0.35 15.90 70.5 0.34 15.92 70.7Grinding-up stage 2 0.38 15.92 68.4 0.36 15.92 69.7


combined effect of the surface oxygen content and a small change in Ra

nearly equals the effect of a large change in Ra on the modification of yreveals that the oxygen surface content does influence y and its effect is lessthan the considerable change in surface roughness. In addition, despite thefact that surface roughness values for the untreated and CO2 laser treatedstainless steel samples are similar after the first fine grinding-down stage,there is a discernible difference in the value of y, being 1.4� lower for the CO2

laser treated sample. Whereas the surface oxygen content of the untreatedstainless steel sample remained around the original value, the surfaceoxygen content of the CO2 laser treated stainless sample reduced to a levelsimilar to that of the untreated sample. As the first fine grinding-down stagedoes not remove the CO2 laser induced microstructure, this difference in ycan, therefore, be taken as an indication that gp

sv is active in determining thewettability characteristics of the stainless steel. Thus it is reasonable toconclude that the wettability characteristics of the stainless steel alloy are,after surface roughness, influenced predominantly by the surface oxygencontent and, to some extent, by the microstructure. The dependency of y onsurface roughness is well known and, moreover, surface roughness has beenidentified as the predominant factor governing changes in wettabilitycharacteristics of steel after surface treatment with various lasers [138, 236].

8.3.4 Generic Effects of CO2 Laser Treatmenton the Wettability Characteristics of Biometals

As one can see from Figure 8.18, CO2 laser surface treatment of both the Ti–6Al–4V and 316 LS stainless steel caused a general reduction in y which wasdecreased further as the laser power density was increased. A similar

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Con

tact

Ang

le,

(deg

)θ

Untreated 1.3 1.6 1.6 1.8CO2 Laser (kW/cm2) CO2 Laser (kW/cm2)

Ti6Al4V alloyStainless steel

Figure 8.18 Contact angle (for glycerol) for the Ti–6Al–4V alloy and the 316 LSstainless steel following CO2 laser treatment


observation was made by Verdier et al. [296], who conducted a study ontopographic and wettability modifications induced by laser treatment on analuminium alloy, a Ti–6Al–4V alloy and a 2C22 ferritic steel. In that work thelaser treatment led to better wetting for the aluminium alloy and the Ti–6Al–4V alloy at certain laser beam energy densities, but did not produce betterwetting for the 2C22 ferritic steel. The reason cited for the lack of improve-ment in the wettability characteristics of the 2C22 ferritic steel after lasertreatment was that the laser power used was limited to 40 W and no oxygenprocessing gas was employed. Additionally, Lawrence and Li [295] success-fully improved the wettability characteristics of carbon steel using an HPDL.It would seem reasonable to infer from such results that laser surfacetreatment of metallic materials is an effective and controllable means forthe modification of wettability characteristics.

As was discussed previously in Sections 6.3.2 and 8.3.2, the reduction in yobserved for the Ti–6Al–4V alloy and 316 LS stainless steel when surfacetreated with the CO2 laser is influenced a great deal by the enhancement ofsurface roughness, which is in accord with Equation (4.1). As shown inTable 8.8, the surface roughness of both the Ti–6Al–4V alloy and 316 LSstainless steel increased with increasing CO2 laser power density. This inturn led to a natural reduction in y. The reasons for such an observationhave been discussed in Sections 6.3.2 and 8.3.2. Such modifications to thesurface roughness for metallic materials have been observed in previousstudies [138, 236, 295, 296]. The improvement in the wettability character-istics of the Ti–6Al–4V alloy and the 316 LS stainless steel after CO2 lasertreatment (see Figure 8.18) would have been influenced by the increase inthe surface oxygen content (see Table 4.8). Indeed, it was reported byLawrence and Li [138, 236, 295] that an increase in the surface oxygencontent was one of the factors influencing the enhanced wettability of acarbon steel following laser surface treatment. The discussions in Sections6.4 and 6.5, along with those of Sections 8.3.2 and 8.3.3, revealed that theincrease in wettability characteristics of the Ti–6Al–4V alloy and the 316 LSstainless steel following the CO2 laser treatment were effected partially bythe increase in gp

sv.

Table 8.8 The surface roughness, surface oxygen content and polar surface energyof the untreated and CO2 laser treated Ti–6Al–4V alloy and 316 LS stainless steel

Ti–6Al–4V alloy Stainless steel

CO2 laser power density Untreated 1.3 1.6 Untreated 1.6 1.8(kW/cm2)

Surface roughness, Ra (mm) 0.35 0.39 0.42 0.19 0.27 0.31Surface oxygen content (at %) 23.0 41.6 49.1 15.9 15.7 27.7gp

sv(mJ/m2) 4.9 9.3 10.4 3.3 6.4 6.7


Following on from this, it would seem reasonable to postulate that the effectsof the CO2 laser surface treatment on the surface roughness, surface oxygencontent and gp

sv of the bioinert metals studied in this work are generic.Furthermore, by extension of this postulation it is possible to maintain thatmodification of the wettability characteristics of the bioinert metals causedby the changes to these factors is also generic. This assertion regarding thegeneric nature of the changes to the wettability characteristics of thebiometals as a result of the CO2 laser surface treatment can be substantiatedsomewhat by considering the information given in Table 4.8. Here, one cansee that the increases in the surface roughness, surface oxygen content andgp

sv of both the Ti–6Al–4V alloy and the 316 LS stainless steel following theCO2 laser treatment appear to have altered in relation to one another. Thissimilar alteration could only have occurred if the changes were generic.

More verification of the assertion that the CO2 laser surface treatmentbrings about generic modification of the wettability characteristics of boththe Ti–6Al–4V alloy and the 316 LS stainless steel can be obtained fromconsideration of the findings developed from the discussions in Sections 6.5and 8.3.3. These findings established that the leading factor governing themodification of the wettability characteristics of both Ti–6Al–4V alloy andthe 316 LS stainless steel was the increase in surface roughness. The surfaceroughness was also identified as the predominant mechanism active ininfluencing the wettability characteristics of other laser treated metallicmaterials [236]. It seems inevitable then to conclude that because surfaceroughness is the primary mechanism in determining the wettability char-acteristics of such a range of metals following surface treatment with anumber of different lasers, the laser induced changes to the wettabilitycharacteristics of these metals were actually generic.

8.3.5 CO2 Laser Induced Effects on Protein Adsorptionand the Cell Response on a 316 LS Stainless Steel

The thickness of the human serum albumin layer on the untreated 316 LSstainless steel sample was found to be higher than on either of the CO2 lasermodified samples, as shown in Figure 8.19. Conversely, Figure 8.19 showsthat the thickness of the human plasma fibronectin layer was less on theuntreated 316 LS stainless steel sample than on both of the CO2 lasermodified samples. What is more, whereas the thickness of the adsorbedhuman serum albumin layer was seen to decrease as the CO2 laser powerdensity applied increased, the thickness of the adsorbed human plasmafibronectin layer increased with increasing laser power density (seeFigure 8.19). In addition, the statistical analysis revealed that the thicknessof the absorbed fibronectin on the untreated 316 LS stainless steel samplewas significantly less than on the sample CO2 laser treated with a power


density of 1.8 kW/cm2, but not significantly less than on the sample CO2 lasertreated with a power density of 1.6 kW/cm2 ðp<0:01Þ. For the untreated316 LS stainless steel sample a disparate situation existed, with the thicknessof the absorbed albumin layer being significantly higher than on either of theCO2 laser modified samples, as shown in Figure 8.19. As stated in Section5.4.2, protein adsorption is influenced by the surface topography [198] andthe surface chemistry (wettability characteristics) [199].

It was found by Deligianni et al. [200] that human serum albumin wasadsorbed preferentially on to a smooth substratum, while the rough sub-stratum bounded a higher amount of total protein (from a culture mediumsupplied with 10 % serum) and fibronectin (10-fold) over a smooth sub-stratum. These findings are reflected by those of the 316 LS stainless steel inthis study so it is possible to declare that the increasing surface roughness ofthe 316 LS stainless steel associated with increases in the CO2 laser powerdensities used brought about directly the observed reduction in the albuminadsorption and enhancement of the fibronectin adsorption.

As one can see from Figure 8.20, as the wettability characteristics of the316 LS stainless steel increased, the adsorbed amounts of albumin decreased.The results of the albumin adsorption tests for the 316 LS stainless steel are

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1.6 kW/cm2 1.8 kW/cm2


1.6 kW/cm2 1.8 kW/cm2

Thi

ckne

ss o

f A

dsor

bed

Pro

tein

Lay

er (

Å)

Fibronectin Albumin

**

Figure 8.19 Thicknesses of the adsorbed fibronectin and albumin layers on theuntreated and CO2 laser treated 316 LS stainless steel with different laser powerdensities. For the fibronectin adsorption, there was a significant statistical differencein thickness between the untreated stainless steel and CO2 laser treated sample at apower density of 1.8 kW/cm2, and no statistical difference between the untreatedstainless steel and CO2 laser treated sample at a power density of 1.6 kW/cm2. For thealbumin adsorption, there was a significant statistical difference in thickness betweenthe untreated stainless steel and CO2 laser treated samples at power densities of 1.6and 1.8 kW/cm2 ð�p < 0:05Þ


consistent with a previous finding by Serro et al. [201], in which the increasein surface hydrophilicity of Ti was seen to result in lower albumin adsorp-tion, implying that the wettability characteristics of the 316 LS stainless steelmight be a factor active in promoting albumin adsorption. Also, Janochaet al. [297] found that the amount of adsorbed BSA protein decreased withincreasing surface energy of the substrate. The interfacial energy of thesolid–liquid interface increases for substrates of high surface energy, whichindicates that the decrease in interfacial energy is not the only driving forceof this adsorption, for expulsion of the proteins from the solution also takesplace (hydrophobic interaction). This would therefore suggest that theenhancement of the surface energy of the 316 LS stainless steel after CO2

laser treatment inhibited the adsorption of the protein. Hence the wettabilitycharacteristics of the material chiefly governed the adsorption of humanserum albumin.

The results of the adsorption of fibronectin show that it increased with theincreasingly wettable characteristics (hydrophilic) of the 316 LS stainlesssteel surface. A previous investigation by Grinnell and Feld [204] on theextent of fibronectin adsorption as compared to its biological activity onhydrophobic and hydrophilic surfaces suggested the possibility that fibro-nectin was more actively adsorbed on the hydrophilic surfaces. The resultsshowed that the antiplasma fibronectin antibody appeared to bind to theconformation of fibronectin adsorbed on hydrophilic surfaces much betterthan the conformation of fibronectin adsorbed on hydrophobic surfaces[204]. It is therefore possible to maintain that for the 316 LS stainless steel,the modification of human plasma fibronectin was influenced predomi-nantly by the wettability characteristics of the steel. It is noticeable that aconsiderable change in gp

sv affected the wettability characteristics of the

−0.05 0.00 0.05 0.10 0.15 0.20 0.250

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Thi

ckne

ss o

f A

bsor

bed

Pro

tein

Lay

er (

Å)


FibronectinAlbumin

Figure 8.20 The relationship between the thickness of the adsorbed fibronectin layerand wettability characteristics (cos y) of the 316 LS stainless steel


316 LS stainless steel after CO2 laser irradiation, signifying that the albuminand fibronectin adsorption on the 316 LS stainless steel surfaces was prob-ably due to the polar and chemical interactions [205].

Generally, cells in contact with the surface of a material will firstly attach,adhere and spread. From Figure 8.21 it is quite clear that the adhesion andspreading of osteoblast cells was influenced by the CO2 laser treatment. Thecells on the untreated and mechanically roughed surfaces shown inFigures 8.21(a) and (b) present the initial stage of the adhesion, withindividual cells covering a small surface area and not spreading. For theCO2 laser treated surfaces the situation was different: the cells showed agood state of adhesion and flattening, leading to coverage of more surfacearea (see Figures 8.21(c) and (d)). This means that the cells showed betteradhesion on the CO2 laser treated samples than on the untreated andmechanically roughened samples, suggesting that the surface propertiesgenerated by the CO2 laser treatment were more favourable for osteoblast

Figure 8.21 Typical SEM images of one-day osteoblast cell adhesion on (a) theuntreated, (b) the mechanically roughed and the CO2 laser treated 316 LS stainlesssteel at a laser power density of (c) 1.6 kW/cm2 and (d) 1.8 kW/cm2


cell adhesion. Moreover, the number of cells on the untreated 316 LSstainless steel sample is nearly the same as on the mechanically roughenedsample, but much lower than on both of the CO2 laser treated samples (seeFigure 8.21). It can therefore be propounded that the CO2 laser treatment ofthe 316 LS stainless steel surface has more of an effect on cell adhesion andgrowth than mechanical roughening of the surface and thus plays a moreimportant role than the surface roughness alone.

From the MTT results given in Figure 8.22, it can be seen that themechanically roughened 316 LS stainless steel sample and both of the CO2

laser treated 316 LS stainless steel samples experienced better cell prolifera-tion than the untreated sample. Further, Figure 8.22 shows that thecell proliferation increased as the CO2 laser power density increased. Inaddition, the statistical analysis revealed that cell proliferation improvedsignificantly on the CO2 laser treated samples compared with the untreatedsample. A comparison of the cell proliferation on the untreated sample withthe mechanically roughened sample revealed that, statistically, mechanicalroughening of the surface produced no significant increase. It thereforeseems apparent from these results that the surface generated by the CO2

laser treatment was more favourable for cell proliferation than either theuntreated or the mechanically roughened surfaces.

In fact, the cells not only adhered better on the surface of the CO2 lasertreated 316 LS stainless steel samples than on either the untreated or the

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0.16

Untreated Mechanically CO2 Laser CO2 Laser

Roughened 1.6 kW/cm2 1.8 kW/cm2

MT

T O

ptic

al D

ensi

ty

Figure 8.22 MTT optical density of osteoblast cells grown on untreated and CO2 lasertreated 316 LS stainless steel after 7 days of cell culture. There was a significantstatistical difference between the untreated sample and samples CO2 laser treatedat laser power densities of 1.6 and 1.8 kW/cm2, and no statistical difference amongthe untreated sample and mechanically roughened samples (�p < 0.05)


mechanically roughened samples (see Figure 8.21), but also grew better (seeFigure 8.22). Additionally, the cell response increased with the powerdensity of the CO2 laser treatment. The observation that the growth ofosteoblast cells was marginally better on the mechanically roughened sur-face compared with the smoother, untreated 316 LS stainless steel surface issimilar to that reported by Feng et al. [213], who noted that a rougher surfacepromotes more osteoblast-like attachment. This demonstrates that surfacetopography does have an effect on osteoblast cell proliferation. However, thesurface generated by the CO2 laser treatment at a power density of 1.8 kW/cm2 exhibited the best cell adhesion and proliferation, yet its surface wassmoother than that of the mechanically roughened sample. The superiorperformance of the CO2 laser treatment 316 LS stainless steel sample in termsof cell adhesion and proliferation is most certainly due to the higherwettability characteristics of this sample. Furthermore, the results revealthat the cell proliferation increased as the wettability characteristics of thesamples increased. Indeed, Hallab et al. [215] demonstrated that the surfacefree energy was a more important surface characteristic than surface rough-ness for cellular adhesion strength and proliferation. Schakenraad et al. [115]found that, despite the great number of parameters interfering with cellularadhesion and spreading, the solid surface energy is apparently a dominatedfactor in cellular attachment to a polymer surface and remains so, even if thesolid surface has been covered by a protein layer. Previous work by Hao andLawrence [247, 248] also showed that enhancement of the wettabilitycharacteristics of the MgO–PSZ after CO2 laser treatment resulted in a betterresponse of human fibroblast and human osteoblast cells. Owing to this, it isreasonable to postulate that the wettability characteristics of the 316 LSstainless steel play a vital role in initiating cell proliferation, and is the keyfactor in producing the observed improved cell adhesion and proliferationover mechanical roughening alone.

8.3.6 Generic Effects of CO2 Laser Treatment on ProteinAdsorption and the Cell Response on Biometals

As one can see from Figure 8.23, the CO2 laser surface treatment resultedin both the Ti–6Al–4V alloy and the 316 LS stainless steel displaying anincrease in cell proliferation (higher MTT optical density). From the discus-sion in Section 8.3.4 it was established that CO2 laser surface treatmentcaused generic enhancement of the surface roughness and, perhaps moreimportantly, the wettability characteristics of these biometals. Furthermore,Chapter 7 and Section 8.3.2 revealed that favourable protein adsorption andimprovements in the cell response of the Ti–6Al–4V alloy and the 316 LSstainless steel after the CO2 laser surface treatment were correlated directlyto the CO2 laser induced enhancement of the surface roughness and


wettability characteristics of these biometals. Because of the generic effects ofthe laser surface treatment on the surface roughness and wettability char-acteristics of other metals, it seems valid to say that improvements in the cellresponse of these biometals after the CO2 laser surface treatment are generic.

8.4 Summary

To ascertain the presence of generic features of the CO2 laser surfaceprocessing technique for improving the biocompatibility of ceramic andmetal biomaterials, surface modifications of the Y–PSZ and the 316 LSstainless steel were conducted. Changes in the surface properties of theY–PSZ and the 316 LS stainless steel following CO2 laser irradiation wereanalysed and the in vitro biological responses of the untreated and the lasermodified materials were evaluated.

It was found that improvement in the wettability characteristics of theY–PSZ as a result of the CO2 laser treatment was primarily due to enhance-ment of the surface energy, particularly gp

sv. This increase in gpsv was

attributed to microstructural changes induced on the surface of the Y–PSZby the melting and re-solidification. An in vitro test using hFOB cellsrevealed that cell adhesion was better on CO2 laser treated samples thanon the untreated Y–PSZ. Although microtopography, especially surfacegrooves, influenced the osteoblast cell spreading, the results suggest thatCO2 laser induced changes in the wettability characteristics of the Y–PSZcould be the main mechanism governing osteoblast cell adhesion.

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0.16

MT

T O

ptic

al D

ensi

ty

Untreated 1.3 1.6 1.6 1.8CO2 Laser (kW/cm2) CO2 Laser (kW/cm2)

Ti6Al4V alloyStainless steel

Figure 8.23 Cell cover density (hFOB osteoblast cell) on the Ti–6Al–4V alloy and the316 LS stainless steel following CO2 laser treatment


For the MgO–PSZ and Y–PSZ bioinert ceramics examined, the CO2 lasersurface treatment of both caused a general reduction in y. This observation,coupled with the findings of others, suggests that not only the CO2 lasersurface treatment but also the surface treatment of a wide range of ceramicmaterials with a variety of lasers is inherently capable of effecting changes inwettability characteristics. Moreover, the findings allow one to postulate thatmodification of the wettability characteristics of the ceramics are generic. It ispossible to claim this because increases in the surface roughness, surfaceoxygen content and gp

sv of both the MgO–PSZ and Y–PSZ following the CO2

laser treatment altered in relation to one another, an occurrence that couldonly take place if the changes were generic. Further support for thisproposition comes from consideration of the predominant factors governingmodification of the wettability characteristics of both the MgO–PSZ andY–PSZ. In each case the predominant factor was identified as the increase ingp

sv. An increase in gpsv for other ceramic materials has been found by others to

be the predominant factor responsible for wettability characteristics mod-ification. Again, the incidence of gp

sv being the primary mechanism indetermining the wettability characteristics of such a wide range of ceramicsfollowing surface treatment with a number of different lasers could only beif the changes were actually generic. Improvements in the cell response ofthe MgO–PSZ and Y–PSZ after the CO2 laser surface treatment werecorrelated directly to CO2 laser induced enhancement of the wettabilitycharacteristics. On account of the generic effects of the laser surface treat-ment on the surface roughness and wettability for other ceramics, theimprovements in the cell response of the bioinert ceramics after the CO2

laser surface treatment appear to be generic.The predominant mechanism active in determining changes in the

wettability characteristics of the 316 LS stainless steel following CO2 lasertreatment was identified as being the increase in surface roughness, whileincreases in the surface oxygen content and gp

sv were shown to be active to alesser extent. An in vitro analysis using osteoblast cells showed betteradhesion on the CO2 laser treated samples than on the untreated andmechanically roughened samples, suggesting that the surface propertiesgenerated by the CO2 laser treatment were more favourable for osteoblastcell adhesion. The proliferation of the osteoblast cells was found to be betteron the mechanically roughened sample and both of the CO2 laser treated316 LS stainless steel samples. Furthermore, cell proliferation increased asthe wettability characteristics of the samples increased. Therefore, the increasedwettability of the 316 LS stainless steel played a vital role in initiating cellproliferation and was the key factor in producing the observed improvedcell adhesion and proliferation over mechanical roughening alone.

The observed changes in the wettability characteristics of the Ti–6Al–4Valloy and the 316 LS stainless steel when surface treated with the CO2 laser

Summary 177

were deemed to be generic as the increases in the surface roughness, surfaceoxygen content and gp

sv of both biometals appear to have altered in relation toone another. Additional verification of the generic nature of thewettability characteristics modification of both biometals comes from thefact that the leading aspect governing the modification of the wettabilitycharacteristics of both biometals was the increase in surface roughness. Sincesurface roughness is the predominant mechanism active in influencing thewettability characteristics of other laser treated metallic materials, then itseems highly likely that the laser induced changes to the wettabilitycharacteristics of most metals will actually be generic. Favourable proteinadsorption and improvements in the cell response of the Ti–6Al–4V alloyand the 316 LS stainless steel after the CO2 laser surface treatment werecorrelated directly to CO2 laser induced enhancement of the surface rough-ness and wettability characteristics of these biometals. Because of the genericeffects of the laser surface treatment on the surface roughness and wett-ability characteristics of other metals, it seems valid to say that improve-ments in the cell response of these biometals after the CO2 laser surfacetreatment are generic.


Conclusions

To investigate the efficacy of the CO2 laser as a means for improving thebioactivity and biointegration of bone implants materials, work was con-ducted to alter the surface properties of two widely used bioinert ceramics,magnesia–partially stabilised zirconia (MgO–PSZ) and yttria–partially sta-bilised zirconia (Y–PSZ), and two established biograde metals, Ti–6Al–4Valloy and 316 LS stainless steel. More specifically, the ability of the CO2

laser to modify the wettability characteristics of the materials and inducefunctional groups, thereby allowing bone-like apatite formation, proteinadsorption and cells to be manipulated, was studied. Valuable inroadshave been made as a result of this work for establishing the CO2 laser as anovel and viable technique for improving the biocompatibility of implantmaterials.

The CO2 laser surface treatment of the MgO–PSZ brought about areduction in the contact angle, y, formed between the MgO–PSZ and thecontrol test liquids, providing a clear indication that the wettability char-acteristics of the material were modified. Moreover, the extent of thiswettability characteristics modification was demonstrated to be variableand controllable by means of manipulation of the CO2 laser operatingparameters. Changes in the wettability characteristics of the MgO–PSZwere attributed to the following factors: (a) an increase in surface roughness,(b) incorporation of oxygen at the MgO–PSZ surface resulting from CO2

laser treatment and (c) the increase in the polar component, gpsv, of the

surface energy resulting from the melting and re-solidification of the MgO–PSZ surface. In addition, gp

sv for the MgO–PSZ was seen to increase as thecrystal size and the presence of the tetragonal phase increased after the CO2

laser treatment. Further analysis revealed that gpsv, by way of the re-solidified

microstructure, was the primary influential factor governing changes in yand hence the wettability characteristics of the MgO–PSZ. Incorporation of


oxygen at the surface was also shown to influence, to a lesser extent, changesin the wettability characteristics, while surface roughness was found to playa very minor role in inducing changes in the wettability characteristics of theMgO–PSZ.

The bioactivity of the untreated and CO2 laser modified MgO–PSZ wasinvestigated in SBF, while protein adsorption and hFOB cells were used toexamine the in vitro biological response. It was demonstrated that the CO2

laser treatment could improve the bioactivity of the MgO–PSZ surface bygenerating functional groups to facilitate the formation of bone-like apatites.The apatite formed readily on the CO2 laser treated MgO–PSZ samples, withrelatively high amounts of hydroxyl groups being generated. In contrast,no apatite formation was observed on the untreated MgO–PSZ samplesand consequently few hydroxyl groups were generated. Further analysisrevealed that the Zr–OH groups on the surface of the CO2 laser treatedMgO–PSZ samples were the functional groups facilitating the apatiteformation; the surface melting on the MgO–PSZ induced by the CO2 laserprocessing provided the Zr4 þ ion and OH� ion and in turn created the Zr–OH group. Compared with the untreated MgO–PSZ, the CO2 laser treatmentbrought about a thinner adsorbed albumin layer and a thicker adsorbedfibronectin layer on the MgO–PSZ surface. Whereas the albumin adsorptiondecreased, the fibronectin increased with increased wettability, indicatingthat the wettability of the MgO–PSZ was the predominant factor governingprotein adsorption. Further, the observed effect of gp

sv on protein adsorptionimplied that protein adsorption on the MgO–PSZ surface was probably dueto polar and chemical interactions. Better hFOB osteoblast cell responseswere witnessed on the CO2 laser treated MgO–PSZ samples in comparisonwith untreated samples. Generally, the cell cover density increased withincreasing CO2 laser power density. The change in topography induced bythe CO2 laser treatment is certain to be one of the factors influencing thehFOB osteoblast, but its role will be minor. The improved wettabilitycharacteristics of the MgO–PSZ due to enhanced surface energy broughtby the CO2 laser treatment, especially gp

sv, played a significant role inprecipitating initial cell attachment and spreading in high numbers, conse-quently enhancing long-term cell adhesion and growth.

The CO2 laser surface treatment of the Ti–6Al–4V alloy brought about areduction in the y formed between the Ti–6Al–4V alloy and the simulatedphysiological liquids, signifying that the wettability characteristics of thematerial were modified. It was found that modification of the surfaceroughness, surface oxygen content and surface energy of the Ti–6Al–4Valloy following the CO2 laser treatment were the factors influencing thewettability characteristics. It was found that the wettability characteristics ofthe Ti–6Al–4V alloy were, after the surface roughness, influenced by thesurface oxygen content and, to some extent, by the microstructure. The

180 Conclusions

reductions in y occasioned by the CO2 laser treatment were found tocontribute to an augmentation in the work adhesion of selected physiologi-cal liquids (SBF and SBF þ BSA) on the Ti–6Al–4V alloy. As both SBF andSBF þ BSA have close chemical compositions to human body fluids, theincrease in the work adhesion of the Ti–6Al–4V alloy surface towards thesefluids would mean better suitability of the Ti–6Al–4V alloy for use as abiomaterial after the CO2 laser treatment.

Apatite formation on the untreated and CO2 laser treated Ti–6Al–4V alloyafter soaking in SBF was used to investigate bioactivity. In addition, proteinadsorption and the hFOB cell response were used to examine the in vitrobiological response on the untreated and CO2 laser treated Ti–6Al–4V alloy.The fact that apatite nuclei were observed only on the CO2 laser modifiedTi–6Al–4V alloy samples demonstrated that the CO2 laser treatment wascapable of improving the bioactivity of the Ti–6Al–4V alloy; no apatite nucleiappeared on the untreated samples. It is believed that the CO laser-inducedoxidised surface layer on the Ti–6Al–4V alloy generated the hydroxide ionsin the water and resulted in the nucleation of the apatite. The CO2 lasertreatment brought about a thinner adsorbed albumin layer and a thickeradsorbed fibronectin layer on the Ti–6Al–4V alloy compared with theuntreated samples. Moreover, the albumin adsorption was seen to decrease,while the fibronectin increased, with increasing wettability of the Ti–6Al–4Valloy. This would suggest that the wettability characteristics of the Ti–6Al–4V alloy are the chief driver for protein adsorption. Further, the observedeffect of gp

sv on protein adsorption implied that the protein adsorption on theTi–6Al–4V alloy surface was most likely due to the polar and chemicalinteractions. One-day cell adhesion tests showed that cells not only adheredand spread better but also grew faster on the CO2 laser treated Ti–6Al–4Valloy sample than on either the untreated or mechanically roughened (thetraditional method of improving cell adhesion) samples. Additionally,compared with the untreated sample, MTT cell proliferation analysisrevealed that mechanical roughening of the surface of the Ti–6Al–4V alloyresulted in only a slight enhancement, while the CO2 laser treatment broughtabout a considerable increase. Although surface roughness is surely one ofthe factors influencing cell adhesion and proliferation, certain other aspectsof wettability characteristics – surface oxygen content and – were found toplay an important role in promoting cell proliferation. Indeed, it was evidentthat the better wettability characteristics of the CO2 laser treated Ti–6Al–4Valloy were responsible for the improved MTT value. Thus it would bereasonable to maintain that a correlation exists between the CO2 laserinduced wettability characteristics of the Ti–6Al–4V and the hFOB osteoblastcell bioactivity. Moreover, it is apparent from the results that the CO2 lasertreatment could be a more effective way to improve osteoblast cell adhesionthan the traditional methods currently available.

Conclusions 181

To determine the presence of generic features of the CO2 laser surfaceprocessing technique for improving the biocompatibility of ceramic andmetal biomaterials, surface modifications of the Y–PSZ and the 316 LSstainless steel were conducted. Changes in the surface properties of theY–PSZ and the 316 LS stainless steel following the CO2 laser irradiation wereanalysed and the in vitro biological responses of the untreated and the lasermodified materials were evaluated.

The CO2 laser surface treatment of both the MgO–PSZ and the Y–PSZbioinert ceramics caused a general reduction in y, suggesting that CO2 laserinduced changes to the wettability characteristics of these bioinert ceramicmaterials are generic. It is possible to claim this as increases in the surfaceroughness, surface oxygen content and gp

sv of both the MgO–PSZ and the Y–PSZ following the CO2 laser treatment altered in relation to one another, anoccurrence that could only take place if the changes were generic. Thepredominant factor governing modification of the wettability characteristicsof both the MgO–PSZ and the Y–PSZ was identified as the increase in gp

sv,further reinforcing the proposition that the changes are generic. Improve-ments in the cell response of the MgO–PSZ and the Y–PSZ after the CO2

laser surface treatment were correlated directly to CO2 laser inducedenhancement of the wettability characteristics. On account of the genericeffects of the CO2 laser surface treatment on the surface roughness andwettability, improvements in the cell response of the bioinert ceramics afterthe CO2 laser surface treatment appear to be generic.

Observed changes in the wettability characteristics of the Ti–6Al–4V alloyand the 316 LS stainless steel when surface treated with the CO2 laser weredeemed to be generic as increases in the surface roughness, surface oxygencontent and gp

sv of both biometals appear to have altered in relation to oneanother. Additional verification of the generic nature of the wettabilitycharacteristics modification of both biometals comes from the fact that theleading aspect governing the modification of the wettability characteristicsof both biometals was the increase in surface roughness. Because surfaceroughness is the predominant mechanism active in influencing the wett-ability characteristics of other laser treated metallic materials, it seemshighly likely that the laser induced changes to the wettability characteristicsof most metals will actually be generic. Favourable protein adsorptionand improvements in the cell response of the Ti–6Al–4V alloy and the316 LS stainless steel after the CO2 laser surface treatment were correlateddirectly to CO2 laser induced enhancement of the surface roughness andwettability characteristics of these biometals. Because of the generic effectsof the laser surface treatment on the surface roughness and wettabilitycharacteristics of other metals, it seems valid to say that improvements in thecell response of these biometals after the CO2 laser surface treatment aregeneric.

182 Conclusions

The major focus of this work was the employment of a CO2 laser forthe surface processing of bioinert ceramics and biograde metals that arewidely used as load-bearing bone implants. This contemporary research, inconjunction with the findings of others, firmly establishes the potential ofthe CO2 laser for improving the biocompatibility of a variety of otherbiomaterials.

Conclusions 183

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Index

Absorption coefficient, 69, 72Adhesion, 5, 23, 24, 61, 82, 114, 127,

128Air incubator, 127Alkaline phosphatase assay, 82Alumina, 14Attractive van der waals force, 30, 31

Bioactivity, 1, 3, 66, 119Biochemical methods, 7, 135Biocompatibility, 2, 65, 117Bioinert ceramic, 13, 142, 144, 149, 150, 153,

157Biointegration, 1, 5, 66, 119Biological Response, 2Biomaterial interface, 29Biomaterials, 2, 11, 12, 17, 18, 19, 23, 29, 31,

141Bio-metals, 141, 158, 168, 175Biomolecule, 8, 9, 12Bonding, 23, 28Bone cell adhesion, 5Bone-implant interface, 3, 6Bone Implants, 3, 141Bone like apatite, 67, 120Bovine serum albumin, 61, 75, 114,

123, 165

Cell adhesion, 82, 128Cell attachment, 83Cell cover density, 90, 91, 93, 154Cell culture, 81, 127

Cell cytotoxicity, 81, 83, 127, 128Cell formation, 51Cell growth, 85Cell morphology, 82Cell proliferation, 82, 128, 130Ceramic materials, 35Chemical bonding, 28Chemical reactions, 33CO2 laser, 37, 39, 41, 65, 77, 89, 99, 102, 103,

117, 121, 124, 131, 141, 142, 150, 153, 157,168, 170, 175

Compressive strength, 38, 143Contact angle, 24, 41, 103Continuous wave (CW), 20, 39Conversion coating, 19Corona discharge, 35Corrosion, 16, 21Cross-section, 39Crystal size, 56Culture polystyrene plate, 81, 82, 128

Dendrite structure, 47Density, 7, 19, 38, 43, 44, 47, 48, 51, 52, 57–59,

61–63, 69–72, 74, 77, 78, 82, 85, 89–93, 101,103–105, 108, 110, 115, 124, 125, 130, 132,143, 144, 146, 150, 151, 154, 157, 159, 161–163, 165, 166, 168–171, 174, 175

Dental implant, 1, 5, 11Diamond rimmed cutting blade, 38, 39, 67,

120Dispersive component of surface energy, 26,

47


Electrowetting, 35Ellipsometer, 76, 123, 159Energy dispersive X-ray analysis, 39Etheneglycol, 40, 102Excimer laser, 20–22, 35, 36, 38, 47, 60, 61,

100, 101, 149, 158

Fetal bovine serum, 81, 127Filopodia, 66, 84, 154Formamide, 40, 102Fourier Transform Infrared Spectrometer

(FTIR), 68, 69Functional group, 4

Generic effect, 141, 142, 150, 157, 168,175

Glycerol, 40–43, 52, 54, 92, 102, 112, 150

Hemocytometer, 82Hexagonal structure, 47, 149High power diode laser, 36, 47, 148Host response, 2, 11Human blood, 61, 62, 114, 115, 165, 166Human blood plasma, 61, 62, 68, 95, 114, 115,

121, 165, 166Human plasma fibronectin, 76, 123, 124, 159,

170, 172Human serum albumin, 75, 76, 78, 123, 124,

159, 170, 171, 172Hydrophilic, 2, 33, 172Hydrophobic, 2, 33Hydroxide ions, 121, 122, 139Hydroxyl group (OH group), 69, 70

In vitro, 2, 65, 117In vivo, 95, 135Infrared (IR), 69Instrument plate reader, 128Interfacial biophysics, 30Ion-beam assisted deposition, 18, 34Ion beam processing, 18, 34Ion implantation, 18, 34, 137

LS stainless steel, 142, 157–159, 166,170

Lactate dehydrogenase, 81, 127Langmuir-blodgett deposition, 19Laser, 19, 20, 21, 22, 35, 36, 37, 39, 41, 67, 89,

99, 103, 117, 121, 124, 131, 141, 142, 150,153, 157, 168, 170, 175

Laser grafting, 22Laser patterning, 20Lasers, 19, 35, 38, 111, 118, 119, 141, 153, 158,

168, 170, 177Magnesia partially stabilised zirconia, 37, 65,

118MAPLE direct write, 21Matrix-assisted pulsed laser evaporation, 21Mechanical bonding, 28Mechanical properties, 1, 16Metallic materials, 36Microstructure, 45, 56Morphology, 82, 128MTT assay, 128, 159

Nd:YAG laser, 58, 69, 136

Ordinary Portland cement, 36Orthopaedic implant, 1, 5, 11Osseointegration, 5, 12, 38, 101, 135Osteoblast cell, 6, 65–67, 80–86, 89–91, 93, 98,

117, 119, 120, 127, 128, 131, 132Oxidation, 34, 107, 108, 121, 137, 144

Parylene coating, 19Phase, 39, 56, 61, 102, 105Phosphate buffered salines, 76, 123Photografting, 17Physical bonding, 28Physicochemical methods, 7Physiological liquids, 61, 114Plasma surface modification, 18, 33Polar component of surface energy, 63, 75,

164Polyglycol 15-200, 40, 102Polyglycol E-200, 40, 102Predominant mechanisms, 52, 111, 149, 166Profilometer, 39Protein adsorption, 75, 76, 123, 170, 175Pulsed laser deposition, 20

Radiation grafting, 17, 34Radiation, 17, 34, 41, 99, 103Rapidly solidified microstructure, 45Repulsive electrostatic force, 24, 31Rutile, 107

Scanning electron microscopy, 12, 66Self-assembled monolayers, 4, 19Sessile drop, 40, 41, 102, 143, 159

210 Index

Silanization, 19Simulated body fluid, 119, 142Specific heat, 38, 101, 143Stainless steel, 159, 166, 170Statistics, 76, 83, 124, 128Stress shielding, 16Surface analysis, 12, 39Surface energy, 25, 47, 56, 61, 109Surface melting, 45Surface modification, 7, 8, 11, 17, 19, 33, 35, 36Surface-modifying additive, 19Surface oxygen content, 42Surface properties, 12, 65, 99, 117, 141Surface roughness, 43, 78, 125Surface roughness, 43, 78, 125

Tensile modulus, 38, 143Tensiometry, 29Test control liquids, 164Thermal conductivity, 38, 101, 143Thermodynamic, 31Titanium alloys (Ti6Al4V), 99, 109, 117, 119,

120, 124, 128Topography, 6, 89, 132

Total interaction energy, 31Transverse electromagnetic multimode, 39Tribological, 17, 100Trypsin-EDTA, 82, 128

Ultraviolet (UV), 17, 18, 34Uniform cell structure, 47

Vickers hardness, 38, 143

Wavelength, 39, 58Wettability characteristics, 35, 37, 40, 41, 45,

52, 71, 79, 91, 99, 102, 103, 105, 111, 125,132, 144, 149, 150, 159, 166, 168

Work of adhesion, 24

X-ray diffraction (XRD), 40, 66X-ray photoemission spectroscopy (XPS), 12,

40

Yttria partially stabilised zirconia, 141, 142,143

Zirconia, 14, 37, 65

Index 211

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Laser Surface Treatment of Bio-implant Materials

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