Interactions between titanium surfaces and biological components

PhD thesis Dissertation

Marta Pegueroles Neyra

Supervisors: Professor Conrado José Aparicio Bádenas, PhD

Professor Francisco Javier Gil Mur, PhD

Departament de Ciència dels Materials i Enginyeria Metal.lúrgica E.T.S. d’Enginyeria Industrial de Barcelona

Universitat Politècnica de Catalunya

July 2009

3

Als meus pares, Fernando i Rosa Maria

Al meu germà, Ferran

Al Ferran i a l’Arnau

5

Agraïments

Arribar fins aquí ha estat un camí llarg i molt gratificant, ara mateix estic en un núvol. Evidentment, res hagués estat tan divertit i satisfactori sense una gran quantitat de persones que han estat al meu costat i als quals vull agrair el seu suport, ajuda i confiança.

En primer lloc, el més sincer agraïment és per als Professors Javier Gil i Conrado Aparicio, directors de la present tesis doctoral. Al Professor Javier Gil vull agrair-li moltes coses ja que sense ell no hagués pogut començar aquesta gran aventura; des de l’inici va depositar en mi una gran confiança, convençut de què ho aconseguiria i aquí estic. També agraeixo el seu bon humor i disposició i ajuda sempre que l’he necessitat tant a nivell professional com personal. Al Professor Conrado Aparicio li agraeixo la seva confiança en mi, la paciència en les correccions i discussions de cada un dels resultats obtinguts al llarg de la tesis; també el fet d’ensenyar-me a ser una bona investigadora, exigent, rigurosa i pragmàtica, espero haver-ho conseguit. L’haver estat amb mi dia rere dia en la feina, deixant-me lliure controladament. He après molt de dos grans mestres i la seva aprovació sempre ha estat i serà molt important per a mi, gràcies.

Igualment voldria agrair de forma especial al Professor Josep Anton Planell , per donar-me l’oportunitat de formar part del grup d’investigació de Biomaterials i per la seva exigència en la qualitat científica del grup. Per la seva admirable dedicació a la ciència que ha permès aconseguir grans fites a partir del grup de Biomaterials, com és la creació de l’Institut de Bioenginyeria de Catalunya del qual n’és direc tor.

Al Professor George Altankov per la seva dedicació en temps i treball i ensenyar-me a realitzar múltiples caracteritzacions proteiques i cel.lulars com a expert en interaccions cèl.lules/substrat. El treball es veu reflexat en algun capítol d’aquesta tesis doctoral, del qual crec que n’està orgullós. Per la seva bona disposició i passió en la feina.

A la Dra. Elisabeth Engel per la seva orientació i discussió dels resultats biològics. Per ajudar-me a entendre el món de les cèl.lules, les integrines i les proteïnes.

També vull agrair de forma especial a la Professora Mª Pau Ginebra, que és un exemple de dedicació i rigor. A més, tots els seus comentaris sobre els meus resultats en diverses reunions, han estat d’inestimable ajuda.

M’agradaria donar les gràcies a la Professora Véronique Migonney, directora del Laboratoire de Biomatériaux et Polymères de Spécialité (LBPS) de l’Université Paris 13, en Villateneuse, França, per acollir-me dins del seu grup d’investigació i donar-me l’oportunitat de descobrir el món cel.lular. A la Dra. Graciela Pavon-Djavid, biòloga, per aportar-me un valuós coneixement del treball amb cèl.lules i per fer-me descobrir el món

6

del RT-PCR. Per la seva constant, bona disposició i companyerisme. Al Josselin, la Sandra, la Kimberly, la Jie i la Lamia pels seus bons moments a Paris.

De la mateix manera voldria agrair a tots els membres del Instituto de Engenheria Biomédica (INEB) d’Oporto, Portugal, especialment als Professors Mario Barbosa i Fernando Monteiro per haver-me permès realitzar uns estudis als seus laboratoris i per acollir-me al seu grup. Voldria agrair especialment a la Dra. Ana Cordeiro per haver-me guiat científicament en els estudis de potencial zeta. Per la seva hospitalitat i per la seva valuosa ajuda durant la meva estada.

Voldria expressar el meu més sincer agraïment a tots aquells amb els quals he compartit de forma més directa aquests anys de treball, als meus companys del Grup de Biomaterials:

Al Dr. José María Manero, per la seva disponibilitat, paciència i interès en l’observació amb el microscopi electrònic. Al Dr. Damien Lacroix, per la seva amistat personal i canya, de tant en tant, professional. Al Dr. Javier Peña i Jordi Casals, que van ser els primers amb els quals vaig estar al laboratori, amables, sincers i incondicionals. Agrair especialment a la gent que va entrar amb mi al principi i que ara ja són doctors: al Sebas, per la seva simpatia i bon humor i pels seus pantalons tacats; a l’Alex, perquè la seva amistat i ajuda a la feina que sempre han estat sincers i valuosos per a mi, no canviïs; al Jérôme, per el seu optimisme i bon humor, espero que la vida et convergirà sempre; a la Melba per la seva tenacitat, per ser un valor segur al qual referenciar-se; a la Carol per la seva simpatia innata i obrir-me les portes de la Pepi i del Pedro; a la Montse, per la seva amistat personal, inicis amb el BCA i exemple en molts sentits; al Sergio, per ser tan exigent tant amb la seva feina com amb la meva i per ajudar-me amb les cèl.lules, interferòmetre i múltiples estudis, per la seva amistat. A la Teresa, per ser una gran companya i amiga, per la seva ajuda i consells. Molt especialment vull agrair les estones passades amb el Pablo, que vam començar pràcticament junts i que encara duren, que segueixi així. Per la seva amistat, converses i ajuda en pràcticament tot dins el laboratori i a fora. A l’Aleix per la seva ajuda i hores d’escalada, no amb mi és clar, per ser un gran company. Al Lluís per facilitar-me la feina aquest últim any i ser un manetes. A la Txell per ajudar-me a què no em passi res ja que tot li passa a ella. A la Noelia, per la seva disposició i més que bon humor. A la Marta G, una excel.lent companya. També vull agrair als futurs doctors i post-docs del grup de Biomaterials: Emiliano, Daniel, Miguel, Montse, Lucía, Nathalia, Martin, Johan, Marta M, Clara, Ana G, Edgar, Roman, Milena, Aitor, Gemma, Andrea, Andy, Xavi, Nuno, Hortensia, Tiziano, Virginia. Sense ells la feina no hagués estat ni de bon tros tan divertida, ni els problemes tan menyspreables. Tampoc m’oblido dels alumnes de projecte final de carrera que han posat tanta dedicació a la feina: Enric, Javi, Guillem, Margherita, Chiara, Manu, col.laboradors d’alguns dels estudis desenvolupats en la tesis doctoral.

Agraeixo també a tots els professors, PAS i doctorands del Departament de Ciència dels Materials i Enginyeria Metal.lúrgica, que d’alguna manera han contribuït a que aquesta

7

tesis doctoral arribi a bon port. Especialment vull agrair l’amistat i bons moments oferts per la Jessica, la Gemma ‘noia gimnàs’, el Carlos, el Nayar, el Gonzalo i el Robert.

A tots els membres de l’Institut de Bioenginyeria de Catalunya i de la Plataforma de Nanotecnologia del Parc Científic de Barcelona. Gràcies a ells i a la seva ajuda en nombroses tècniques, la tesis ha pogut avançar com calia. També als serveis Científico-Tècnics de la Universitat de Barcelona, especialment als serveis d’XPS, real time RT-PCR i microscopia confocal i de fluorescència pels assaigs realitzats i suport tècnic.

A l’empresa Materias Primas Abrasivas, S. L., especialment a l’Agustin García, per l’amable cessió de màquines de granallat i per tot el material abrasiu utilitzat en aquesta tesis doctoral. A més de la seva ajuda tècnica en tot moment.

A l’empresa fabricant d’implants Klockner S.A., especialment a Mercedes Roldán y al Dr. Alejandro Padrós, pel seu suport en aquesta investigació amb la financiació de part dels estudis realitzats.

També vull agrair a la Universitat Politècnica de Catalunya per concedir-me una beca de finalització d’estudis i, en l’actualitat, gaudir d’un contracte com a professor ajudant.

Finalment vull agrair a tots els meus amics i familia, per ser-hi. Als meus companys d’escola: Cristina, Claudia, Kenneth, Elisa i Virginia. I als meus companys d’universitat: Georgina, Modest, Eli, Jaume, Ana, Xavi, Glòria, Coke, Congui, Marta G, Marta C, Anna i Jordi. I també a la Berta, la Neus, la Cristina de Xerta i la Mireia, per l’amistat sincera.

La tesis està principalment dedicada als meus pares Fernando i Rosa Mª, perquè sempre han confiat en mi i m’han guiat durant tota la vida. Sóc com sóc i he arribat fins aquí gràcies a ells. Pel seu amor incondicional, sé que estan orgullosos, contents i feliços de mi, per mi i amb mi. Gràcies.

Al meu germà Ferran, pel profund amor que li tinc i per ser carinyós i encantador.

A les meves àvies, per a l’abuelita i la iaia, per adorar-me i ser senzilles. Pel meu iaio i abuelito que ja no estan, sé que estarien molt contents.

Finalment al Ferran, la vida va canviar quan et vaig conèixer i arribar fins aquí amb tu és un honor. Gràcies pel teu amor, suport i canya. Per confiar més amb mi que jo amb mi mateixa. Per ser-hi sempre. I per a l’Arnau, per ser la meva i la nostra alegria constant; per ser tendre, innocent i dolç. Ets una passada.

A tots, gràcies.

9

Index ABSTRACT ...................................................................................................................................................................15

RESUMEN.....................................................................................................................................................................17

1. INTRODUCTION ...........................................................................................................................................21

1.1. SCOPE ..........................................................................................................................................................21 1.2. OBJECTIVES.................................................................................................................................................23 1.3. IMPLANTABLE MATERIALS ......................................................................................................................24

1.3.1. Biocompatibility ...................................................................................................................................24 1.3.2. Osseointegration ..................................................................................................................................25 1.3.3. Titanium as a biomaterial ...................................................................................................................25

1.3.3.1. Titanium ...................................................................................................................................... 25 1.3.3.1.1. Chemical composition and microstructure ...................................................................... 25 1.3.3.1.2. Surface charges on hydrated titanium oxide surfaces .................................................... 27 1.3.3.1.3. Mechanical properties.......................................................................................................... 27 1.3.3.1.4. Corrosion properties ............................................................................................................ 28 1.3.3.1.5. Biological properties ............................................................................................................ 29

1.3.3.2. Biomedical applications ............................................................................................................ 29 1.3.3.3. Surface modification of titanium for biomedical applications............................................ 30

1.3.3.3.1. Physical methods.................................................................................................................. 31 1.3.3.3.2. Chemical methods ................................................................................................................ 32 1.3.3.3.3. Biochemical methods ........................................................................................................... 33

1.4. BONE-IMPLANT INTERFACE ....................................................................................................................34 1.4.1. Bone ........................................................................................................................................................34

1.4.1.1. Bone structure ............................................................................................................................ 34 1.4.1.1.1. Macroscopic structures ........................................................................................................ 34 1.4.1.1.2. Microscopic structures......................................................................................................... 35

1.4.1.2. Bone cells properties and functions ........................................................................................ 36 1.4.1.2.1. Osteoblasts............................................................................................................................. 37 1.4.1.2.2. Osteocytes and Bone-lining cells........................................................................................ 38 1.4.1.2.3. Osteoclasts ............................................................................................................................. 38

1.4.1.3. Mechanical properties of bone................................................................................................. 39 1.4.2. Events at the bone-implant interface ................................................................................................40

1.4.2.1. Material response....................................................................................................................... 42 1.4.2.2. Host response ............................................................................................................................. 42 1.4.2.3. In vitro studies ............................................................................................................................ 43

1.5. BIOMATERIAL-CELL INTERACTIONS THROUGH ADSORBED PROTEINS ...........................................44 1.5.1. Role of adhesion proteins in cell behaviour......................................................................................44

1.5.1.1. Cell adhesion .............................................................................................................................. 44 1.5.1.2. Cell spreading and migration .................................................................................................. 45 1.5.1.3. Cell proliferation and differentiation...................................................................................... 46

1.5.2. Cell response to specific surface characteristics through adsorbed proteins ..............................46 1.5.2.1. Surface chemistry....................................................................................................................... 47 1.5.2.2. Surface topography and roughness ........................................................................................ 47 1.5.2.3. Surface wettability and surface energy .................................................................................. 48 1.5.2.4. Surface charge/polarity ............................................................................................................ 49

1.5.3. Responses to biomaterials in vivo......................................................................................................50 1.5.3.1. Topographic/Chemical surface modifications ..................................................................... 51 1.5.3.2. Bioactive ceramic coating ......................................................................................................... 51 1.5.3.3. Biological molecules surface functionalisation...................................................................... 52

1.6. CONCLUSIONS ...........................................................................................................................................54 1.7. BIBLIOGRAPHY...........................................................................................................................................56

10

2. TOPOGRAPHICAL AND SURFACE ENERGY CHARACT ERIZATION OF DIFFERENT GRIT-BLASTED AND STERILIZED TI SURFACES ..................................................................................65

2.1. INTRODUCTION .........................................................................................................................................65 2.2. OBJECTIVES.................................................................................................................................................68 2.3. TOPOGRAPHY, ROUGHNESS AND SURFACE CHEMISTRY ...................................................................69

2.3.1. Topography/roughness and cell/protein response ..........................................................................69 2.3.2. Surface chemistry and cell/protein response ...................................................................................69 2.3.3. Materials................................................................................................................................................70

2.3.3.1. Grit-blasting................................................................................................................................ 71 2.3.3.2. Sterilisation ................................................................................................................................. 71

2.3.4. Methods .................................................................................................................................................71 2.3.4.1. Surface observation ................................................................................................................... 71 2.3.4.2. Surface roughness ...................................................................................................................... 71 2.3.4.3. Surface chemistry....................................................................................................................... 73 2.3.4.4. Statistics....................................................................................................................................... 74

2.3.5. Results....................................................................................................................................................74 2.3.5.1. Surface observation ................................................................................................................... 74 2.3.5.2. Surface roughness ...................................................................................................................... 75 2.3.5.3. Surface chemistry....................................................................................................................... 76

2.3.6. Discussion .............................................................................................................................................78 2.3.7. Conclusions ...........................................................................................................................................80

2.4. WETTABILITY AND SURFA CE ENERGY ...................................................................................................81 2.4.1. Wettabilty and cell/protein response ................................................................................................81 2.4.2. Wettability.............................................................................................................................................82 2.4.3. Surface tension and surface free energy ...........................................................................................84

2.4.3.1. Young’s equation ....................................................................................................................... 85 2.4.3.2. The Neumann equation of state .............................................................................................. 86 2.4.3.3. The polar-dispersion approach................................................................................................ 86 2.4.3.4. The electron donor-electron acceptor/ Acid-Base/ Good-van Oss approach.................. 87 2.4.3.5. Surface roughness and surface heterogeneity effect on droplet wetting .......................... 88

2.4.4. Materials................................................................................................................................................89 2.4.5. Methods .................................................................................................................................................90

2.4.5.1. Wettability and surface free energy ........................................................................................ 90 2.4.5.1.1. Test conditions ...................................................................................................................... 91

2.4.5.2. Time-related contact angle ....................................................................................................... 92 2.4.5.2.1. Test conditions ...................................................................................................................... 92

2.4.5.3. Statistical analysis ...................................................................................................................... 93 2.4.6. Results....................................................................................................................................................94

2.4.6.1. Wettability and surface free energy ........................................................................................ 94 2.4.6.1.1. Smooth surfaces .................................................................................................................... 94 2.4.6.1.2. Rough surfaces...................................................................................................................... 96

2.4.6.2. Time-related contact angle ..................................................................................................... 100 2.4.7. Discussion .......................................................................................................................................... 103 2.4.8. Conclusions ........................................................................................................................................ 106

2.5. SURFACE CHARGE PROPERTIES............................................................................................................ 108 2.5.1. Surface charge and cell/protein response ...................................................................................... 108 2.5.2. Electrokinetic effects ......................................................................................................................... 108

2.5.2.1. Electrochemical Double Layer and Zeta Potential.............................................................. 108 2.5.2.2. Isoelectric point ........................................................................................................................ 109 2.5.2.3. Methods to determine electrokinetic potential.................................................................... 110 2.5.2.4. Geometrical heterogeneity effects: surface roughness ....................................................... 111

2.5.3. Materials............................................................................................................................................. 111 2.5.4. Methods .............................................................................................................................................. 111

2.5.4.1. Test conditions ......................................................................................................................... 111 2.5.4.2. Streaming Potential ................................................................................................................. 113

2.5.4.2.1. Surface conductivity correction........................................................................................ 114 2.5.4.2.2. PMMA correction ............................................................................................................... 115

2.5.5. Results................................................................................................................................................. 116

11

2.5.6. Discussion .......................................................................................................................................... 117 2.5.7. Conclusions ........................................................................................................................................ 119

2.6. BIBLIOGRAPHY........................................................................................................................................ 121

3. KINETICS AND TWO-STEP COMPETITIVE PROTEIN ADSORPTION CHARACTERIZATION ON TIO2 CRYSTALS USING QCM-D......................................................... 129

3.1. INTRODUCTION ...................................................................................................................................... 129 3.2. OBJECTIVES.............................................................................................................................................. 132 3.3. PROTEIN ADSORPTION .......................................................................................................................... 133

3.3.1. Protein structure and functions ..................................................................................................... 133 3.3.2. Proteins adsorption studies in TiO2 surfaces............................................................................... 135

3.3.2.1. Fibronectin ................................................................................................................................ 136 3.3.2.2. Fibrinogen ................................................................................................................................. 137 3.3.2.3. Bovine serum albumin ............................................................................................................ 137

3.3.3. Surfaces ............................................................................................................................................... 138 3.3.4. Interactions between proteins and surfaces .................................................................................. 140

3.3.4.1. Kinetics of protein adsorption ............................................................................................... 143 3.4. THE QCM-D TECHNIQUE .................................................................................................................... 144

3.4.1. Introduction ....................................................................................................................................... 144 3.4.2. Fundamental principles of the QCM-D technique ..................................................................... 144

3.4.2.1. The dissipation factor .............................................................................................................. 145 3.4.2.2. Viscoelastic properties of the adlayer ................................................................................... 147 3.4.2.3. Overtone sensitivity ................................................................................................................ 150 3.4.2.4. Data interpretation .................................................................................................................. 151 3.4.2.5. Data analysis............................................................................................................................. 153

3.4.2.5.1. Rigid films, Sauerbrey’s approximation ......................................................................... 153 3.4.2.5.2. Viscoelastic films, Voigt’s approximation ...................................................................... 153

3.4.2.6. Other techniques for protein adsorption studies ................................................................ 155 3.4.3. The QCM-D system ......................................................................................................................... 156 3.4.4. TiO2-crystals...................................................................................................................................... 157 3.4.5. Characterization of TiO2-crystals .................................................................................................. 158

3.4.5.1. Morphological and topographical characterization ........................................................... 159 3.4.5.2. Chemical characterization ...................................................................................................... 159 3.4.5.3. Structure characterization....................................................................................................... 161 3.4.5.4. Roughness characterization ................................................................................................... 162 3.4.5.5. Wettability characterization ................................................................................................... 163 3.4.5.6. Statistics..................................................................................................................................... 164

3.4.6. Conclusions ........................................................................................................................................ 164 3.5. SINGLE-PROTEIN ADSORPTION............................................................................................................ 165

3.5.1. Materials............................................................................................................................................. 165 3.5.1.1. TiO2 crystals.............................................................................................................................. 165 3.5.1.2. Protein solutions ...................................................................................................................... 165

3.5.2. Methods .............................................................................................................................................. 165 3.5.2.1. Solution concentration and temperature.............................................................................. 165 3.5.2.2. The experimental assay........................................................................................................... 166

3.5.3. Results................................................................................................................................................. 167 3.5.3.1. Fibronectin adsorption on TiO2-crystals .............................................................................. 167 3.5.3.2. Fibrinogen adsorption on TiO2-crystals ............................................................................... 170 3.5.3.3. Bovine serum albumin adsorption on TiO2-crystals .......................................................... 172 3.5.3.4. Comparison of Fn, Fbg and BSA protein adsorption on TiO2-crystals............................ 174 3.5.3.5. Effect of temperature............................................................................................................... 176

3.5.4. Discussion .......................................................................................................................................... 177 3.5.4.1. Fibronectin adsorption on TiO2-crystals .............................................................................. 177 3.5.4.2. Fibrinogen adsorption on TiO2-crystals ............................................................................... 179 3.5.4.3. BSA adsorption on TiO2-crystals ........................................................................................... 181 3.5.4.4. Comparison of Fn, Fbg and BSA protein adsorption on TiO2-crystals............................ 182 3.5.4.5. Temperature ............................................................................................................................. 184

12

3.5.5. Conclusions ........................................................................................................................................ 184 3.6. TWO-STEP PROTEIN ADSORPTION STUDIES ....................................................................................... 185

3.6.1. Materials............................................................................................................................................. 185 3.6.1.1. TiO2-crystal ............................................................................................................................... 185 3.6.1.2. Protein solutions ...................................................................................................................... 185

3.6.2. Methods .............................................................................................................................................. 185 3.6.2.1. Test conditions ......................................................................................................................... 185 3.6.2.2. QCM-D tests ............................................................................................................................. 186

3.6.3. Results................................................................................................................................................. 187 3.6.3.1. Interaction between BSA, fibrinogen, and TiO2-crystals ................................................... 188

3.6.3.1.1. First protein BSA-Second protein fibrinogen ................................................................. 188 3.6.3.1.2. First protein fibrinogen - Second protein BSA ............................................................... 190

3.6.3.2. Interaction between BSA and fibronectin ............................................................................ 192 3.6.3.2.1. First protein BSA-Second protein fibronectin ................................................................ 192 3.6.3.2.2. First protein fibronectin -Second protein BSA ............................................................... 194

3.6.4. Discussion .......................................................................................................................................... 196 3.6.5. Conclusions ........................................................................................................................................ 200

3.7. BIBLIOGRAPHY........................................................................................................................................ 201

4. DEVELOPMENT OF EARLY FIBRONECTIN MATRIX AND ADHESION OF MG63 OSTEOBLASTS-LIKE CELLS ON C.P. ROUGH TI COATED WITH DIFFERENT BIOLOGICAL PRE-COATINGS....................................................................................................................... 209

4.1. INTRODUCTION ...................................................................................................................................... 209 4.2. OBJECTIVES.............................................................................................................................................. 212 4.3. MATERIALS AND METHODS ................................................................................................................ 213

4.3.1. Materials............................................................................................................................................. 213 4.3.1.1. Surfaces...................................................................................................................................... 213 4.3.1.2. BSA, Fn and Fbg....................................................................................................................... 213 4.3.1.3. Cells............................................................................................................................................ 214

4.3.2. Methods .............................................................................................................................................. 214 4.3.2.1. Protein adsorption ................................................................................................................... 214

4.3.2.1.1. Fluorescent FITC- BSA, Fn, and Fbg preparation.......................................................... 214 4.3.2.1.2. Quantification of FITC-labelled BSA, Fn and Fbg adsorption..................................... 214 4.3.2.1.3. Visualization of adsorbed FITC-fibronectin ................................................................... 215

4.3.2.2. Initial cell adhesion.................................................................................................................. 215 4.3.2.3. Fibronectin matrix ................................................................................................................... 216 4.3.2.4. Statistical analysis .................................................................................................................... 216

4.4. RESULTS ................................................................................................................................................... 217 4.4.1. Protein adsorption............................................................................................................................. 217 4.4.2. Cell adhesion and overall morphology........................................................................................... 219 4.4.3. Fibronectin matrix formation ......................................................................................................... 222

4.5. DISCUSSION ............................................................................................................................................. 225 4.6. CONCLUSIONS ........................................................................................................................................ 229 4.7. BIBLIOGRAPHY........................................................................................................................................ 230

5. EFFECT OF BLASTING-TREATMENT AND FN PRE-COATING ON MG63 ADHESION AND DIFFERENTIATION ON C.P. TI. A GENE EXPRESSION STUDY USING QRT-PCR.. 235

5.1. INTRODUCTION ...................................................................................................................................... 235 5.1.1. Cell adhesion ...................................................................................................................................... 237 5.1.2. Osteogenic cell differentiation markers ......................................................................................... 237 5.1.3. Cell line ............................................................................................................................................... 239

5.2. OBJECTIVES.............................................................................................................................................. 240 5.3. MATERIALS.............................................................................................................................................. 241

5.3.1. Surfaces ............................................................................................................................................... 241 5.3.2. Fn pre-coating.................................................................................................................................... 241

13

5.3.3. Cells ..................................................................................................................................................... 241 5.4. METHODS ................................................................................................................................................ 243

5.4.1. Real time qRT-PCR technique........................................................................................................ 243 5.4.1.1. Gene expression of integrin subunit α5, ALP and OC protocol........................................ 245

5.4.2. Cell differentiation markers ............................................................................................................. 246 5.4.2.1. Total Protein Concentration ................................................................................................... 246 5.4.2.2. Alkaline phosphatase activity................................................................................................ 246 5.4.2.3. Osteocalcin production ........................................................................................................... 247

5.4.3. Statistical analysis ............................................................................................................................ 247 5.5. RESULTS ................................................................................................................................................... 248

5.5.1. Cell adhesion ...................................................................................................................................... 248 5.5.1.1. α5 integrin subunit gene expression...................................................................................... 248

5.5.2. Cell differentiation ............................................................................................................................ 249 5.5.2.1. Alkaline phosphatase gene expression................................................................................. 249 5.5.2.2. Alkaline phosphatase activity................................................................................................ 249 5.5.2.3. Osteocalcin gene expression................................................................................................... 251 5.5.2.4. Osteocalcin production ........................................................................................................... 251

5.6. DISCUSSION ............................................................................................................................................. 253 5.6.1. Cell adhesion ...................................................................................................................................... 253 5.6.2. Cell differentiation ............................................................................................................................ 255

5.7. CONCLUSIONS ........................................................................................................................................ 259 5.8. BIBLIOGRAPHY........................................................................................................................................ 260

6. CONCLUSIONS ........................................................................................................................................... 263

15

Abstract The understanding of cell/protein/biomaterial interactions is critical to the engineering of substrates for numerous biomedical and biotechnological applications and to the clinical success of implants. The final biological response induced by implants is strongly influenced by the biological-components/synthetic-material surface interactions. It is well accepted that the physical and chemical surface properties of a biomaterial rather than its bulk properties will influence the protein adlayer and then the cell response to it, both in vitro and in vivo.

The aim of this PhD thesis is to gain an increased understanding of the material-biosystem interactions, with an emphasis on establishing correlations between surface properties of titanium surfaces and its in vitro biological response.

Commercially pure titanium (c.p. Ti) is being widely and successfully used implant biomaterial in bone surgery over many years. Its excellent biocompatibility is based in its appropriate mechanical properties and, more importantly, in its excellent corrosion resistance, which is mainly due to the presence of a naturally-occurring thin protective titanium oxide film. Modification of titanium surface topography has been a subject of research in the past with the purpose of improving its osseointegration. Grit blasting is one of the most used technologies to roughen titanium surfaces for this purpose. The optimal roughness and type of abrasive blasting-particles for a better in vitro and in vivo response was previously determined in our lab. However, which and how different relevant surface properties of the blasted titanium surfaces induce that optimal biological behavior is still poorly understood.

Smooth/polished and rough c.p. Ti surfaces obtained by blasting with abrasive particles of different chemical composition (Al2O3 and SiC) and different sizes (212–300µm; 425–600µm; 1000-1400µm) were studied.

The comprehensive characterization of physical and chemical surface properties, including roughness, chemical composition, wettability/free energy and electrical charge of the tested surfaces led to a series of relevant conclusions. Among them, it is worth noting that a) the chemical composition of the grit-blasting particles as well as the method of sterilization were found the main factors influencing wettability and surface free energy of the titanium surfaces; b) the sterilization method changed the electron donor character of the surfaces by changing the amount/nature of physisorbed substances on the surfaces, and c) the chemical composition of the blasting particles did not influence on the electrical charge at physiological pH and the isoelectric point of the surfaces.

16

A second step consisted in the use of a quartz crystal microbalance with monitoring of the energy dissipation to study the adsorption kinetics (amount and conformation) and adsorption competition processes of three proteins of special interest in the healing processes of bone -bovine serum albumin (BSA), fibrinogen (Fbg), and fibronectin (Fn)- on smooth TiO2-coated sensors. Different patterns of adsorption with processes in one, two or multiple distinguishable steps were determined depending of the protein in solution. The BSA adlayers showed the most significant changes in their mechanical properties/conformation/incorporation of water until steady protein-adsorption conditions were reached. BSA, the smallest of the tested proteins, displaced Fn and Fbg when in competition for adsorption, which is an indication of its higher affinity for TiO2 surfaces. Fluorescent labelling techniques where used to study protein adsorption on blasted rough surfaces. Most significantly, the amount of Fn and BSA adsorbed on blasted surfaces was positively correlated with their surface energy. The adsorption of fibronectin from solution on shot-blasted rough titanium surfaces resulted in an irregular pattern of adsorption with a higher amount of protein adsorbed on peaks than on valleys of the topography.

Further, the spatial organization of the osteoblast extracellular matrix, ECM, on smooth and rough Ti surfaces was evaluated by visualizing fluorescently-stained Fn-fibrils. Osteoblast-like cells deposited Fn- fibrils in a specific facet-like pattern that was organized within the secreted total matrix . It appeared as a film overlying the top of the different rough titanium surfaces. Interestingly, the thickness of this layer increased with the roughness of the underlying topography, but no more than half of the total maximum peak-to-alley distance was covered.

Finally, taking into consideration the differences in ECM organization and Fn adsorption on the tested Ti surfaces a qRT-PCR study was carried out to elucidate the influence of titanium surface properties with and without Fn-precoatings on the osteoblast response. The expression of α5 integrin subunit gene, as a marker for cell adhesion, was increased in SiC-blasted surfaces compared to alumina-blasted surfaces. This was related to the higher amount of adhesive-protein Fn adsorbed caused by the higher surface energy of SiC-blasted surfaces. The increase of roughness as well as the presence of alumina particles on blasted surfaces increased ALP activity and ALP mRNA gene expression by osteoblasts, and so their differentiation.

This research work contribute to increase our knowledge on the interactions taking place at the bio/non-bio interface between different biological components –water, proteins, cells- and materials of clinical relevance, such as rough titanium. The intertwined effects of the different properties of the synthetic surfaces appear as a challenge to unravel the ultimate causes that determine the fate of cells on synthetic biomaterials.

17

Resumen El conocimiento de las interacciones entre célula/proteína/biomaterial es fundamental para la ingeniería de superficies debido a las numerosas aplicaciones biomédicas y biotecnológicas que se están desarrollando así como al éxito clínico que han alcanzado muchos implantes. La respuesta biológica final inducida por los implantes está fuertemente influenciada por las interacciones superficiales entre los componentes biológicos y el material sintético. Las propiedades físicas y químicas de la superficie de un biomaterial, en lugar de las propiedades en su masa, influyen directamente en la capa de proteínas que se adsorben sobre el biomaterial y, como consecuencia de ello, en la respuesta celular a la misma, tanto in vitro como in vivo.

El objetivo de esta tesis doctoral es profundizar en el conocimiento de las interacciones material-biosistema, con el énfasis en el descubrimiento de relaciones entre las propiedades superficiales de las superficies de titanio y su respuesta biológica in vitro.

El titanio comercialmente puro (Ti c.p.) está siendo ampliamente utilizado con éxito durante muchos años como biomaterial para implantes en cirugía ósea. Su excelente biocompatibilidad se basa en sus adecuadas propiedades mecánicas y, con mayor importancia, en su excelente resistencia a la corrosión. Esta última se debe principalmente a la formación espontanea de una fina película de óxido de titanio que le confiere protección natural contra los ataques degradativos. La modificación de la topografía de la superficie del titanio ha sido objeto de investigación en el pasado con el fin de mejorar la osteointegración. El granallado de partículas es una de las tecnologías más utilizadas para conferir rugosidad a las superficies del titanio. La rugosidad óptima y el tipo de partículas abrasivas del granallado para una respuesta óptima in vitro e in vivo fue previamente determinada en nuestro laboratorio. Sin embargo, todavía están por determinar cuáles son las causas últimas que llevan al biomaterial a su exitosa respuesta biológica.

En este trabajo se han estudiado superficies pulidas y rugosas de Ti c.p. obtenidas mediante el granallado con partículas abrasivas de diferente composición química (Al2O3 y SiC) y diferentes tamaños (212-300µm; 425-600µm; 1000-1400µm).

La completa caracterización de las propiedades física y química de la superficie, incluyendo la rugosidad, la composición química, la mojabilidad/energía libre y la carga eléctrica de las superficies ensayadas ha llevado a una serie de relevantes conclusiones. Entre ellas, cabe destacar que a) la composición química de las partículas de granallado, así como el método de esterilización fueron los principales factores que influyeron en la mojabilidad y la energía libre superficial de las superficies de titanio estudiadas, b) el método de esterilización cambió en la energía superficial el carácter de donante de electrones de las superficies mediante el cambio de la cantidad y la naturaleza de las

18

sustancias adsorbidas, y c) la composición química de las partículas de granallado no influyó en la carga eléctrica a pH fisiológico ni en el punto isoeléctrico de las superficies.

Un segundo paso consistió en el uso de una microbalanza de cristal de cuarzo con monitorización de la energía de disipación, para el estudio de la cinética de adsorción (cantidad y conformación) y de los procesos de adsorción competitiva de tres proteínas de especial interés en los procesos de curación del hueso - la albúmina de suero bovino (BSA), el fibrinógeno (Fbg), y la fibronectina (Fn)- en sensores lisos recubiertos de TiO2. Se determinaron diferentes modelos de procesos de adsorción con una, dos o múltiples pasos distinguibles en función de las proteínas en solución. La capa adsorbida de BSA mostró los cambios más significativos en sus propiedades mecánicas, de conformación y de incorporación de agua hasta que se alcanzaron las condiciones estables de adsorción de proteínas. La BSA, la más pequeña de las proteínas ensayadas, desplazó la Fn y el Fbg cuando se ensayó en condiciones de la competencia por la adsorción, indicando su mayor afinidad por las superficies de TiO2. También se emplearon técnicas de marcaje fluorescente para el estudio de la adsorción proteica en superficies rugosas granalladas. En este estudio, por un parte, se pudo determinar que la cantidad de Fn y BSA adsorbidas en las superficies granalladas está directamente correlacionada con su energía superficial. Por otra parte, se visualizó la adsorción de fibronectina en solución sobre muestras granalladas rugosas de Ti. La Fn formó un patrón irregular de adsorción con una mayor cantidad de proteína adsorbida en los picos que en los valles de la topografía.

También se evaluó la organización espacial de la matriz extracelular de los osteoblastos, ECM, sobre superficies de Ti lisas y rugosas por medio de la visualización de las fibrillas de Fn teñidas con marcador fluorescente. Las células osteoblásticas depositaron las fibrillas de Fn con un determinado patrón organizado dentro de la matriz total secretada. Aparecen como una película que cubre la parte superior de las diferentes superficies rugosas de titanio. Un resultado relevante es que el espesor de esta capa aumentó con la rugosidad de la topografía subyacente. Sin embargo no más de la mitad de la máxima distancia pico-valle se cubrió con la proteína secretada y/o reorganizada.

Por último, teniendo en cuenta las diferencias en la organización de la ECM y la adsorción de Fn en las superficies ensayadas de Ti, se realizó un estudio de qRT-PCR para determinar la influencia de las propiedades superficiales del titanio, con y sin pre-adsorción de Fn, en la respuesta osteoblástica. La expresión génica de la subunidad α5 de la integrina celular, como marcador de la adhesión celular, se incrementó en las superficies granalladas con SiC en comparación con las granalladas con alúmina. Este resultado fue correlacionado con la mayor cantidad de Fn adsorbida debido a la mayor energía superficial de las superficies granalladas con SiC. El aumento de la rugosidad, así como la presencia de partículas de alúmina en las superficies rugosas incrementó la actividad de ALP y la expresión génica de ALP mRNA por los osteoblastos, y por lo tanto su diferenciación.

19

Este trabajo de investigación contribuye a aumentar el conocimiento sobre las interacciones que tienen lugar en la interfaz bio/no-bio entre los distintos componentes biológicos -agua, proteínas, células- y materiales de relevancia clínica, como el titanio rugoso. Las interrelaciones que existen entre las diferentes propiedades de las superficies sintéticas constituyen un desafío para la definición de las últimas causas que determinan el comportamiento de las células sobre los biomateriales sintéticos de interés.

Chapter 1

21

1. Introduction

1.1. Scope

A key issue in most applications of biomaterials is how the material influences on, and is influenced by the biological response that results from the contact between the biomaterial and the biological system at their interface. There is a noticeable lack of understanding about material-biosystem interactions. The initial events include the adsorption of water molecules, ions, and biomolecules, which form a conditioning film on the biomaterial surface. It is trough this hydrated biomolecule coating that cells of the host tissue interact with the biomaterial. Then, the original surface properties of an implanted biomaterial constitute an important starting condition for the dynamics of the biological components-biomaterial interface.

Many authors have demonstrated that the most relevant properties of the biomaterial surface should be structural and chemical properties from the macroscopic scale down to the atomic level. Then, the surface properties should be regarded as important factors underlining the biological response to biomaterials. One key question in biomaterials

Introduction

22

research is then: what is the relationship between different surface properties (chemical composition, morphology, topography, microstructure, etc.) of biomaterials and the biological response to them? The overall objective of this work is to gain an increased understanding of the material-biosystem interactions, with an emphasis on establishing correlations between different surface properties and the response of different biological components when get in contact with the synthetic material.

Commercially pure titanium (c.p. Ti) is being widely and successfully used implant biomaterial in bone surgery over many years in the fields of osteosynthesis, oral implantology, and certain applications in joint prosthetics. Its excellent biocompatibility and corrosion resistance, mainly due to its thin protective oxide film, are their most important properties. Titanium is used in this work as a model material to explore the aforementioned bio/non-bio interactions. This metal has demonstrated ability to allow a close bone-implant apposition, and it has an extensive clinical documentation in dental applications. Moreover, it is a relatively simple material and its surface (oxide) properties can be varied over a relatively wide range without altering other relevant properties.

In previous works [1;2], the optimal roughness and type of abrasive particles for a better in vitro and in vivo response have been determined in our lab. However, it is poorly understood which surface properties mainly induce this optimal biological behavior due to the lack of exhaustive surface characterization. Further biological studies using carefully prepared and systematically varied titanium surfaces are necessary to unravel key parameters that allow us to understand the paths of biological response leading to the biocompatibility of titanium and other materials.

Chapter 1

23

1.2. Objectives

The overall objective of this PhD thesis is to explore and identify fundamental relationships between surface properties of titanium surfaces and its in vitro biological response. The interactions between those metallic surfaces and different biological components –water, proteins, and cells- are the central subject of the research.

The Introduction guides the reader through the main aspects of the cell/protein/material interactions from the macro down to the nano-scale level. First section is a review of the available implantable materials, which is focused on describing titanium and its main properties. Next section describes bone-implant interface where special attention is paid to bone cells. Finally, events at the micro and nano-scale level that occur in biomaterial-cell interactions and adsorbed proteins are presented in the third section.

Introduction

24

1.3. Implantable materials

A biomaterial is a synthetic material used to replace a part of a living system or to function in intimate contact with living tissue. It is formally defined as a “systematically and pharmacologically inert substance designed for implantation within or incorporation with living systems” [3].

1.3.1. Biocompatibility

The success of a biomaterial or an implant is highly dependent on three major factors:

1. the properties and biocompatibility of the implant

2. the health condition of the patient

3. the competency of the surgeon

This section wills focus on the first factor, since it is the only one that can be ameliorated from an engineering point of view. The requirements for a successful implant, are [3]:

a) Biocompatibility

b) Pharmacological acceptability (non-toxic, non-allergenic, non-immunogenic, non-carcinogenic…)

c) Chemically inert and stable

d) Adequate mechanical strength

e) Adequate fatigue life

f) Proper weigth and density

g) Relatively inexpensive, reproducible, and easy to fabricate and process for large-scale production

Metallic, polymeric, ceramic and composite biomaterials always show a minimal reaction in humans or animals, over a reasonable period of time. The response of an implant material in contact with tissues can be defined in terms of biocompatibility as proposed by Williams: “…the ability of a material to perform with an appropriate host response in a specific application” [4].

Among the metallic implants, most material scientists and medical-device engineers favour the titanium-based biomaterials as being generally more inert, more corrosion resistant and more biocompatible in more implant applications than other groups of metallic biomaterials such as cobalt-based alloys and stainless steels. The use of titanium

Chapter 1

25

and its alloys have found its main application in dentistry and orthopaedics as bone-replacing implant.

1.3.2. Osseointegration

The concept of osseointegration was originally introduced by Branemark et al. [5] and the importance of the implant surface properties for a successful osseointegration was first addressed by Albrektsson and co-workers in 1981 [6]. Osseointegration is defined as “a direct structural and functional connection between ordered, living bone and the surface of a load-carrying implant ” [7]. This can be understood as the process and condition for having a stable and loaded biomaterial surface in direct contact with the host bone tissues without intervening connective tissue [5;8]. It is a clinical state that provides long-term stability of prosthesis.

1.3.3. Titanium as a biomaterial

Titanium was first discovered in England in 1790 [9], and since the mid 50’s of the past century has been used as a long-term implantable material [10]. The widespread and successful application of titanium (Ti) in medical implants is due to the combination of a high corrosion resistance and an appropriate mechanical performance, which in turn makes it biocompatible.

1.3.3.1. Titanium

1.3.3.1.1. Chemical composition and microstructure

Titanium is one of the transition elements in group IV and period 4 of Mendeleev’s periodic table. It has an atomic number of 22 and atomic weight of 47,9. Being a transition element, titanium has an incompletely filled d shell in its electronic structure [9]. The incomplete shell enables titanium to form solid solutions with most substitutional elements. In the elemental form, titanium has a high melting point (1668 ºC) and possesses a hexagonal closely packed crystal structure (hcp), also called the alpha phase, up to a temperature of 882,5 ºC [11]. Titanium transforms into a body centred cubic structure (bcc) β above this temperature.

Many of the favourable properties of titanium and titanium alloys arise from the presence of a spontaneously-formed few-nm thick oxide-layer covering the surface. The native oxide film grows spontaneously on the surface upon exposure to air. The excellent

Introduction

26

chemical inertness, corrosion resistance, repassivation ability, and biocompatibility of titanium are thought to result from the chemical stability, high corrosion resistance, a thermodynamically stable state and structure of the titanium oxide film mainly composed of the stable oxide TiO2 at physiological pH values [9;12-14] [15]. In addition, titanium and its oxides have a low ion-formation tendency in aqueous environments [16].

Titanium dioxide (TiO2) is a solid with a melting point of 1855ºC. TiO2 is polymorphous, it exists in multiple modifications or crystal structures [17]. The following crystal structures have been recognized for TiO2: rutile (tetragonal), anatase (tetragonal), and brookite (orthorhombic). TiO2 also exists in an amorphous state [18]. Regarding the surface oxide layer on c.p. Ti dental implant systems spectroscopic studies suggest that the oxide is amorphous and its thickness is approximately 2 to 17nm [14;17;18]. Generally the thinnest oxide layers on dental implants, less than 10nm, are amorphous and morphologically homogeneous. As the oxide thickness is increased up to 40nm, the oxide becomes increasingly more crystalline [15]. The composition and oxide thickness of mechanically polished c.p. Ti surfaces characterized by X-ray photoelectron spectroscopy (XPS) are summarized in Table 1-1 and the characteristics of films grown at room temperature on pure titanium are schematically shown in Figure 1-1.

Table 1-1 Typical XPS oxide film composition and oxide layer thickness of mechanically polished c.p. Ti

surfaces (adapted from [13]) Figure 1-1 Schematics of the oxide film formed on pure titanium [13]

Chapter 1

27

1.3.3.1.2. Surface charges on hydrated titanium oxide surfaces

Hydroxide ions (OH-) attached to metal cations show acid/base properties depending on the type of the metal cations and the coordinate bonds with the cations. Hydroxides or hydro-complexes of multivalent cations are generally “amphoteric”, that is, exhibiting both acid and base (or alkaline) properties. The underlining hydrolysis equations of titanium in an aqueous solution can be shown as follows:

Reaction 1-1 [ ] ++−⇔+− OHO-TiOHOHTi 32

Reaction 1-2 [ ] −++⇔+− OHOH-TiOHOHTi 22

Reaction 1-1 leads to the formation of negative charges on the surface and Reaction 1-2 yields to positive charges [9;13]. A number of papers have been published suggesting that the titanium oxide surface has two hydroxide groups: acidic and basic types [13;16]. As a quantitative measure, the isoelectric point (IEP) is often used to investigate the surface charges and it indicates the pH at which the charge on the solid arising from all sources is zero, i.e. adsorption of positively charged species equals that of negatively charged [16]. The IEP values of titanium vary from 5 to 6,7 [9;13;19].

1.3.3.1.3. Mechanical properties

A distinction is made between four grades of commercially pure titanium, c.p. Ti, characterized by different degrees of purity and mechanical properties. The mechanical properties of the titanium depend very much on the amount of trace elements present. Thus, 4 grades of titanium are distinguished for technical as well as biomedical applications. As the number of the grade goes up, the percentage of trace elements in the metal, such as iron and oxygen, increase (Table 1-2) and, as a consequence the corresponding mechanical properties also increase (Table 1-3).

Table 1-2 Chemical composition for various c.p. Titanium grades corresponding to ASTM F67-06. Maximum compositional limits in weight-%

Nitrogen Carbon Hydrogen Iron Oxygen Titanium

Ti grade 1 0,03 0,08 0,015 0,20 0,18 Balance

Ti grade 2 0,03 0,08 0,015 0,30 0,25 Balance

Ti grade 3 0,05 0,08 0,015 0,30 0,35 Balance

Ti grade 4 0,05 0,08 0,015 0,50 0,40 Balance

Introduction

28

Table 1-3 Mechanical properties of various c.p. Titanium grades corresponding to ASTM F67-06

Yield strength, 0,2% min. (MPa)

Tensile strength min. (MPa)

Elongation in 4D min. (%)

Ti grade 1 170 240 24

Ti grade 2 275 345 20

Ti grade 3 380 450 18

Ti grade 4 483 550 15

The mechanical properties of titanium and its alloys are summarized in Table 1-4 and compared to those of stainless steel and Co-Cr based alloys, which are the major families of selected metals in the manufacture of dental, maxillofacial, orthopaedic, cardiac and cardiovascular implants. Titanium is good for orthopaedics due to its high specific strength and low elastic modulus compared to other metallic materials. Because the Young’s modulus is smaller, less stress shielding can be expected [20]. However, titanium has low wear and abrasion resistance because of its low hardness. Titanium is mainly used to fabricate endosseous dental implants. Although c.p. Ti is not strong enough material to be used in orthopaedic devices, the addition of aluminium (Al) and vanadium (V) creates an alloy (Ti6Al4V) with significantly greater strength. This improved strength, together with its low elastic modulus and excellent biocompatibility, makes Ti6Al4V a preferred orthopaedic material.

Table 1-4 Mechanical properties of c.p. Ti grade II and Ti-6Al-4V. Adapted from [3;21;22]

Material Density (kg/m3)

Young’s modulus, E (GPa)

Yield strength, 0.2%(MPa)

Tensile strength, min

(MPa) c.p. Ti grade II 4200 100-110 275 345

Ti-6Al-4V 4500 100-130 870-895 950

Stainless steel 316L 7800 200 280 580-650

Co-Cr alloys 8500 210−230 600-625 880-910

1.3.3.1.4. Corrosion properties

Materials implanted in vivo initially come in contact with extracellular body fluids, such as blood and interstitial fluids. The chloride ion concentration in blood plasma is sufficiently high to corrode metallic materials [23]. Body fluids also contain amino acids and proteins that tend to accelerate corrosion [24]. The pH of normal blood and interstitial fluid is 7,35-7,45. However, the pH decreases to about 5,2 in hard tissue after implantation and recovers to 7,4 within 2 weeks [25]. The pH of the body fluid around the materials surface may change based on the isoelectric point of the biomolecules, such as proteins. Whatever the cause, toxicity and allergy occur in vivo if metallic materials are

Chapter 1

29

corroded by the body fluids, leading to release of metal ions into the body fluid for a prolonged period of time and ions combining with biomolecules, such as proteins and enzymes. Thus, the corrosion resistance of metallic biomaterials is important. It is commonly accepted that titanium exhibits high stability and corrosion resistance in vitro and in vivo [13].

1.3.3.1.5. Biological properties

Titanium and titanium alloys typically do not suffer from significant corrosion in a biological environment. Titanium readily adsorbs proteins from biological fluids such as albumin [26;27], fibronectin [28], and fibrinogen [29] among others. Afterwards, titanium can also support cell growth and differentiation. Much work has been devoted to the investigation of cell and protein interactions with titanium surfaces.

In bones, titanium heals in close apposition to the mineralized tissues under the proper conditions. However, titanium and bones are generally separated by a thin soft-tissue layer [30] as a result of a weak foreign body reaction that prevents titanium to be in direct contact with bone. Different surface modification methods have been proposed to improve osseoconductivity of titanium.

1.3.3.2. Biomedical applications

Earlier applications of titanium in medical, surgical and dental devices were a result of post-World War II advances in manufacturing processes. Increased use of titanium and its alloys as biomaterials was basically due to their lower Young’s modulus, superior biocompatibility and better corrosion resistance when compared to stainless and cobalt-based alloys. The applications of titanium and its alloys can be classified according to their biomedical functionalities.

• Hard tissue replacements

Hard tissues are often damaged due to accidents, aging, and other causes. It is a common practice to surgically substitute the damaged hard tissues with artificial replacements. Depending on the regions in which the implants are inserted and the functions to be provided, the requirements of the substituting biomaterials are different.

Because of the mentioned desirable properties, titanium and titanium alloys are widely used as hard tissue replacements in artificial bones, joints, and dental implants. As a hard tissue replacement, the low elastic modulus of titanium and its alloys is generally viewed as a biomechanical advantage because the smaller elastic modulus can result in smaller stress shielding [10].

Introduction

30

• Cardiac and cardiovascular applications

Titanium and titanium alloys are used in cardiovascular applications such as prosthetic heart valves and artificial hearts among others, because of their unique properties. Artificial hearts entirely made of titanium have not been clinically successful mainly due to problems with blood-clotting occurring on the device surface. The use of titanium in cardiac applications relies mainly on its inertness. Therefore, it is necessary to improve the antithrombogenic properties of titanium implants for cardiovascular applications.

• Other applications

Besides artificial bones, joints replacements, and dental implants, titanium and titanium alloys are often used in osteosynthesis, such as bone fracture-fixation. It is a method of treating the bone fracture by surgical means. Typical implants for osteosynthesis include bone screws, bone plates, maxillofacial implants, etc.

1.3.3.3. Surface modification of titanium for biomedical applications

Numerous surface modifications have been developed and are currently used with the aim of enhancing clinical performance of titanium-based implants. The main properties of biomaterials, such as non-toxicity, corrosion resistance or controlled degradability, modulus of elasticity, and fatigue strength have long been recognized to be highly relevant in terms of the selection of the right biomaterials for a specific biomedical application. The events after implantation are focused mainly at the interactions between the biological environment and the implant material surface. The material surface plays an extremely important role in the response of the biological environment to the artificial medical devices.

A reason for conducting surface modification to titanium medical devices is that specific surface properties, different from those in the bulk, are often required. Proper surface modification techniques not only will retain the excellent bulk attributes of titanium, but also will improve specific surface properties required by different clinical applications.

Increasing the rate of bone formation along implant surfaces is an important issue, particularly for endosseous dental implants designed for immediate or early functional loading. Implant surface modifications such as acid etching in combination with grit blasting [31] and calcium phosphate coatings formed directly by plasma spraying [32], sputtering [33], wet chemical deposition (sol-gel) [34], or precipitation from solution on alkali- and heat-treated titanium substrates to form biomimetic coatings have been investigated as ways to increase osteoconductivity [35;36].

Increasing osseointegration by these surface design strategies has been proposed to be related to the altered implant topography that results from surface modification arising from enhanced interface zone fibrin clot retention during healing [37]. Alternatively, it

Chapter 1

31

may be due to the formation of surfaces that, as a result of their chemical constitution, allow preferential adhesion of bioactive molecules involved in osteoblasts adhesion, thereby accelerating bone formation. However the associated roughening resulting from current methods and the rapid alteration of prepared smooth surfaces due to heterogeneous coating have confused the effects of topography and surface chemistry among other surface properties.

As an example of the complexity and alternatives available, it has been estimated that dentists can choose from more that 1300 types of implants that vary in form, material, dimension, surface properties, and interface geometry [38]. It is therefore important to know whether certain surface modifications or particular materia ls will improve clinical results and provide the best available treatment. In this section the surface modification technologies [39;40] will be summarized. These methods are classified into three main categories: physical, chemical and biochemical treatments according to the mechanism of formation of the modified layer on the material surface.

1.3.3.3.1. Physical methods

Physical methods involve the external action by the application of physical forces to modify the surface characteristics. Common mechanical surface modification methods, such as machining [41], polishing [42;43], and grit-blasting [44;45], involve physical treatment, shaping, or removal of the materials surface. The main objective of mechanical modification is to obtain specific surface topographies and roughness.

In the case of metallic materials, machining usually produce deformations: crystalline grains disappear, surface properties are changed and, in general, surface hardness increases. Typical average roughness parameter (Ra) values vary from 0,3 up to 0,6 µm [40]. In order to obtain finer finishing degrees, surfaces can be exposed to smoothing process, by means of polishing. Due to the mechanical properties of titanium and its alloys, abrasive particle (generally carborundum and industrial diamond) could be embedded into the metallic matrix. Representative average roughness values (Ra) of smoothed surfaces vary from 0,1 up to 1 µm [40].

Blasting process consists of abrasive particles forced to impact against the surface by compressed air. In addition to increasing surface roughness, surface blasting could increase the surface reactivity of the metal and blasting particles can induce abrasive pollution on the implant surface. Average roughness obtained by blasting treatments are in a wide range of values: from 0,5-1,5 [41] up to 2-6 µm [44], depending on the particle size and the applied pressure.

Introduction

32

1.3.3.3.2. Chemical methods

Chemical methods are based on chemical reactions occurring at the interface between titanium and a solution which will produce modifications in the chemical composition, structure, and/or stoichiometry of native materials surface. Among the available methods, the most widely used will be presented.

a) Acid and alkaline etching

Chemical etching imply the immersion of the metallic samples into different acid (e.g. HCl, H2SO4, HNO3 and HF) [40;46;47] or alkaline (e.g. NaOH) [6;48] chemical solutions. Depending on the chemical solution, titanium and titanium alloys can exhibit different reactivity. Thus, it is possible to induce the formation of a protecting layer by passivation, the dissolution and the flattening of surface defects.

Since titanium and its alloys normally react with oxygen, only a few acids, presented above, are able to react with the oxide. The alkaline attack improves the formation of –OH group on the TiO2 surface due to the high pH solution. Moreover titanate salts form and dissolve so that the surface is progressively covered by a hydrogel that can locally precipitate [35]. The underlying metal topography is then modified by either acid or alkali etching. When the base attack is performed an elevated temperatures, the chemical reaction kinetic is accelerated and sub-micrometric defects are normally produced.

Nevertheless, chemical treatments might reduce the mechanical properties of titanium. For instance, acid-etching can lead to hydrogen embrittlement of the titanium, creating micro cracks on its surface that could reduce the fatigue resistance of the implants [49].

b) Electrochemical methods

The electrochemical processes are performed connecting the metallic device to the positive pole of an electrical circuit and immersing the entire device into an electrolyte solution which can contain ionic substances or oxidants. This methodology allows incorporating some ions onto the material surface, also the possibility of changing the surface finishing. One method that can be identified is anodization which is obtained by using strong acids (e.g. H2SO4, H3PO4, HNO3, and HF) as electrolyte solution. The result of the anodization is the thickening of the oxide layer from the usual 5-10 nm of atmospheric oxidation up to 40µm on titanium [40;50] and modifications in the microstructure and the crystallinity of the titanium oxide layer are also produced [51].

c) Deposition methods

Some processes are based on the deposition of foreign chemical substances on the surface under treatment. Metallic compounds (e.g. Ti), oxides and ceramics (e.g. HA) are applied by being volatilised due to the exposure to an energy supplying source, and then transported to the surface. Concerning to that, HA, which is considered one of the most

Chapter 1

33

effective bioactive materials, is usually applied by means of magnetron sputtering techniques [33]. Titanium plasma spraying method and HA coating has been extensively studied [52;53]. Both surfaces have defect dimensions but HA-coating modifies surface composition [52]. However, partcile realeasing has sometimes been found in the adjacent bone to these implants [54].

1.3.3.3.3. Biochemical methods

The biochemical methods allow the possibility of adding and bonding specific biomolecules at the implants surface. These treatments are aimed at controlling and guiding the complex sequence of biochemical phenomena that take place at the interface between implanted devices and biological tissues, such as osseointegration. But the biomechanical performances of an implanted device largely depend on the properties of its surface, in terms of both chemical composition and roughness [40;55;56].

The anchored or adsorbed biomolecules are normally present on cell membrane and extra-cellular matrix [57;58]. Among other proteins, many studies have been focused on the family of transforming growth factors beta (TGF-β) and bone morphogenetic proteins (BMPs). However, the use of a protein by itself is not very reproducible since it is going to have a low chemical stability and solubility into the biological environment. As an alternative, smaller biologically-active sequences, peptides, which are part of the total aminoacid sequence of a protein have been isolated/synthesized and attached to the desired substrate. One of the most investigated peptides is the Arg-Gly-Asp (RGD) amino acidic sequence, which is known to be the minimal cell-recognisable sequence in many adhesive proteins [59].

Basically there are three methods available for the biochemical treatment of a metallic surface: a) physico-chemical adsorption; b) covalent binding of a biomolecule on the surface; and c) peptide inclusion into a carrier material.

Adsorption of an active molecule to the surface [60] consists of an immersion of the sample into a bioactive peptide containing solution. But this method does not allow a controlled deposition of the peptide. This is important in controlling the interactions with biological tissues. Also, adsorbed molecules can be displaced and diffused from the site. Covalent attachment uses the chemical functionalities of the material surface to covalently bind the bioactive molecule. For that goal, different strategies are focused on improving the number of -OH reactive groups [60]. This technique rarely controls surface density of the peptides, which can affect biological response [61]. The advantage of peptide inclusion to carrier materials is the control of the amount of bioactive peptide introduced and used to coat the implant surface. Carriers, mostly polymers, can be either simply impregnated with the active biomolecules or covalently bound to the carrier structure.

Introduction

34

1.4. Bone-Implant interface

1.4.1. Bone

Bone is the main constituent of the body system and differs from other connective tissues in rigidity and hardness. These characteristics of bone result from the mineral phase constituted by hydroxyapatite crystals, (Ca10(PO4)6(OH)2), permeating the matrix, which consists 90% of collagen fibers oriented in a preferential direction and about 10% of various non-collagenous proteins [62].

The main bone functions are the following:

• Biological functions

1. Mineral storage: Bones act as reserves of minerals important for the body, such as calcium, sodium, phosphorus, hydrogen and magnesium.

2. Haematopoiesis: The marrow, located within the medullary cavity of long bones and the interstices of cancellous bone, produces blood cells.

• Mechanical functions

3. Protection: Bones can serve to protect internal organs, such as the skull protecting the brain or the ribs protecting the heart and lungs.

4. Shape: Bones provide a frame to keep the body supported.

5. Movement: Bones, skeletal muscles, tendons, ligaments and joints function together to generate and transfer forces so that individual body parts or the whole body can be manipulated obtaining movement of the body.

1.4.1.1. Bone structure

1.4.1.1.1. Macroscopic structures

Anatomically, two types of bones can be distinguished in the skeleton: flat bones (skull bones, scapula, mandible …) and long bones (tibia, femur, humerus, …). Long bones serve as the classical model for the macroscopic structure of bone. A typical adult long bone consists of a central cylindrical part, diaphysis, and two wider and rounded ends, the epiphysis. Conical regions called the metaphysis connect the diaphysis with each epiphysis.

Chapter 1

35

The diaphysis is composed mainly of cortical bone, while the epiphysis and metaphysis contain mostly cancellous bone with a thin shell of cortical bone.

The main macroscopic structures of bone are [62;63]:

• Cortical, compact or dense bone

Cortical bone is a dense, solid mass with microscopic channels. Approximately 80% of the skeletal mass in the adult human skeleton is cortical bone, which forms the outer wall of all bones and is largely responsible for the supportive and protective function of the skeleton.

• Cancellous or trabecular bone

Filling the interior of the organ, the remaining 20% of the bone mass is cancellous bone, an open cell porous network (also called cancellous or spongy bone) which is comprised of a network of rod- and plate-like elements that make the overall organ lighter and allowing room for blood vessels and marrow.

• Periosteum

The outer surface of most bone is covered by the periosteum, a sheet of fibrous connective tissue and an inner cellular layer of undifferentiated cells. The periosteum has the potential to form bone during growth and fracture healing.

1.4.1.1.2. Microscopic structures

Mammalian cortical or cancellous tissue bone is of two main types, woven or lamellar.

Bone tissue in the developing embryo is the woven type. Woven bone consists of a matrix of interwoven coarse collagen fibers with osteocytes distributed more or less randomly, is less organized and shorter-lived than lamellar bone [62]. It is provisional material that is eventually resorbed and replaced by lamellar bone. In human long bones, woven bone is replaced by lamellar bone at age 2 and 3 years.

Lamellar bone is built up of layers (lamellae) of approximately 3 to 7µm thick and contains parallel and well-defined fine collagen fibers [63]. The lamellae can be parallel to each other if deposited along a flat surface, or concentric if deposited surrounding a longitudinally vascular channel that together form a structural cone, the osteon or Haversian system (Figure 1-2).

Introduction

36

Figure 1-2 Hierarchical structural organization of bone: (a) cortical and cancellous bone; (b) osteons with Haversian systems; (c) lamellae; (d) collagen fiber assemblies of collagen fibrils; (e) bone mineral crystals,

collagen molecules, and non-collagenous proteins [64]

1.4.1.2. Bone cells properties and functions

Bone consists both of cells and intercellular matrix. Twenty percent of bone is water. The dry weight consists of approximately 30 to 35% organic matrix and 65 to 70% mineral [62;65] (Figure 1-3).

BONE

Organic Matrix (30-35%)

Cells (2%)

Osteoblasts

Osteocytes

Osteoclasts

Matrix (98%) Collagen related proteins (90%)

Non-collagenous proteins (10%)

Mineral (65-70%)

Hydroxyapatite (98%)

Magnesium, sodium, potassium and fluoride

Figure 1-3 Bone composition. Adapted from [62]

The organic matter is concentrated in the bone matrix, which consists mainly of 90% collagen fibres and other non-collageneous proteins. Collagen is the most abundant protein in the body and collagen fibers constitute the shape-forming structural

Chapter 1

37

framework of bone in which the hydroxyapatite is inserted and provides rigidity to the tissue (Table 1-5). The role of the non-collagenous proteins is not yet well understood but is thought to play an important role in the calcification process and the fixation of the hydroxyapatite crystals to the collagen [62]. The most abundant ones are osteonectin, osteocalcin, and bone sialoprotein, Table 1-5.

Table 1-5 Collagen-related and non-collagenous proteins and their functions. Adapted from [62]

Collagen related proteins

Protein Function

Type I (COL I) Serve as scaffolding; may be nucleator of mineral deposition

Non-collagenous proteins

Protein Function

Alkaline phosphatase (ALP)

A phosphotransferase potential Ca2+ carrier; hydrolyzes inhibitors of mineral deposition such as pyrophosphates

Osteonectin May mediate deposition of hydroxyapatite; binds to growth factors; may influence cell-cycle antiadhesive protein

Osteocalcin (bone gla protein, BGP)

May regulate activity of osteoclasts and their precursors, may mark turning point between bone resorption and formation; regulate mineral maturation

Fibronectin Binds to cells, fibrin, heparin, gelatin, collagen

Vitronectin Cell attachment protein, binds to collagen , plasminogin and plasminogen activator inhibitor, and to heparin

Osteopontin Binds to cell, inhibits mineralization and nitric oxide synthase; may regulate proliferation, tissue repair, and initiate mineralization

Bone sialoprotein Binds to cell, binds Ca2+; may initiate mineralization

Albumin Transports proteins; inhibits hydroxyapatite crystal growth

Bone is a self-repairing structural material; it is capable of adapting its mass, shape and properties to the changes in mechanical and physiological requirements. The unique physical and biological properties of bone are the direct results of actions by its constituent cells. There are three major types of bone cells: osteoblasts, osteoclasts, and osteocytes. Osteoblasts and osteocytes and bone lining cells are closely related and represent different stages in the maturation of a single type [66].

1.4.1.2.1. Osteoblasts

Osteoblasts are bone forming cells, this process occurs in two stages. First, an organic and unmineralized bone matrix called osteoid is laid down, after which mineral salts are precipitated throughout that matrix from ions, principally calcium and phosphate, in the

Introduction

38

extracellular fluid [63]. The osteoid serves as a template whose organization determines that of the subsequently formed mineral phase. Osteoblasts synthesize and secrete the major structural elements of osteoid, regulate their assembly into an organized matrix, and finally facilitate mineral deposition.

Osteoblasts are derived from mesenchymal stem cells in the mesenchyme or from the marrow stroma progenitors. They synthesize and secrete the unmineralized bone matrix which consists of 90% collagen and about 10% non-collagenous proteins (Figure 1-3). Osteoblasts do not synthesize all bone matrix proteins uniformly. For example, proteins like fibronectin and osteonectin are expressed early in osteoblastic cell cultures undergoing mineralization in vitro, while osteocalcin is expressed only after the development of an established matrix [62].

Control of bone mineral crystal growth and proliferation is governed by the spatial limitation of the collagen fibrils, as well as by the absorption of matrix proteins.

Once bone matrix constituents are secreted, they proceed to form an organized matrix. Osteoblasts, particularly at the early stages of differentiation, similarly secrete fibronectin. Moreover, disrupting the binding of fibronectin to integrines with antibodies results in production of a disorganized matrix, diminished mineralization, and impaired differentiation of osteoblasts.

How osteoblasts regulate the deposition of mineral within osteoid is still not fully understood. In addition to producing an appropriately organized matrix, osteoblasts express high levels of alkaline phosphatase (ALP). Moreover, ALP expression rises very early in osteoblastic differentiation, although its activity remains high in mature osteoblasts [62].

1.4.1.2.2. Osteocytes and Bone-lining cells

Osteocytes are former osteoblasts that became trapped within the mineralized bone matrix. They are the most abundant cell type in mature bone [62]. Osteocytes and lining cells are believed to comprise a sensory network that controls mechanical load and tissue damage. The osteocytes may: (1) stabilize bone mineral by maintaining an appropriate local ionic milieu; (2) detect microdamage; and (3) respond to the amount and distribution of strain within bone tissue. Prior to bone resorption, osteocytes near the lesions underwent changes consistent with apoptosis, programmed cell death.

1.4.1.2.3. Osteoclasts

Osteoclasts, the bone-resorbing cells, are multinucleated giant cells that contain from 1 to more than 20 nuclei and range in diameter from 20 to over 100µm [62]. They are

Chapter 1

39

specialized to resorb bone [62]. The surface of osteoclasts adjacent to bone surface has a ruffled border, an area of extensive membrane enfolding which secretes products leading to bone destruction. Resorption, like formation, occurs in two steps essentially simultaneous: dissolution of mineral and enzymatic digestion of organic macromolecules [63].

Osteoclasts dissolve both the mineral phase and organic component of the matrix of bone by secreting H+ ions into the extracellular resorption space, lowering the pH within this limited environment down to around 3,5. The H+ ions neutralize negative charges on hydroxyl ions and PO4 3- ions in the highly basic bone mineral hydroxyapatite, whose formal structure is (Ca10(PO4)6(OH)2); this leads to the release of Ca2+ and phosphate as HPO42- and H2PO4 - [62].

The development of osteoblast and osteoclasts are inseparably, moreover, osteoblast differentiation is a pre-requirement for osteoclast development.

1.4.1.3. Mechanical properties of bone

Bone has a varied arrangement of material structures at many length scales which work in concert to perform diverse mechanical, biological and chemical functions; such as structural support, protection and storage of healing cells, and mineral ion homeostasis.

Mechanical properties of bone vary at different structural levels. These levels and structures have been explained in section 1.4.1.1. This hierarchically organized structure has an irregular, but optimized, arrangement and orientation of the components, making the material of bone heterogeneous and anisotropic (Figure 1-2).

The mechanical properties of cortical bone have been well documented. They can be measured via traditional testing techniques such as: uniaxial compressive or tensile testing, or three or four-point bending. They can also be tested using ultrasound techniques or micro or nanoindentation. Cortical bone exhibits high degree of anisotropy and values of mechanical properties vary between animal species, bone location and testing conditions, age and disease. Testing conditions may vary between the different treatment of specimens (wet or dry samples), the different age and health of donors and the differences between bones and sites in the bones (longitudinal or transverse direction). That’s the reason why, in most cases, it can not be specified a unique value for strength and elastic modulus while a range of values is detected [22;64;67].

Table 1-6: Mechanical properties of human cortical and cancellous bone. Adapted from [22;64;67;68]

Cortical bone Strength range (MPa) Elastic modulus range (GPa)

Compression test 80-150 17-20

Tensile test 131-224 7-18

Introduction

40

Cancellous bone Strength range (MPa) Elastic modulus range (GPa)

Compression test 0,1-30 0,01-2

Measuring the properties of cancellous bone is more complex compared to cortical bone. The complexity is due to the small dimensions of the individual trabeculae. It is speculated that differences in moduli between cortical and cancellous bone are entirely due to the bone density. Thus, as shown in Table 1-4, some authors find values of elastic modulus of cancellous bone as high as those for cortical bone [64].

As previously mentioned, biocompatibility was defined in 1987 as “the ability of a material to perform with an appropriate host response in a specific application” [4]. Where the two key expressions are: the appropriate host response and specific application which will be discussed in the followings subsections.

1.4.2. Events at the bone-implant interface

Biomaterials incorporated into medical devices are implanted into tissues and organs. As soon as a material is placed within the body it comes into contact with tissue fluids, the surface of the biomaterial is covered with a layer of proteins and other biomolecules. The kinetics and extent of this process will vary from material to material and it will be a dynamic phenomenon with adsorption and desorption processes continuously taking place. Because cells depend on specific proteins for achieving anchorage and get extracellular instructions, the composition of the adsorbed layer is a key mediator of cell behaviour. In this manner, the required proteins, correctly presented, i.e. in the appropriate conformation and orientation, can stimulate a constructive cell response, favouring wound repair and tissue integration. On the contrary, proteins in an unrecognizable state, for instance in an denatured state, may indicate a foreign material to be removed or isolated [69].

When cells approach to an implant material, they will not make direct contact with its surface. Rather, the molecular events at the material-tissue interface involve water molecules which can dissociate to OH- groups or bind to the surface by hydrogen bonds. Then, the rapid adsorption of proteins from blood (or serum) effectively translates the structure and composition of the foreign surface into a biological language. Cells will respond to protein signals contributing to the ultimate outcomes in both implantation and tissue culture situations (Figure 1-4). Once the cells reach the implant site, they start to examine the protein-covered surface, looking for activation factors and places to attach to [22]. Following to the attachment, a complex series of cell reactions is initiated, which will lead to either the desired wound healing or the foreign body reaction.

Chapter 1

41

Figure 1-4 Schematic illustration of the material-tissue interaction at different levels after implantation of a

medical implant [70]

Wound healing encompasses a cascade of events that our body translate into actions to resolve injury. Repairing tissue involves restoring both cells and their protein matrix. The protein matrix defines tissue structure and supports cellular adhesion and migration. Cells influence the composition and activity of the matrix, and vice versa. Generally, these events are grouped into three phases: 1) early inflammatory stage; 2) the proliferative/repair stage; and 3) the late remodeling/maturation stage [59;71;72].

The inflammatory phase begins at the time of injury and continues for a few days at the most. Inflammatory cells (macrophages, monocytes, lymphocytes, and polymorphonuclear cells) and fibroblasts infiltrate the bone, which results in the formation of granulation tissue, ingrowth of vascular tissue, and migration of mesenchymal cells. The primary nutrient and oxygen supply of this early process is provided by the exposed cancellous bone and muscle [72].

Introduction

42

In the repair phase, a collagen matrix is laid down while osteoid is secreted and subsequently mineralized, which leads to the formation of a soft callus around the repair site [65]. This occurs a few days to a few weeks post-injury.

Finally, fracture healing is completed during the remodelling phase in which the healing bone is restored to its original shape, structure and mechanical strength by continuously maturation of the matrix and cellular content [72]. Remodelling of bone occurs slowly over months to years and is facilitated by mechanical stress placed on the bone.

The foreign body reaction begins as wound healing, including accumulation of exudates at the site of injury, infiltration of inflammatory cells, and the formation of granulation tissue. However, the persistent presence of a biomedical implant or other foreign bodies inhibits full healing. Rather than the resorption and reconstruction that occurs in wound healing, the foreign body reaction is characterized by the formation of foreign body giant cells, encapsulation of the foreign object, and chronic inflammation. It is not clear what turns the balance towards encapsulation or regeneration but the direction taken is highly affected by (1) the in situ conditions, referring to the proliferative capacity of cells in the tissue receiving the implant; (2) the biomaterial properties, i.e. the size, shape, topography, chemistry, physics and mechanics of the implant; (3) the surgical technique; and (4) systemic factors, e.g. age, sex, general health status.

1.4.2.1. Material response

In addition to protein adsorption on the implant’s surface, significant changes also occur in the materials surface. One of the main changes described, is the oxidation of metallic implants both in vivo and in vitro assays [73]. Although metallic biomaterials are selected because of their stable oxide films, they still undergo electrochemical changes in the physiological environment. Another consequence of these events is the release of metal ions into surrounding tissues, [74], which can accumulate locally, but they also spread systemically. Trace metals are found in health, but they can also be toxic or cause hypersensitivity reactions depending on the quantity.

1.4.2.2. Host response

The host response to implants placed in bone involves a series of cell and matrix events, ideally culminating in intimate apposition of bone to biomaterial, i.e., “osseointegration”. Whenever a material is implanted into the tissues of the body, there has to be a degree of trauma associated with the insertion process. This will inevitably establish an acute inflammatory response, which is the body’s natural defence

Chapter 1

43

mechanism to any injury. The inflammation is totally desirable and helpful since it is the precursor to the second phase of the response, which is that of tissue repair.

While this intimate contact occurs, unfavourable conditions will disrupt the newly forming tissue, leading to formation of a fibrous capsule [22;71]. On the other hand, if the material is able to react with the tissues, chemically, mechanically or any other way, it will act as a continuous stimulus to inflammation. In the context of the definition of biocompatibility, it is important that the interaction between the material and the tissues is one which leads to an acceptable balance between inflammatory and repair.

Morphological studies have revealed the heterogeneity of the bone-implant interface. One feature often reported is an interfacial zone, rich in non-collagenous proteins. This zone forms with variations on thickness and appearance, regardless of the type of biomaterial implanted, including cpTi, stainless steel and hydroxiapatite [75;76]. The absence or low quantity of serum proteins may indicate a selective accumulation/deposition of molecules at the interface. Even though several biomolecules have been identified at the interface, they are likely not the only ones present. These other biomolecules may have essential role in directing the bone response to the implant, and further work is needed to identify them and determine their functions at the interface.

The nature of the host response will vary very considerably from one host to another and from one location to another within a particular host. Sometimes, it is forgotten that host variables are as important as the material variables in the determination of biocompatibility.

1.4.2.3. In vitro studies

Bone cell culture models are increasingly employed to study bone-biomaterial interactions. Most of the cultures have utilized osteoblastic cells [77]. Primary and passaged cells from several species and anatomical locations have been used, as well as several osteosarcoma, clonal, and immortalized cell lines [71]. Substrate-dependent differences have been reported, however, the variety of models used makes it difficult to determine general conclusions. Nonetheless, in vitro models have the potential to help elucidate events at the bone-implant interface by providing morphological, biochemical, and molecular information regarding osteoblastic development and synthesis of matrix at the interface with various biomaterials. Moreover, the information obtained reflects in vivo events [71].

Introduction

44

1.5. Biomaterial-cell interactions through adsorbed proteins

An appropriate cellular response to implanted surfaces is essential for tissue regeneration and integration. It is well described that implanted materials are immediately coated with proteins and other biomolecules from blood and interstitial fluids, and it is through this adsorbed layer that cells interrogate foreign surfaces. Hence, it is the adsorbed proteins, rather than the surface itself, to which cells initially react to.

To achieve osseointegration of titanium implant the recruitment on the surface of a sufficient number of osteoblasts cells is required, and their subsequent differentiation and biosynthetic activity is needed. Consequently, proteins encountered by cells must support or actively promote those activities, including the ones happening after cell attachment and proliferation. In other words, a selective protein adsorption, which enhances osteoblasts differentiation on the implant surface, will be desired and pursued.

1.5.1. Role of adhesion proteins in cell behaviour

1.5.1.1. Cell adhesion

Cell adhesion is a fundamental process directly involved in cell growth, migration and cell differentiation. Cells in their natural environment are anchored by discrete attachments to proteins in the extracellular matrix, ECM. Similarly, cell attachment to culture surfaces in vitro is usually mediated by adhesion proteins contained in serum-supplemented culture medium. Those can serve as ligands for specific receptors on the cell surface by providing points of anchorage as well as inducing internal signalling pathways leading to cell progression and differentiation [69]

Osteoblastic cells (and various other cell types) cultured in vitro have been shown to depend primarily on adsorbed vitronectin, fibronectin, and osteopontin, [78], for initial adhesion and spreading on materials, including tissue culture polystyrene, titanium, stainless steel, and hydroxyapatite [69;79;80]. Vitronectin is frequently dominant over fibronectin in cell adhesion processes, because of its greater ability to adsorb competitively over adhesion-inhibiting proteins [81-83]. Thus, the ability of materials to adsorb such proteins, in an active state from serum, determines their ability to support cell adhesion and spreading. Osteoblasts, in culture, show substantially improved adhesion if the cultured surface is pre-coated, or the medium supplemented, with fibronectin or vitronectin [84;85].

Chapter 1

45

The primary interaction between cells and adhesion proteins occurs via integrins (heterodimeric receptors in the cell membrane), as demonstrated by the decrease in cell attachment observed when antibodies are introduced to prevent these interactions [85;86]. Integrins are also involved in intracellular signalling and, thus, a diverse range of cell functions. The functions of integrins in osteoblast responses to biomaterials have been reviewed by Siebers et al [87]. This work determined the relevant surface characteristics of the substrate such as, composition and topography, and their effects on integrins. Different conclusions were reported depending on the working substrate and the studied cell line. As a result, osteoblasts express the integrin subunits α2, α3, α4, α6, αv, β1, and β3 in titanium alloy surfaces; while the subunits α3, α6 and β3 where not always expressed on cobalt-chrome-molybdenum (CoCrMo) samples [86]. Concerning on topography effects, MG63 cells cultured on titanium substrates with different roughness showed an increase in the expression of integrin subunits α2, α3, α5, β1, and β3 on rougher titanium compared to polished titanium [88].

Addition to the initial states of the surface and the cells, the interface between them are critical, because it is a dynamic system. Cells may alter their environments and adhesion mechanisms by secreting fibronectin [89] or by manipulating the extracellular matrix. Furthermore, cells adapt to their environment [69].

Many of the adhesion proteins associated with osteoblasts in vitro are similarly observed in vivo. Vitronectin and/or fibronectin have been detected among the proteins adsorbed from whole blood and plasma in vitro, and to implanted surfaces [90]. However, there will be differences in the molecular and cellular composition of implant sites and cell culture interfaces.

1.5.1.2. Cell spreading and migration

Cell spreading is a process related to adhesion and usually involves similar extracellular proteins [91]. Cell migration requires a dynamic interaction between the cell, its substrate and its cytoskeleton. Migration rates and cell adhesion strength on various surfaces often show opposite trends. Apparently, maximal cell motility requires adhesion strength sufficient to maintain substratum contact [92]. Although adhesion is clearly critical in osteoblast development, excessive adhesion strength may inhibit subsequent cell activities. Some antiadhesive extracellular matrix proteins such as tenascin, thrombospondin laminin, muscin and proteoglycans also play a roles in cell migration [93].

Introduction

46

1.5.1.3. Cell proliferation and differentiation

Cell adhesion and morphology influence subsequent activity such as proliferation and differentiation [69]. Interactions of ECM molecules with cells can also modulate cell proliferation.

With respect to osteoblasts and adsorbed proteins, fibronectin (and its receptors) appears to become progressively more important as the cells differentiate [94-96]. Osteoblasts produce their own fibronectin which may indicate a proper environment for cells. Moreover, cells can reorganize the fibronectin bounded to the surface for fibroblast proliferation [97]. The fibronectin α5β1 integrin subunit is predominant in the differentiation of mesenchymal stem cells toward osteoblasts [94;98]. The vitronectin binding αvβ3 appears to be involved in mineralization, too.

1.5.2. Cell response to specific surface characteristics through adsorbed

proteins

Surface characterization and controlled surface preparation are very important to improve our understanding of the material-tissue interface.

The key determinants of protein-mediated cell responses are the adsorption profile (quantification and identification of the adsorbed proteins) and protein bioactivity (functionality, determined by protein conformation and orientation). The surface properties of the implant may influence the interface evolution by determining the nature of the water layer, which in turn determines the protein-surface and cell surface interactions. These interactions eventually condition the ultimate success or failure of the implant.

Many parameters describe a surface (Table 1-7). The more parameters measured, the better the description of the surface. A complete characterization requires the use of many techniques to compile all the information needed. Unfortunately, the most important parameters for the understanding of the biological responses to surfaces have not been identified so far. Studies have been published on the importance of roughness, wettability, surface energy, surface mobility, chemical composition, electrical charge, crystallinity, and heterogeneity to biological reaction. Since it is not known which surface factors are predominant in each situation, the controlling variable or variables must be independently ascertained.

Chapter 1

47

Table 1-7 Properties and required information to describe surfaces. Adapted from [99]

Property Type of physico-chemical information

Chemical composition Atomic composition

Chemical state of elements

Structure

Order

Disorder

Crystallinity

Inclusions

Vacancies Crystal boundaries

Morphology Texture

Roughness Form

2D and 3D form of surface features Specific surface area

Porosity

Surface energy Wettability Surface energy

Adsorptivity

Electric properties Surface potential Surface charges

Mechanics Elasticity/plasticity Tensile and compressive forces in surface layers

1.5.2.1. Surface chemistry

Cells can discriminate between markedly different chemistries [100]. They are sensitive to surface chemistry by affecting cell attachment and proliferation. Incorporation of ions and other substances leading to changes of surface chemistry can change the orientation of binding proteins, and ultimately, the binding of cells to the material.

1.5.2.2. Surface topography and roughness

Cells are sensitive to the physical characteristics of the materials they interact with. One of the most studied characteristics of materials is the three-dimensional morphology of the substrate. In orthopaedics and oral/maxillofacial applications, surface roughness appears to promote osseointegration. When examining the ability of osteoblast-like cells to attach to c.p. Ti surfaces, a higher percentage of cells attached to rougher surfaces was determined [100]. Cells use the morphology of the substrate for orientation and migration, too.

A comparison of various studies [101] revealed a few consistent trends in the effects of surface topography on initial cell adhesion. Interpretation and comparison of results is frequently complicated because of differences in methods and parameters used to

Introduction

48

characterize the surface topography as well as the magnitude of surface roughness [87]. Furthermore, methods used to produce surface textures -sandblasting, grinding, and plasma-spraying- on titanium may alter surface chemistry and/or physicochemical properties [44;53;88;102]. The type of cells used also influences these studies.

Surface topography is frequently seen to have a significant effect on both proliferation and differentiation of osteoblastic cells. The most commonly observed trend is that an increase in roughness and disorder of the surfaces produces and increase in cell differentiation and/or extracellular matrix synthesis. This is accompanied with a reduction in cell proliferation [1;102-104]. However, once again, these trends are not consistent [8].

Although cellular responses to differences in surface chemistry and wettability may be attributed to the composition and the bioactivity of the adsorbed protein layer; the involvement of surface topography in directing cell reactions is less clear. There is some evidence suggesting that changes in the nature of the focal adhesions used to anchor the cell to its substrate. This may play a relevant role since the morphology of osteoblastic cells varies considerably as surface roughness increases [104;105].

Other studies measured an increase in fibronectin adsorption (from single-protein buffer) on rougher surfaces [43;101], but when François et al. [43] corrected their results for the increased in surface area, this trend was reversed.

Apart from increasing the surface area, topographic features may also create confined spaces which may interfere with wetting on hydrophobic surfaces [43] or restrict protein exchange between the surface and the solution. This suggests that rough surfaces may affect cell behaviour by limiting cell-substratum contact.

It has been suggested that topographic effects on protein adsorption relates to an increase in surface energy with roughness [101;106]. However, the reported results lead to a cautious comparison of contact angles and related parameters when measured on differently textured/structured surfaces.

1.5.2.3. Surface wettability and surface energy

It is widely believed that water is the first molecule to contact biomaterials in any clinical application. This is because water is the major component in biological fluids as well as a small and very mobile molecule. Consequently, the behaviour of water near surfaces and the role of water in biology are very important [107].

One of the most important and extensively studied elements of protein adsorption is its variation with surface wettability (hydrophilia or hydrophobia). Water does not readily interact with non-polar solutes and surfaces because water is a polar solvent. Hence,

Chapter 1

49

contact between water and these hydrophobic surfaces increases self-association (by hydrogen bonding) within the neighbouring water molecules [108].

For a protein to adsorb, both adsorbate and surface must dehydrate at least partially [69]. This is thermodynamically favourable for similarly hydrophobic sorbents and adsorbates, as it increases the entropy of water [109;110]. While a hydrophilic surface is unfavourable due to the displacement of water molecules which presents a substantial energy barrier to protein adsorption [108]. However, adsorption does occur on hydrophilic materials [111], meaning that charge interactions and protein conformation changes provide the necessary driving force to proteins to adsorb. Presumably because of a greater number of possible adsorption-promoting interactions, hydrophobic surfaces are usually reported to adsorb more proteins than hydrophilic surfaces [112].

The movement of atoms and molecules near the surface in response to the outside environment is often highly significant. In response to a hydrophobic environment (e.g., air), more hydrophobic (lower energy) components may migrate to the surface of a material by reducing interfacial energy. Responding to an aqueous environment, the surface may reverse its structure and point polar (hydrophilic) groups outward to interact with the polar water molecules. Again, energy minimization drives this process [113].

One mechanism of enhanced cell response on hydrophilic surfaces is the difference in the array of proteins adsorbed to hydrophilic and hydrophobic materials from serum, plasma, or whole blood [114;115]. In particular, the ability of vitronectin to successfully compete for adsorption sites is clearly apparent on hydrophilic surfaces [82;83;116].

The functionality of vitronectin does not vary with the wettability of the surface. In contrast, fibronectin adsorbed to hydrophobic surfaces shows a marked reduction in cell-adhesive function [89]. The ability of fibronectin to retain its functionality on hydrophilic surfaces is therefore considered another contributor to improved cell responses.

It is worth mentioning that cleaning and sterilization methods, which can induce important changes in contamination, surface chemistry, surface energy and ion release [117], can also affect cellular activity [8;118].

1.5.2.4. Surface charge/polarity

Interfaces are charged in aqueous solution, and cell membranes carry negative charge which means that electrostatic interactions will play a role in the biological response to implant materials. Generally, opposite charges will be mutually attracted. The situation is much more complicated in aqueous solutions, though. Surface charges are shielded by water hydration and counterbalanced by small ions [109]. This creates an electrical double layer at charged surfaces. The charge on a surface is minimized near the pH

Introduction

50

indicated by its isoelectric point (IEP). When the working pH is minor than the IEP, the surface will be positively charged. Conversely, when the working pH is higher than the IEP the surface will be negatively charged.

Davies et al. [75] pointed that these interactions alone are not sufficient to explain the observed differences in adhesion and migration of cells on both positively and negatively charged surfaces.

The mode of cell adhesion is distinct for positive and negative charges. Cell membranes, that carry a negative charge, adhere closely to positively charged surfaces, whereas contact occurs only at distinct points on negatively or neutral charged surfaces [109;119]. These differences can be attributed primarily to electrostatic attraction and repulsion, respectively, but it is also demonstrated quantitative and qualitative patterns in protein adsorption and functionality [120].

Fibronectin showed affinity for positively charged surfaces [112]. However MacDonald et al. [112] suggested that hydrophobicity was more dominant parameter for adsorption than charge. Other studies [121] got opposite conclusion using immunofluorescence techniques, which therefore may be susceptible to conformation changes.

1.5.3. Responses to biomaterials in vivo

In vitro models are, in general, static and do not take into account the dynamics at the implant/biological components in vivo. Most of the information presented is based on in vitro studies. However, it is important to address how relationships observed in vitro translate into in vivo applications.

Factors such as the anatomical position of the material, load, fatigue of the material, fretting of the material, changes in the material in response to the biological fluids, changes in the surrounding tissue, and age of the patient are all important considerations in vivo, which are either irrelevant or uncontrollable in vitro [122;123]. Despite significant deviations in vivo from cell culture models, the relationship between surface characteristics and protein adsorption is likely to be significant throughout the tissue regeneration process. Conversely, in vitro experiments provide useful information in isolating specific components of the biological response to materials.

Titanium and its alloys, mainly Ti6Al4V, have become the metals of choice for many biomedical applications such as hip and knee prostheses [124], dental [125;126], cardiac and cardiovascular [127] applications among others [20]. The clinical results presented in this section will be focused mainly in dental implant applications since implant dentistry is one of the most successful treatment modalities in medicine and dentistry [122].

To evaluate the different titanium surfaces in in vivo studies, different parameters are considered [126]:

Chapter 1

51

a) Histological analysis to evaluate qualitatively the healing response of bone following implant placement

b) Histomorphometric analysis to evaluate the percentage of bone-implant contact (BIC)

c) Biomechanical testing by measuring removal torque values of osseointegrated threaded dental implants

Most of the bibliographic studies are focused on the comparison between rough and smooth surfaces, from one side; and on the chemical composition of the surface by the use of bioceramic coatings from the other side. A novel and less studied in vivo approach is based on the incorporation of biological molecules at the implant surface.

1.5.3.1. Topographic/Chemical surface modifications

The most common techniques providing roughness to surfaces are titanium plasma spray-coating blasting and chemical etching. Roughness was classified by Wenneberg et al. [128] who suggested a classification considering the terms minimally rough (0,5 to 1 µm), intermediately rough (1 to 2 µm), and rough (2 to 3 µm). However the majority of literature reports based on the average surface roughness, Ra, surfaces with a Ra = 1 µm are considered smooth, and those with Ra > 1 µm are described as rough. Machined c.p. Ti is a smooth surface with an Ra value of 0,53 to 0,96 µm depending on the manufacture protocols [122].

Several studies have demonstrated that modification of roughness, especially if Ra values range from 0,5 to 2 µm, tends to not only increase the bone-implant contact but also the biomechanical interaction of the interface between them at early implantation times [6]. Those range of roughness increase the functional area 1,5 to 2 times the initial surface area. A higher increase of the functional area may not be favourable for bone growth and apposition because it also becomes a factor when there is an exposure of the implant surface to the oral fluids and bacteria [122]. Bacterial contamination could cause loss of metallic substrate and bone attachment.

In vivo studies have shown the importance of surface roughness on osseointegration improvement and acceleration [6;122;129]. Rough surfaces have shown torque to failure values significantly higher than implants with machined profiles [6;122;129].

1.5.3.2. Bioactive ceramic coating

Among all engineering based surface modifications for dental and orthopaedic implants, the addition of calcium- and phosphorous-based materials as coatings have received significant attention [130-132]. This interest is in part, because these elements are the

Introduction

52

same basic components of natural bone and coatings can be applied along the implant surfaces by various industrial processing methods [122;130].

Substantially enhanced in vivo bone-to-bioceramic bonding and bone-to-implant contact magnitudes have been observed at early implantation times for plasma sprayed hydroxyapatite (PSHA) coated implants [122;129;131]. However, this type of implant has fallen out of favour in dental practice as studies have shown that coatings do not uniformly dissolve/degrade after long periods in function. Moreover, it is also noted compromised coating and bone-coating interface mechanical properties [129;131].

There are other techniques for incorporation of calcium and phosphorous at the implant surface such as ion beam assisted deposition (IBAD) and discrete crystalline deposition (DCD) method. Animal studies at early implantation times including sputtering-coated [133;134] and IBAD-coated [135] Ca- and P-based thin films on titanium implants have demonstrated a higher biomechanical fixation [133-135], bioactivity [135], and BIC [133;134] when compared with non-coated implants. A potential drawback of the novel processing techniques for thin film deposition is its relatively high cost for large scale production.

DCD method incorporates nanometer-size crystals of CaP onto a previously treated surface, i.e. grit-blasted [122]. Studies in rat models have shown that the surface chemistry and topography of the DCD treatment was beneficial when compared with a dual acid-etched surface [136]. Moreover, studies in humans with DCD surfaces have shown a higher BIC after 2 months in vivo [137].

Another attempt was made to coat metal implants with layers of calcium phosphate under more physiological conditions of temperature and pH. The biomimetic coating technique involves the nucleation and growth of bone-like crystals upon pre-treated substrate by immersing this in a supersaturated solution of calcium phosphate under physiological conditions developed by Kokubo [35]. An in vivo study has shown an accelerated osseointegration compared to grit-blasted, acid etched and as machined implants. These implants developed in vivo onto their surfaces a layer of a calcium phosphate, which is probably apatite [2].

1.5.3.3. Biological molecules surface functionalisation

Another promising field is the incorporation of biological factors to implant surfaces. Liu et al. [138] obtained implants coated with a crystalline latticework of the inorganic components into which BMP-2 was incorporated which is suggested to stimulate the alkaline phosphatase activity of bone-marrow stromal cells. In vivo studies on rats for 5 weeks showed that bone tissue was apparent two weeks after implantation and the BIC parameter increased significantly. Stadlinger et al. [139] evaluated different coatings of implants by components of the extracellular matrix, such as RGD and collagen, on the

Chapter 1

53

osseointegration of dental implants over a period of 6 months on minipigs. The study showed that the influence of biological surface coatings seemed to be especially helpful during the early phase of bone formation promoting faster and better implant integration. The incorporation of RGD molecules to implant dental surfaces was also considered by Germanier et al. [140] by implantation on minipigs and results analysed after 2 and 4 weeks of healing. And finding a significant enhancement of new bone apposition to the RGD-functionalised dental implant surfaces compared tot non-coated ones.

The majority of the in vivo results applied on the dental implants field, consist on the surface functionalisation with extracellular matrix components, especially RGD containing sequences. Nonetheless, a low amount of in vivo studies of functionalised surfaces has been found.

It is expected that future improvements in implant surface engineering will present a combination between surface roughness, chemistry, and incorporated biological factors that will further enhance the host-to-implant response. Current limiting factors for such technology remain its large-scale fabrication and the lack of regulatory standards [122].

Introduction

54

1.6. Conclusions

The fate of implants is determined by the response of cells to the surface of the biomaterial. In both cell culture and implantation, the initial response of a foreign material to cells is the spontaneously adsorption of a protein layer, whose composition and bioactivity provide a biological interpretation of the underlining physicochemical properties.

Since all aspects of osteoblast response are dependent on their mode of adhesion and their morphology, the availability and distribution of adhesion proteins are critical in the process of bone formation. Very specific signal pathways are followed to heal wounds in our body. When a biomaterial is implanted, the complex signals that derive from the non-specifically adsorbed protein layer triggers the immunologically-defensive foreign body reaction in our system. Hence, the appropriate adsorption of proteins such as vitronectin and fibronectin is a key factor in processes such as bone formation at an implant surface. The interaction of fibrinogen with implant surfaces is also likely to be critical in vivo, because of its involvement in clot formation and platelet activation.

Several general conclusions can be derived from the diverse literature on in vitro cell/materials interactions:

• Moderately hydrophilic surfaces appear to induce more favourable cell responses than hydrophobic materials, by favouring adsorption of vitronectin and/or fibronectin and better preserving the bioactivity of the adsorbed proteins

• Net charges on hydrophilic surfaces offer a further driving force for protein adsorption, and may influence the selectivity and orientation of proteins bound.

• Both positively and negatively charged surfaces are able to bind adhesion proteins and support osteoblast functions, however, an optimum charge has not been found.

• Titanium and hydroxyapatite, two of the most relevant biomaterials used in osseointegrative implants, exhibit a degree of hydrophilicity and negative surface potential at physiological pH.

• Cells respond to topography, but the mediating mechanism leading the specific response has not been unravelled yet. Changes in the nature of the focal adhesions used to anchor the cell to its substrate due to surface topography are the most plausible cause of it.

The design of materials that induce specific and appropriate protein adsorption -amount, orientation, conformation-, which could trigger normal wound healing and tissue integration is a challenging and ultimate goal in materials science. The protein recruitment can be either spontaneous or directed by surface modifications.

Chapter 1

55

The understanding of the formation and function of the adsorbed protein layer as well as the most relevant surface properties influencing those responses would provide an important link between biomateria l surface properties and biological responses.

Although the complexity of the interface and difficulties in its characterization continue to restrict understanding of cell-biomaterial interactions, this work is focused on identifying a number of key elements in the system.

Introduction

56

1.7. Bibliography

[1] Aparicio C, Gil FJ, Planell JA, Engel E. Human-osteoblast proliferation and differentiation on grit-blasted and bioactive titanium for dental applications. Journal of Materials Science-Materials in Medicine 2002; 13: 1105-1111.

[2] Aparicio C, Gil FJ, Thams U, Munoz F, Padros A, Planell JA. Osseointegration of grit-blasted and bioactive titanium implants: Histomorphometry in minipigs. Key engineering materials 2004; 254-2: 737-740.

[3] Park J, Lakes RS. Introduction. In Biomaterials. An introduction. Springer; 2007: 2-13.

[4] Williams DF. Definitions in biomaterials. Amsterdam: Elsevier; 1987.

[5] Branemark PI, Adell R, Albrektsson T, Lekholm U, Lundkvist S, Rockler B. Osseointegrated Titanium Fixtures in the Treatment of Edentulousness. Biomaterials 1983; 4: 25-28.

[6] Albrektsson T, Branemark PI, Hansson HA, Lindstrom J. Osseointegrated Titanium Implants - Requirements for Ensuring A Long-Lasting, Direct Bone-To-Implant Anchorage in Man. Acta Orthopaedica Scandinavica 1981; 52: 155-170.

[7] Branemark PI, Albrektsoon T. Introduction to osseointagration. In: Branemark PI, Zarb G, Albrektsoon T (eds.), Tissue Integrated Prostheses. Chicago: Quintessence; 1985.

[8] Stanford CM, Keller JC, Solursh M. Bone Cell Expression on Titanium Surfaces Is Altered by Sterilization Treatments. Journal of Dental Research 1994; 73: 1061-1071.

[9] Liu XY, Chu PK, Ding CX. Surface modification of titanium, titanium alloys, and related materials for biomedical applications. Materials Science & Engineering R-Reports 2004; 47: 49-121.

[10] Williams D.F. Titanium for medical applications. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Springer; 2001: 14-23.

[11] Freese H, Vola M, Wood R. Metallurgy and technological properties of titanium and titanium a alloys. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Springer; 2001: 26-49.

[12] Williams D.F. Corrosion of implant materials. In. 1976: 237-266.

[13] Textor M, Sittig C, Frauchiger V, Tosatti S. Properties and biological significance of natural oxide films on titanium and its alloys. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Springer; 2001: 172-224.

[14] Lausmaa J. Surface spectroscopic characterization of titanium implant materials. Journal of Electron Spectroscopy and Related Phenomena 1996; 81: 343-361.

[15] Sul YT, Johansson CB, Petronis S, Krozer A, Jeong Y, Wennerberg A, Albrektsoon T. Characteristics of the surface oxides on turned and electrochemically oxidized pure titanium implants up to dielectric breakdown: : the oxide thickness, micropore configurations, surface roughness, crystal structure and chemical composition. Biomaterials 2002; 23: 491-501.

[16] Tengvall P, Lundström I. Physico-chemical considerations of titanium as a biomaterial. Clinical Materials 1992; 9: 115-134.

[17] Paine M, Wong J. Osteoblast response to pure titanium and titanium alloy. In: Ellingsen J, Lyngstadaas S (eds.), Bio-implant interface. Improving biomaterials and tissue reactions. CRC Press; 2003.

Chapter 1

57

[18] Wälivaara B, Aronsson B, Rodahl M, Lausmaa J, Tengvall P. Titanium with different oxides: in vitro studies of protein adsorption and contact activation. Biomaterials 1994; 15: 827-834.

[19] Kurrat R, Walivaara B, Marti A, Textor M, Tengvall P, Ramsden JJ, Spencer ND. Plasma protein adsorption on titanium: comparative in situ studies using optical waveguide lightmode spectroscopy and ellipsometry. Colloids and Surfaces B-Biointerfaces 1998; 11: 187-201.

[20] Perren S, Pohler O, Schneider E. Titanium as implant material for osteosynthesis applications. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Springer; 2001: 772-823.

[21] Boyer R, Welsch G, Collings EW. Materials properties handbook. Titanium Alloys, 3rd ed. ASM International; 2003.

[22] Black J., Hastings G. Handbook of biomaterial properties. Thomson Science; 1998.

[23] Ratner BD. A perspective on titanium biocompatibility. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Springer; 2001: 1-12.

[24] Williams RL, Brown SA, Merritt K. Electrochemical Studies on the Influence of Proteins on the Corrosion of Implant Alloys. Biomaterials 1988; 9: 181-186.

[25] Hench LL, Ethridge EC. Biomaterials. The interfacial problem. Advances in Biomedical Engineering 1975; 5: 35.

[26] Klinger A, Steinberg D, Kohavi D, Sela MN. Mechanism of adsorption of human albumin to titanium in vitro. Journal of Biomedical Materials Research 1997; 36: 387-392.

[27] Serro AP, Fernandes AC, Saramago B, Lima J, Barbosa MA. Apatite deposition on titanium surfaces - The role of albumin adsorption. Biomaterials 1997; 18: 963-968.

[28] Tengvall P. Proteins at titanium interfaces. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Springer; 2001: 458-480.

[29] Sundgren JE, Bodo P, Ivarsson B, Lundstrom I. Adsorption of Fibrinogen on Titanium and Gold Surfaces Studied by Esca and Ellipsometry. Journal of Colloid and Interface Science 1986; 113: 530-543.

[30] Thomsen P, Larsson C, Ericson LE, Sennerby L, Lausmaa J, Kasemo B. Structure of the interface between rabbit cortical bone and implants of gold, zirconium and titanium. Journal of Materials Science-Materials in Medicine 1997; 8: 653-665.

[31] Cochran DL, Schenk RK, Lussi A, Higginbottom FL, Buser D. Bone response to loaded and unloaded titanium implants with sandblasted and acid-etched surface: a histometric study in the canine mandible. Journal of Biomedical Materials Research 1998; 40: 1-11.

[32] Klein CP, Wolke JG, de Blieck-Hogervorst JM, de Groot K. Calcium phosphate plasma-sprayed coatings and their stability: an in vivo study. Journal of Biomedical Materials Research 1994; 28: 909-917.

[33] Jansen JA, Wolke JG, Swann S, Van der Waerden JP, de Groot K. Application of magnetron sputtering for producing ceramic coatings on implant materials. Clinical Oral Implants Research 1993; 4: 28-34.

[34] Montenero A, Gnappi G, Ferrari F, Cesari M, Salvioli E, Mattogno L, Kaciulis S, Fini M. Sol-gel derived hydroxyapatite coatings on titanium substrate. Journal of Materials Science 2000; 35: 2791-2797.

Introduction

58

[35] Kokubo T, Kim HM, Kawashita M, Nakamura T. Bioactive metals: preparation and properties. Journal of Materials Science-Materials in Medicine 2004; 15: 99-107.

[36] Aparicio C. Tratamientos de superficie sobre titanio comercialmente puro para la mejora de la osteointegración de los implantes dentales. Universitat Politècnica de Catalunya; 2005. Thesis.

[37] Davies JE, Hosseini MM. Histodynamics of endosseous wound healing. In: Vacanti J, Davies JE (eds.), Bone engineering. Toronto: 2000: 1.

[38] Esposito M, Worthington H, Coulthard P, Wennerberg A, Thomsen P. Role of implant surface properties on the clinical outcome of osseointegrated oral implant therapy: an evidence-based approach. In: Ellingsen J, Lyngstadaas S (eds.), Bio-implant interface. Improving biomaterials and tissue reactions. CRC Press; 2003: 3-28.

[39] Lausmaa J. Mechanical, thermal, chemical and electrochemical surface treatment of titanium. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Springer; 2001: 232-258.

[40] Bagno A, Di Bello C. Surface treatments and roughness properties of Ti-based biomaterials. Journal of Materials Science-Materials in Medicine 2004; 15: 935-949.

[41] Orsini G, Assenza B, Scarano A, Piatelli M, Piatelli A. Surface analysis of machined versus sandblasted and acid-etched titanium implants. International journal of oral and maxillofacial implants 2000; 15: 779-784.

[42] Taborelli M, Jobin M, Francois P, Vaudaux P, Tonetti M, SzmuklerMoncler S, Simpson JP, Descouts P. Influence of surface treatments developed for oral implants on the physical and biological properties of titanium .1. Surface characterization. Clinical Oral Implants Research 1997; 8: 208-216.

[43] Francois P, Vaudaux P, Taborelli M, Tonetti M, Lew DP, Descouts P. Influence of surface treatments developed for oral implants on the physical and biological properties of titanium .2. Adsorption isotherms and biological activity of immobilized fibronectin. Clinical Oral Implants Research 1997; 8: 217-225.

[44] Aparicio C, Gil FJ, Fonseca C, Barbosa M, Planell JA. Corrosion behaviour of commercially pure titanium grit blasted with different materials and sizes of grit particles for dental implant applications. Biomaterials 2003; 24: 263-273.

[45] Gotfredsen K, Karlsson U. A prospective 5-year study of fixed partial prostheses supported by implants with machined and TiO2-blasted surface. Journal of Prosthodontics 2001; 6: 2-7.

[46] Zinger O, Anselme K, Denzer A, Habersetzer P, Wieland M, Jeanfils J, Hardouin P, Landolt D. Time-dependent morphology and adhesion of osteoblastic cells on titanium model surfaces featuring scale-resolved topography. Biomaterials 2004; 25: 2695-2711.

[47] Massaro C, Rotolo P, De Riccardis F, Milella E, Napoli A, Wieland M, Textor M, Spencer ND, Brunette DM. Comparative investigation of the surface of commercial titanium dental implants. Part 1: chemical composition. Journal of Biomedical Materials Research. Materials in Medicine 2002; 13: 535-548.

[48] Adell R, Eriksson B, Lekholm U, Brånemark PI, Jemt T. Long-term follow-up study of osseointegrated implants in the treatment of totally edentulous jaws. International Journal of Oral and Maxillofacial Implants 1990; 5: 347-359.

[49] Yokoyama K, Ichikawa T, Miyamoto Y, Asaoka K. Fracture mechanisms of retrieved titanium screw thread in dental implants. Biomaterials 2002; 23: 2459-2465.

[50] Le Guéhennec L, Soueidan A, Layrolle P, Amouriq Y. Surface treatments of titanium dental implants for rapid osseointegration. Dental Materials 2007; 23: 844-854.

Chapter 1

59

[51] Sul Y.T., Johansson CB, Roser K, Albrektsson T. Qualitative and quantitative observations of bone tissue reactions to anodised implants. Biomaterials 2002; 23: 1809-1817.

[52] Buser D, Schenk RK, Steinemann S, Fiorellini JP, Fox CH, Stich H. Influence of surface characteristics on bone integration of titanium implants. A histomorphometric study in miniature pigs. Journal of Biomedical Materials Research 1991; 25: 889-902.

[53] Martin JY, Schwartz Z, Hummert TW, Schraub DM, Simpson J, Lankford J, Dean DD, Cochran DL, Boyan BD. Effect of Titanium Surface-Roughness on Proliferation, Differentiation, and Protein-Synthesis of Human Osteoblast-Like Cells (Mg63). Journal of Biomedical Materials Research 1995; 29: 389-401.

[54] Urban RM, Jacobs JJ, Tomlinson MJ, Gavrilovic J, Black J, Peoch M. Dissemination of wear particles to the liver, spleen and abdominal lymph nodes of patients with hip or knee replacement. The Journal of Bone and Joint Surgery 2000; 82: 457-477.

[55] Anselme K. Osteoblast adhesion on biomaterials. Biomaterials 2000; 21: 667-681.

[56] MacDonald DE, Markovic B, Boskey AL, Somasundaran P. Physico-chemical properties of human plasma fibronectin binding to well characterized titanium dioxide. Colloids and Surfaces B-Biointerfaces 1998; 11: 131-139.

[57] Lee J, Saavedra S. Molecular orientation in heme protein films adsorbed to hydrophilic and hydrophobic glass surfaces. Langmuir 1996; 12: 4025-4032.

[58] Healy KE, Thomas CH, Rezania A, Kim JE, McKeown PJ, Lom B, Hockberger PE. Kinetics of bone cell organization and mineralization on materials with patterned surface chemistry. Biomaterials 1996; 17: 195-208.

[59] Dee KC, Rueger DC, Andersen TT, Bizios R. Conditions which promote mineralization at the bone-implant interface: a model in vitro study. Biomaterials 1996; 17: 209-215.

[60] Nanci A, West JD, Peru L, Brunet P, Sharma V, Zalzal S, MCkee D. Chemical modification of titanium surfaces for covalent attachment of biological molecules. Journal of Biomedical Materials Research Part A 1998; 40: 324-335.

[61] Rezania A, Healy K. The effect of peptide surface density on mineralization of a matrix deposited by osteogenic cells. Journal of Biomedical Materials Research 2000; 52: 595-600.

[62] Webster SS. Integrated bone tissue physiology: anatomy and physiology. In: Cowin S. (ed.), Bone mechanics handbook. Boca Raton; 2001: 1-1-1-34.

[63] Baron R. Anatomy and biology of bone matrix and cellular elements. In: Favus M (ed.), Primer on the metabolic bone diseases and disorders of mineral metabolism. Lippincott-Raven; 1996: 3-35.

[64] Rho JY, Kuhn-Spearing L, Zioupos P. Mechanical properties and the hierarchical structure of bone. Medical Engineering & Physics 1998; 20: 92-102.

[65] Pilitsis J, Lucas D, Rengachary S. Bone healing and spinal fusion. Neurosurgery Focus 2002; 13: Article 1.

[66] Majeska R. Cell biology of bone. In: Cowin S. (ed.), Bone mechanics handbook. Boca Raton; 2001: 2-1-2-14.

[67] Rho JY, Ashman RB, Turner CH. Youngs Modulus of Trabecular and Cortical Bone Material - Ultrasonic and Microtensile Measurements. Journal of Biomechanics 1993; 26: 111-119.

[68] Rho JY, Tsui TY, Pharr GM. Elastic properties of human cortical and trabecular lamellar bone measured by nanoindentation. Biomaterials 1997; 18: 1325-1330.

Introduction

60

[69] Wilson CJ, Clegg RE, Leavesley DI, Pearcy MJ. Mediation of biomaterial-cell interactions by adsorbed proteins: A review. Tissue Engineering 2005; 11: 1-18.

[70] Kasemo B. Biological surface science. Surface Science 2002; 500: 656-677.

[71] Puleo DA, Nanci A. Understanding and controlling the bone-implant interface. Biomaterials 1999; 20: 2311-2321.

[72] Kalfas I. Principles of bone healing. Neurosurgery Focus 2001; 10: Article 1.

[73] Sundgren JE, Bodo P, Ivarsson B, Lundstrom I. Adsorption of Fibrinogen on Titanium and Gold Surfaces Studied by Esca and Ellipsometry. Journal of Colloid and Interface Science 1986; 113: 530-543.

[74] Winter GD. Tissue reaction to metallic wear and corrosion products in human patients. Journal of Biomedical Materials Research 1974; 8: 11-26.

[75] Davies JE, Lowenberg B, Shiga A. The Bone Titanium Interface Invitro. Journal of Biomedical Materials Research 1990; 24: 1289-1306.

[76] Nanci A, McKee MD, Zalzal S, Sakkal S. Ultrastructural and immunocytochemical analysis of the tissue response to metal implants in the rat tibia. In: Davidovitch Z, Mah J (eds.), Biological Mechanisms of Tooth Eruption, Resorption and Replacement by Implants. Boston: Harvard Society for the Advancement of Orthodontics; 1998: 487-500.

[77] Cooper LF, Masuda T, Yliheikkila PK, Felton DA. Generalizations regarding the process and phenomenon of osseointegration. Part II. In vitro studies. International Journal of Oral & Maxillofacial Implants 1998; 13: 163-174.

[78] Boyan BD, Lohmann CH, Dean DD, Sylvia VL, Cochran DL, Schwartz Z. Mechanisms involved in osteoblast response to implant surface morphology. Annual Review of Materials Research 2001; 31: 357-371.

[79] Howlett CR, Evans MDM, Walsh WR, Johnson G, Steele JG. Mechanism of Initial Attachment of Cells Derived from Human Bone to Commonly Used Prosthetic Materials During Cell-Culture. Biomaterials 1994; 15: 213-222.

[80] Kilpadi KL, Chang PL, Bellis SL. Hydroxylapatite binds more serum proteins, purified integrins, and osteoblast precursor cells than titanium or steel. Journal of Biomedical Materials Research 2001; 57: 258-267.

[81] Schneider G, Burridge K. Formation of Focal Adhesions by Osteoblasts Adhering to Different Substrata. Experimental Cell Research 1994; 214: 264-269.

[82] Fabriziushoman DJ, Cooper SL. Competitive Adsorption of Vitronectin with Albumin, Fibrinogen, and Fibronectin on Polymeric Biomaterials. Journal of Biomedical Materials Research 1991; 25: 953-971.

[83] McFarland CD, Thomas CH, DeFilippis C, Steele JG, Healy KE. Protein adsorption and cell attachment to patterned surfaces. Journal of Biomedical Materials Research 2000; 49: 200-210.

[84] Schneider G, Burridge K. Formation of Focal Adhesions by Osteoblasts Adhering to Different Substrata. Experimental Cell Research 1994; 214: 264-269.

[85] Degasne I, Basle MF, Demais V, Hure G, Lesourd M, Grolleau B, Mercier L, Chappard D. Effects of roughness, fibronectin and vitronectin on attachment, spreading, and proliferation of human osteoblast-like cells (Saos-2) on titanium surfaces. Calcified Tissue International 1999; 64: 499-507.

[86] Sinha RK, Tuan RS. Regulation of human osteoblast integrin expression by orthopedic implant materials. Bone 1996; 18: 451-457.

Chapter 1

61

[87] Siebers MC, ter Brugge PJ, Walboomers XF, Jansen JA. Integrins as linker proteins between osteoblasts and bone replacing materials. A critical review. Biomaterials 2005; 26: 137-146.

[88] Lange R, Luthen F, Beck U, Rychly J, Baumann A, Nebe B. Cell-extracellular matrix interaction and physico-chemical characteristics of titanium surfaces depend on the roughness of the material. Biomolecular Engineering 2002; 19: 255-261.

[89] Grinnell F, Feld MK. Fibronectin Adsorption on Hydrophilic and Hydrophobic Surfaces Detected by Antibody-Binding and Analyzed During Cell-Adhesion in Serum-Containing Medium. Journal of Biological Chemistry 1982; 257: 4888-4893.

[90] Fabriziushoman DJ, Cooper SL. A Comparison of the Adsorption of 3 Adhesive Proteins to Biomaterial Surfaces. Journal of Biomaterials Science-Polymer Edition 1991; 3: 27-47.

[91] Schneider G, Burridge K. Formation of Focal Adhesions by Osteoblasts Adhering to Different Substrata. Experimental Cell Research 1994; 214: 264-269.

[92] Webb K, Hlady V, Tresco PA. Relationships among cell attachment, spreading, cytoskeletal organization, and migration rate for anchorage-dependent cells on model surfaces. Journal of Biomedical Materials Research 2000; 49: 362-368.

[93] Chiquet-Ehrismann R. Inhibition of cell adhesion by anti-adhesive molecules. Current Opinion in Cell Biology 1995; 7: 715-719.

[94] Garcia AJ, Boettiger D. Integrin-fibronectin interactions at the cell-material interface: initial integrin binding and signaling. Biomaterials 1999; 20: 2427-2433.

[95] Globus RK, Doty SB, Lull JC, Holmuhamedov E, Humphries MJ, Damsky CH. Fibronectin is a survival factor for differentiated osteoblasts. Journal of Cell Science 1998; 111: 1385-1393.

[96] Stephansson SN, Byers BA, Garcia AJ. Enhanced expression of the osteoblastic phenotype on substrates that modulate fibronectin conformation and integrin receptor binding. Biomaterials 2002; 23: 2527-2534.

[97] Altankov G, Groth T. Reorganization of Substratum-Bound Fibronectin on Hydrophilic and Hydrophobic Materials Is Related to Biocompatibility. Journal of Materials Science-Materials in Medicine 1994; 5: 732-737.

[98] Gronthos S, Stewart K, Graves SE, Hay S, Simmons PJ. Integrin expression and function on human osteoblast-like cells. Journal of Bone and Mineral Research 1997; 12: 1189-1197.

[99] Vöros J, Wieland M, Ruiz-Taylor L, Textor M, Brunette DM. Characterization of Titanium Surfaces. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Ed.Springer; 2001: 87-144.

[100] Boyan BD, Hummert TW, Dean DD, Schwartz Z. Role of material surfaces in regulating bone and cartilage cell response. Biomaterials 1996; 17: 137-146.

[101] Deligianni DD, Katsala N, Ladas S, Sotiropoulou D, Amedee J, Missirlis YF. Effect of surface roughness of the titanium alloy Ti-6Al-4V on human bone marrow cell response and on protein adsorption. Biomaterials 2001; 22: 1241-1251.

[102] Anselme K, Linez P, Bigerelle M, Le Maguer D, Le Maguer A, Hardouin P, Hildebrand HF, Iost A, Leroy JM. The relative influence of the topography and chemistry of TiAl6V4 surfaces on osteoblastic cell behaviour. Biomaterials 2000; 21: 1567-1577.

[103] Lincks J, Boyan BD, Blanchard CR, Lohmann CH, Liu Y, Cochran DL, Dean DD, Schwartz Z. Response of MG63 osteoblast-like cells to titanium and titanium alloy is dependent on surface roughness and composition. Biomaterials 1998; 19: 2219-2232.

Introduction

62

[104] Anselme K, Bigerelle M, Noel B, Dufresne E, Judas D, Iost A, Hardouin P. Qualitative and quantitative study of human osteoblast adhesion on materials with various surface roughnesses. Journal of Biomedical Materials Research 2000; 49: 155-166.

[105] Boyan BD, Batzer R, Kieswetter K, Liu Y, Cochran DL, Szmuckler-Moncler S, Dean DD, Schwartz Z. Titanium surface roughness alters responsiveness of MG63 osteoblast-like cells to 1 alpha,25-(OH)(2)D-3. Journal of Biomedical Materials Research 1998; 39: 77-85.

[106] MacDonald DE, Markovic B, Allen M, Somasundaran P, Boskey AL. Surface analysis of human plasma fibronectin adsorbed to commercially pure titanium materials. Journal of Biomedical Materials Research 1998; 41: 120-130.

[107] Saenger W. Structure and dynamics of water surrounding biomolecules. Annual Review of Biophysics and Biomolecular Structure 1987; 16: 93-114.

[108] Vogler EA. Structure and reactivity of water at biomaterial surfaces. Advances in Colloid and Interface Science 1998; 74: 69-117.

[109] Haynes CA, Norde W. Globular proteins at solid/liquid interfaces . Colloids andSurfaces B: Biointerfaces 1994; 2: 517-566.

[110] Norde W. Adsorption of proteins from solution at the solid-liquid interface. Advances in Colloid and Interface Science 1986; 25: 267-340.

[111] Webb K, Hlady V, Tresco PA. Relationships among cell attachment, spreading, cytoskeletal organization, and migration rate for anchorage-dependent cells on model surfaces. Journal of Biomedical Materials Research 2000; 49: 362-368.

[112] MacDonald DE, Deo N, Markovic B, Stranick M, Somasundaran P. Adsorption and dissolution behavior of human plasma fibronectin on thermally and chemically modified titanium dioxide particles. Biomaterials 2002; 23: 1269-1279.

[113] Carew E, Cooke F, Lemons J, Ratner B, Vesely I, Vogler E. Properties of materials. In: Ratner BD, Hoffman AS, Schoen FJ, Lemons JE (eds.), Biomaterials Science, 2nd ed. Elsevier; 2004: 23-66.

[114] Nygren H. Initial reactions of whole blood with hydrophilic and hydrophobic titanium surfaces. Colloids and Surfaces B-Biointerfaces 1996; 6: 329-333.

[115] Faucheux N, Schweiss R, Lutzow K, Werner C, Groth T. Self-assembled monolayers with different terminating groups as model substrates for cell adhesion studies. Biomaterials 2004; 25: 2721-2730.

[116] Schneider G, Burridge K. Formation of Focal Adhesions by Osteoblasts Adhering to Different Substrata. Experimental Cell Research 1994; 214: 264-269.

[117] Kilpadi DV, Raikar GN, Liu J, Lemons JE, Vohra Y, Gregory JC. Effect of surface treatment on unalloyed titanium implants: Spectroscopic analyses. Journal of Biomedical Materials Research 1998; 40: 646-659.

[118] Vezeau PJ, Koorbusch GF, Draughn RA, Keller JC. Effects of multiple sterilization on surface characteristics and in vitro biologic responses to titanium. Journal of Oral and Maxillofacial Surgery 1996; 54: 738-746.

[119] Shelton RM, Rasmussen AC, Davies JE. Protein Adsorption at the Interface Between Charged Polymer Substrata and Migrating Osteoblasts. Biomaterials 1988; 9: 24-29.

[120] Salgueiredo E, Vila M, Silva MA, Lopes MA, Santos JD, Costa FM, Silva RF. Biocompatibility evaluation of DLC-coated Si3N4 substrates for biomedical applications. Diamond and Related Materials 2008; 17: 878-881.

[121] Scotchford CA, Ball M, Winkelmann M, Voros J, Csucs C, Brunette DM, Danuser G, Textor M. Chemically patterned, metal-oxide-based surfaces produced by

Chapter 1

63

photolithographic techniques for studying protein- and cell-interactions. II: Protein adsorption and early cell interactions. Biomaterials 2003; 24: 1147-1158.

[122] Coelho P, Granjeiro J, Romanos G, Suzuki M, Cardaropoli G, Silva N, Thompson vP, Lemons J. Basic research methods and current trends of dental implant surfaces. Journal of Biomedical Materials Research Part B: Applied Biomaterials 2009; 88B: 579-596.

[123] Sykaras N, Iacopino A, Marker V, Tripplet R, Woody R. Implant materials, designs, and surface topographies: their effect on osseointegration. A literature review. International Journal of Oral & Maxillofacial Implants 2000; 15: 675-690.

[124] Windler M, Klabunde R. Titanium for hip and knee prostheses. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Springer; 2001: 704-737.

[125] Esposito M. Titanium for dental applications (I). In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Berlin: Springer; 2001.

[126] Buser D. Titanium for dental applications (II): Implants with roughened surfaces. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Berlin: Springer; 2001.

[127] Olin C. Titanium in cardiac and cardiovascular applications. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Berlin: Springer; 2001.

[128] Wenneberg A, Albrektsson T. Suggested guidelines for the topographic evaluation of implant surfaces. International journal of oral and maxillofacial implants 2000; 15: 331-344.

[129] Ong JL, Carnes DL, Bessho K. Evaluation of titanium plasma-sprayed and plasma-sprayed hydroxyapatite implants in vivo. Biomaterials 2004; 25: 4601-4606.

[130] Lacefield WR. Current status of ceramic coatings for dental implants. Implant Dentistry 1998; 7: 315-322.

[131] Yang Y, Kim KH, Ong JL. A review on calcium phosphate coatings produced using sputtering process-an alternative to plasma spraying. Biomaterials 2005; 26: 327-337.

[132] Albrektsson T. State of the art in oral implants. Journal of Clinical Periodontology 1991; 18: 474-481.

[133] Vercaigne S, Wolke JG, Naert I, Jansen JA. A mechanical evaluation of TiO2-gritblasted and

Ca-P magnetron sputter coated implants placed into the trabecular bone of the goat: Part 1. Clinical Oral Implants Research 2000; 11: 305-313.

[134] Vercaigne S, Wolke JG, Naert I, Jansen JA. A histological evaluation of TiO2-gritblasted and

Ca-P magnetron sputter coated implants placed into the trabecular bone of teh goat: part 2. Clinical Oral Implants Research 2000; 11: 314-324.

[135] Coelho PG, Suzuki M. Evaluation of an IBAD thin-film process as an alternative method for surface incorporation of bioceramics on dental implants. A study in dogs. Journal of Applied Oral Science 2005; 13: 87-92.

[136] Mendes VC, Moineddin R, Davies JE. The effect of discrete calcium phosphate nanocrystals on bone-bonding to titanium surfaces. Biomaterials 2007; 28: 4748-4755.

[137] Orsini G, Piatelli M, Scarano A, Petrone G, Kenealy J, Platelli A, Caputi S. Randomized, controlled histologic and histomorphometric evaluation of implants with nanometer-scale calcium phosphate added to the dual acid-etched surface in the human posterior maxilla. Jounal of Periodontology 2006; 77: 1984-1990.

Introduction

64

[138] Liu Y, Li J, Hunziker EB, de Groot K. Incorporation of growth factors into medical devices via biomimetic coatings. Philosophical transactions of the Royal Society A 2006; 364: 233-245.

[139] Stadlinger B, Pilling E, Huhle M, Khavkin E, Bierbaum S, Scharnweber D, Kuhlisch E, Eckelt U, Mai R. Suitability of differently designed matrix-based implant surface coatings: an animal study on bone formation. Journal of Biomedical Materials Reseearch Part B: Applied Biomaterials 2008; 87B: 524.

[140] Germanier Y, Tosatti S, Broggini N, Textor M, Buser D. Enhanced bone apposition around biofunctionalized sandblasted and acid-etched titanium implant surfaces. A historphometric study in miniature minipigs. Clinical Oral Implant Research 2006; 17: 251-257.

Chapter 2

65

2. Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

2.1. Introduction

Biological tissues mainly interact with the outermost atomic layers of an implant. Consequently, much effort is being devoted to methods of modifying surfaces of existing biomaterials to achieve desired biological responses [1]. Until now, implantology research has studied numerous additive and subtractive surface modification techniques aimed on altering the surface topography of Ti implants for enhanced bone-to-implant contact and increased biomechanical interlocking with bone, [2], and then, reducing healing time. A more detailed explanation of the different modifying surface treatments on Ti can be found at the introduction chapter section 1.3.3.3.

Obtaining rough surfaces has been crucial for this goal and nowadays it is widely used since it has been correlated to an increasing direct-contact between bone and implant as confirmed by numerous in vivo studies [3-10]. In addition, grit-blasting has been widely used to roughen titanium implants because surface roughness of this metal is known to

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

66

have a considerable effect on the in vitro attachment, adhesion, proliferation and differentiation of the osteoblasts [11-14].

In general, cells cultured on rougher surfaces tend to exhibit attributes of more differentiated osteoblasts compared to smoother surfaces [10;15;16]. Then, roughness results in an improvement of bone anchorage of titanium implants [17] and increased removal torque forces [17]. Studies using electron micromachined surfaces show that micron-scale roughness contributes to cell attachment, spreading and differentiation, and superposition of submicron scale enhances local factor production [19].

In previous works, the optimal roughness, 4 µm, and type of abrasive particles, SiC and Al2O3, for a better in vitro and in vivo response in the system studied here have been determined [20;21]. Surface roughness depends on the particle size but does not significantly depend on the nature of the particles [22]. Considering macro-roughness, above 50µm that is the typical dimension of human osteoblasts, different authors have considered 0,72 [23], 1,32 [24] and 4 µm [21], among others, as the best roughness for enhancing adhesion and differentiation properties of the implants. Differences can be explained as the different abrasive particles, cell lines and cell culture conditions used. In general, silicon carbide (SiC) [20], alumina (Al2O3) [25;26], rutile (TiO2) [27], which are inert ceramic materials in the biological environment, and hydroxyapatite, the natural mineral composing bone, of different size are the grit-blasting particles mostly used for Ti and Ti alloys implants [28]. Grit-blasting also causes the incrustation of parts of the abrasive particles on the substrate, which leads to chemical heterogeneities on the implant surface [29;30]. The use of rutile avoids the grit-blasting contamination, silicon carbide and aluminium oxide residues can be partially removed in acidic solutions, and hydroxyapatite particles can enrich the surface in Ca and P [28]. Also, silicon carbide and amorphous alumina are cytocompatible for human fibroblasts and osteoblasts [31].

Cell behaviour on biomaterial surfaces depends on implant-cell interactions, correlated with surface properties. Surface hydrophilicity, roughness, texture, chemical composition, charge and morphology strongly affect cellular responses in contact with the implants by means of protein response. Moreover, the interactions between adsorbed proteins and the implant surface, depending on the binding strength, may lead to changes in conformation of a protein and then, changes in cellular response [32-34]. However, it is poorly understood which of the surface properties mainly induce an optimal biological behavior due to the lack of exhaustive surface characterization.

It is expected that the wettability of a surface, which is directly related to its free energy, polar character and surface charge, will significantly influence on the proteins/synthetic surface interactions [35]. Those interactions will eventually have a strong influence on the in vitro and in vivo cell behaviour. Roughness and chemical composition also affect surface wettability. Consequently, the exhaustive characterization of these relevant surface characteristics can help to better understand the influence of grit-blasting on the appropriate biological response.

Chapter 2

67

Moreover, sterilization is the final step when manufacturing any implant device. Taking into consideration the previous comments, the effect of sterilization processes on the modification of the first surface layers of the biomaterials surfaces must be clearly understood and characterized [36], specially because some of the sterilization processes affect surface finishing, causing “contamination” [37] and increasing the thickness of the oxide surface layer, which in turn may affect the biologic response [38]. Up to now, it has not been thoroughly investigated whether contamination due to sterilization and manufacturing processes on titanium surface, have consequences for the surface properties and hence, for the biological response [39;40].

The characterization, before and after the surface treatments, includes scanning electron microscopy (SEM), white light interferometer microscopy (LIM), X-ray photoelectron spectroscopy (XPS), static contact angle analysis (SCA), and zeta potential (ZP), so as to determine surface morphology, roughness, contamination, wettability and energetic properties, and surface charge respectively.

Time-related contact angle measurements (Time-CA) have been also performed, which allows monitoring the changes in the wetting of the surface during time. According to this, an additional objective of this work is to characterize the influence of protein adsorption on the wettability of the studied materials by using different culture media as testing fluid in Time-CA tests.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

68

2.2. Objectives

The main objectives of the present chapter are:

1. Evaluating and characterizing roughness and morphology of different rough grit-blasted Ti surfaces. Studying the influence of grit-blasting and sterilization procedures on surface chemical composition of smooth and rough Ti surfaces.

2. Characterizing wettability and surface free energy of smooth and rough Ti surfaces and determining the influence of roughness and chemistry on surface free energy.

3. Studying the influence of protein adsorption on the wettability of smooth and rough Ti surfaces.

4. Characterizing surface charge of smooth and rough Ti series and determining the influence of roughness and chemistry on zeta-potential.

Chapter 2

69

2.3. Topography, roughness and surface chemistry

2.3.1. Topography/roughness and cell/protein response

Different studies have demonstrated the influence of topography and roughness on cell response, which will determine the exit or the failure of an implant [41-43]. Cell-material interactions occur in two steps, the first step involves the attachment, adhesion and spreading of the cells. The quality of this first step is influencing the outcome of second step, which is the capacity of cells to proliferate and differentiate when in contact with the implant [44].

Bone interlocking or micro-mechanical anchorage at the interface is not a common feature to all surfaces; to achieve it, a certain level of roughness is required. Altering the surface topography of titanium implants has been determined as a method to enhance the bone healing process [45]. Bone forms more readily on rough surfaces whereas fibrous connective tissue is found more frequently on smooth surfaces. Martin et al. [13] found that proliferation and differentiation of MG63 cells on titanium surfaces were affected by surface roughness. Kieswetter et al. [46] observed that cell proliferation was reduced but not blocked by the surface topography; and phenotypic differentiation was enhanced by rougher surface topographies. Moreover, Aparicio et al. [10] suggested that rough and bioactive titanium surfaces enhances adhesion and differentiation activity of human osteoblasts cells. This can lead to a rapid formation of the extracellular matrix and, consequently, to an accelerated short-term osseointegration.

Differences in surface roughness seem to influence the amount of adsorbed proteins. Higher amounts of fibronectin and lower amounts of albumin were found on rough Ti, mean-average roughness (Ra) under 1 µm, alloy compared to smoother samples [8]. In contrast, in another study showed that increased surface micro-roughness of Ti partly decreased the in vitro adsorption of fibronectin [26]. Much of the research on Ti roughness and its effect on bone cells has been performed on randomly roughened surfaces [8;10;47].

2.3.2. Surface chemistry and cell/protein response

Surface chemistry has also been considered an influencing parameter on the biological response of an implant. The composition of the outermost Ti implants surface is mainly TiO2. The oxide surface has a monolayer of carbon-containing species adsorbed on it, probably mainly hydrocarbons from the preparation procedure. In addition there are small amounts, less than 5%, of some inorganic impurities. Surfaces contain unsaturated bonds which lead to the formation of surface reactive layers and adsorbed contamination

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

70

layers. Then, implant surfaces can be affected by sterilisation procedures. Several authors have demonstrated the crucial effect of the sterilization methods of commercially pure titanium (c.p. Ti) on in vitro subsequent cell adhesion [48;49]. These effects may be related to sterilization- induced surface chemical modifications. Moreover, the grit-blasting particles can remain on the surface after cleaning procedures; then, the effect on protein adsorption and cell behaviour of these contaminants particles has also to be considered.

Variations in chemical composition of the surface produce different types of bonding to the biomolecules, which may also then affect their biological response. There are several types of bonding by which biomolecules may stick to the implant surface. These include long-range but weak van der Waals interactions and the short-range, strong chemical bonding such as ionic and covalent bonding [50]. In summary, an oxidized metal implant surface offers a wide range of bond strengths and chemical bonding types. And then, some bonding situations will lead to irreversible bonding and some will imply a continuous exchange of molecules pursuing the dynamic equilibrium. The surface chemistry composition also determines the chemical stability/reactivity of the implant material.

2.3.3. Materials

10×10 mm2 plaques were cut from sheets of cp Titanium Grade II (c.p. Ti). Five different series of c.p. Ti surfaces were obtained depending on the surface treatments applied to the samples:

Smooth Polished c.p. Ti. The samples were abraded subsequently with 400, 600 and 1200 grit silicon carbide abrasive paper and finally polished with a water suspension of 1µm alumina powder

S3 c.p. Ti blasted with SiC-particles of 212–300 µm in size

S6 c.p. Ti blasted with SiC-particles of 425–600 µm in size

S9 • c.p. Ti with SiC-particles of 1000-1400 µm in size

A3 c.p. Ti blasted with Al2O3-particles of 212–300 µm in size

A6 c.p. Ti blasted with Al2O3-particles of 425–600 µm in size

A9 c.p. Ti blasted with Al2O3-particles of 1000-1400 µm in size

Moreover, each of the previous series was finally sterilized with different methods:

WS Without sterilization

SA Steam autoclaving

EO Ethylene oxide

Chapter 2

71

GR Gamma radiation

2.3.3.1. Grit-blasting

Grit-blasting was carried out with a laboratory grit-blasting machine at 0,25 MPa-pressure during the time required for saturation of the roughness of the samples. The particles used for each sample series were those mentioned above.

Prior to the sterilization procedure, all polished and blasted specimens underwent a cleaning protocol consisting of 15 min with acetone and 15 min with bidistilled water.

2.3.3.2. Sterilisation

Five samples of each series were sterilized with the following methods: steam autoclaving (SA) at 121 °C, during 30 min at 1 bar; ethylene oxide (EO) at 37°C during 5 h at 760 mmbar, 18 h forced aeration, 24 h natural aeration; and γ-radiation (GR) using a 25 kGy dose. A series without sterilization (WS) was also used as a control. The samples were appropriately stored in order to maintain sterilization properties until the different characterizations were carried out.

2.3.4. Methods

2.3.4.1. Surface observation

The surfaces were studied by means of scanning electron microscopy (SEM) (6400, Jeol, Japan) with an acceleration voltage of 8,0 keV in order to evaluate qualitatively the effect of the different surface treatments.

2.3.4.2. Surface roughness

White light interferometry (LIM) has been used for many years as a reliable non-contact optical profiling system for measuring step heights and surface roughness in many precision engineering applications. Recent developments in both instrumentation and in measurement software for this technique have increased the vertical (i.e. height) resolution of these instruments to give a capability better than 0,01 nm (i.e. 0,1 Angstrom). The ultimate lateral resolution is limited to around 0,35µm.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

72

Figure 2-1 Schematic diagram of an interference microscope [51]

Roughness measurements were performed with an interferometer microscope (Optical Profiling System, Wyko NT1100, Veeco, USA). Three samples of each of the different grit-blasting and sterilization series were analysed and for every sample three measurements of roughness were carried out.

Surface analysis area was of 189,2×248,7 µm2 for smooth and 459,9×604,4 µm2 for rough series. Data analysis was performed with Wyko Vision 232TM software (Veeco, USA). A Gaussian filter was used to separate waviness and form from roughness of the surface. By applying a cut-off value (?c) of 2,5 mm for roughened samples (A3, A6, A9, S3, S6 and S9) and ?c=0,25 mm for smoothed-ones according to ISO 11562:1996 standard.

Several amplitude, spacing and hybrid roughness parameters exist to describe surface topography [28;52-55]. In this study, the following parameters to describe the surface have been used: (1) two amplitude parameters, which are descriptive in height: Ra, Ry.; (2) two spacing parameters, which describe the spacing between the topographical irregularities: Pc and Sm.; and (3) an hybrid parameter which combines amplitude and spacing characteristics: index area (Figure 2-2) [28;52;56].

• Ra (Arithmetic average roughness): Arithmetic average of the absolute values of all points of the profile.

∑=

=N

1iia Y

N1

R

• Ry (Maximum peak-to-valley height): Describes the distance between the height of the maximum peak Yp and the depth from the maximum valley Yv within a sample length.

vpy YYR +=

Chapter 2

73

z

xMean line

Max. peak

Max. valley

S1 S2 S3 S4 Si

Ry

z

xMean line

Max. peak

Max. valley

S1 S2 S3 S4 Si

Ry

Figure 2-2 Graphical representation of roughness parameters Ra, Ry and Sm.

• Pc (Peak density): Number of pairs peak-valley by unit length within a sample length. • Sm (Average mean line spacing): Arithmetic average spacing between the falling

flanks of peaks on the mean line.

∑=

=N

1iim S

N1

S

Mentioned parameters were obtained from the 2D-profiles extracted from 3D images obtained by LIM. Index areas were calculated from the ratio of real surface area/geometric surface area, which were measured after obtaining 3D-topographical images of the surfaces by LIM.

2.3.4.3. Surface chemistry

The chemical analysis of the surface, before and after the sterilization procedures, was investigated by X-ray photoelectron spectroscopy (XPS) (PHI 5500 ESCA System, Perkin- Elmer Physical Electronics, USA).

The XPS spectra were obtained using a monochromatic X-ray source, Al anode with a Ka radiation energy of 1486,6 eV, the analyses were conducted at a take-off angle of 45°. The elements present on the surface were qualitatively evaluated by low-resolution (pass energy=187,85 eV) survey spectra, whereas high-resolution (pass energy=23,5 eV) spectra were taken to establish the binding energy and peak area for quantitative analysis. The positions of the peaks were referenced to C 1s at 285,0 ± 0,2 eV.

The determination of the atomic concentration provides the ratio element/sum of all the elements present in the acquired data. The calculation is based on the evaluation of the area of the peak, using the sensitivity factors provided by Perkin-Elmer. All data calculations (peak fitting, integration and background subtraction) were performed with appropriate software (Multipak 6.0, Perkin-Elmer Physical Electronics, USA).

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

74

2.3.4.4. Statistics

ANOVA tables and Fisher’s or Tukey’s multiple comparisons tests were performed to assess statistically significant differences between average values for different samples. Tukey's test was used, instead of that of Fisher, for the comparison of several groups of samples when the samples had different sizes. A p-value<0,05 was chosen to determine significance. These tests were carried out with MinitabTM Release 14 (Minitab Inc.,USA).

2.3.5. Results

2.3.5.1. Surface observation

Figure 2-3 shows SEM-pictures that illustrate the different surface finishing for polished and blasted titanium surfaces. Grit-blasting causes the plastic deformation of the metal, resulting in a surface with randomly distributed topographic features. Some abrasive particles are embedded on the samples (Figure 2-3 c-e) due to the experimental procedure, which causes chemical heterogeneities on the surface. These particles have remained on all different blasted-surfaces even after their ultrasonically cleaning. In previous works [20], an approximately 10% surface covering with abrasive particles was determined.

Figure 2-3 SEM images of: a) cpTi polished sample; b) S6 sample; c) SiC particle embedded into a S6

surface; d) A6 sample; e) Al2O3 particle embedded into an A6 surface

Chapter 2

75

2.3.5.2. Surface roughness

Values of roughness parameters and images for the different surfaces studied are shown in Table 2-1 and Figure 2-5. Sterilization process does not statistically affect the roughness of the surface (Figure 2-4), and the greater the size of the grit-blasting particles, the higher the Ra, with statistically significant differences (Table 2-1). Statistically significant differences between samples blasted with different chemical composition of the particles and the same particle size have been also determined. The titanium surfaces treated with SiC-particles show lower Ra-values than those blasted with Al2O3-particles.

An inverse correlation between the amplitude (Ra, Ry) and the Pc spacing parameter has been obtained (Table 2-1). Index area, which is the ratio between the real and the projected area, increases with the particle size.

Table 2-1 Mean-values ± standard deviation of roughness parameters for all titanium surfaces studied

Surface Ra (µm) ± SD Ry (µm) ± SD Sm (µm) ± SD Pc (1/mm) ± SD Index Area ± SD

S3 2,58 ± 0,06 16,63 ± 0,62 33,72 ± 0,69 8,41 ± 0,27 1,30 ± 0,01

S6 4,49 ± 0,25 25,46 ± 1,71 52,84 ± 3,21 5,18 ± 0,56 1,37 ± 0,02

S9 7,09 ± 0,08 43,20 ± 2,84 69,61 ± 2,98 4,37 ± 0,40 1,52 ± 0,06

A3 2,92 ± 0,08 18,29 ± 0,49 42,13 ± 2,55 6,79 ± 0,46 1,27 ± 0,01

A6 4,80 ± 0,25 26,89 ± 1,39 53,29 ± 2,71 5,21 ± 0,60 1,42 ± 0,03

A9 8,64 ± 0,61 50,63 ± 2,22 71,43 ± 5,96 3,82 ± 0,35 1,78 ± 0,20

Smooth 0,034 ± 0,009 0,236 ± 0,063 9,48 ± 9,87 43,39 ± 18,72 1,00 ± 0,00

Surface Ra (µm) ± SD

Smooth-WS 0,034 ± 0,009

Smooth-SA 0,037 ± 0,013

Smooth-EO 0,041 ± 0,008

Smooth-GR 0,030 ± 0,015

Smooth-WS Smooth-SA Smooth-EO Smooth-GR0,000

0,005

0,010

0,015

0,020

0,025

0,030

0,035

0,040

0,045

0,050

0,055

0,060

Ave

rage

roug

hnes

s R

a(µ

m)

Figure 2-4 Influence of sterilization procedure on roughness parameters for smooth titanium surfaces

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

76

A3 A6 A9

Smooth

S3 S3 S6S6 S9 S9

Figure 2-5 3D-images of the different smooth and rough Ti surfaces obtained by white light interferometry

2.3.5.3. Surface chemistry

XPS analysis revealed the presence of O, C and N elements on all different titanium surfaces (Figure 2-6). The quantity of C was significantly higher on the EO- and WS-treated surfaces than on the surfaces treated with the other types of sterilization processes (Figure 2-7).

The atomic ratios of the most significant elements (C, N, O, Al, Si) on the surfaces with different surface qualities, are shown in Figure 2-8.

Higher values of carbon and oxygen atomic ratio were found on blasted surfaces than on polished ones. The amount of nitrogen was statistically equal for all the working samples. Significant amounts of Si and Al were found on SiC and Al2O3 grit-blasted surfaces, respectively.

Other elements, such as Na, Cl and Ca were found in the different surfaces but in lower amounts (minor than 1%).

Chapter 2

77

Figure 2-6 XPS survey spectrum from the surface of: a) Smooth-WS; b) A6-WS and c) S6-WS Ti series. The spectrum exhibits peaks at binding energies that are characteristic for the elements from which they

originate. The spectra are dominated by strong titanium and oxygen peaks, showing that surface consists of a titanium oxide. A relative large carbon signal, as well as smaller nitrogen and calcium signals, caused by

contamination, is also observed. Alumina and silicon from grit-blasting process is also detected on b) A6 and c) S6 spectra, respectively

Figure 2-7 Overlay XPS surveys spectrum from the surface of Smooth-WS (red), Smooth-EO (turquoise),

Smooth-GR (violet) and Smooth-SA (blue)

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

78

C/Ti N/Ti O/Ti Al/Ti Si/Ti0

1

2

3

4

5

6

7

8

9

10

SmoothWS A6WS S6WS SmoothSA A6SA S6SA SmoothEO A6EO S6EO SmoothGR A6GR S6GR

Surface C/Ti N/Ti O/Ti Al/Ti Si/Ti Others

Smooth

WS 4,59 0,16 3,73 Na,Cl,Ca SA 3,20 0,09 3,74 Na,Cl,Ca

EO 6,89 0,10 4,05 Na,Al GR 4,35 0,06 5,06 <0.01

A6

WS 8,41 0,06 6,52 1,22 <0.01 SA 4,39 0,05 6,48 1,50 Na, Cl

EO 8,63 0,15 7,07 1,16 Na, Cl GR 4,44 0,14 6,55 1,31 Cl, Ca

S6

WS 5,60 0,11 4,19 1,83 <0.01 SA 5,48 0,11 4,46 1,94 Cl

EO 6,85 0,12 5,24 2,08 Cl GR 5,54 0,16 5,42 2,12 Na,Al,Cl

Figure 2-8 Atomic percent with respect to Ti of the more significant elements on surfaces with different surface qualities

2.3.6. Discussion

The Ra-values (Table 2-1) show that, as expected, statistically significant differences in surface roughness are due to the different sizes of the grit-blasting particles used. Furthermore, there is a statistically significant difference in the roughness of titanium

Chapter 2

79

blasted with Al2O3 or SiC particles. This result may be explained taking into consideration different causes. Firstly, size distribution of the particles is a main influencing factor. The distribution of sizes for a standard mean size can significantly change for commercially available products. Secondly, the mechanical and physical properties (hardness, fracture toughness and density) of the grit particles can influence surface roughness because the energy transfer when the particle impact depends on its facility for breaking into small parts and its mass as well. As a result, the surface roughness of the surfaces blasted with Al2O3-particles is higher than that obtained when grit-blasting with SiC-particles.

A consequence of treating metallic surfaces with a grit-blasting method is that some grit-blasting particles remain adhered in the surface, obtaining up to 16% of the surface covered with ‘contaminant’ particles (Figure 2-3) [20]. Consequently, alumina and silicon carbide grit-blasting particles are used so that the surface biocompatibility is not compromised.

Some authors have concluded that the increase of the oxide thickness due to different sterilization procedures can be significant enough to increment surface roughness [36-38]. Thierry et al. [36] demonstrated that sterilized materials have a higher Ra trend, particularly for steam autoclave processing. It is clear that sterilization procedure can only affect roughness at the nano-level, but our results carried out to smooth (i.e. polished) surfaces, demonstrate that sterilization process do not statistically increase the surface roughness of smooth series (Figure 2-4). Then, since rough surfaces are at the micro-level, – Ra ranging from 2,58 up to 8,64 µm-, sterilisation processes are not significant in terms of roughness.

Different studies have determined the chemical composition of the surface after sterilization procedures. Keller et al. [38] found that the autoclaved surfaces were contaminated with various ion species. Doundoulakis [37] suggested that autoclave treatment of surfaces resulted in the deposition of contaminating films, and consequently obtaining a lower surface energy [49]. Lausmaa et al. [57] reported that on cpTi and on other Ti alloys, the oxide was basically TiO2 with the usual O, C, and N contaminants. Additional inorganic contaminants including Na, Cl, Si, Ca, P and S were occasionally observed [58].

Moreover, an interesting parameter to study is the O-Ti ratio. As known, the cp Ti oxide film is mainly composed of the stable oxide TiO2. The TiO2/Ti interface has an O to Ti concentration ratio that varies gradually from 2 down to 1 since a lower ratio is found at the bulk [59]. But this ideal O/Ti ratio can be seen disturbed due to contamination. Some organic species like hydrocarbons adsorb and metal-organic species, such as alkoxides or carboxylates of titanium also exist on the outmost surface layer whose concentrations depend on not only the surface conditions, such as cleanliness but also the exposure time to air as well as the quality of the atmosphere during storage. In our study, where samples have been exposed continuously to air atmosphere, the O/Ti ratio is far from

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

80

the ideal value of 2. Nevertheless, smooth and polished surfaces, without any sterilisation value, is the series which approximates more to the value. As seen, for all the studied series (smooth, A6 and S6), the SA does not increase the contamination in oxygen levels; while EO and GR do. But is the effect of the roughness procedure of grit-blasting which mainly increases the O/Ti ratio and the levels of contaminants, probably due to the increase of real surface area.

The series of surface characterization techniques used in this study also demonstrate that the implant surfaces were slightly altered as a result of the sterilization procedures (Figure 2-8). The levels of contamination, as determined by XPS, and the increased wetting angles, as discussed at the following section 2.4.6.1, were probably related. Carbon was on the surface of all specimens, probably resulting in part from compounds adsorbed from the air. This organic contamination is more significant for EO series which have been found the sterilization procedure that leaves more residues at the surface probably due to the sterilization procedure itself. The GR process is a physical and not chemical treatment, as occurs for EO treatment, where the treated surface is exposed to an ionizing radiation. This could explain the less polluted titanium surfaces obtained with this sterilization method, as generally confirmed by our XPS results.

2.3.7. Conclusions

v Grit-blasting increases roughness of Ti surface samples to the micro-level. Differences are found depending on the particle size and, also, the particle nature. The larger the particles, the rougher the blasted surface and the higher the real surface area.

v Sterilization processes did not statistically increase the surface roughness of the different studied series.

v Chemistry of the smooth and rough Ti surfaces was slightly changed due to the different sterilization procedures.

v Grit-blasting procedure embedded abrasive particles on the Ti surfaces. Then, the chemistry of the surface was changed on rough surfaces.

Chapter 2

81

2.4. Wettability and surface energy

2.4.1. Wettabilty and cell/protein response

The bioadhesive phenomena, such as the protein adsorption, the cell and bacterial adhesion, or the integration of a biomedical device in a tissue, are the ones characterized by the formation of a new interphase between the biological system and the implanted material. The thermodynamic laws that explain the interactions between biomolecules and atoms imply that the creation of this new interphase leads to a favourable energy balance. Consequently, the study and quantification of the interfacial free energies involved in bioadhesive phenomena could explain and be useful to predict the biological behaviour of a biomaterial.

For that reason, many authors have focused their works on the study of possible correlations between the wettability and the surface free energy (SFE) of different biomaterials and the biological response, generally in vitro. However, the obtained results vary from one study compared to another depending on the experimental conditions: substrate, proteins and cell lines.

In general, cellular proliferation [60;61] and adhesion [60;62;63] is favoured on hydrophilic surfaces. Fibroblasts react in front of wettability changes, the spreading of these cells increased with wettability when going from hydrophobic to hydrophilic surfaces [64]. An study [65] showed that, despite the great number of parameters interfering with cellular adhesion and spreading, the solid SFE apparently is a dominant factor in cellular attachment to a polymer surface, even if the solid surface has been covered by a protein layer.

Concerning to surface free energy and its components different conclusions have been pointed out by different authors. Valk et al., [66], studied the influence of SFE on cell spreading and cell division rates of transformed mouse fibroblasts on various polymers and it was found that cell spreading was dependant on the polar component of SFE. The spreading was low when the polar component was less than 5mJ/m2, while marked spreading occurred when the polar component was greater than 15mJ/m2. Some authors demonstrated that cell adhesion increased with increasing polarity [67;68]. In particular Lee et al., [68] observed that cells adhered, spread and grew on surfaces with moderate hydrophilicity, also confirmed by Webb et al [60], with a maximum adhesion of cells at a water contact angle of approximately 57º, regardless the cell type used on modified polyethylene surfaces. Bren et al., [69], studied the influence of SFE and electron donor (base) and electron acceptor (acid) parameters on the attachment of osteoblast cell on stainless steel and Ti alloys. Materials with a higher SFE produced increased

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

82

differentiation compared to substrates with lower SFE. The acidic part of SFE was found to have a strong correlation to alkaline phosphatase activity, encouraging differentiation.

As mentioned before, wettability and surface charge both play an important role in protein adsorption to titanium (Ti) dioxide and different approaches have been found favouring hydrophilic or hydrophobic surfaces to increase protein adsorption. Some authors [43;70] confirmed that the hydrophobic character of a surface generally favours protein adsorption. The reason could be the displacement of water molecules at the biomaterial surface, which is an important energetic barrier to protein adsorption [71]. Nevertheless, this behaviour depends on the protein of study [72]. Fibronectin and fibrinogen are two protein which prefer hydrophobic surfaces [[62;73], but they can also adsorb to hydrophilic surfaces [73;74]. Vitronectin is also an important adhesive protein which also likes better hydrophilic surfaces [75;76]. Additionally, it has been demonstrated that the adhesive functions of fibronectin decrease when it is adsorbed onto hydrophobic surfaces, while they are maintained onto hydrophilic surfaces [73;74] being the different protein conformation on the different substrates, the explanation for such different results. On the other hand, protein binding itself changes the wettability of a surface In fact, it is said after protein adsorption, the surface is more hydrophilic due to a decrease of the Lifshitz van der Waals component, dispersive component, which ameliorates the cellular adhesion [77].

2.4.2. Wettability

Wettability or wetting is the actual process when a liquid spreads on (wets) a solid substrate. The wettability of a solid with a liquid is usually determined from their contact angle, CA, or wetting angle. The shape of a liquid front in contact with a solid substrate can be explained by the relative strength of the cohesive (Liquid/Liquid) and adhesive (Solid/Liquid) (Figure 2-9) and is determined by the interfacial forces of the participating phases (Figure 2-12). The possible interfaces can be summarized in terms of the three states of matter solid, liquid and vapor as follows:

• Vapor/liquid • Vapor/solid • Liquid/liquid • Liquid/solid • Solid/solid

Strong adhesion with weak cohesion produces very low contact angles with nearly complete wetting. As the solid/liquid interactions weaken and the liquid/liquid interactions strengthen, wetting diminishes and contact angle increases. Wetting (hydrophilic) or non-wetting (hydrophobic) of a solid by water is a qualitative criterion since all surfaces of condensed-phase materials attract water molecules to a considerable

Chapter 2

83

degree [78]. As shown on Figure 2-10, an elevated contact angle defines a hydrophobic drop while a low value is considered a hydrophilic drop.

Figure 2-9 Schematic representation of the adhesive and cohesive forces distribution inside a drop

Figure 2-10 Hydrophobic and a hydrophilic drop with its respective contact angle

The hydrophilicity of a surface is related to its affinity to water [79] and by extension to its capability to form hydrogen bonds. Some authors [80] suggested defining the hydrophobicity as a repulsive force between non-polar molecules and water, also called the hydrophobic effect. Actually, the hydrophilicity and hydrophobicity concepts are simplified by grouping the different interactions between molecules. In different terms: dispersion (van der Waals including London as well Debye and Keesom) or polar (acid-base Lewis). From this viewpoint, the interaction between a solid surface and a fluid molecule appears as a particular case of the intermolecular forces problem (Figure 2-11). Consequently, the overall interaction depends on (1) the chemical groups on the surface, which are mainly responsible for both polar and electrostatic interactions, (2) the surface typology (porosity, roughness, and crystallography) and, (3) the structure and functionality of the fluid components.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

84

1. Ion-Dipole = Electrostatic forces (ion to polar molecule)2. Hydrogen bonds (polar to polar, must involve hydrogen)3. Dipole-dipole forces (polar to polar)4. Van der Waals = Dispersion forces (non-polar to non-polar)

i. London forcesii. Debye forcesiii. Keesom forces

1. Covalent forces2. Ionic forces3. Metallic forces

Intermolecular forces are the forces of attractions that exist between atoms within a molecule

Intramolecular forces are the forces of attractions that exist between molecules in a compound

Stro

nges

t

Stro

nges

t

Stro

nges

t

Figure 2-11 Schematic representation of the intramolecular and intermolecular forces and their compounds

2.4.3. Surface tension and surface free energy

Capillarity concerns interfaces that are sufficiently mobile to assume an equilibrium shape. The most common examples are meniscuses and drops formed by liquids in air or in another liquid. The basic equation of capillarity was given in 1805 by Young and Laplace. In general, it is necessary to have two radii of curvature to describe a curved surface, R1 and R2, where in mechanical equilibrium the fundamental equation of capillarity is the following:

Equation 2-1

+⋅γ=∆

21 R1

R1

P

where ∆P is the pressure difference across the fluid interface, ? is the surface tension and R1 and R2 are the principal radii of curvature.

Although referred to as a free energy per unit area, surface tension may equally be thought as a force per unit length. The corresponding SI units are joules per square meter (J/m2) or newtons per meter (N/m). Equilibrium surfaces may be treated mathematically, using either the concept of surface tension or the equivalent concept of surface free energy. The first term will be used to define the surface tension of a liquid, while surface free energy will be referred to solids.

Solid surface free energy, SFE, can be estimated using different independent approaches, i.e. direct force measurement, contact angles, capillary penetration into columns of

Chapter 2

85

powders, sedimentation of particles, etc., the contact angle, CA, is believed to be the simplest and hence widely used [35].

2.4.3.1. Young’s equation

Thomas Young proposed contact angle of a liquid as mechanical equilibrium of the drop resting on a plane solid surface at the three-phase boundary (Figure 2-12). The starting point for the calculation of interfacial energetic of solids by wettability measurement is the well known Young equation, one of the basic equations of capillarity and surface science [39;79;81]:

Equation 2-2 θ⋅γ+γ=γ cosLVSLSV

where θ is the contact angle, γSV is the surface free energy of the solid, γLV is the surface tension of the liquid and γSL is the interfacial free energy between the liquid and the solid.

Figure 2-12 Contact angle ?, surface free energy of the solid γSV, surface tension of the liquid γLV and

interfacial free energy between the liquid and the solid γSL [35]

The Young equation, Equation 2-2, relates the contact angle formed by a droplet on a substrate to interfacial free energy at the solid-vapour (SV), solid-liquid (SL) and liquid-vapour (LV) interface. Normally the solid-vapour (γSV) and liquid-vapour (γLV) interfacial free energy correspond to the intrinsic solid (γS) or liquid (γL) surface free energy [35;39;82;83].

It must be appreciated that only two of the four variables contained in Equation 2-2 can be directly measured: the surface tension of the liquid phase γLV, which is obtained through several different methods such as pendant drop, [79], and the contact angle θ of the liquid on the solid surface. The calculation of the solid SFE requires a further equation that links the interfacial free energy γSL to the other variables.

Different approaches to interfacial thermodynamics have arisen; among them three approaches will be presented:

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

86

2.4.3.2. The Neumann equation of state

The starting point of the Neumann theory is the demonstration of the existence of a function of the type:

Equation 2-3 ),(f LSSL γγ=γ

An empirical approach was used to obtain the relationship described in Equation 2-3 whose combination with the Young equation resulted in:

Equation 2-4 )1)(015.0(

)()2015.0(cos 2/1

LSL

L2/1

LSS

−γγγγ+γγ⋅−γ

=θ

Neumann theory is based, among other assumptions, on the impossibility of negative interfacial tensions and on the impossibility of dividing surface and interfacial free energies into components, each of them describing a particular intermolecular force [39].

From the practical point of view, it must be noted that Neumann’s equation of state approach only requires a single contact angle measurement, only one liquid, to obtain the quantitative value of the surface and interfacial free energy. Many papers have been published where the equation of state approach is questioned or refuted on experimental or rigorous theoretical grounds [39;84].

2.4.3.3. The polar-dispersion approach

If two liquids or a liquid and a solid are in contact, there are interactions at the interface which are due to different microscopic phenomena. In compliance with the fundamental interaction forces, the energy could be divided into different contributions.

Consequently, a completely different approach to interfacial interactions was introduced by Fowkes, [39], who suggested dividing surface tension into different components and to consider the total surface tension as a sum of independent terms, each representing a particular intermolecular force. In case of water, the main components arise from van der Waals, London or ‘dispersion’ forces (d) and the non-dispersive or polar (p) part mainly due to hydrogen bonding.

Equation 2-5 pS

dSS γ+γ=γ

where γSd is the dispersive part of the solid SFE and γSP is the polar part of the solid SFE.

Chapter 2

87

In case of dispersive forces, the term can be calculated by a geometric mean approach. Several researches extended the geometric mean approach including the polar part. These attempts resulted in the so called Owens and Wendt equation [85]:

Equation 2-6 ))()((2)cos1( 2/1pS

pL

2/1dS

dLL γγ+γγ=θ+γ

where γLd is the dispersive part of the liquid surface tension and γLP is the polar part of the liquid surface tension.

2.4.3.4. The electron donor-electron acceptor/ Acid-Base/ Good-van Oss

approach

Recently, a clear distinction between non-polar or dispersive Lifshitz-van der Waals interactions comprising the dispersion London forces; orientation Keesom forces and the induction Debye interactions in the condensed state on the one hand, and polar interactions on the other hand has been introduced [81;84]. The polar interactions under consideration are of the hydrogen-bonding type and can be designated as (Lewis) acid/base (AB) or electron-acceptor/electron-donor interactions, respectively. In the previously described polar-dispersion approach, London forces are contained in the γsd component, while the γSP component accounts for the other dipolar interactions. In the most recent theories on SFE components, the dispersive components is either still called γsd, with the understanding that this term contains all interactions arising from non-permanent dipoles, or γsLW, where LW stands for Lifshitz-van der Waals [35]. The γsP component is also named γsAB corresponding to acid/base interactions.

Acidic sites, in the Lewis sense, are all those sites that can act as electron acceptors. Lewis bases, on the other hand, are electron donors. Many compounds contain both acidic and basic sites: water, for instance, is a self-associated compound because of the balanced basicity of the oxygens and the acidity of hydrogens. It is correct to define water as ‘polar’ compound, because of the polarity of the –OH bond.

One of the most recent theories is the Acid-Base or Good-van Oss approach [35;39;86;87]. In this approach the dispersive component contains the interactions from instantaneous dipoles, γsLW, and the polar component is divided into an acid (electron acceptor γs+) and a basic (electron donor γs-) component [88].

Equation 2-7 ABS

LWSS γ+γ=γ

Equation 2-8 −+γγ=γ SSABS 2

Equation 2-9 2/1SL

2/1SL

2/1LWS

LWLL )(2)(2)(2)cos1( +−−+ γγ⋅+γγ⋅+γγ⋅=θ+γ

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

88

where γsLW is due to the interactions arising from induced dipoles (Lifshitz-van der Waals (LW)), γs+ is the Lewis-acid component and γs- is the Lewis-base component [86;89;90] of a solid (S) or a liquid (L) as indicates the subindex.

From the practical point of view, the LW and the electron donor and electron acceptor parameters of solid surfaces can be calculated by contact angle measurement with liquids of known values of the different components of surface tensions, through a system of three equations and three unknowns. Thus, contact angle measurement with, at least three liquids, is required. The approximation will be accurate if the surface tension and components of the liquids are significant different. If they are similar, the calculated values of the three parameters for the solid will be sensitive to small errors in the values of the parameters of the liquid, and in the measured contact angles [91].

The electron donor-electron acceptor approach is, up to now, the most advanced theory and the one that uses the presently accepted physical knowledge to account for interfacial interactions.

2.4.3.5. Surface roughness and surface heterogeneity effect on droplet

wetting

As explained in section 2.4.3.1, the equilibrium contact angle of a liquid drop on a solid surface is linked to the surface free energy of the solid trough the Young equation (Equation 2-2). The Young’s equation rests on several hypotheses, among them it is required that the surface is chemically homogeneous and smooth. Real solids are rough and heterogeneous and thus, their wettability will be significantly influenced by the geometrical structure of the roughness. Wenzel and later Cassie and Baxter studied the wetting of rough substrates. Wenzel, [92], assumed that the liquid filled up the grooves on the rough surface and generalized Young’s equation to obtain the intrinsic or Young contact angle (θi):

Equation 2-10 )cos(r)cos( iθ⋅=θ

where r is the ‘roughness factor’ and defined as the ratio of the actual area of a rough surface to a geometric, projected area on the horizontal plane; also called ‘index area'. Since the roughness factor is always greater than 1, the effect of roughness results in the improvement of the wetting for θ<90º but enhances hydrophobicity for θ>90º.

Cassie and Baxter, [93], considered the wettability of a composite surface, porous surfaces, composed two types of homogeneous patches that have different solid-fluid interfacial tensions. Then, the intrinsic contact angle is:

Equation 2-11 2211i cosfcosfcos θ⋅+θ⋅=θ

Chapter 2

89

where fi and θi represent the surface area fraction and the contact angle of patch i, respectively. When the surface is rough but not porous, f2 is zero and Equation 2-11 educes to Wenzel’s equation Equation 2-10 for the apparent Ca of a rough surface with the roughness factor f1 [93].

Figure 2-13 Schematic of the liquid droplets on rough surfaces: (a) Wenzel stat, the liquids fills the entire

surface; (b) Cassie-Baxter situation, the liquids forms a line of contact on the rough surface with air trapped below the contact line [94]

To decide which theory should be employed when a criteria is that the wetting state corresponds to the one with a lower free energy. Thus in general, the Wenzel theory prevails for θ<90º, while the Cassie-Baxter state dominates for θ>90º [95;96]. Topographical features of the type obtained in grit-blasted surfaces with contact angles below 90º results, in rough surfaces, under a wetting state in the Wenzel regime.

2.4.4. Materials

Different smooth and rough (S3, S6, S9, A3, A6 and A9) titanium surfaces prepared as explained previously in section 2.3.3.1 were studied. Also, each of the series was sterilized by different treatments (WS, SA, EO, GR) as explained in section 2.3.3.2.

Contact angles have been measured with three different liquids on each material (Table 2-2): ultrapure distilled water (MilliQ) [97], formamide [87], which are both polar liquids, and di-iodomethane [97], which is a non-polar liquid; following recommendations by [98].

Table 2-2 Surface tension and their components for the three liquids used in static contact angle measurements [87;97]

Liquid γL (mN/m) γLLW(mN/m) γL

+(mN/m) γL–(mN/m)

Ultrapure distilled water 72,80 21,80 25,50 25,50

Formamide 58,00 39,00 2,28 39,60

Di-iodomethane 50,80 50,80 0 0

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

90

At least three samples of each group were analysed. For each group, 20 to 40 measurements were carried out.

2.4.5. Methods

2.4.5.1. Wettability and surface free energy

Small drops or bubbles tend to be spherical because surface forces depend on the area, which decreases as the square of the linear dimension, whereas distortions due to gravitational effects depend on the volume, which decreases as the cube of linear dimension [99]. When gravitational and surface tensional effects are comparable, then one can determine the surface tension from measurements of the shape of the drop or bubble Figure 2-14.

Figure 2-14 Shapes of sessile and pendant liquid drops [209]

The general procedure is to form the drop or bubble under correct conditions such as, not to be object of disturbances, and make certain measurements of its profile from a photograph obtained with the help of a digital camera.

The sessile drop method involves the deposition of a drop of liquid on a horizontal surface and observing it in cross section. The angle formed between the liquid/solid interface and the liquid/vapor interface is the contact angle which is determined optically (Figure 2-12). Commercially available systems, often called goniometers, use the sessile drop method. However, goniometers typically require that an operator optically dial in a measurement and read the contact angle; this process will introduce errors of approximately 1°–2°. The sessile drop method can accommodate more varied surfaces and geometries, and its accuracy can be improved by increasing the number of observations through automated image processing techniques.

Chapter 2

91

2.4.5.1.1. Test conditions

The most widely used method for measuring contact angle, θ, is sessile drop. Imaging the sessile drop through a microscope fitted with a goniometer scale, allows the direct measurement of the angle. The contact angle may be obtained from a photograph of the drop profile either by measuring the angle or by calculating it from the entire drop profile.

The apparent static contact angles, θ, of three reference liquids were measured with the Sessile Drop method. Drops were generated with a micrometric syringe and deposited on the substrate surface. The wettability studies were performed with a contact angle video based system (Contact Angle System OCA15plus, Dataphysics, Germany) and analysed with the SCA20 software (Dataphysics, Germany). To account for the effect of surface roughness on surface wettability the so-called Wenzel equation (Equation 2-10) was applied.

Figure 2-15 Contact Angle System OCA15plus, Dataphysics, Germany

The surface free energies of the different samples were calculated using the Acid-Base or Good-van Oss. The Acid-Base model gives the long-range dispersion (Lifshitz-van der Waals) (γsLW) and the polar acid-base (γsAB) components, the latter divided into two parts, acidic (γs+) and basic (γs-) [100] according to Equation 2-9. For comparative reasons between the different grit-blasted surfaces, the total SFE and its LW, acid and base components for all series were calculated after apparent contact angles were determined and intrinsic contact angles calculated from Equation 2-10.

For both methods the spreading pressure was not taken into account. This pressure gives the contribution to SFE of the adsorption of an external layer from the atmosphere and

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

92

has to be calculated for SFE calculation correction if SFE is higher than 60 mJ/m2 [43]. In the present study, SFE values are lower than this limit and the spreading pressure can be neglected.

A series of at least four drops (3 µl/drop) were used in five trials, for a total of 15 per liquid on a particular c.p.Ti surface using the Sessile Drop method. CA’s were determined semi-manually from the image of the drop with an accuracy of ± 1º. The measurements were obtained at room temperature (T=25 ºC) in an environmental chamber previously saturated with the liquid of study. Samples were ultrasonically cleaned for 20 min, firstly in acetone, secondly in 70% V/V ethanol before a final rinsing with distilled water before air drying for all non-sterilized series.

2.4.5.2. Time-related contact angle

2.4.5.2.1. Test conditions

Time-related contact angle measurements (Time-CA) have been also performed, which allows monitoring the changes in the wetting of the surface during time. According to this, an additional objective of this work is to characterize the influence of protein adsorption on the wettability of the studied materials by using different culture media as testing fluid in Time-CA tests.

After the deposition of the drop onto the surface, the contact angle can continuously change as a function of time depending on the surface and the liquid of study. The study of the time-related contact angles (Time-CA) for water and two organic liquids, one of them containing proteins, has been performed to mainly study the influence of the organic molecules, especially proteins, on the wettability of the surface.

Time-CA tests were done with three different fluids: (1) ultrapure distilled water (MilliQ, Millipore, Germany); (2) Dulbecco’s Modified Eagle Medium (DMEM); and (3) DMEM supplemented with a 10% of Foetal Calf Serum (FCS) on Smooth, S3, S6, A3 and A6 surfaces.

The pendant drop method was applied to study the liquid surface tensions using the same apparatus and software. The surface tension values were taken when no further time variation was detected. All the measurements were carried out at 25 ºC. The equilibrium values for the surface tension of DMEM was 44,9 ± 1,6 mN/m.

Chapter 2

93

2.4.5.3. Statistical analysis

For the statistical study of the obtained results ANOVA-Tables with multiple comparison Fisher’s test were used to determine statistically significant (p-value<0.05) differences between the means of the different groups.

Two experimental designs were carried out to find out whether the effects of the different properties (sterilization, particle size and particle nature) are statistically significant in the surface energy response (instrinsic CA, SFE and their components: Lifshitz-van der Waals, acid and base). The two experimental designs were carried out to separate the analysis of the smooth-samples from the rough-ones. Sterilization is the only property studied in the experimental design performed with the smooth-samples, whereas the effect of the size and the nature of the abrasive particles were added in the experimental design with rough-samples. A third analysis considering non-sterilized rough samples was performed.

Table 2-3 Factors (properties) of study, levels, and levels codes for roughened samples resulting a 41·22 experimental design. Sterilization is the only factor analysed in the case of smooth samples, 41 experimental

design

Factors Level code Level

Grit-blasting particle size (C-factor)

– 300 µm

+ 600 µm

Grit-blasting particle nature (D-factor)

– SiC

+ Al2O3

Sterilization technique (X -factor)

– – Without sterilization

– + Steam autoclave

+ – Ethylene oxide

+ + Gamma radiation

All different factors (properties), levels, and codes for the levels used in the experimental design with rough-samples are shown in Table 2-3. Factors are expressed in a mixed-factorial design with a one four-level factor, accounting for sterilization effect (X.-factor), and two two-level factors, accounting for the effects of the nature (D-factor) and the size (C-factor) of the abrasive particles, i.e., a 41·22 mixed-design. The objective is to estimate all the main effects and the interaction effects involving these variables. This can be done with 16 runs if the 41-design is converted to a 22-design (A- and B- factors) (Table 2-4), which results in an easily-defined 24-design [101]. In the first step, this design is analysed and solved as a 24 experimental-design, calculating the sums of squares for the A-, B-, C-,

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

94

and D-factors. In the second step, the sums of squares for the three main factors (X, C, and D) and their interactions are solved following Montgomery et al.[101]:

Equation 2-12 SSX = SSA+SSB+SSAB (3 degrees of freedom) SSC = SSC (1 degrees of freedom) SSD = SSD (1 degrees of freedom) SSCD = SSCD (1 degrees of freedom) SSXC = SSAC+SSBC+SSABC (3 degrees of freedom) SSXD = SSAD+SSBD+SSABD (3 degrees of freedom) SSXCD = SSACD+SSBCD+SSABCD (3 degrees of freedom)

An analysis of variance is performed to later obtain the regression equation of the model.

Table 2-4 Experimental design of a four-level factor expressed as two two -level factors

Two-Level Factors Four-Level Factor

A B X

– – x1

+ – x2

– + x3

+ + x4

2.4.6. Results

2.4.6.1. Wettability and surface free energy

2.4.6.1.1. Smooth surfaces

The influence of sterilization process is introduced by studying the results obtained with smooth samples. Figure 2-16 confirms the increase of hydrophobicity of ethylene oxide, EO, and steam autoclaving, SA, sterilization procedures founding statistical differences to non-sterilized and gamma radiation, GR, surfaces. GR series present a statistical equal CA respect to control surfaces (non-sterilized).

The SFE and their components for smooth series with all the sterilization treatments applied are shown in Figure 2-17. Higher values of SFE are calculated from lower values of CA, i.e., from more hydrophilic surfaces, e.g. WS and GR, to the most hydrophobic surfaces, as for SA- and EO-treated surfaces (Figure 2-16 and Figure 2-17). For all studied surfaces, the samples have, mainly a dispersive character. Statistically-significant differences in electron-donor/acceptor character between the smooth surfaces sterilized

Chapter 2

95

with different processes have been determined. WS and GR are electron-donor surfaces, whereas SA and EO have a bipolar character. This differences are caused by the base-Lewis (electron donor) component of the surface free energy because the acid-Lewis and the dispersive components do not significantly change whatever the sterilization process applied (Figure 2-17).

Surface CA (º) ± SD Smooth-GR 65,5 ± 7 Smooth-WS 66,3 ± 5 Smooth-EO 86,2 ± 2 Smooth-SA 95,3 ± 2

GR WS EO SA0

20

40

60

80

100

CA

(º)

Smooth Figure 2-16 CA of smooth Ti surfaces with different sterilisation processes. Bar joins CA mean-values with

non-statistically significant differences

SA EO WS GR

0

10

20

30

40

50

60

SFE

(m

J/m

2)

Smooth

SFE LW Acid Base

Figure 2-17 SFE and their different components (Lifshitz-van der Waals, acid-Lewis and base-Lewis) of

smooth Ti surfaces with different sterilisation treatments. Bar joins SFE mean-values with non-statistically significant differences

Table 2-5 Regression model of smooth, sterilized titanium surfaces obtained by experimental factor design. Study of method of sterilization, ST, contribution on CA, SFE and its components

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

96

Response of smooth series

Regression model equation

Contribution of significant factors (SSfactor /SStotal×l00)

CA (º) Y =78,9-11,2ST ST: 92,8%

SFE (mJ/m2) Y = 36,9+4,6ST ST: 85,9% LW (mJ/m2) Y = 32,8+1,1ST ST: 53,8%

Acid (mJ/m2) Y = 1,1+0,7ST ST: 88,6% Base Y = 7,7+4,0ST ST: 82,7%

The regression equation of the smooth series and the contribution of significant factors, are shown in Table 2-5. The obtained results confirm the influence of the sterilization process on the wettability and SFE response of smooth series. Specifically, sterilization process contributes a 92,8 and a 85,9% to the change of CA and SFE, respectively, of smooth and sterilized samples.

2.4.6.1.2. Rough surfaces

The influence of size and nature of abrasive particles on wettability and SFE of grit-blasted rough Ti surfaces is shown in Table 2-6 and Figure 2-18 pectively. The apparent CA, and their intrinsic values, iCA, obtained with the three working liquids for all tested surfaces are shown in Table 2-6. Overall, the grit-blasting treatment decreased surface wettability with water i.e. increased contact angles. This effect was more pronounced for those surfaces grit-blasted with Al2O3. For the other studied liquids, di-iodomethane and formamide, no relevant tendencies were found for both apparent and intrinsic contact angles. Indeed, those results are reflected at the analysis of SFE and its components (Figure 2-18).

Table 2-6 Apparent, CA, and intrinsic, iCA, contact angles for the three liquids used on the different c.p. Ti surfaces. Values are mean ± standard deviation. Statistical differences vs. smooth surfaces for each

column are indicated by single and double asterisk-symbols (p<0,05)

Surface Water Di-iodomethane Formamide

Apparent CA (º)

Intrinsic iCA (º)

Apparent CA (º)

Intrinsic iCA (º)

Apparent CA (º)

Intrinsic iCA (º)

Smooth 66,3 ± 5 66,3 ± 5 51,5 ± 3 * 51,5 ± 3 * 51,8 ± 1 51,8 ± 1 *

S3 66,8 ± 7 72,4 ± 5 * 38,5 ± 4 ** 53,1 ± 3 * 35,0 ± 5 * 51,0 ± 3 *

S6 63,3 ± 7 70,9 ± 5 * 44,5 ± 1 58,7 ± 1 ** 44,5 ± 3 58,7 ± 2 **

S9 77,4 ± 3 * 81,8 ± 2 ** 39,5 ± 2 ** 59,5 ± 1 ** 38,5 ± 2 * 59,0 ± 1 **

A3 76,9 ± 3 * 79,7 ± 2 ** 52,4 ± 1 * 61,3 ± 1 ** 35,2 ± 5 * 50,1 ± 3 * A6 75,4 ± 5 * 79,8 ± 4 ** 58,2 ± 2 68,3 ± 1 59,3 ± 2 69,0 ± 2

A9 82,1 ± 5 * 85,7 ± 3 37,3 ± 3 ** 63,5 ± 1 35,7 ± 1 * 62,8 ± 1

Chapter 2

97

Moreover the calculated SFE values and their different components following the Acid-Base approach (section 2.4.3.4) from the intrinsic CA are represented in Figure 2-18. The higher the roughness the lower the SFE for each particular type of blasting particles used, with statistically significant differences, respect to smooth surfaces. However, no apparent correlation between roughness and the components of the SFE was determined.

Smooth S3 S6 S9 A3 A6 A90

5

10

15

20

25

30

35

40

45

50

SF

E (

mJ/

m2)

SFE LW Acid Base

Surface Surface free energy (mJ/m2)

Total Lifshitz-van der Waals, ? sLW Acid, ? s+ Base, ? s-

Smooth 39,4 ± 1 * 33,4 ± 1 * 0,6 ± 1 * 16,3 ± 5 * S3 38,9 ± 2 * 32,5 ± 1 * 1,3 ± 1 ** 09,4 ± 5

S6 34,7 ± 1 ** 29,4 ± 1 ** 0,5 ± 1 * 15,1 ± 4 *,•

S9 33,6 ± 1 ** 28,9 ± 1 ** 1,3 ± 1 ** 04,3 ± 1 ** A3 34,1 ± 1 ** 27,8 ± 1 **,• 3,7 ± 1 03,1 ± 2 **

A6 27,6 ± 1 23,8 ± 1 0,4 ± 1 * 11,6 ± 5 • A9 30,4 ± 1 26,6 ± 1 • 1,4 ± 1 ** 02,8 ± 2 **

Figure 2-18 Surface free energy and their different components (Lifshitz-van der Waals, acid-Lewis and base-Lewis) of non-sterilised smooth and rough Ti. Bar joins SFE mean-values with non-statistically

significant differences. Statistical differences for each column are indicated by point, single-asterisk and double-asterisk symbols (p<0,05)

If comparing samples blasted with the same particle size, a statistically significant effect of the chemical nature of the particles used on SFE and the Lifshitz-van der Waals component was determined. That is, grit-blasting with SiC-particles results in a higher SFE of the c.p. Ti surfaces than grit-blasting with Al2O3-particles (Figure 2-19, right). Moreover the increasing particle size leads to a decrease of the SFE (Figure 2-19, left).

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

98

300um 600um 900um0

5

10

15

20

25

30

35

40

45

50

SiC Al2O3

SF

E (

mJ/

m2 )

SiC Al2O30

5

10

15

20

25

30

35

40

45

50

300um 600um 900um

SF

E (

mJ/

m2 )

Figure 2-19 Effect of particle size, left, and nature, right, on SFE of rough and non-sterilized Ti surfaces

The statistical study by using an experimental design allowed separating and studying the influence of different surface factors on the wettability and surface free energy of the solid. The results of the experimental design study are shown in Table 2-7, for rough and non-sterilized series, and in Figure 2-18 for rough and sterilized series. Focusing the analysis on rough and non-sterilized surfaces, Table 2-7, a major effect, 62,3%, of chemical nature of the grit particles, PN, on the intrinsic water CA was determined, but the size of the particles, PS, had no effect. However, both PS and PN contributed to the values of SFE with a 30,5 and 62,4%, respectively. The LW component of SFE was influenced mainly by the particle nature while Acid and Base components were affected mostly by the particle size.

Secondly, it was also considered a regression model for rough and sterilized samples estimating all the parameters that can influence the SFE and their interactions (Figure 2-18). The main factor, for rough surfaces, on SFE results is the particle nature of the grit-blasting particles, achieving a contribution of nearly the 52% while the sterilization process was a 32% and the particle size only contributed a 6%. There is no significant effect for the interaction between the working parameters. Lifsithz-van der Waals component of SFE was mainly influenced by PN compared to the other factors; whereas no clear effect was found for Acid and Base components. Figure 2-20 confirms that the effect of sterilization on smooth samples is the same than for rough samples: after SA and EO treatments surfaces are more hydrophobic compared to GR and non-sterilized surfaces.

Chapter 2

99

Table 2-7 Regression model of rough and non- sterilized series obtained by experimental factor design. Study of particle nature (PN) and size (PS) contribution on iCA, SFE and its components

Response of smooth series

Regression model equation Contribution of significant factors (SSfactor /SStotal×l00)

iCA (º) Y = 76,4-76,4PN+0,8PS/PN PN: 62,3% Interaction PS/PN: 26,0%

SFE (mJ/m2) Y = 34,3+2,6PS-34,3PN PS: 30,5% PN: 62,4%

LW (mJ/m2) Y = 28,3+2,4PS-28,3PN PS: 29,4%, PN: 67,4%

Acid (mJ/m2) Y = 1,4-0,3PS-1,4PN+0,8PS/PN PS: 84,5% PN: 9,4% Interaction PS/PN: 4,0%

Base (mJ/m2) Y = 9,2+2,6PS-9,2PN-3,5PS/PN PS: 50,3% PN: 28,6% Interaction PS/PN: 13,4%

Table 2-8 Regression model of rough and sterilized series obtained by experimental factor design. Study of sterilization (ST), particle nature (PN) and size (PS) contribution on iCA, SFE and its components

Response of rough series

Regression model equation Contribution of significant factors (SSfactor /SStotal×l00)

iCA (º) Y = 80,2-6,3ST+2,8PS-4,1PN-4,1ST/PN ST: 26,3% PS: 7,1% PN: 15,1% Interaction PS/PN: 21,0%

SFE (mJ/m2) Y = 40,6+6,3ST-1,6PS+4,8PN ST: 33,2% PS: 6,0% PN: 51,6%

LW (mJ/m2) Y = 53,7+5,8ST-1,6PS+4,3PN ST: 29,4% PS: 5,7% PN: 38,5%

Acid (mJ/m2) Y = 0,3-0,2ST-0,2PN-0,2PS ST: 15,6% PS: 9,6% PN:9,2% Interaction ST/PS: 16,5% Interaction ST/PN: 18,0% Interaction PS/PN: 4,4% Int. ST/PS/PN: 18,0%

Base (mJ/m2) No significant factors

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

100

SA EO WS GR0

20

40

60

80

100

120

S3 S6 A3 A6 Smooth

intr

insi

c C

A(º

)

Figure 2-20 Effect of sterilization treatment on intrinsic water CA of different smooth and rough Ti

surfaces

2.4.6.2. Time-related contact angle

Time-CA results show abrupt changes in the wettability of the samples, mainly on the blasted ones (Figure 2-21). Time-CA decreased with time in all surfaces with all the fluids used i.e., in all cases the surfaces became more hydrophilic. This effect was significantly greater with the organic liquid containing proteins. Figure 2-22 shows Time-CA plots for smooth and grit-blasted surfaces with all the studied liquids in this section: water, DMEM and complete culture medium (DMEM+10%FCS). The differences between the initial and the final CA in Time-CA plots are higher when using protein containing medium compared to the others studied liquids (Figure 2-23). Generally, the increasing roughness led to an increase of the Initial CA-Final CA difference for all working liquids.

Chapter 2

101

-1 0 1 2 3 4 5 6 7 8 9 10 11 12 1350

55

60

65

70

75

80

85

DMEM+10%FCS

DMEM

Water

CA

(º)

Time(s)

a) Smooth

0 5 10 15 20 25 30

55

60

65

70

75

80

DMEM+10%FCS

DMEM

Water

CA

(º)

Time (s)

b) S6

0 5 10 15 20 25 30

70

75

80

85

DMEM+10%FCS

DMEM

Water

CA

(º)

Time (s)

c) A6

Figure 2-21 Evolution of CA with time for smooth, S6 and A6 on non- sterilized samples using water, DMEM and culture medium (DMEM+10%FCS) as the liquids of study. Arrows point sudden changes of

wettability that could indicate either changes in conformation of the adsorbed proteins or evolution of metastable states of the drop

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

102

Water DMEM DMEM+10%FCS

Smoo

th

0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (º

)

Time (s) 0 5 10 15 20

40

50

60

70

80

90

100

110

120

Tim

e-C

A (

º)

Time (s) 0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (

º)

Time (s)

S3

0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (º

)

Time (s) 0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (

º)

Time (s) 0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (

º)

Time (s)

S6

0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (

º)

Time (s) 0 5 10 15 20

40

50

60

70

80

90

100

110

120

Tim

e-C

A (º

)

Time (s) 0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (

º)

Time (s)

A3

0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (

º)

Time (s) 0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (

º)

Time (s) 0 5 10 15 20

40

50

60

70

80

90

100

110

120

Tim

e-C

A (º

)

Time (s)

A6

0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (º

)

Time (s) 0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (º

)

Time (s) 0 5 10 15 2040

50

60

70

80

90

100

110

120

Tim

e-C

A (º

)

Time (s)

Figure 2-22 Time-CA plots of droplets of water, DMEM, DMEM+10%FCS on smooth and rough non-sterilized samples

Chapter 2

103

Smooth S3 S6 A3 A60

5

10

15

20

(Ini

tial C

A-F

inal

CA

) (º

)

Water DMEM DMEM+10% FCS

Figure 2-23 Initial–Final CA difference obtained by time-related CA measurements with different working

liquids on smooth and grit-blasted samples

2.4.7. Discussion

Concerning to the wettability and surface free energy results some cautions must be considered since the surface finishes tested here are far from being ideal surfaces to fulfil conditions to directly apply Young’s equation (Equation 2-2). On rough surfaces two main effects occur. The first one affects to contact angles/wettability since for the same nominal area; the total real area is higher for rougher surfaces. That means that if the surface were ideally plane and nothing else than roughness would change on that surface, the hydrophobic/hydrophilic character of the surface will be reinforced. This effect on the contact angles that are measured is taken into consideration by applying the Wenzel correction factor and so, obtaining the intrinsic contact angle of each working series from the apparent contact angles measure directly from the contact angle meter. The second main effect is contact-angle hysteresis. Roughness induces a range of metastable states [102] separated by free energy barriers. Then depending on the height of the energetic barriers and the macroscopic vibrational energy of the droplet, the CA on rough surfaces can vary continuously between two extremes: the highest allowed value is the advancing contact angle; the lowest is termed the receding contact angle. Thus, a smooth surface can be defined by a single or equilibrium Young angle while a rough surface shows a continuum of allowed values. Then, it is difficult to determine which is the true equilibrium apparent contact angle, CA, to be corrected by the roughness factor obtaining the intrinsic CA, iCA, and placed in the Young equation to calculate the SFE. In our case, the apparent contact angles were measured after the droplet was set on the surface to account for the first advancing angle. This process is an attempt trying to avoid the dynamics of the drop due to surface roughness.

The wettability and surface free energy results calculated by taking into account the previously mentioned considerations are used in terms of comparison between the

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

104

different surface finishes. Even though theoretical limitations do not allow calculations of SFE on rough surfaces that are far away from ideal smooth and homogeneous conditions, we have considered using the iCA to gain some insight in the energetic state of that surfaces, being aware that the results must be analysed and discussed carefully and be regarded as an approach.

The analysis and discussion of the obtained results will be performed separating the smooth and rough surfaces. First, the influence of the sterilization process is introduced by studying the results obtained with the polished samples. The second step includes the analysis of the influence of the size and the nature of the abrasive particles on the blasted-rough titanium surfaces. This separated analysis is because roughness has a significant influence on the final drop shape obtained that is not only explained by the energetic properties of the surface but also because the drop cannot freely spread on the surface, reaching various metastable states.

For smooth surfaces the analysis of the SFE can follow Young’s equation without any correction factor since metastable states produced in grit-blasted surfaces when the drop of the liquid can be temporary or definitively retained by the topographical features of the surface are not produced. The statistically significant differences in SFE depending on the sterilization process where due to the difference in the Lewis-basic component, i.e., in the electron-acceptor character of the surface. SA and EO sterilization methods showed a dual character whereas WS and GR showed a clear electron-acceptor character (Figure 2-16). The chemical contamination on EO surfaces and the increase of the thickness of the TiO2 surface layer leads to the decrease in the hydrophilicity of the EO and SA samples as reported [48;103]. This could explain the relatively high contact angles (sometimes greater than 90º) for the surfaces produced by SA and EO. While EO samples become more hydrophobic due mainly to the increase in the carbon contamination, as observed from XPS results (section 2.3.5.3, Figure 2-8), the GR samples were highly hydrophilic and less contaminated. A reason for such behaviour found for gamma radiated series could be the activation of hydroxyls groups at the surface, a similar process that occurs with plasma cleaning. As reported previously, [104], the clean surface has a high surface energy while the contaminated one has a lower surface energy.

According to the Wenzel approach [103], an increase in the surface roughness should lead to a decrease in the values of the contact angles < 90º due to the increase in the real surface area. According to our findings a significant difference in the effect of roughness on wettability is reached if roughness of the surface is in the nano- or in the micro-level, and an increase of surface hydrophobicity is obtained when testing a micro-rough surface (A3, A6, S3, and S6). This is because of the metastability states that reach the drop on these surfaces, as explained before.

Surface energy is an indicator of potential cellular adhesion. In general, materials with SFE values from 20 up to 30 mJ/m2 exhibit minimal cell adhesion, whereas materials

Chapter 2

105

with surface energies above these values tend to present a greater bioadhesion [37]. Generally, the titanium surfaces studied here have SFE values above 30mJ/m2, and among them, the rough SiC-blasted surfaces present the highest surface energy, which may indicate a better biological adhesion and then probably a better differentiation (Figure 2-18 and Figure 2-1, right). However, previous in vitro studies performed in our group, [10], showed non-significant differences in the osteoblasts differentiation between SiC and Al2O3 grit-blasted surfaces. Nevertheless, the SiC-particles had mean ALP–activity higher than those blasted with alumina. This suggests that there is a different influence in cell response depending on the material of the grit, which could be related to differences in the SFE because of most likely differences in the chemical surface heterogeneities. In fact, alumina-blasted and SiC-blasted titanium surfaces have similar topographies, as confirmed in this work in section 2.3. The experimental design analysis confirmed those speculations because the chemical composition o fthe grit-blasting particles and the sterilization method were the major effects explaining differences in the surface energy of the studied titanium surfaces.

Samples blasted with SiC particles have a higher SFE than the ones blasted with Al2O3. This may be attributed to the more significant contribution of the LW component to the value of SFE in the blasted samples because SiC has a significant higher LW component that Al2O3. [80] (Table 2-9). The influence of the particle nature was more relevant at the SFE and Ca response compared to the size of the grit-blasting particles as indicated by the statistical study by applying the experimental design. Consequently, changes in the chemical composition of the blasted and/or sterilized titanium surfaces are more influencing on its wettability than the topographical changes.

Table 2-9 Surface energy components of some mineral solids derived from immersion enthalpies into water, formamide, and n-heptane [80]

Solids LWSγ (mJ/m2) +γS (mJ/m2) −γS (mJ/m2) Sγ (mJ/m2)

SiC 218 30,6 30,8 279

Alumina 126 143 390 599

Wettability measurements not only provides contact angle values which are important parameters in adhesion and adsorption processes, but also allows the detection of minor changes in the surface composition or morphology. In particular, time-CA analysis may lead to information about the time-dependent phenomena occurring during the contact of a liquid to a solid. The dynamics of a liquid drop spreading spontaneously on a substrate is significantly influenced by the surface state of the substrate [80]. The former objective of studying time-related CA of the different titanium surfaces with organic and inorganic liquids was to elucidate the sensitivity of the technique to detect adsorption, desorption and adsorption/desorption protein processes.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

106

According to the random occurrence of the abrupt changes along the different surfaces at any time shown in time-CA plots, the changes might be attributed to processes of protein adsorption and/or desorption (Figure 2-21). However, the non-reproducibility of this occurrence and the random type of the changes lead to the conclusion that these variations are due to the different drop metastable states throughout the experiments. The local heterogeneities of the surface induced a stick and slip movement of the three phase line, which results in contact angle fluctuations in order to get the equilibrium.

Time-CA studies demonstrate that proteins, while adsorbing on titanium surfaces, change the wettability of the surface (Figure 2-22). All the series studied showed a decrease in contact angle values with time of contact with DMEM and completed culture-medium drops. This effect was strongest on DMEM+10%FCS protein containing liquid probably due to rearrangements of adsorbed proteins monolayers. It is probable that proteins adsorb on titanium with their more hydrophobic side facing the metal surface and exposing to the medium their more hydrophilic cues.

The initial contact angle is similar within the different liquids studied (Figure 2-23). This may indicate that water, which is the main component of all the liquids studied, provide the initial contact angle. An effect of surface roughness on the contact angle was previously reported, and an increased or decreased value with time depending on its initial value was showed [77]. However, as explained before, no statically significant differences in the initial contact angles have been determined in this study depending on the roughness of the substrate.

2.4.8. Conclusions

v For all the samples of study, smooth, rough and sterilized Ti surfaces, the surfaces had a prevalent dispersive character (Lifshitz-van der Waals (LW) component).

v Sterilization affected the energetic properties of titanium surfaces, with a relevant difference in their electron donor/acceptor character. Non-sterilized surfaces and gamma-radiated surfaces has electron-donor character, whereas steam-autoclaved and othelyne-oxide treated surfaces had bipolar character. The later so, were more hydrophilic.

v The chemical nature of the blasting particles had a significant effect on SFE of rough titanium surfaces. SiC-blasted surfaces had higher SFE than alumina-blasted ones.

v Significant difference in the effect of roughness on wettability is reached if roughness of the surface is in the nano- or in the micro-level, and an increase of surface hydrophobicity is obtained when testing surfaces with roughness in the microlevel.

Chapter 2

107

v Chemical composition of the grit-blasting particles and the sterilization method were the main factors affecting wettability and surface free energy of titanium surfaces.

v Time-CA decreased with time in all surfaces with all the fluids used i.e., in all cases the surfaces became more hydrophilic. This effect was significantly greater with the organic liquid which contained proteins.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

108

2.5. Surface charge properties

2.5.1. Surface charge and cell/protein response

Cell membrane is negatively charged and proteins have also a global charge at a specific pH so electrostatic interactions have a specific role on the biological response of implant materials. The air-formed passive layer of near stoichiometric TiO2 always present on titanium is thought to be responsible for the biocompatibility behaviour of titanium implants [105]. According to the literature, this oxide layer is of defective nature, showing n-type semi-conducting properties and has an amorphous structure [40]. Due to the short range order in amorphous structures, a capacity to incorporate electrolyte ions is expected, which may have consequences for the formation and composition of the electrochemical double layer at the oxide electrolyte interface. However, the potential distribution over the electrochemical double layer is considered to be of specific relevance for materials used in biological fluids [106].

As cell membrane is negatively charged, positively charged surfaces will be favoured by cell adhesion [42;107;108]; but also negatively charged surfaces have been found to favour osteoblast cells adhesion [108]. Any study has determined the optimal charge for cell adhesion processes, but titanium and hydroxyapatite, which are well-known biomaterials for osseointegration, have a net negative charge at physiological pH [109].

While the surface charge and surface potential cannot be measured, it is possible to experimentally determine the zeta potential, which depends on the concentration of electrolyte or ionic liquid phase. If the ionic strength of the solution is low, the zeta potential is a good approximation of the potential surface. In this way, the potential provides information on the electrical characteristics and behavior of the surface of the material studied and ions interacting with the surface

2.5.2. Electrokinetic effects

2.5.2.1. Electrochemical Double Layer and Zeta Potential

Most interfaces in aqueous solution are electrically charged. This charge can result from the association or dissociation of covalent bonds between groups of the surface and/or specific adsorption of ions of low molecular weight present in the aqueous solution [110].

Under equilibrium conditions, the electro-neutrality requires the balance between the surface charges and the net charge of the liquid solution next to it. This will form an

Chapter 2

109

electric double layer (EDL) at the interface solid / aqueous solution (Figure 2-24). EDL structure is described by the classical Gouy-Chapman Stern model in which the double layer is divided in an immobile layer, called the Stern layer, and a Gouy or diffuse layer. The immobile layer is further divided in an inner Helmholtz plane (IHP), which is contact with the surface, and an outer Helmholtz plane (OHP). Ions from the bulk liquid phase are specifically adsorbed on the solid surface by attractive interaction. These ions form the inner Helmholtz plane and their charge is compensated by counter ions which are partly arranged in the outer Helmholtz plane of the immobile layer and partly contained in the diffuse layer. The potential of the inner Helmholtz plane decreases linearly within the immobile layer. In the diffuse layer the potential is decreasing exponentially. A schematic representation of the EDL model is given in Figure 2-24 where the parameters characterizing its structure are marked [111].

At the boundary between the immobile layer and the diffuse layer defines the so-called slip plane. This plane is the boundary between the fixed charges of the solid and the mobile charges of the liquid phase. The configuration of the EDL allows defining a surface charge density and an electric potential at the slip plane. The zeta-potential value is the potential at the slip plane [111;112].

Distance x

IHP OHP Slip plane

0

Pot

entia

l

Ψ0

Ψs

Ψd

ζ

σ0 σs σd

Figure 2-24 Schematic representation of the electric double layer formed with an aqueous solution at a positively charged interface following the Gouy-Chapman Stern model. Surface (electric potential Ψ0;

charge density σ0), the inner Helholtz plane (IHP) or Stern plane (electric potential Ψ s; charge density σs) and the diffuse layer (electric potential Ψ d; and charge density σd). The potential at the slip plane is the

electrokinetic or zeta -potential, ζ [111;112]

2.5.2.2. Isoelectric point

The formation of the electrochemical double layer depends on the characteristics of the electrolyte solution and, among other properties, its pH. The iso-electric point is an important quantity that may be determined from the pH dependence of zeta potential. The isoelectric point, IEP, corresponds to the pH where zeta-potential becomes zero (ζ=0); i.e. the pH at which the net charge is zero [113].

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

110

2.5.2.3. Methods to determine electrokinetic potential

The application of an external force parallel to the solid/liquid interface leads to a relative motion between the solid and the liquid phases. The immobile layer remains on the solid surface, while the diffuse layer of the electrochemical double layer is sheared off and moves with the bulk liquid phase. The position of the shear plane (or slip plane) between the immobile and diffuse layers is determined by the equilibrium between the attractive surface forces and the external force (Figure 2-24). The electrokinetic or zeta potential, which is assigned to this shear plane, is therefore an indicator for the structure and mechanism of formation of the electrochemical double layer [111]. A modification of the liquid phase changes the structure of the electric double layer, which is reflected by the variation in zeta potential. The electrokinetic potential or zeta-potential (ζ) of the solid surface, of macromolecules, particles, or larger surfaces is the potential measured at the slipping plane by electrokinetic methods such as electrophoresis, electroosmosis, streaming potential, etc (Figure 2-25) [111].

SEDIMENTATION POTENTIAL

ZETA POTENTIAL

ELECTROPHORESIS ELECTRO-OSMOSIS

STREAMING POTENTIAL

Disperse phasemovesLiquid phase stationary

Disperse phasestationaryLiquid phase moves

Disperse phase stationaryLiquid phase in motion

Disperse phase in motionLiquid phase stationary

Applied electricfield

Applied force

Figure 2-25 Relationship between the four electrokinetic phenomena where an electric field can be applied (electrophoresis and electro-osmosis) or a force (streaming potential and sedimentation potential) to cause

movement of the liquid phase (streaming potential and electro-osmosis) or of the solid phase (electrophoresis and sedimentation potential) [114]

The streaming potential method is the most convenient method for determining the zeta-potential of macroscopic solid surfaces owing to its versatility that allows working with planar surfaces [115;116].

Chapter 2

111

2.5.2.4. Geometrical heterogeneity effects: surface roughness

It must be remembered that the position of the slip plane determines zeta potential value as seen in Figure 2-24. The slip plane along a solid surface is related to the structure of the liquid adjacent to a solid [111]. Then, it is important to know the exact shape and position of the slip plane to obtain the correct zeta potential measurements. For rough surfaces, the exact shape of the liquid velocity profile in the closest proximity of a solid surface is not known.

Geometrical morphology of the solid/liquid interface influences the factors determining electrokinetic phenomena, i.e., hydrodynamic flow and the electric field pattern. Then, surface roughness affects the electrokinetic slipping process. The liquid within cavities at rough surface is hydrodynamically immobilized as simulated by Schnitzer et al. [117].

Despite of the roughness influence on zeta potential measurements, there is no influence on the isoelectric point or algebraic sign of zeta potential in corresponding pH regimes [117].

2.5.3. Materials

Different smooth and rough (Smooth, S3, S6, A3 and A6) titanium surfaces prepared as explained previously in section 2.3.3.1 were studied. Also, each of the series was sterilized by different treatments (WS, SA, EO, GR) as explained in section 2.3.3.2.

The measurements were carried out with 1mM KCl-solution as the electrolyte and 0,1 M HCl as titration liquid. The pressure ramp was run to a maximum pressure of 600 mbar. The electrolyte solution was adjusted to pH 9,0 with 0,1 M NaOH followed by automatic titration to pH 3,0 with 0,1 M HCl. All electrolyte solutions (potassium chloride, potassium hydroxide and hydrochloric acid) were prepared using deionized water.

2.5.4. Methods

2.5.4.1. Test conditions

A commercial streaming potential instrument is available which has been used for all the measurements reported (Electrokinetic Analyzer, EKA, Anton Paar, Graz, Austria) (Figure 2-26). This apparatus includes a titration unit and allows performing measurement series at varied solution composition automatically.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

112

Figure 2-26 Streaming potential measurement apparatus (Electrokinetic Analyzer, EKA, Anton Paar)

To obtain the ζ potential for our surfaces, it has been used an asymmetric clamping cell. A schematic representation of the asymmetric clamping cell is shown in the Figure 2-27.

Figure 2-27 Schematic representation of the asymmetric clamping cell where the upper image is a plan

view and the bottom image is a lateral view [115]

The cell comprises several parallel rectangular capillary channels. At the two ends, the channels merge into two common manifolds that act as the electrode compartments. Each rectangular channel has a length L, width W, and a height h. The height-to-width ratio in each channel is <<1, allowing us to assume a two-dimensional flow and neglect any influence of the side walls on the polarization of the electrolyte and the flow field [115]. Streaming potential measurements with the asymmetric cell were taken eight times at each pH, with alternating flow direction of the solution (i.e., 4 measurements in each direction). ζ potentials of the test surfaces were calculated from Equation 2-14.

Chapter 2

113

A piece of PMMA was used as a supporting medium (back-plate) for all the samples in the asymmetric clamping cell. The back-plate ensured that the sample remained flat over the PMMA grooved portion of the cell. After loading the sample in the asymmetric cell with the clamped PMMA back-plate, the entire system was flushed thoroughly with deionized water for 20 min. Following the deionized water rinse, the streaming potential analyzer system and measuring cell were flushed with the electrolyte solution (1mM KCl) for 10min. This time was sufficient to completely displace deionized water in the system tubing. Once thoroughly flushed, the electrolyte solution was recirculated through the cell for 20min. The system was equilibrated and the pH adjustment prior to the subsequent streaming potential measurement was done. All electrokinetic experiments were performed at ambient temperature.

2.5.4.2. Streaming Potential

Streaming potential occurs when counter-ion displacement caused by hydrodynamic flow along a charged surface is balanced by reverse conduction current developed due to potential difference appearing along liquid stream. The mechanism is represented schematically in Figure 2-28.

Figure 2-28 Diagrammatic representation of the mechanism leading to the development of the streaming

potential [114]

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

114

In fact, the solid sample is mounted between two electrodes for detection of the potential difference. The liquid phase, which is aqueous electrolyte solution, is pumped through this measuring cell thereby generating a pressure difference across the sample plug. The differential pressure causes a relative movement of the charges in the electrochemical double layer and the streaming potential can be obtained by measuring the resulting electrical potential with electrodes in the bulk electrolyte solution [116].

The relation between streaming potential and zeta potential was first derived by Smoluchowski [111;114;118]:

Equation 2-13 RQ

LdpdU

0 ××

ε×εη

×=ζ

Where dU/dp is the slope of streaming potential versus pressure, ? is the viscosity of the electrolyte solution, ε is the dielectric constant of the electrolyte solution, ε0 is the vacuum permitivity, L and Q are dimensions of the fluid flow and R is the cell resistance. Then, zeta potential is not only determined by the nature and charge of the solid surface but also by the properties of the electrolyte solution.

Since it is impossible to determine the geometry of the fluid flow at the solid interface, the Smoluchowski equation needs some modification to be applicable in practice. Fairbrother and Mastin, have developed an approach in 1925, the surface conductivity correction, which is still accepted. By introducing the conductivity of the electrolyte solution, the geometric factor of the sample may be eliminated. The electrical conductivity is experimentally accessible.

2.5.4.2.1. Surface conductivity correction

The surface conductivity is the excess electrical conductivity tangential to a charged surface compared to the conductivity of the bulk electrolyte. The influence of surface conduction on the zeta potential should be considered for samples with sufficiently high conductive surfaces or samples that exhibit a strong swelling of the surface.

To obtain the surface conductivity correction, the electrolyte solution used for the measurement of the streaming potential is replaced by a 0,1 M KCl solution. Due to the high electrical conductivity of such concentrated solution, the surface conductivity may be neglected. The conductivity of the 0,1 M KCl solution is available from the literature data.

With this method of surface conductivity correction, zeta potential is determined from streaming potential measurement by introducing this term in the the Smoluchowski equation:

Chapter 2

115

Equation 2-14 corr0

kdpdU

×ε×ε

η×=ζ

NKCl001,0

NKCl1,0NKCl1,0corr R

KRk

×=

where K0,1MKCl and R0,1MKCl are the specific electrical conductivity and cell resistance using 0,1 M KCl solution.

For a dilute solution, the electrolyte concentration, the contribution of surface conductivity to the overall conductivity decreases until it becomes negligible at high electrolyte concentration. In our work a dilute solution has been used to perform the measurements and then a surface conductivity correction has been accomplished.

2.5.4.2.2. PMMA correction

Due to the construction of the clamping cell, the fluid goes through the studied sample and a PMMA insert, the data of the measured zeta potential have to be further corrected. A mixed zeta potential value is obtained since the PMMA insert of the measuring unit of the clamping cell contributes one measured surface while the other is the sample surface. In order to determine the zeta potential of the sample, the zeta potential of PMMA at the respective measuring conditions has to be known.

The PMMA insert of the clamping cell consists of a row of parallel rectangular channels where the height of the channels is much smaller than their width. The measured zeta potential may therefore be derived half from the sample, half from the PMMA [115].

Equation 2-15 )(21

SPMMAapp ζ+ζ=ζ

where ζapp is the apparent zeta potential of sample plus PMMA, ζPMMA is the zeta potential of PMMA, and ζS is the net zeta potential of the sample.

Since, our surfaces are rough and non-ideal the results will be used in comparative purposes. Then the PMMA correction will not be performed and the results will be given as apparent zeta potential which is a sum of both contributions of the sample and PMMA. Moreover, other studies [40;118] done with conductive samples or with the clamping cell, only compare between the different treatments but don’t provide the absolute values of zeta potential because of the complex theory around this model.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

116

2.5.5. Results

The apparent ζ potentials, which consist of the contribution of both sample and PMMA, for all the modified Ti surfaces have been determined. First, the results from smooth samples will be shown to study the effect of the sterilization process (Figure 2-29). For all tested surfaces the apparent zeta-potential becomes more negative with increasing pH and their IEP range from 3,5, for non-sterilized; followed by 3,6, for EO; 3,7, for GR; up to 3,9 for SA sterilized surfaces.

2 3 4 5 6 7 8 9 10

-40

-35

-30

-25

-20

-15

-10

-5

0

5

10

15

20

App

aren

t Zet

a P

oten

tial (

mV

)

pH

WS SA EO GR

Surfaces IEP WS 3,5 SA 3,9 EO 3,6 GR 3,7

Figure 2-29 Apparent zeta-potential (ζ) of smooth and sterilized samples as a function of pH. (WS)

without sterilization, (SA) steam autoclaving, (EO) ethylene oxide, (GR) gamma radiation

Secondly, the values for rough samples will be added to the smooth and non-sterilized results to study the influence of particle nature and size of the grit-blasting procedure on zeta potential determinations. Figure 2-30 shows the evolution of ζ-values depending on pH for rough Ti surfaces. As expected, the higher the pH, the lower the ζ. The values and evolution are similar for all the rough samples; the decrease of zeta-potential is moderate all along the studied pH compared to smooth samples. Starting at pH=6,5 plateau in the ζ vs pH plot is obtained.

The IEP values for the S3, S6, A3 and A6 are at pH 2,9; 3,5; 3,2; 3,0; respectively.

Chapter 2

117

2 3 4 5 6 7 8 9-35

-30

-25

-20

-15

-10

-5

0

5

10

App

aren

t Zet

a po

tent

ial (

mV

)

pH

A6 A3 S6 S3 Smooth

2 3 4 5-10

-5

0

5

10

App

aren

t Zet

a po

tent

ial (

mV

)

pH

A6 A3 S6 S3 Smooth

Surfaces IEP Smooth 3,5 S3 2,9 S6 3,5 A3 3,2 A6 3,0

Figure 2-30 Apparent zeta-potential (ζ) of smooth and rough surfaces: a general graph (upper image) and a

zoom of the IEP (bottom image)

2.5.6. Discussion

The ζ potential, formally defined as the electrical potential at the electrokinetic plane of shear or slip plane, is a very important property of charged solid-liquid interfaces. Streaming potential and streaming current measurements are known to be useful methods for the investigation of charge formation processes at solid/liquid interfaces [40;115;118].

The obtained zeta-potential curves showed a steadily more negative ζ potential as the solution was more alkaline. This is an effect more clearly showed on the smooth samples. The charge formation process is here due to the preferential adsorption of

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

118

negatively-charged electrolyte anions, such as OH¯, Cl¯. This is a conclusion which is further supported by the rather large negative values of the ζ potential in the basic zone.

In our study, the highest IEP values were obtained for steam-autoclaving surfaces,. Taking into consideration this result, one could agree with Werner et al. [118] in that surfaces with higher IEP are more hydrophobic (lower contact angles). However, the fact that the other type of surfaces do not correlate their IEP with either their hydrophobicty, or their SFE, or any of the components of the SFE, lead us to agree with van Oss et al. [181] in that the ζ -potential of a surface is rarely directly related to its γSAB and then, contact angle measurements usually are not affected by the ζ -potential of the solid surface. The possible influence of the hydrophilicity/hydrophobicity of the surface on their IEP can be shown if a) the differences in the values of contact angles with water are higher than the ones obtained for the surfaces tested here; and/or b) roughness of the surfaces tested is low enough to be able to directly relate the surface charge density and the surface potential.

Comparison between our results and bibliography has been used as a reference since the assay conditions are not the same. In our case it must be remembered the contribution of PMMA zeta potential at the sample ζ-potential giving the presented apparent zeta potentials at the results section. Electrokinetic data have been reported in many more works for titanium oxide powder (anatase, rutile) than for different crystalline modifications of titanium. For anatase and rutile powder, isoelectric points of 4-6,6 and 4,7-7,0, respectively are found in the literature [40;113]. The IEP’s obtained for the smooth and blasted titanium surfaces in this work, in the range 3,0-4,0, are slightly lower than those values. They are also lower than the values reported by Roessler al. [40] in a streaming potential electrokinetic analysis on sputtered Ti and Ti6Al4V surfaces i.e. 4,5 and 4,4, respectively.

However, several other works have shown that the IEP can be easily changed by structural or adsorbed impurities [119], thus resulting in IEP’s between 3 and 6,6 for anatase and 3,5-7,0 for rutile. This effect has also to be considered in the PMMA also which may also adsorb contaminants from one measurement to another, allowing concluding that PMMA insert might slightly change from measurement to measurement as also occurs for sample measurements.

On one hand, the adsorption of dissociating species can shift the IEP to or toward that of the adsorbate [119]. Both, smooth and blasted titanium surfaces are hydrophobic (iCA ≈ 50-90o) which, taking into consideration the XPS analysis, is probably due to the result of adsorbed hydrocarbons. These adsorbed hydrocarbons, which possess not dissociating groups, can also have a significant influence on the position of the IEP of the Ti surfaces because electrokinetic investigations on polymers with no dissociating groups revealed an IEP at pH 4 [118].

Chapter 2

119

On the other hand, as Rossler et al. pointed out [40], the shift to more acidic values of the IEP in the amorphous oxide layers on Ti surfaces compared to the crystalline titanium oxide phases may be also explained by differences in the structure of the phase. The short range order of the amorphous layers may alter the stability and density of surface hydroxyl groups, and so the IEP of the surface. However, the fact that the values of IEP obtained for the surfaces studied here are even lower than those reported by Rossler et al. [40] lead to the conclusion that hydrocarbons may seriously influence the IEP as discussed above.

The effect of roughness on the zeta potential curves in the basic zone is remarkable. Compared to smooth surfaces, the blasted surfaces showed less negative values of ζ-potential at values of pH above the IEP. The ζ-potential in blasted surfaces reaches an stable value for pH above 6,5-7,0. These results seem to indicate that roughness is influencing the values of ζ-potential by using the streaming potential methodology. Streaming potential depends on the total channel conductivity, which includes the streaming channel dimensions and the surface topography. In this respect, the electrolyte fluid velocity is lower at the valleys of the topography than at the peaks and at the fluid channel. This results in a possible retention of charge on the valleys and so, not all the displaced charges are transported to the electrodes at the end of the channel for detection. The increase in concentration of charge carriers inside the valleys might occur, which would even lead to an opposite flow [117]. The fact that roughness (1) retains charges inside the valleys preventing their detection at the end of the channel; and (2) possibly generates an opposite flow; lead to a decrease in measured streaming potential and hence to a decrease in zeta potential. Roughness did not influence on the IEP or algebraic sign of zeta potential in corresponding pH regimes, though.

According to the results obtained, all surfaces tested will be negatively-charged at physiological pH (7,4). This occurrences and the fact that the obtained IEP’s for rough Ti surfaces is around pH of 3-3,5 with not significant deviations of the values obtained for smooth Ti surfaces is of significance. The IEP is 8-9 for Al2O3 [120] and 2-3,5 for SiC [121]. Thus, one would expect some changes due to the presence of the blasting particles on the rough surfaces, especially for the alumina-blasted ones. No shift in the direction of the IEP was recorded, which means that the alumina blasted surfaces will not influence protein adsorption or cell behaviour because of a change in the electrostatic properties of the surface.

2.5.7. Conclusions

v At physiological pH, smooth samples presented a zeta potential considerably more negative than blasted-rough surfaces.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

120

v Sterilization methods presented similar electrostatic properties with slightly different IEP values that can not be directly correlated with their corresponding wettability properties.

v Different blasted-rough surfaces showed similar electrokinetic curves and IEPs, which indicates that the remaining blasting particles on the titanium surfaces did not significantly alter their surface electrical charge.

Chapter 2

121

2.6. Bibliography

[1] Puleo DA, Nanci A. Understanding and controlling the bone-implant interface. Biomaterials 1999; 20: 2311-2321.

[2] Cooper LF. A role for surface topography in creating and maintaining bone at titanium endosseous implants. Journal of Prosthetic Dentistry 2000; 84: 522-534.

[3] Wennerberg A, Albrektsson T, Johansson C, Andersson B. Experimental study of turned and grit-blasted screw-shaped implants with special emphasis on effects of blasting material and surface topography. Biomaterials 1996; 17: 15-22.

[4] Ronold HJ, Ellingsen JE. Effect of micro-roughness produced by TiO2 blasting--tensile testing of bone attachment by using coin-shaped implants. Biomaterials 2002; 23: 4211-4219.

[5] Buser D, Schenk RK, Steinemann S, Fiorellini JP, Fox CH, Stich H. Influence of surface characteristics on bone integration of titanium implants. A histomorphometric study in miniature pigs. Journal of Biomedical Materials Research 1991; 25: 889-902.

[6] Aparicio C. Tratamientos de superficie sobre titanio comercialmente puro para la mejora de la osteointegración de los implantes dentales. Universitat Politècnica de Catalunya; 2005. Thesis.

[7] Jayaraman M, Meyer U, Buhner M, Joos U, Wiesmann HP. Influence of titanium surfaces on attachment of osteoblast-like cells in vitro. Biomaterials 2004; 25: 625-631.

[8] Deligianni DD, Katsala N, Ladas S, Sotiropoulou D, Amedee J, Missirlis YF. Effect of surface roughness of the titanium alloy Ti-6Al-4V on human bone marrow cell response and on protein adsorption. Biomaterials 2001; 22: 1241-1251.

[9] Anselme K, Bigerelle M. Topography effects of pure titanium substrates on human osteoblast long-term adhesion. Acta Biomaterialia 2005; 1: 211-222.

[10] Aparicio C, Gil FJ, Planell JA, Engel E. Human-osteoblast proliferation and differentiation on grit-blasted and bioactive titanium for dental applications. Journal of Materials Science-Materials in Medicine 2002; 13: 1105-1111.

[11] Anselme K, Linez P, Bigerelle M, Le Maguer D, Le Maguer A, Hardouin P, Hildebrand HF, Iost A, Leroy JM. The relative influence of the topography and chemistry of TiAl6V4 surfaces on osteoblastic cell behaviour. Biomaterials 2000; 21: 1567-1577.

[12] Anselme K, Bigerelle M, Noel B, Dufresne E, Judas D, Iost A, Hardouin P. Qualitative and quantitative study of human osteoblast adhesion on materials with various surface roughnesses. Journal of Biomedical Materials Research 2000; 49: 155-166.

[13] Martin JY, Schwartz Z, Hummert TW, Schraub DM, Simpson J, Lankford J, Dean DD, Cochran DL, Boyan BD. Effect of Titanium Surface-Roughness on Proliferation, Differentiation, and Protein-Synthesis of Human Osteoblast-Like Cells (Mg63). Journal of Biomedical Materials Research 1995; 29: 389-401.

[14] Lincks J, Boyan BD, Blanchard CR, Lohmann CH, Liu Y, Cochran DL, Dean DD, Schwartz Z. Response of MG63 osteoblast-like cells to titanium and titanium alloy is dependent on surface roughness and composition. Biomaterials 1998; 19: 2219-2232.

[15] Bachle M, Kohal RJ. A systematic review of the influence of different titanium surfaces on proliferation, differentiation and protein synthesis of osteoblast-like MG63 cells. Clinical Oral Implants Research 2004; 15: 683-692.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

122

[16] Boyan BD, Lossdorfer S, Wang L, Zhao G, Lohmann CH, Cochran DL, Schwartz Z. Osteoblasts generate an osteogenic microenvironment when grown on surfaces with rough microtopographies. European cells and materials 2003; 6: 22-27.

[17] Citeau A, Guicheux J, Vinatier C, Layrolle P, Nguyen T, Pilet P, Daculsi G. In vitro biological effects of titanium rough surface obtained by calcium phosphate grid blasting. Biomaterials 2005; 26: 157-165.

[18] Wennerberg A, Albrektsson T, Johansson C, Andersson B. Experimental study of turned and grit-blasted screw-shaped implants with special emphasis on effects of blasting material and surface topography. Biomaterials 1996; 17: 15-22.

[19] Zinger O, Zhao G, Schwartz Z, Simpson J, Wieland M, Landolt D, Boyan BD. Differential regulation of osteoblasts by substrate microstructural features. Biomaterials 2005; 26: 1837-1847.

[20] Aparicio C, Gil FJ, Fonseca C, Barbosa M, Planell JA. Corrosion behaviour of commercially pure titanium grit blasted with different materials and sizes of grit particles for dental implant applications. Biomaterials 2003; 24: 263-273.

[21] Aparicio C, Gil FJ, Thams U, Munoz F, Padros A, Planell JA. Osseointegration of grit-blasted and bioactive titanium implants: Histomorphometry in minipigs. Key engineering materials 2004; 254-2: 737-740.

[22] Wennerberg A, Albrektsson T, Johansson C, Andersson B. Experimental study of turned and grit-blasted screw-shaped implants with special emphasis on effects of blasting material and surface topography. Biomaterials 1996; 17: 15-22.

[23] Mustafa K, Wennerberg A, Wroblewski J, Hultenby K, Lopez BS, Arvidson K. Determining optimal surface roughness of TiO2 blasted titanium implant material for attachment, proliferation and differentiation of cells derived from human mandibular alveolar bone. Clinical Oral Implants Research 2001; 12: 515-525.

[24] Wennerberg A, Albrektsson T, Johansson C, Andersson B. Experimental study of turned and grit-blasted screw-shaped implants with special emphasis on effects of blasting material and surface topography. Biomaterials 1996; 17: 15-22.

[25] Taborelli M, Jobin M, Francois P, Vaudaux P, Tonetti M, SzmuklerMoncler S, Simpson JP, Descouts P. Influence of surface treatments developed for oral implants on the physical and biological properties of titanium .1. Surface characterization. Clinical Oral Implants Research 1997; 8: 208-216.

[26] Francois P, Vaudaux P, Taborelli M, Tonetti M, Lew DP, Descouts P. Influence of surface treatments developed for oral implants on the physical and biological properties of titanium .2. Adsorption isotherms and biological activity of immobilized fibronectin. Clinical Oral Implants Research 1997; 8: 217-225.

[27] Mustafa K, Wennerberg A, Wroblewski J, Hultenby K, Lopez BS, Arvidson K. Determining optimal surface roughness of TiO2 blasted titanium implant material for attachment, proliferation and differentiation of cells derived from human mandibular alveolar bone. Clinical Oral Implants Research 2001; 12: 515-525.

[28] Bagno A, Di Bello C. Surface treatments and roughness properties of Ti-based biomaterials. Journal of Materials Science-Materials in Medicine 2004; 15: 935-949.

[29] Darvell BW, Samman N, Luk WK, Clark RKF, Tideman H. Contamination of Titanium Castings by Aluminum-Oxide Blasting. Journal of Dentistry 1995; 23: 319-322.

[30] Orsini G, Assenza B, Scarano A, Piatelli M, Piatelli A. Surface analysis of machined versus sandblasted and acid-etched titanium implants. International journal of oral and maxillofacial implants 2000; 15: 779-784.

Chapter 2

123

[31] Naji A, Harmand M-F. Cytocompatibility of two coating materials, amorphous alumina and silicon carbide, using human differentiated cell cultures. Biomaterials 1991; 12: 690-694.

[32] Puleo DA, Nanci A. Understanding and controlling the bone-implant interface. Biomaterials 1999; 20: 2311-2321.

[33] Tengvall P, Lundström I. Physico-chemical considerations of titanium as a biomaterial. Clinical Materials 1992; 9: 115-134.

[34] Boyan BD, Hummert TW, Dean DD, Schwartz Z. Role of material surfaces in regulating bone and cartilage cell response. Biomaterials 1996; 17: 137-146.

[35] Sharma PK, Rao KH. Analysis of different approaches for evaluation of surface energy of microbial cells by contact angle goniometry. Advances in Colloid and Interface Science 2002; 98: 341-463.

[36] Thierry B, Tabrizian M, Savadogo O, Yahia LH. Effects of sterilization processes on NiTi alloy: Surface characterization. Journal of Biomedical Materials Research 2000; 49: 88-98.

[37] Doundoulakis JH. Surface-Analysis of Titanium After Sterilization - Role in Implant-Tissue Interface and Bioadhesion. Journal of Prosthetic Dentistry 1987; 58: 471-478.

[38] Keller JC, Draughn RA, Wightman JP, Dougherty WJ, Meletiou SD. Characterization of sterilized cp Titanium implant surfaces. International Journal Oral Maxillofacial Implants 1990; 5: 360-367.

[39] Morra M, Cassinelli C. Bacterial adhesion to polymer surfaces: A critical review of surface thermodynamic approaches. Journal of Biomaterials Science-Polymer Edition 1997; 9: 55-74.

[40] Roessler S, Zimmermann R, Scharnweber D, Werner C, Worch H. Characterization of oxide layers on Ti6Al4V and titanium by streaming potential and streaming current measurements. Colloids and Surfaces B-Biointerfaces 2002; 26: 387-395.

[41] Hallab NJ, Bundy KJ, O'Connor K, Moses RL, Jacobs JJ. Evaluation of metallic and polymeric biomaterial surface energy and surface roughness characteristics for directed cell adhesion. Tissue Engineering 2001; 7: 55-71.

[42] Healy KE, Thomas CH, Rezania A, Kim JE, McKeown PJ, Lom B, Hockberger PE. Kinetics of bone cell organization and mineralization on materials with patterned surface chemistry. Biomaterials 1996; 17: 195-208.

[43] Ponsonnet L, Reybier K, Jaffrezic N, Comte V, Lagneau C, Lissac M, Martelet C. Relationship between surface properties (roughness, wettability) of titanium and titanium alloys and cell behaviour. Materials Science & Engineering C-Biomimetic and Supramolecular Systems 2003; 23: 551-560.

[44] Anselme K. Osteoblast adhesion on biomaterials. Biomaterials 2000; 21: 667-681.

[45] Ismail FSM, Rohanizadeh R, Atwa S, Mason RS, Ruys AJ, Martin PJ, Bendavid A. The influence of surface chemistry and topography on the contact guidance of MG63 osteoblast cells. Journal of Materials Science-Materials in Medicine 2007; 18: 705-714.

[46] Kieswetter K, Schwartz Z, Hummert TW, Cochran DL, Simpson J, Dean DD, Boyan BD. Surface roughness modulates the local production of growth factors and cytokines by osteoblast-like MG-63 cells. Journal of Biomedical Materials Research 1996; 32: 55-63.

[47] Lincks J, Boyan BD, Blanchard CR, Lohmann CH, Liu Y, Cochran DL, Dean DD, Schwartz Z. Response of MG63 osteoblast-like cells to titanium and titanium alloy is dependent on surface roughness and composition. Biomaterials 1998; 19: 2219-2232.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

124

[48] Vezeau PJ, Koorbusch GF, Draughn RA, Keller JC. Effects of multiple sterilization on surface characteristics and in vitro biologic responses to titanium. Journal of Oral and Maxillofacial Surgery 1996; 54: 738-746.

[49] Stanford CM, Keller JC, Solursh M. Bone Cell Expression on Titanium Surfaces Is Altered by Sterilization Treatments. Journal of Dental Research 1994; 73: 1061-1071.

[50] Kasemo B. Biocompatibility of Titanium Implants - Surface Science Aspects. Journal of Prosthetic Dentistry 1983; 49: 832-837.

[51] Blunt RT. White Light Interferometry – a production worthy technique for measuring surface roughness on semiconductor wafers . CS MANTECH Conference. 2006. Canada.

Ref Type: Conference Proceeding

[52] Wieland M, Hanggi P, Hotz W, Textor M, Keller BA, Spencer ND. Wavelength-dependent measurement and evaluation of surface topographies: application of a new concept of window roughness and surface transfer function. Wear 2000; 237: 231-252.

[53] Vöros J, Wieland M, Ruiz-Taylor L, Textor M, Brunette DM. Characterization of Titanium Surfaces. In: Brunette DM, Tengvall P, Textor M, Thomsen P (eds.), Titanium in Medicine. Ed.Springer; 2001: 87-144.

[54] Annarelli CC, Fornazero J, Cohen R, Bert J, Besse JL. Colloidal protein solutions as a new standard sensor for adhesive wettability measurements. Journal of Colloid and Interface Science 1999; 213: 386-394.

[55] Stout KJ, Blunt L. Three Dimensional Surface Topography, Second Edition ed. Jessica Kingsley publishers; 1994.

[56] vonRecum AF, Brown CE, Shannon CE, LaBerge M. Surface topography. In: Andreas F.von Recum (ed.), Handbook of biomaterials evaluation. 2000: 227-240.

[57] Lausmaa J, Kasemo B, Hansson S. Accelerated Oxide-Growth on Titanium Implants During Autoclaving Caused by Fluorine Contamination. Biomaterials 1985; 6: 23-&.

[58] Keller JC, Stanford CM, Wightman JP, Draughn RA, Zaharias R. Characterizations of Titanium Implant Surfaces .3. Journal of Biomedical Materials Research 1994; 28: 939-946.

[59] Liu XY, Chu PK, Ding CX. Surface modification of titanium, titanium alloys, and related materials for biomedical applications. Materials Science & Engineering R-Reports 2004; 47: 49-121.

[60] Webb K, Hlady V, Tresco PA. Relative importance of surface wettability and charged functional groups on NIH 3T3 fibroblast attachment, spreading, and cytoskeletal organization. Journal of Biomedical Materials Research 1998; 41: 422-430.

[61] Georgi A, Grinnel F, Groth T. Studies on the biocompatibility of materials: fibroblast reorganisation of substratum-bound fibronectin on surfaces varying in wettability. Journal of Biomedical Materials Research 1996; 30: 385-391.

[62] Altankov G, Groth T. Reorganization of Substratum-Bound Fibronectin on Hydrophilic and Hydrophobic Materials Is Related to Biocompatibility. Journal of Materials Science-Materials in Medicine 1994; 5: 732-737.

[63] Webb K, Hlady V, Tresco PA. Relationships among cell attachment, spreading, cytoskeletal organization, and migration rate for anchorage-dependent cells on model surfaces. Journal of Biomedical Materials Research 2000; 49: 362-368.

[64] Ruardy TG, Schakenraad JM, Vandermei HC, Busscher HJ. Adhesion and Spreading of Human Skin Fibroblasts on Physicochemically Characterized Gradient Surfaces. Journal of Biomedical Materials Research 1995; 29: 1415-1423.

Chapter 2

125

[65] Schakenraad JM, Busscher HJ, Wildevuur CRH, Arends J. Thermodynamic Aspects of Cell Spreading on Solid Substrata. Cell Biophysics 1988; 13: 75-91.

[66] Van der Valk P, van Pelt AW, Busscher HJ, de Jong HP, Wildevuur CR. Interaction of fibroblasts and polymer surfaces: relationship between surface free energy and fibroblast spreading. Journal of Biomedical Materials Research 1983; 17: 807-817.

[67] Birdi K. Cell adhesion on solids and the role of surface forces. Journal of Theoretical Biology 1981; 93: 1-5.

[68] Lee JH, Khang G, Lee JW, Lee HB. Interaction of different types of cells of polymer surfaces with wettability gradient. Journal of Colloid and Interface Science 1998; 205: 323-330.

[69] Bren L, English L, Fogarty J, Policoro R, Zsidi A, Vance J, Drelich J, Istephanous N, Rohly K. Hydrophilic/electron-acceptor surface properties of metallic biomaterials and their effect on osteoblast cell activity. Journal of Adhesion Science and Technology 2004; 18: 1711-1722.

[70] MacDonald DE, Deo N, Markovic B, Stranick M, Somasundaran P. Adsorption and dissolution behavior of human plasma fibronectin on thermally and chemically modified titanium dioxide particles. Biomaterials 2002; 23: 1269-1279.

[71] Haynes CA, Norde W. Globular proteins at solid/liquid interfaces . Colloids andSurfaces B: Biointerfaces 1994; 2: 517-566.

[72] Wilson CJ, Clegg RE, Leavesley DI, Pearcy MJ. Mediation of biomaterial-cell interactions by adsorbed proteins: A review. Tissue Engineering 2005; 11: 1-18.

[73] Absolom DR, Zingg W, Neumann AW. Protein Adsorption to Polymer Particles - Role of Surface-Properties. Journal of Biomedical Materials Research 1987; 21: 161-171.

[74] Grinnell F, Feld MK. Fibronectin Adsorption on Hydrophilic and Hydrophobic Surfaces Detected by Antibody-Binding and Analyzed During Cell-Adhesion in Serum-Containing Medium. Journal of Biological Chemistry 1982; 257: 4888-4893.

[75] Fabriziushoman DJ, Cooper SL. Competitive Adsorption of Vitronectin with Albumin, Fibrinogen, and Fibronectin on Polymeric Biomaterials. Journal of Biomedical Materials Research 1991; 25: 953-971.

[76] McFarland CD, Thomas CH, DeFilippis C, Steele JG, Healy KE. Protein adsorption and cell attachment to patterned surfaces. Journal of Biomedical Materials Research 2000; 49: 200-210.

[77] Lampin M, WarocquierClerout R, Legris C, Degrange M, SigotLuizard MF. Correlation between substratum roughness and wettability, cell adhesion, and cell migration. Journal of Biomedical Materials Research 1997; 36: 99-108.

[78] van Oss CJ, Giese RF. The Hydrophilicity and Hydrophobicity of Clay-Minerals. Clays and Clay Minerals 1995; 43: 474-477.

[79] Adamson AW. The solid-liquid interface, contact angle. In Physical chemistry of surfaces. John Wiley&sons; 1990: 379-414.

[80] Medout-Marere V, El Ghzaoui A, Charnay C, Douillard JM, Chauveteau G, Partyka S. Surface heterogeneity of passively oxidized silicon carbide particles: Hydrophobic-hydrophilic partition. Journal of Colloid and Interface Science 2000; 223: 205-214.

[81] van Oss CJ, Chaudhury MK, Good RJ. Interfacial Lifshitz-van der Waals and polar interactions in macroscopic systems. Chemical Reviews 1988; 88: 927-941.

[82] Chibowski E. Surface free energy of a solid from contact angle hysteresis. Advances in Colloid and Interface Science 2003; 103: 149-172.

Topographical and surface energy characterization of different grit-blasted and sterilized Ti surfaces

126

[83] Gu Y. Contact angle measurement techniques for determination of wettability. In: Marcel Dekker (ed.), Encyclopedia of Surface and Colloid Science. 2002: 1213-1227.

[84] van Oss CJ, Good RJ, Chaudhury MK. Additive and nonadditive surface tension components and the interpretation of contact angles. Langmuir 1988; 4: 884-891.

[85] Owens DK, Wendt RC. Estimation of Surface Free Energy of Polymers. Journal of Applied Polymer Science 1969; 13: 1741-&.

[86] Good RJ. Contact-Angle, Wetting, and Adhesion - A Critical-Review. Journal of Adhesion Science and Technology 1992; 6: 1269-1302.

[87] van Oss CJ, Ju L, Chaudhury MK, Good RJ. Estimation of the Polar Parameters of the Surface-Tension of Liquids by Contact-Angle Measurements on Gels. Journal of Colloid and Interface Science 1989; 128: 313-319.

[88] Vanoss CJ, Chaudhury MK, Good RJ. Interfacial Lifshitz-Vanderwaals and Polar Interactions in Macroscopic Systems. Chemical Reviews 1988; 88: 927-941.

[89] Vogler EA. Structure and reactivity of water at biomaterial surfaces. Advances in Colloid and Interface Science 1998; 74: 69-117.

[90] Drelich J. Static contact angles for liquids at heterogeneous rigid solid surfaces. Polish Journal of Chemistry 1997; 71: 525-549.

[91] Wu W, Giese RF, Vanoss CJ. Evaluation of the Lifshitz-Van der Waals Acid-Base Approach to Determine Surface-Tension Components. Langmuir 1995; 11: 379-382.

[92] Wenzel RN. Resistance of solid surfaces to wetting by water. Industrial and Engineering Chemistry 1936; 28: 988-994.

[93] Cassie ABD, Baxter S. Wettability of porous surfaces. Transactions of the Faraday Society 1944; 546.

[94] Balaur E, Macak JM, Tsuchiya H, Schmuki P. Wetting behaviour of layers of TiO2

nanotubes with different diameters. Journal of Materials Chemistry 2005; 15: 4488-4491.

[95] Sheng Y-J, Jiang S, Tsao H-K. Effects of geometrical characteristics of surface roughness on droplet wetting. The journal of chemical physics 2007; 127: 234704-1-234704-7.

[96] Bico J, Marzolin C, Quéré D. Pearl drops. Europhysics Letters 1999; 47: 220-226.

[97] Erbil Y. Surface Tesion of Polymers. In: CRC Press (ed.), CRC Handbook of Surface and Colloid Chemistry. Boca Raton; 1997: 292.

[98] Combe EC, Owen BA, Hodges JS. A protocol for determining the surface free energy of dental materials. Dental Materials 2004; 20: 262-268.

[99] Adamson AW. Capillarity. In Physical chemistry of surfaces. John Wiley&sons; 1990: 4-49.

[100] Vanoss CJ, Chaudhury MK, Good RJ. Interfacial Lifshitz-Vanderwaals and Polar Interactions in Macroscopic Systems. Chemical Reviews 1988; 88: 927-941.

[101] Montgomery DC. Three-Level and Mixed-Level Factorial and Fractional Factorial Designs. In: John Wiley & Sons (ed.), Design and analysis of experiments, 1997 ed. 2005: 460-474.

[102] Johnson RE, Dettre RH. Wettability and contact angles. New York: Wiley; 1969: 85-153.

[103] Serro AP, Saramago B. Influence of sterilization on the mineralization of titanium implants induced by incubation in various biological model fluids. Biomaterials 2003; 24: 4749-4760.

Chapter 2

127

[104] Kasemo B, Lausmaa J. Biomaterial and Implant Surfaces - on the Role of Cleanliness, Contamination, and Preparation Procedures. Journal of Biomedical Materials Research-Applied Biomaterials 1988; 22: 145-158.

[105] Albrektsson T, Branemark PI, Hansson HA, Lindstrom J. Osseointegrated Titanium Implants - Requirements for Ensuring A Long-Lasting, Direct Bone-To-Implant Anchorage in Man. Acta Orthopaedica Scandinavica 1981; 52: 155-170.

[106] Lyklema J, Rovillard S, De Coninck J. Electrokinetics: The properties of the stagnant layer unraveled. Langmuir 1998; 14: 5659-5663.

[107] Shelton RM, Rasmussen AC, Davies JE. Protein Adsorption at the Interface Between Charged Polymer Substrata and Migrating Osteoblasts. Biomaterials 1988; 9: 24-29.

[108] Qiu Q, Sayer M, Kawaja M, Shen X, Davies JE. Attachment, morphology, and protein expression of rat marrow stromal cells cultured on charged substrate surfaces. Journal of Biomedical Materials Research 1998; 42: 117-127.

[109] Kim HM, Himeno T, Kawashita M, Lee JH, Kokubo T, Nakamura T. Surface potential change in bioactive titanium metal during the process of apatite formation in simulated body fluid. Journal of Biomedical Materials Research Part A 2003; 67A: 1305-1309.

[110] Haynes CA, Sliwinsky E, Norde W. Structural and Electrostatic Properties of Globular-Proteins at A Polystyrene Water Interface. Journal of Colloid and Interface Science 1994; 164: 394-409.

[111] Zembala M. Electrokinetics of heterogeneous interfaces. Advances in Colloid and Interface Science 2004; 112: 59-92.

[112] Delgado AV, Gonzalez-Caballero E, Hunter RJ, Koopal LK, Lyklema J. Measurement and interpretation of electrokinetic phenomena - (IUPAC technical report). Pure and Applied Chemistry 2005; 77: 1753-1805.

[113] Kosmulski M. The significance of the difference in the point of zero charge between rutile and anatase. Advances in Colloid and Interface Science 2002; 99: 255-264.

[114] McFadyen P. Zeta Potential of Macroscopic Surfaces from Streaming Potential Measurements. 1995.

Ref Type: Unpublished Work

[115] Walker SL, Bhattacharjee S, Hoek EMV, Elimelech M. A novel asymmetric clamping cell for measuring streaming potential of flat surfaces. Langmuir 2002; 18: 2193-2198.

[116] Christoforou CC, Westermannclark GB, Anderson JL. The Streaming Potential and Inadequacies of the Helmholtz-Equation. Journal of Colloid and Interface Science 1985; 106: 1-11.

[117] Schnitzer C, Ripperger S. Influence of surface roughness on streaming potential method. Chemical Engineering & Technology 2008; 31: 1696-1700.

[118] Werner C, Konig U, Augsburg A, Arnhold C, Korber H, Zimmermann R, Jacobasch HJ. Electrokinetic surface characterization of biomedical polymers - a survey. Colloids and Surfaces A-Physicochemical and Engineering Aspects 1999; 159: 519-529.

[119] Parks GA. The isoelectric points of solid oxides, solid hydroxides, and aqueous hydroxo complex systems. Chemical Reviews 1965; 65: 177-198.

[120] Wang XH, Hirata Y. Electrical and rheological properties of colloidal alumina suspensions. Journal of Ceramic Processing Research. 2000; 1: 64-68.

[121] Hyrata Y, Miyano K, Sameshima S, Kamino Y. Reaction between SiC surface and aqueous solutions containing A1 ions. Colloids and Surfaces A-Physicochemical and Engineering Aspects 1998; 133: 183-189.