JAIRO MATOZINHO CORDEIRO
CARACTERIZAÇÃO ESTRUTURAL, MECÂNICA, QUÍMICA E
ELETROQUÍMICA DE LIGAS EXPERIMENTAIS DE TITÂNIO PARA
IMPLANTES DENTÁRIOS
STRUCTURAL, MECHANICAL, CHEMICAL AND
ELECTROCHEMICAL CHARACTERIZATION OF EXPERIMENTAL
TITANIUM ALLOYS FOR DENTAL IMPLANTS
Piracicaba 2017
CARACTERIZAÇÃO ESTRUTURAL, MECÂNICA, QUÍMICA E
ELETROQUÍMICA DE LIGAS EXPERIMENTAIS DE TITÂNIO PARA
IMPLANTES DENTÁRIOS
STRUCTURAL, MECHANICAL, CHEMICAL AND
ELECTROCHEMICAL CHARACTERIZATION OF EXPERIMENTAL
TITANIUM ALLOYS FOR DENTAL IMPLANTS
Dissertação apresentada à Faculdade de Odontologia de Piracicaba da Universidade Estadual de Campinas como parte dos requisitos exigidos para a obtenção do título de Mestre em Clínica Odontológica, na Área de Prótese dental.
Dissertation presented to the Piracicaba Dental School of the University of Campinas in partial fulfillment of the requirements for the degree of Master in Dental Clinic, in Dental Prosthesis area.
Orientador: Prof. Dr. Valentim Adelino Ricardo Barão
Este exemplar corresponde à versão final da dissertação defendida pelo aluno Jairo Matozinho Cordeiro e orientada pelo Prof. Dr. Valentim Adelino Ricardo Barão.
Piracicaba 2017
Ficha catalográfica
Universidade Estadual de Campinas
Biblioteca da Faculdade de Odontologia de Piracicaba Marilene Girello - CRB 8/6159
Cordeiro, Jairo Matozinho,
C811c CorCaracterização estrutural, mecânica, química e eletroquímica de ligas experimentais de titânio para implantes dentários / Jairo Matozinho Cordeiro. – Piracicaba, SP : [s.n.], 2017.
CorOrientador: Valentim Adelino Ricardo Barão.
CorDissertação (mestrado) – Universidade Estadual de Campinas, Faculdade de Odontologia de Piracicaba.
Cor1. Implantes dentários. 2. Ligas de titânio. 3. Corrosão. I. Barão, Valentim Adelino Ricardo,1983-. II. Universidade Estadual de Campinas. Faculdade de Odontologia de Piracicaba. III. Título.
Informações para Biblioteca Digital
Título em outro idioma: Structural, mechanical, chemical and electrochemical
characterization of experimental titanium alloys for dental implants
Palavras-chave em inglês:
Dental implants Titanium alloys Corrosion
Área de concentração: Prótese Dental Titulação: Mestre em Clínica Odontológica Banca examinadora:
Valentim Adelino Ricardo Barão [Orientador] Wirley Gonçalves Assunção
Antônio Pedro Ricomini Filho
Data de defesa: 17-02-2017
Programa de Pós-Graduação: Clínica Odontológica
Dedico este trabalho à minha família, cujo amor, carinho e educação foram a base da minha formação. Em todas as instâncias vocês permitiram a concretização deste sonho, me protegeram e ampararam mesmo na distância. Muito obrigado por serem a inspiração contínua dos meus anseios.
Ao Prof. Dr. Valentim Adelino Ricardo Barão, pela orientação inestimável e oportunidades que me foram dadas, permitindo o meu crescimento e amadurecimento. Sou profundamente grato pela atenção, inspiração, honestidade e qualidade de orientação. Você tem sido um exemplo brilhante de pessoa e profissional o qual quero me tornar. Muito obrigado!
A Deus, por ter me fornecido sabedoria e me guiado pelas escolhas corretas. À Universidade Estadual de Campinas – UNICAMP, na pessoa do Magnífico Reitor, Prof. Dr. José Tadeu Jorge, pelo meu mestrado nesta instituição.
À Faculdade de Odontologia de Piracicaba – UNICAMP, na pessoa do seu Diretor Prof. Dr. Guilherme Elias Pessanha Henriques, pela oportunidade da realização do Programa de Pós-Graduação em Clínica Odontológica.
À Coordenadora Geral da Pós-Graduação, Profa. Dra. Cinthia Pereira Machado Tabchoury, e à Coordenadora do Programa de Pós-Graduação em Clínica Odontológica, Profa. Dra. Karina Gonzales Silvério Ruiz pela excelência em suas funções.
Aos docentes Prof. Dr. Marcelo Ferraz Mesquita, Prof. Dr. Rafael Leonardo Xediek Consani, Profa. Dra. Altair Del Bel Cury, Profa. Dra. Renata Cunha Matheus Rodrigues Garcia, Profa. Dra. Célia Rizzati Barbosa, Prof. Dr. Wander José da Silva, Prof. Dr. Mauro Antônio de Arruda Nóbilo, Prof. Dr. Frederico Andrade e Silva e Prof. Dr. Wilkens Aurélio Buarque e Silva, pela disponibilidade e por todo conhecimento passado.
Aos professores Drs. Nilson Cristino da Cruz, Elidiane Cipriano Rangel e Ana Lúcia R. Ribeiro pela parceria no desenvolvimento deste trabalho e ensinamentos compartilhados.
Aos técnicos Rafael Parra, Jéssica Gonçalves e Jamille Altheman pelo apoio na realização de análises cruciais para o desenvolvimento deste trabalho.
À Sra. Eliete A. Ferreira Lima Marim, secretária do Departamento de Prótese e Periodontia da FOP-UNICAMP, e ao Eduardo Pinez, técnico do Laboratório de Prótese Total, pela competência e solicitude constantes.
Aos amigos, Bruno Rodrigues, Herick Fernando, Luziany Sene, Bianca Rodrigues e Rívia Juliana, que torceram por mim a cada nova conquista e me encorajaram a caminhar sempre em frente, muito obrigado pelo amparo e amor eterno. Cada um de vocês me inspiraram até aqui, e me passaram a certeza de que sempre estaremos juntos, e qualquer obstáculo daqui para a frente é pequeno diante de nossa amizade.
Àqueles que tive o prazer de compartilhar esta jornada de perto, Mariana Barbosa, Heloísa Navarro, Camilla Fraga, Louise Dornelas, Rodrigo Lins e Giancarlo De la Torre, só posso agradecer pelo companheirismo e amizade, que foram esteio na exaustão, ânimo nas incertezas e impulso de novos anseios. Nos méritos desta conquista, há muito da presença de vocês. Muito obrigado.
Oliveira, Guilherme Machado, Ricardo Caldas, Conrado Caetano, e amigos conquistados durante estes dois anos, em especial, Olívia Figueiredo, Mayara Abreu, Talita Malini, Edmara Tatiely, Victor Muñoz, Bruna Ximenes, Vinícius Aguiar, Elis Lira, Amanda Bandeira, Rahyza Freire, e a todos aqueles que participaram de alguma forma, sou grato pelos momentos e apoio partilhados que facilitaram a concretização dessa etapa.
Este trabalho objetivou fornecer um resumo de vários aspectos relevantes de ligas de titânio (Ti) para uso como implantes dentários. As informações existentes sobre as propriedades mecânicas, químicas, eletroquímicas e biológicas das principais ligas desenvolvidas nos últimos anos foram profundamente revisadas para fornecer evidências científicas a favor da substituição do titânio comercialmente puro (Ticp) por suas ligas na prática clínica. Posteriormente, foi conduzido um estudo in vitro com a finalidade de avaliar as características estruturais, mecânicas, químicas, eletroquímicas e biológicas de ligas binárias e ternárias de Ti contendo zircônio (Zr) e nióbio (Nb). Ligas experimentais foram desenvolvidas (% em massa): Ti-5Zr, Ti-10Zr, Ti-35Nb-5Zr e Ti-35Nb-10Zr, e usinadas em forma de discos com 10 mm de diâmetro e 2 mm de espessura. Discos de Ticp e da liga Ti-6Al-4V foram usados como controles. Análise microestrutural foi realizada por meio de difratometria de raios-X (XRD) e microscopia eletrônica de varredura (MEV). As propriedades mecânicas de microdureza Vickers e módulo de elasticidade foram avaliadas. As características de superfície foram analisadas através da espectroscopia de energia dispersiva (EDS), espectroscopia de fotoelétrons excitados por raios X (XPS), microscopia de força atômica (AFM), rugosidade de superfície (Ra, Rq, Rt, Rz) e energia livre de superfície. O ensaio eletroquímico constou de testes padrões conduzidos em solução de fluido corpóreo (pH 7,4) para simular o plasma sanguíneo. Adsorção de albumina foi mensurada através do método do ácido bicinconínico. Os dados foram avaliados por meio do teste ANOVA de um fator e teste de Tukey (α = 0,05). As ligas de Ti podem ser modificadas por processos termomecânicos afetando sua microestrutura e consequentemente suas propriedades mecânicas. O Ticp e as ligas Zr apresentaram em sua microestrutura apenas a fase α. As ligas 6Al-4V e 35Nb-5Zr apresentaram uma estrutura cristalina α+β. A fase β foi detectada para a liga Ti-35Nb-10Zr. Foi observado na literatura melhores propriedades para as ligas de Ti, tais como baixo módulo de elasticidade, alta resistência à tração, biocompatibilidade satisfatória, adequada resistência à corrosão e segurança e sobrevida in vivo semelhantes ao Ticp. No estudo in vitro, todas as ligas apresentaram microdureza superior à do Ticp (p < 0,05). O módulo de elasticidade foi estatisticamente inferior para as ligas Ti-Nb-Zr (p < 0,05). Todos os materiais apresentaram em sua constituição os elementos específicos de cada liga acrescidos de carbono e oxigênio. De uma forma geral, todos os grupos apresentaram uma camada de óxido nativa em sua superfície formada principalmente por TiO2.Ticp apresentou
a maior rugosidade de superfície dentre os grupos (p < 0,05) enquanto a liga Ti-10Zr apresentou menor energia livre de superfície. Esta liga apresentou ainda o melhor comportamento eletroquímico devido à combinação de maiores valores de resistência à polarização e menores valores de capacitância (p < 0,05). As ligas Ti-Nb-Zr apresentaram a menor estabilidade eletroquímica. Não houve diferenças significativas entre os grupos com relação à adsorção de albumina. A superioridade das ligas de Ti em relação ao Ticp é evidente em vários aspectos, apesar disso não há evidências científicas que assegurem a substituição total deste material in vivo. Estudos in vivo com as novas ligas de Ti devem ser encorajados com a finalidade de consolidar o seu uso como substitutos do Ticp. As ligas do tipo β tem demonstrado serem mais promissoras para aplicações biomédicas. Apesar disso, a liga binária α Ti-Zr exibiu a melhor combinação de propriedades mecânicas e eletroquímicas, sendo forte candidata para uso em implantes dentários.
This study aimed to provide a summary of several aspects of titanium (Ti) alloys for using as dental implants. Existing information about the mechanical, chemical, electrochemical and biological properties of the main alloys developed over the past few years were deeply reviewed to provide scientific evidence in favor of using Ti-based alloys as alternative of commercially pure titanium (cpTi) with its alloys in the clinical scenario. Subsequently, an in
vitro study was carried out aiming to evaluate the structural, mechanical, chemical,
electrochemical and biological properties of binary and ternary Ti alloys containing zirconium (Zr) and niobium (Nb). The experimental alloys were developed (in %wt): Ti-5Zr, Ti-10Zr, Ti-35Nb-5Zr and Ti-35Nb-10Zr, and machined into discs with 10 mm in diameter and 2 mm in thickness. CpTi and Ti-6Al-4V alloy discs were used as controls. Microstructural analysis was performed by means of X-ray diffraction (XRD) and scanning electron microscopy (SEM). The mechanical properties such as Vickers microhardness and elastic modulus were evaluated. The surface characteristics were analyzed by dispersive energy spectroscopy (EDS), X-ray excited photoelectron spectroscopy (XPS), atomic force microscopy (AFM), surface roughness (Ra, Rq, Rt, Rz) and surface free energy. The electrochemical assessment consisted of standard tests conducted in a body fluid solution (pH 7.4) to simulate the blood plasma. The albumin adsorption was measured by the bicinchoninic acid method. Data were evaluated through one-way ANOVA and Tukey test (α = 0.05). Ti alloys can be modified by thermomechanical processes affecting their microstructure and consequently their mechanical properties. CpTi and Ti-Zr alloys presented only α phase in their microstructure. Ti-6Al-4V and 5Zr alloys showed a α + β structure. The β phase was detected for the Ti-35Nb-10Zr alloy. It has been observed in the literature better properties for Ti alloys, such as low elastic modulus, high tensile strength, satisfactory biocompatibility, adequate corrosion resistance, and safety. Additionally, Ti alloys exhibited similar in vivo survivability when compared to cpTi. In the in vitro study, all alloys had a superior microhardness when compared to cpTi (p <0.05). The elastic modulus was statistically lower for Ti-Nb-Zr alloys (p <0.05). All materials presented in their constitution specific elements of each alloy in addition to carbon and oxygen. In general, all groups had a native oxide layer on their surface formed mainly by TiO2. CpTi presented the highest surface roughness (p < 0.05). Ti-10Zr
alloy had lower surface energy than the control group (p < 0.05). Ti-10Zr alloy presented the best electrochemical behavior due to the combination of higher values of polarization resistance and lower values of capacitance (p < 0.05). Ti-Nb-Zr alloys had the lowest electrochemical stability (p <0.05). All materials exhibited similar albumin adsorption (p > 0.05). The superiority of Ti alloy when compare to cpTi is evident in many ways. However, there is no scientific evidence to ensure the full replacement of this material in vivo. In vivo studies with the new Ti alloys should be encouraged in order to consolidate their use as cpTi substitutes. β-type alloys have been shown to be more promising for biomedical applications. Despite this, the binary α Ti-Zr alloy showed the best combination of mechanical and electrochemical properties and may be considered an important candidate to fabricate dental implants.
1 INTRODUÇÃO ... 12
2 ARTIGOS ... 16
2.1 Artigo: Is there scientific evidence favoring the substitution of commercially pure titanium with titanium alloys for the manufacture of dental implants? ... 16
2.2 Artigo: Development of binary and ternary titanium alloys for dental implants .. 57
3 DISCUSSÃO ... 93
4 CONCLUSÃO ... 95
REFERÊNCIAS ... 96
ANEXOS ... 103
Anexo 1 – Permissão para utilização do artigo 1 ... 103
1 INTRODUÇÃO
A população mundial está ficando cada vez mais idosa, e com o aumento da expectativa de vida (Niinomi et al., 2012; Al-Nawas et al., 2012), a reabilitação de pacientes parcialmente e totalmente desdentados com próteses implanto-suportadas tornou-se uma importante opção de tratamento nas últimas décadas, com resultados confiáveis a longo prazo (Lee e Chen, 2013; Tolentino et al., 2014). Para o sucesso dos implantes é necessário que as propriedades dos materiais utilizados sejam adequadas às situações à que os mesmos serão submetidos. Para isso, os materiais de escolha para implantes dentários necessitam apresentar baixo módulo de elasticidade, resistência à fadiga e à tração elevada (Datta et al., 2016), boa resistência à corrosão, biocompatibilidade e serem livres de elementos tóxicos (Elias et al., 2015).
O titânio comercialmente puro (Ticp) é o material de escolha para a fabricação de implantes dentários (Quirynen et al., 2014), e tem sido utilizado por mais de 30 anos (Barter et al., 2012), principalmente, devido a suas excelentes propriedades, como a resistência à corrosão, tanto no ar quanto em fluídos biológicos, a densidade relativamente baixa e a biocompatibilidade (Kobayashi et al., 1998; Faria et al., 2011; Grandin et al., 2012). Contudo, a sua utilização é restrita quando alta resistência é necessária, devido uma fraca resistência ao desgaste e resistência à tração e fadiga insuficientes (Grandin et al., 2012; Mishnaevsky et al., 2014; Qiu et al., 2015). Problemas de fadiga podem ocorrer devido o titânio (Ti) ser um material relativamente macio (Gottlow et al., 2010), principalmente quando usado na reabilitação com implantes de pequeno diâmetro (< 3,5mm), que sobretudo devem cumprir os requisitos de alta estabilidade mecânica para evitar a sobrecarga e fratura do implante (Al-Nawas et al., 2012). Além disso, um elevado módulo de elasticidade e a dificuldade para melhorar as propriedades mecânicas, sem reduzir a biocompatibilidade, são tidas como características que limitam o seu uso como material para implantes dentários (Elias et al., 2015).
A utilização de ligas de Ti, obtidas pela moagem de Ti com outros metais, é uma alternativa para se obter melhores propriedades mecânicas e melhor resistência ao desgaste dos implantes (Gottlow et al., 2012). Os elementos constituintes das ligas de Ti são classificadas em três categorias microestruturais: α-estabilizadores [alumínio (Al), oxigênio (O), nitrogênio (N) e carbono (C)], β-estabilizadores [molibdênio (Mo), vanádio (V), nióbio (Nb), tântalo (Ta) ferro (Fe), cromo (Cr), níquel (Ni) e cobalto (Co)] e neutro [zircônio (Zr)]
(Ikarashi et al., 2007). As propriedades das ligas de Ti irão variar de acordo com os elementos que as compõe.
Comumente, o Ti é ligado com Al e V formando Ti-6Al-4V (Barter et al., 2012; Saulacic et al., 2012). A adição de Al e V foi considerada com a finalidade de proporcionar uma estrutura bifásica α-β, desejável devido aos efeitos estabilizadores α e β, respectivamente. Além disso, a liga com Al fornece uma alta taxa de endurecimento para a matriz de Ti (Kobayashi et al., 1998). A liga Ti-6Al-4V apresenta uma maior resistência mecânica (Ikarashi et al., 2007; Quirynen et al., 2014), relativo baixo módulo de elasticidade e propriedades de resistências equivalentes com os aços inoxidáveis e as ligas de Co-Cr (Kobayashi et al., 1998). No entanto, a adição de Al e V diminui a biocompatibilidade do implante quando comparado com aqueles produzidos a partir de Ticp (Ikarashi et al., 2007; Barter et al., 2012), tornando a utilização da liga Ti-6Al-4V indesejável (Kobayashi et al., 1998). O V tem sido classificado como um elemento tóxico (Okazaki et al., 1998; Okazaki e Gotoh 2005), e apesar de não ter sido comprovado, o Al tem sido relacionado a efeitos neurotóxicos significativos, principalmente ao considerar os recentes relatos de seu envolvimento na doença de Alzheimer (Okazaki e Gotoh 2005; Niinomi et al., 2012).
A necessidade de maior resistência mecânica e adequada biocompatibilidade dos biomateriais utilizados em implantes dentários e aplicações ortopédicas tem motivado a busca por ligas alternativas de Ti livres de substâncias tóxicas (Grandin et al., 2012). Tendo em conta a toxicidade suspeita de Al e V (Ribeiro et al., 2013), uma das soluções encontrada foi adicionar Nb e Zr ao Ti (Ribeiro et al., 2009). Os elementos Nb e Zr são considerados como não-tóxicos e com excelente biocompatibilidade (Okazaki et al., 1998; Niinomi et al., 2012; Ribeiro et al., 2013). O primeiro atua como um β-estabilizador, o que melhora as propriedades mecânicas da liga (Ribeiro et al., 2009) e resulta em uma alta resistência à corrosão (Niinomi et al., 2012); enquanto o segundo tem propriedades químicas similares às do Ti (Kobayashi et al., 1998; Chen et al., 2008; Gottlow et al., 2012), sem toxidade local ou sistêmica, e é usado para induzir endurecimento da matriz de Ti, aumentando a resistência da liga (Ribeiro et al., 2009; Al-Nawas et al., 2012; Niinomi et al., 2012; Ribeiro et al., 2013).
As ligas Ti-Zr apresentam predominantemente uma estrutura monofásica α (Barter et al., 2012), que lhe confere maior dureza (Gottlow et al., 2012) e melhor resistência ao desgaste (Ikarashi et al., 2007). Ti-Zr tem apresentado comparável ou melhor resistência mecânica, além de uma melhor biocompatibilidade do que o Ticp e a liga Ti-6Al-4V (Chen et al., 2008; Barter et al., 2012). Consoante a isso, essa liga mostrou um perfil biológico in vitro melhor do que Ticp, o que foi relacionado com as propriedades de superfície do material, tais
como a rugosidade, a energia de superfície e a composição do substrato (Sista et al., 2011). A composição da superfície também pareceu favorecer a estabilidade mecânica do implante no osso para a liga de Ti-Nb-Ta-Zr, aumentando a quantidade de osso em torno do implante ao longo do tempo, sem alterar em implantes de Ticp. Estas características tornam-se vantajosas, ao passo que a topografia e a química de superfície provaram ser cruciais na determinação do sucesso dos implantes, uma vez que desempenham papel importante na interação do implante com o tecido ósseo circunjacente (Sista et al., 2011; Grandin et al., 2012).
Uma preocupação importante sobre a segurança das ligas utilizadas em implantes refere-se a sua biocompatibilidade e possível liberação de íons metálicos (Ikarashi et al., 2007; Sista et al., 2011; Al-Nawas et al., 2012). É bem reportado que os biomateriais metálicos sofrem corrosão quando em contato com o ambiente fisiológico (Barão et al., 2012; Ribeiro et al., 2013; Faverani et al., 2015). Os metais utilizados na produção dos implantes podem ser liberados no tecido circunjacente por vários mecanismos, incluindo corrosão, desgaste, e processos mecânicos e eletroquímicos acelerados, como a corrosão sob tensão, corrosão por fadiga e corrosão por atrito (Okazaki e Gotoh 2005).
O Ti e suas ligas podem sofrer este processo de degradação em condições fisiológicas, podendo em alguns casos apresentar reatividade exagerada (Mishnaevsky et al., 2014). O processo de degradação do material pode resultar em efeitos adversos sobre o corpo (ex.: osteólise, reabsorção óssea, infecções) e sobre os implantes (ex.: afrouxamento e falha) (Bayón et al., 2015). A corrente elétrica anormal produzida durante o processo corrosivo pode converter o implante dentário em um eletrodo, afetando o desenvolvimento e o reparo ósseo e, consequentemente, o processo de osseointegração dos implantes (Gittens et al., 2011).
Tanto a biocompatibilidade quanto a resistência à corrosão do Ti são influenciadas pela formação espontânea de uma película de proteção passiva (camada de óxido) sobre sua superfície (Okazaki e Gotoh 2005; Grandin et al., 2012; Faria et al., 2014), sendo que quanto mais forte/estável for o filme passivo, melhor será a resistência à corrosão do implante (Okazaki e Gotoh 2005; Gabriel et al., 2012). As mudanças estruturais nos filmes passivos das ligas de Ti podem alterar sua resistência à corrosão (Cvijovic-Alagic et al., 2011), sendo que menores níveis de corrosão podem ser conseguidos pela adição de alguns elementos ao Ti. Foi demonstrado que ligas contendo paládio (Pd), Zr, Nb e tântalo (Ta) apresentam maior resistência à corrosão do que a liga Ti-6Al-4V (Okazaki et al., 1998). Ligas contendo Nb e Zr em diferentes proporções observou uma cinética de corrosão destas ligas muito semelhante ao observado para a liga Ti-6Al-4V quando em contato com saliva artificial, com a resistência à
corrosão levemente superior para as ligas Ti-Nb-Zr (Ribeiro et al., 2013). A diferença pode ser atribuída a distintas composições química e microestruturais dos óxidos que se formam na superfície do material, bem como pelo aumento da camada superior do filme passivo nos implantes das ligas Ti-Nb-Zr (Gabriel et al., 2012; Ribeiro et al., 2013).
A biofuncionalidade de biomateriais metálicos é atualmente insuficiente, e precisa ser melhorada (Niinomi et al., 2012). Para prolongar a aplicação clínica de ligas de Ti, é necessário desenvolver novas ligas que sejam suficientemente fortes e estáveis em um ambiente corrosivo, apresentem biocompatibilidade satisfatória e segurança para o uso in vivo (Kobayashi et al., 1998; Faria et al., 2014). Embora várias ligas sejam projetadas para aplicações biomédicas, muitos estudos são inconclusivos quanto à possibilidade de substituir o Ticp por esses novos materiais (Cordeiro e Barão, 2016). Além disso, poucos estudos testaram ligas experimentais in vivo para consolidar seu uso (Stenlund et al., 2015). As ligas de Ti contendo Nb e Zr aparecem como promissoras candidatas para a aplicação em implantes dentários (Ribeiro et al., 2009; Ribeiro et al., 2013, Calderon-Moreno et al., 2014; Meng et al, 2014). Estudos extensivos têm sido conduzidos sobre o comportamento químico, físico, mecânico e eletroquímico do Ticp e da liga Ti-6Al-4V (Cvijovic-Alagic et al ., 2011; Faverani et al., 2015; Shah et al., 2016; Filho et al., 2016 Yan et al., 2017), mas estudos comparativos das recém desenvolvidas ligas contendo Nb e Zr com o Ticp e a liga Ti-6Al-4V são limitados.
Dessa forma, este trabalho objetivou fornecer um resumo de vários aspectos relevantes das ligas de Ti para uso como implantes dentários. As informações existentes sobre as propriedades mecânicas, químicas, eletroquímicas e biológicas das principais ligas desenvolvidas nos últimos anos foram profundamente revisadas para fornecer evidências científicas a favor da substituição do Ticp por suas ligas na prática clínica. Posteriormente, foi conduzido um estudo in vitro com a finalidade de caracterizar a microestrutura e as propriedades mecânicas, químicas e eletroquímicas de ligas binárias e ternárias de Ti contendo Zr e Nb. Objetivou-se ainda realizar um estudo comparativo com os materiais amplamente utilizados para implantes dentários: cpTi e liga Ti-6Al-4V. O aspecto biológico de tais ligas foi investigado utilizando um ensaio de adsorção de proteínas.
2 ARTIGOS
2.1 Is there scientific evidence favoring the substitution of commercially pure titanium with titanium alloys for the manufacture of dental implants?
#
Jairo M. Cordeiro a,b, Valentim A. R. Barão a,b,*
aDepartment of Prosthodontics and Periodontology, Piracicaba Dental School, University of
Campinas (UNICAMP), Av Limeira, 901, Piracicaba, São Paulo, Brazil, 13414-903.s
bIBTN/Br - Institute of Biomaterials, Tribocorrosion and Nanomedicine – Brazilian Branch,
Brazil.
*Corresponding author:
Av. Limeira, 901, Piracicaba, SP, Brazil 13414-903, Tel.: + 55-19-2106 5719; Fax: +55-19-2106 5218;
E-mail address: [email protected]
#Artigo publicado na revista Materials Science and Engineering C (IF = 3.420). Mater Sci
Eng C Mater Biol Appl. 2017; 71:1201-1215. doi: 10.1016/j.msec.2016.10.025. Reprinted
from Mater Sci Eng C Mater Biol Appl, 2017; 71:1201-1215, Cordeiro JM, Barao VA. Is there scientific evidence favoring the substitution of commercially pure titanium with titanium alloys for the manufacture of dental implants? Copyright (2016), with permission from Elsevier (Anexo 1).
Abstract
The development of Ti alloys to manufacture dental implants has emerged in recent years due to the increased failure of commercially pure titanium (cpTi) implants. Thus, this study reviews existing information about the mechanical, chemical, electrochemical, and biological properties of the main Ti alloys developed over the past few years to provide scientific evidence in favor of using Ti-based alloys as alternative to cpTi. Ti alloys may be considered viable substitutes in the fabrication of dental implants. Such evidence is given by the enhanced properties of alloys, such as a low elastic modulus, high tensile strength, satisfactory biocompatibility, and good corrosion and wear resistances. In addition, Ti alloys may be modified at the structural, chemical, and thermomechanical levels, which allows the development of materials in accordance with the demands of several situations encountered in clinical practice. Although several in vitro studies have established the superiority of Ti alloys over cpTi, mainly in terms of their mechanical properties, there is no scientific evidence that supports the total replacement of this material in vivo. This review demonstrates the superiority of β-type alloys. However, it is evident that in vivo studies are encouraged to test new alloys to consolidate their use as substitutes for cpTi.
Contents
1. Introduction
2. Titanium alloy classification
3. Thermomechanical process of titanium alloys 4. Mechanical properties
4.1 Strength and ductility 4.2 Fatigue performance
4.3 Hardness and elastic modulus 5. Surface properties 6. Electrochemical properties 7. Biological properties 7.1 Biocompatibility 7.2 Osseointegration 7.3 Antimicrobial properties 8. Surface treatments 9. Final considerations
1. Introduction
The durability of rehabilitative treatments depends on the availability of materials capable of minimizing the risk of mechanical failure, especially in applications involving the treatment of large defects subject to high loads or where it is necessary to reduce the implant dimensions [1]. Progress in this area has achieved improvements in treatment performance and longevity; nevertheless, failures still occur [2]. In order to overcome these failures, dental implant materials should present high fatigue strength, low elastic modulus, high strength [3], and good corrosion resistance and biocompatibility [4].
Commercially pure titanium (cpTi) is the material of choice for the manufacture of dental implants [5]. However, its use is limited in areas subjected to high wear and tensile and fatigue strength [2,6,7]. Because Ti is a relatively soft material [8], fatigue may occur, particularly when it is used in small-diameter implants, which must fulfill high requirements for mechanical stability to avoid overload and implant fracture [9]. In addition, high-elasticity modulus and difficulty in improving its mechanical properties without any reduction in biocompatibility have been considered characteristics that limit the use of cpTi as a material of dental implants [4]. The use of Ti alloy made by grinding Ti with other metals is an alternative option to obtain better mechanical properties [8].
Several elements may be combined with Ti, resulting in alloys with distinct properties and patterns closer to the ideal for use as dental implants. Ti-6Al-4V alloy is widely used owing to its excellent mechanical performance [2]. On the other hand, this alloy showed negative effects on cell viability by the release of Al and V [10], with a consequent adverse influence on implant biocompatibility [11]. Indeed, Al has been linked to significant neurotoxic effects, especially when considering reports of its association with Alzheimer's disease, bone fragility [12], and potential causes of local inflammation [13]. These reports have discouraged the use of Ti-6Al-4V and stimulated the development of alloys free of toxic elements that are inert in the oral environment.
To extend their clinical application, experimental alloys must exhibit satisfactory mechanical properties, with sufficient strength and stability in a corrosive environment, besides being biocompatible and safe for in vivo use [14,15]. Ti alloys have proven to be of great interest for biomedical applications due to their excellent strength and superior biocompatibility [3] associated with properties such as high tensile strength, good corrosion resistance [16], and elastic modulus comparable to that of bone tissue [17,18]. These outstanding properties have pointed to Ti alloys as viable options to be used as an alternative
to cpTi in the manufacture of dental implants, and in many cases, for use as the first choice in treatment [19].
Although several alloys are designed for biomedical applications, many studies are inconclusive concerning the possibility of using these new materials as substitutes to cpTi. In addition, few studies have tested experimental alloys in vivo to consolidate their use. In this article, we provide a summary of several relevant aspects of Ti alloys for use as dental implants. Existing information about the mechanical, chemical, electrochemical, and biological properties of the main alloys developed over the past few years is deeply reviewed to provide scientific evidence in favor of using Ti-based alloys as alternative of cpTi with its alloys in the clinical scenario.
2. Classification of Ti alloys
Ti can take on two different crystal forms in a temperature-dependent manner. The α phase has a hexagonal closed-packed (HCP) structure and is stable from room temperature to 882 °C. The β phase has a body center cubic (BCC) structure and is stable at temperatures higher than those mentioned above [8,19,20]. Ti also presents metastable phases, such as the hexagonal martensite α′ and orthorhombic α″ phases [21]. The transition temperature between the α and β phases can be changed by combining elements with Ti, which consequently modifies its microstructure. Besides the constitution of the alloy, the processing approach and heat treatment conditions affect the material’s microstructure [22].
The microstructure of Ti alloys is defined according to the type and concentration of the alloying elements, as well the crystalline phases present at room temperature [20,23]. Elements that may constitute Ti alloys are classified into three categories: α-stabilizers (Al, O, N, C) tend to stabilize the α phase by increasing the transition temperature; β-stabilizers (Mo, V, Fe, Cr, Ni, Co, Nb) depress the transition temperature by stabilizing the β phase; and elements such as Zr and Sn exhibit no effect on the stability of any phase, being considered neutral elements [11,19]. Table 1 summarizes this information for better understanding. To understand such mechanism, the phase diagram of Ti as a function of stabilizers constituents is shown in Figure 1. The effect of elements addition at the transition temperature between α and β phases can be clearly seen.
Table 1. Main components of Ti alloys and their influence on the transition temperature and
Ti matrix.
Element Influence on the transition temperature
Main effect on Ti matrix
α-stabilizer Al, O, N, C Increase Hardening
β-stabilizer Mo, V, Fe, Cr, Ni, Co, Nb
Decrease Grain Refiners
Neutral Zr and Sn No significant effect Hardening
Figure 1. Phase diagram of Ti according to the stabilizers constituents.
Depending on the proportion of each phase, Ti can be further classified as near α, α, α + β, near β, and β phases [21]. The near-α alloys contain approximately 12% of β-stabilizers and approximately 510% of β phases; alloys that present in their constitution higher amounts of β- stabilizers, resulting in 1030% of β phases in the microstructure, are classified as α + β alloys; the near β and β alloys have higher amounts of β-stabilizers and predominantly β phase in their microstructures [19]. Figure 2 shows the relationship between the concentration of stabilizer elements incorporated into the Ti and its microstructural phases.
Figure 2. Relationship between the concentration of stabilizer elements incorporated into
titanium and its microstructural phases.
It is known that microstructure has a great influence on the physical and chemical properties of the material [22] and is widely affected by the volume fraction, morphology, distribution, and size of the α phase precipitates within the matrix [24]. Knowing the elements that influence the microstructure of Ti and understanding how this relationship occurs has driven many researchers to incorporate elements to pure Ti to produce implants with greater performance than those made of cpTi.
The addition of V and Al to Ti forming Ti-6Al-4V, for example, was considered to provide a biphasic structure (α + β) because of the stabilizing effects of α and β. The alloys that exhibit α + β structure are characterized by higher strength, higher ductility and higher low-cycle fatigue [19]. Furthermore, alloys containing Al provide a high rate of solid solution hardening to the Ti matrix [14], with Ti-6Al-4V being the most common Ti alloy used for biomedical applications where high strength is required [25].
Similar to Al, Zr and Bi have been used to induce a solid solution hardening effect [7,26,27]. When Zr is cast only with Ti, it can form α alloys of various proportions, which usually increase the mechanical strength (such as tensile strength, hardness and flexural strength) and improve the corrosion potential and wear resistance of Ti [21]. In contrast, β alloys have in their constitution β-stabilizers acting as grain refiners [3]. Among them, Nb, Mo and Ta have received emphasis for forming β alloys characterized by a combination of improved mechanical properties and excellent biocompatibility [7,22]. They also present low
elastic modulus [22,28], superior corrosion resistance [22], good plasticity, high yield strength [28], hardenability, fracture toughness, and reasonable ductility [24]. These characteristics have made them the most promising alloys for the manufacture of implants [20,22]. The relationship among the chemical elements, microstructure, and alloy properties is be better discussed in the following topics.
3. Thermomechanical process of alloys
Good mass proportion of the elements involved and a melting method associated with an appropriate sequence of heat treatments is necessary to obtain alloys with enhanced mechanical, physical, and chemical characteristics [26]. An appropriate thermomechanical process depends on the time, temperature, strain rates, cooling rates, and the alloy’s chemical composition [29]. There are elements that are unlikely to produce alloys by conventional grinding and melting processes [30]. Additionally, some microstructural phases are not stabilized at room temperature if there are no changes in the thermomechanical process.
Thermomechanical processing affects the nature and ratio of microstructural phases [19] and modifies the size, shape, and amount of microstructural constituents [22]. Cold-rolled alloys showed increased dislocation density, microstrain, and size changes of the grains and sub-grains, which vary depending on the percentage reduction in alloy and annealing time. Also, it directly influences the strength, elongation, and elastic modulus of the alloy [31,32]. Annealed materials generally exhibit better ductility than castings or those worked cold; however, they show decreases in tensile strength [4].
A previous study [22] evaluated the effect of the cooling process (furnace cooling (FC), air cooling (AC), or water quenching (WQ)) on the microstructure and mechanical properties of a Ti-Nb-Zr-V alloy. Air-cooled samples had a relative grain refinement with the presence of α microstructure distributed in a β matrix, ensuring higher values of hardness of the alloy. The cooling time is crucial to the presence of the α phase (i.e., water quenching is a fast process that is insufficient for the precipitation of this phase in the microstructure).
Sintering is another step that can be modified to obtain improved properties of alloys. Ideally, it should retain the developed microstructure while preventing or minimizing undesirable grain growth [33]. It has been reported that the fracture strength of Ti-Nb-Ag alloy fabricated by a process that involves the combination of amperage and pressure (Spark Plasma Sintering - SPS) was nearly three times higher than that of a vacuum furnace-sintered sample, and promoted a dense structure without any pores [34]. SPS proved to be a promising approach to producing grain refinement. It has been further observed that increasing the
sintering temperature leads to pore reduction and grain size decrease, which has a positive effect on raising the hardness of the alloy [33].
The process that obtains fine structures with smaller grain size has been shown to be of greater interest to increase the mechanical strength and toughness of the alloy; in addition, the microstructure and the general properties obtained are more isotropic [35]. The possibility of modifying the thermomechanical process during the development of Ti alloys is a wide field of study, with the possibility of changing the properties of the alloys using simple techniques. However, there is a certain difficulty in standardizing the procedures that are effective for all alloys.
4. Mechanical properties
Mechanical properties and biocompatibility are the primary considerations in the design of a new alloy [3]. Dental implants require good mechanical properties because they are exposed to loads and fatigue cycles in function [36]. To improve the mechanical properties of Ti and its alloys, mechanisms such as solid solution strengthening by interstitial and substitutional atoms, grain refinement, precipitation, work hardening, and dispersion strengthening, including lamellar and dispersed phases, can be used [4].
Biomaterials are expected to exhibit a combination of high strength and low elastic modulus [29]. However, these characteristics seem to be interexclusive for solid materials, particularly for metals and alloys [18]. Elements such as Mo, Nb, Ta, and Zr are generally used to increase strength and reduce the elastic modulus of the Ti alloy [7]. The search for a balance between strength and elastic modulus is important to improve the performance of implants compared with those made from cpTi. Therefore, it is necessary to understand the properties of the biomaterials and to predict their behavior when fixed to the bone [37]. Table 2 shows the mechanical properties of Ti alloys developed for biomedical applications. Conclusions from the analysis in Table 2 must be drawn with caution, as it is a summary of several studies in which the methodologies were not standardized.
Table 2. Mechanical properties of Ti alloys used for biomedical applications.
Alloy
type Alloy Thermomechanical process
Ultimate tensile strength (MPa)
Yield strength (MPa)
Elastic modulus
(GPa) Elongation (%) Hardness (HV) Ref.
α
CpTi Conventional 465 320 110 36 220 [25]
Ti*-5Zr Conventional or heat treated - - ≈86 - ≈246-357.40 [21,38]
Ti-10Zr Conventional or heat treated - - ≈95 - ≈257-384.00 [21,38]
Ti-15Zr Conventional or heat treated - - ≈113 - ≈250-412.0 [21,38]
Ti-2Bi Conventional ≈360 ≈310 102.0 ≈25 ≈210 [7] Ti-5Bi Conventional ≈370 ≈310 100.1 ≈20 ≈225 [7] Ti-10Bi Conventional ≈520 ≈425 ≈97.9 ≈15 ≈300 [7] Ti-20Bi Conventional ≈585 ≈535 94.3 ≈3 ≈365 [7] Ti-1Ag Conventional - ≈985-1105 - - ≈380-435 [39] Ti-3Ag Conventional - ≈975-1045 - - ≈360-370 [39] Ti-5Ag Conventional - ≈960-1040 - - ≈350-370 [39]
Ti-47Zr-7.3Al Conventional or hot rolled 1144-1564 992-1447 106-117 3.08- 10.08 - [23]
Ti-10Zr-15Si Conventional - 1231 150 2.7 - [40]
α + β
Ti–35Nb Hot rolled and water quenched 561 160 73.9 59 - [29]
Ti-8Fe Conventional ≈1560 ≈130 ≈567 [41]
Ti-6Al-4V Conventional 1000 910 114 18 400 [25]
Ti-6Al-7Nb Annealed 1000-1100 - - 10-15 - [10]
Ti-13Nb-131Zr ST† and MPCR¶ 896-1027 840-935 - 7,8-14 - [42]
Ti-35Nb-5Zr Heat treated 486 - - - 240 [26]
Ti–35Nb-2.5Sn Hot rolled and water quenched 566 139 65.8 55 - [29]
Ti–7Ta–5Fe Conventional - 1250 110 - 430 [43] Ti-8Fe-5Ta Conventional ≈1045 - ≈120 - ≈387 [41] Ti-8Fe-8Ta Conventional ≈1130 - ≈107 - ≈380 [41] Ti-9Fe-2Ta Conventional ≈1075 - ≈110 - ≈388 [41] Ti-9Fe-5Ta Conventional ≈1075 - ≈103 - ≈370 [41] Ti-9Fe-9Ta Conventional ≈1010 - ≈99 - ≈350 [41] Ti-10Fe-10Ta Conventional ≈1005 - ≈92 - ≈342 [41]
Ti-2.0Al-12.8Sn-18.2Cr Cooling in different ways 1005-1075 45.1-72.8 - - - [3]
Ti-6.9Al-4.1Zr-7.0Mo-9.9Sn-10.1Cr Cooling in different ways 1065-1113 46.2-72.8 - - - [3]
Ti-15Zr-4Nb-4Ta-0.2Pd-0.1O-0.005N Annealed or ST and aged 714-919 - 94-99 18-28 243-289 [10]
Ti-15Zr-4Nb-4Ta-0.2Pd-0.2O-0.005N Annealed or ST and aged 881-1026 - 97-100 14-27 301-317 [10]
Ti-15Sn-4Nb-2Ta-0.2Pd-0.05O-0.005N Annealed or ST and aged 860-1109 - 89-103 10-21 292-351 [10]
Ti-15Sn-4Nb-2Ta-0.2Pd-0.2O-0.005N Annealed or ST and aged 966-1189 - 86-98 14-20 336-361 [10]
β
Ti-30Nb Conventional - - 93 - 248 [44]
Ti-38Nb Cold rolled and Annealed 1020 850 56 - - [17]
Ti-40Nb Conventional - 544 62 28 - [40]
Ti-45Nb Conventional or heat treated 480-527 460-438 64-65 21 185-233 [25,45]
Ti-26.88Fe Conventional 2627 2028 - 7.5 - [46]
Ti-12Mo-3Nb ST and water quenched - 450 105 41 - [47]
Ti-30Nb-2Sn Conventional - - 89 - 246 [44]
Ti-30Nb-4Sn Conventional - - 86 - 235 [44]
Ti-33Nb-4Sn Cold rolled and annealed ≈853 ≈ 763 ≈36 - - [18]
Ti-35Nb-5Sn Hot rolled and water quenched 563 130 68.2 36 - [29]
Ti-35Nb-7.5Sn Hot rolled and water quenched 478 152 86.4 38 - [29]
Ti-13Nb-13Zr Conventional 680 515 80 28 265 [25]
Ti-35Nb-10Zr Heat treated 546 - - - 185 [26]
Ti-41.1Nb-7.1Zr Heat treated and water quenched 490 490 65 16 182 [48]
Ti-26.88Fe-2Ta Conventional 2560 2300 - 1.0 - [46]
Ti-26.88Fe-4Ta Conventional 2531 2215 175 5.0 - [46]
Ti-12Nb-5Fe Cold crucible levitation melting - 740 90 - 293 [43]
Ti-10Ta-4Fe Cold crucible levitation melting - 1360 121 - 410 [43]
Ti-20.6Nb-13.6Zr-0.5V ST and cooling by different ways 741-785 520-572 77-81 15-19 260-280 [22]
Ti-35.3Nb-7.1Zr-5.1Ta Heat treated and water quenched 550 550 63 21 173 [48]
Ti-23.72Nb-4.83Zr-1.74Ta Spark Plasma Sintering 2793 1143 35 - - [49]
Ti-10Zr-15-Nb-15Si Conventional - 1185 120 3 - [40]
Ti-31Nb-6Zr-5Mo Hot rolled and heat treated 672-747 610-712 44-48 20.6-26.7 - [50]
Ti-13.6Nb-6.5Al- 6Cu-5.1Ni Conventional 1145 1050 82 3.7 - [51]
Ti-14.1Nb-6.7Al-4Cu-3.4Ni Conventional 880 820 72 14 - [51]
Ti-23.72Nb-4.83Zr-1.74Ta-2Si Spark Plasma Sintering 3263 1296 37 - - [49]
Ti-23.72Nb-4.83Zr-1.74Ta-5Si Spark Plasma Sintering 3267 1347 40 - - [49]
Ti-22.3Nb-4.6Zr1.6Ta-6Fe Spark Plasma Sintering 2650 2425 75 - - [52]
Ti-10Zr-31Cu-10Pd-4Sn Spark Plasma Sintering 2060 - 114 - - [53]
Ti-35Nb-4Sn-6Mo-9Zr ST and water quenched 834 802 65 11.0 - [31]
Ti-35Nb-5Sn-3Mo−3Zr ST and water quenched 690 668 70 18.0 - [31]
Ti-35Nb-5Sn-6Mo-3Zr ST and water quenched 770 729 85 10.5 - [31]
Ti-35Nb-3Sn-9Mo-9Zr ST and water quenched 833 781 70 10.2 - [31]
Ti-35Nb-4Sn-3Mo-6Zr ST and water quenched 666 654 55 8.9 - [31]
Ti-35.3Nb-7.3Zr-5.7Ta Conventional or ST ≈550 ≈450 ≈62-66 ≈21 - [54] Ti-35.3Nb-7.3Zr-5.7Ta-1Si Conventional or ST ≈620 ≈560 ≈70-72 ≈12 - [54] Ti-35.3Nb-7.3Zr-5.7Ta-2Fe Conventional or ST ≈720 ≈590 ≈79-81 ≈22 - [54] Ti-13.6Nb-6Co-5.1Cu-6.5Al Conventional 1550 1110 92 - - [55] Ti-13.6Nb-6Co-5.1Cu-6.5Al-0.5B Conventional 1570 1150 97 - - [55] Ti-13.6Nb-6Co-5.1Cu-6.5Al-1B Conventional 1680 1200 100 - - [55] Ti-35.3Nb-7.3Zr-5.7Ta-0.5Si+1Fe Conventional or ST ≈685 ≈630 ≈75-78 ≈14 - [54] Ti-35.3Nb-7.3Zr-5.7Ta-0.5Si+2Fe Conventional or ST ≈840 ≈720 ≈82-84 ≈15 - [54] Ti-35.3Nb-7.3Zr-5.7Ta-1Si+1Fe Conventional or ST ≈610 ≈650 ≈79-84 ≈7 - [54]
Ti-2.1Al-3.9Zr-10.1Mo-4.9Nb-5.1Ta-8.0Sn-5.8Cr Cooling in different ways - 722-767 41.7-100.9 - - [3] *The values of Ti mass% are proportional to the other elements of the alloy.
†ST - Solution treated
4.1. Strength and ductility
The alloy strength should be high enough to bear the load to which the implant is subjected [3], including tension, compression, bending, and torsion [22]. Tensile strength and fracture toughness are essential properties of materials that are used as hard tissue substitutes. They are responsible for protecting the integrity of the implant and preventing plastic deformation during insertion, thus assuring stability between the implant and the prosthetic components [4]. On the other hand, ductility facilitates several manufacturing processes [29], which is important because implants have complex geometry. It is possible to enhance the strength properties and the plasticity of the alloy by controlling the microstructure and grain size of the material.
The tensile strength of Ti-Al-V is significantly higher than that of cpTi; this is due to the addition of Al and V, which generates changes in solid solution alloy with particle precipitation and transformation from the α to the α + β phase [4]. Similarly, the addition of Al to Ti-Nb and Si to Ti-Nb-Zr-Ta resulted in solid solution strengthening and grain refinement [54,56]. Grain growth suppression promoted by the addition of Si was caused by underpinning of the grain boundaries by silicide intermetallic particles [54]. Likewise, grain refinement by adding Nb to Ti enhances the yield strength and tensile strength 1.5-1.6 times over those of pure Ti [45].
The incorporation of Ta to Ti-Nb-Zr is also shown to be responsible for increasing the ultimate tensile strength and elongation of the alloy, which features an elastic and perfectly plastic material [48]. In a similar alloy composition, the use of a high-pressure torsion process was effective at increasing the tensile strength of the alloy by increasing the density of dislocations and resulted in refinement of the microstructure [16]. On other hand, Ta was responsible for decreasing the alloy strength of Ti-Fe, even though the alloy exhibited high hardness and compressive yield [41].
The incorporation of elements into the alloy will not always result in improved mechanical properties. In the case of Ti-Nb-Sn, the presence of high concentrations of Sn incorporated into the β phase of the alloy tends to reduce both ductility and tensile strength of the alloy [29]. A study conducted by Datta et al. [3] showed that the presence of a larger amount of β-stabilizers might be able to reduce the resistance of the material. This was initially observed in a predictive model, and subsequently verified by manufacturing a Ti-Al-Zr-Mo-Nb-Ta-Sn-Cr alloy, which showed lower strength than Ti-6Al-4V, while two other alloys with lower amounts of β-stabilizers had higher resistance levels. Further studies are
needed to consolidate a thermomechanical alloy process to achieve a β microstructure with smaller fractions of β-stabilizers.
The interstitial content of Ti and its alloys (N, O and C) also is directly related to the strength and ductility of the material [4,51]. The increase in interstitial solutes is deleterious to cpTi toughness reducing the elongation values; nevertheless, the interstitial solutes increase the strength of the material [4]. Intermetallic phases generated by substitute elements, such as Fe, Ni and Cu, are responsible for the same behavior in alloys. This was noted in a study that evaluated two alloys with different concentrations of Cu-Ni, in which the decrease in the fraction of these elements improved the ductility of the alloy, and its increase raised the tensile strength, reaching values of 1050 MPa [51].
The manufacturing process also influences the alloy's characteristics. Cold rolling, aging treatment, annealing, and incorporation of small amounts of ceramic particles into the matrix promote high strength and good ductility of the Ti alloy [16,22,54]. This will ensure the material's biomechanical compatibility with the bone, which is extremely important for the long-term success of dental implants, mainly because they are subject to loads or constant cyclic stresses [19].
The tensile strength of Ti alloys developed by different process ranges from 360 to 3267 MPa (Table 2). As an overall trend, high values of ultimate tensile strength and yield strength were found for the α + β type alloys. It is clear that Zr, Ta, Si, Fe, Al, and Mo have a tendency to increase alloy strength. Elements such as Nb, Sn, and Bi have little influence on this property. The processing method also affects alloy strength, where highest values were found using SPS.
4.2. Fatigue performance
As described in subchapter 4.1., most of the studies investigated the mechanical behavior of Ti alloys using compressive tests rather than fatigue tests. The use of cyclic loading to evaluate the fatigue behavior of dental implant materials is required because it represents a more realistic scenario than pure monotonic loading [57]. Besides, the environment that implants are inserted can influence the fatigue performance by accelerating the initiation of a surface flaw and propagating it to a critical size, leading implant to fracture [58]. Nevertheless, there are few studies evaluating the fatigue behavior of Ti alloys in human-like medium simulating this conditions .
The number of fatigue cycles to failure the Ti–24Nb–4Zr–7.6Sn in 0.9% NaCl solution was larger than that in air, mostly due to the cooling effect of this medium, which
can suppress material softening originated from the temperature increasing during the fatigue test [59]. Additionally, the oxide film formed in NaCl solution may improve the corrosion fatigue resistance of such alloy. Such alloy exhibited a much higher fatigue resistance in the strain-controlled fatigue test than Ti–6Al–4V ELI [59]. A corrosion fatigue test with Ti alloys was carried out in Eagle’s solution. The fatigue strength of Ti-Zr-Nb-Ta-Pd-O-N and Ti-Sn-Nb-Ta-Pd-O at 108 cycles was ≈600 MPa, while for Ti-Al-Nb-Ta was ≈700 MPa. Ti-Mo-Zr-Al alloy showed worse results than previous alloys at 107 cycles [10]. Ti-15Zr discs and
dental implants showed a significantly better fatigue performance when compared to Ti-Grade IV. The fatigue endurance limit was 560 MPa for the alloy and 435 MPa for the cpTi [60]. Meanwhile, Ti–6Al–4V and cpTi exhibited higher fatigue strengths than those of Ti–7.5Mo and Ti–13Nb–13Zr. Nevertheless, Ti–7.5Mo had the best fatigue performance, especially when the strain-controlled fatigue resistance is taken into consideration [61].
Improved fatigue behavior can be achieved by a combination of the material properties, surface properties and design optimization of implants [60]. In this context, the deposition of hard thin coatings, the use of different heat treatments and mechanical processing operations must be used [58]. Cold rolling plays an important role in rising the fatigue resistance by the production of much finer microstructures [59]. Aging treatment was responsible to increase the fatigue limit of Ti–29Nb–13Ta–4.6Zr (≈320 to ≈700 MPa) [62] and the Ti–24Nb–4Zr– 7.6Sn (≈250 to ≈425 MPa) [59] owing to the combination of strength and ductility, which improved the resistance against small fatigue crack initiation and propagation [62]. The surface porosity generated by the casting process may be the main cause of fatigue cracks in alloys [61]. Similarly, the roughness created by surface treatments (e.g. SLA) proved to have a negative effect in the materials fatigue behavior, leading to a higher susceptibility to crack initiation [60] In general, new studies must be developed assessing the fatigue behavior of the Ti alloys under cyclic loading in an human-like medium, such as artificial saliva and simulated body fluid. This is required because the combination of a corrosive environment and cyclic mechanical loading is more prejudicial to most materials used in dental implants than the sum of the individual effects [57]. In addition, thermomechanical processes and surface treatments need to be improved to avoid manufacturing defects that can be stress concentration sites, which may lead to failures initiation.
4.3. Hardness and Elastic modulus
Hardness is defined as the resistance of the material to permanent deformation [26]. It should be low during implant manufacturing to ensure good machinability but at the same time, should present adequate stiffness to avoid shielding the bones from stress [4]. The elastic modulus is an important property for biomechanical interaction between the implant and the bone. Lower elastic modulus values improve stress distribution at the implant–bone interface and decrease bone atrophy [4]. Therefore, the elastic modulus should be as close as possible to the bone to avoid the problem of “stress shielding” [3,37], which is associated with bone resorption around the implant [2,50] and osteoporosis [50]. However, the elastic modulus should not be too low, because it would lead to micro-motion of the metallic implant, resulting in implant loosening, and even failure of the prosthesis [45].
While higher values of alloy hardness are related to the presence of an α phase, its presence can be undesirable for the elastic modulus. The addition of α-stabilizers increases hardness because they act as a substitutional solute, which decreases the atomic mobility of the material [21] or promotes the precipitation of the α phase when the alloy is aged at low temperature [19]. With regard to the elastic modulus, a predictive model concluded that its value should initially increase with rising β content, but will decrease dramatically at higher concentrations [3]. Lower concentrations of β-stabilizers usually generates α + β alloys, which present higher values of elastic modulus due to the presence of the α phase. The suppression of the α phase of an alloy microstructure by the increase of β-stabilizers may be responsible for lower hardness values [26,29], and consequently, the decrease in elastic modulus.
The addition of Zr and Bi to Ti increased the values of hardness when compared with cpTi [7,21,38]; however, with the increase in Zr concentration, the hardness decreased slightly [21,38]. Zr works differently for the elastic modulus. The addition of 5%10% in mass of Zr initially decreased the values; however, concentrations higher than 15% increased the value of this property above that of cpTi [21]. The authors explained the variation by a change in the distance between atoms of the alloy components, which is caused by the higher atomic radius of Zr compared with Ti. This leads to a change in the binding force between atoms that determine the elastic modulus. The incorporation of Nb into a Ti-Zr alloy also showed an increase in hardness, which was raised even more by aging heat treatment, responsible for the formation of a biphasic α + β microstructure [14]. However, this treatment usually increases the elastic modulus to above 100 GPa [54].
The addition of Sn to Ti-30Nb alloy culminated in an alloy with higher amounts of β phase generated by the diffusion of Nb in the Ti matrix, which, besides reducing the elastic modulus, did not change the material's hardness [44]. A Ti-33Nb-4Sn alloy subjected to the cold-rolled and annealing process, resulted in one of the lowest values of elastic modulus in the literature (≈36 MPa), which was possible due to the formation of a β microstructure with low levels of β-stabilizers [18]. Most of these alloys with low elastic modulus have been obtained using d electron alloy theory [32,50,63,64], which takes into consideration the covalent bond strength between Ti and alloying elements and the d-orbital energy, relating to the radius and electronegativities of the elements [64].
Although the cpTi and its alloys present a much lower elastic modulus than stainless steel and Co-Cr alloy [2,16], their values are still very high when compared with those of human bone (10-30 GPa) [4,65]. Regarding elastic modulus, β-type alloys have shown more promising results than cpTi and α and α + β alloys [37,66]. This may be related to the fact that the β phase has a body-centered cubic structure, with the density of atoms in the lattice lower than the hexagonal closed-packed structure found in the α phase [16]. This ensures a higher plastic deformability of the alloy [50]. These different structures have distinctive distances between the atoms, which lead to variations in the atomic-bonding force and, consequently, changes in elastic modulus [21].
Alloy elastic modulus may vary extensively with thermomechanical processing [3,16]. For example, water-quenched samples showed lower elastic modulus values [3,22], which is in accordance with Table 2. This decrease in elastic modulus is related to the inhibition of transformation from β to α phase, as a result of rapid cooling that limits the growth of the α phase [22]. According to the literature, the elastic modulus of Ti alloys varies between 35 and 175 MPa, with emphasis on the β-type alloys, which have values closer to that of bone (Table 2). We highlight the Ti-23.72Nb-4.83Zr-1.74Ta(-Si) and Ti-36Nb-4Sn alloys, which presented lower elastic modulus values. Regarding hardness, cpTi has one of the lowest values, which increases considerably with the addition of the elements Zr, Mo, Fe, and Al. Ti-Fe(-Ta), Ti-Al-V, and Ti-Zr have the highest hardness values.
5. Surface properties
Surface characteristics, such as composition and topography, exhibit high influence on the success of implant treatment. Implants require a surface that promotes osteogenic differentiation and proper mineralization during the initial integration stage [67]. Surface characteristics of the implant material (e.g., surface roughness, surface energy, and substrate
composition) directly or indirectly influence the proliferation and differentiation of bone cells and, consequently, cell adhesion to bone apatite [68,69].
For dental implants, it is desirable that the alloy preserves the topographic micro-roughness while maintaining the hydrophilic surface [70]. Implant surface micro-roughness fits into micrometer (0.11100 µm) and nanometer (1100 nm) scales [71], which also applies for surface morphology (Figure 3). There are indications that micrometer and nanometer levels influence bone response [72], ensuring proper cell attachment, higher bone/implant contact, and greater resistance to torque removal [67,73]. It is expected that nanometric proteins present in the biological environment can easily penetrate into a nanotopography, with positive effects on cellular responses [74].
Rough topographies were found to be associated with the differentiation of osteoblasts while a smooth surface was related with cell proliferation. It is believed that the contacts between cells and rough surface may induce autophagic process by mechanically stimulation, which could lead cells toward differentiation [75]. An implant surface with defined nanoscale features and rough surface results in enhanced secretion of osteogenic markers and specific proteins related with the osteoblast maturation (e.g. osteocalcin, osteopontin, bone sialoprotein, osterix, nictric oxide, transcription factor Runx2, alkaline phosphatase), in which can provide faster and more reliable bone-to-implant contact [74-76]. A surface that induces the maturation of cells is crucial to improve and to accelerate osseointegration.
A previous study produced a Ti-6Al-4V alloy with different levels of surface roughness [71]. The roughness values ranged from 0.25 to 0.87 μm and the adhesion and cell proliferation were sensitive to changes in topography, where increased roughness produced better results. Furthermore, a greater quantity of total protein adhesion was observed on a rough surface, with fibronectin adsorbing up to 10 times more than on the smooth surface. Still need better explanation concerning the role of small differences in surface roughness (Ra < 0.50 μm) and its influence on the cellular response and protein adsorption [69].
Besides, an improved wettability is important to achieve better cell proliferation and adhesion on surfaces of Ti alloy implants [77]. Chemical and physical properties of the surface can influence the wettability of the alloy. The surface is considered hydrophilic when the water contact angle is less than 90◦, whereas the angle above of 90◦ represents an hydrophobic surface [78]. Ti-45Nb was shown to have an enhanced hydrophilicity with a lower contact angle (81.75°) when compared with cpTi (96.46°) [45]. Nevertheless, the Ti-50Zr alloy presented the higher surface energy and the lowest surface roughness (≈37 mN/m,
0.17 µm) than cpTi (≈34 mN/m, 0.20 µm) and the Ti-50Nb alloy (≈32 mN/m, 0.46 µm) [69]. This properties associated with the substrate composition assured a better in vitro biological profile (cells adhesion and proliferation) for Ti-Zr alloy. The Ti showed a slightly rougher topography than Ti–Ta–Nb–Zr, despite the alloy presented a more enlarged surface area. Both materials reveal similar water contact angles, which were characteristic of hydrophobic surfaces [1].
Figure 3. Schematical representation of the micro and nanoscale of implants and their effect
in topographic features. Representative micrographs of microscale surfaces can be seen for machined Ti–6Al–4V–ELI (M) and acid etched in HCl (Ac) samples. Nanoscale features can be observed on the acid etched combined with alkaline treatment (AcAk) surfaces, in which a homogenous sponge-like structure is noticed. The nano-scale topography revel a decrease in water contact angle and consequently improvement in surface energy and hydrophilicity. SEM micrographs were reprinted from Mater Sci Eng C, 51, Oliveira DP, Palmieri A, Carinci F, Bolfarini C, Gene expression of human osteoblasts cells on chemically treated surfaces of Ti – 6Al – 4V – ELI / Results, 250, Copyright (2015), with permission from Elsevier.
Surface modifications by various treatment methods have also shown excellent results. Treatment by mechanical friction achieved different grain sizes on the Ti-25Nb-3Mo-3Zr-2Sn alloy [79]. The nano-grained alloy presented the lowest surface roughness associated
with better wettability (7.0 nm; 50.4°) than ultrafine-grained (7.4 nm; 64.1°) and coarse-grained alloy (7.1 nm; 67.8°). Nanoscale grains (30 nm) showed better cell responses and higher protein adsorption as compared with grain sizes of 90 μm and 180 nm [79]. Two acid-etched surfaces associated with alkaline treatment was responsible for changing the topography of Ti–6Al–4V–ELI from micro- to nano-scale with a surface micro-roughness (0.14-0.48 µm), improving the wettability (≈70° to ≈35°) and surface composition of the material [80]. Some of the surfaces obtained are shown in Figure 2. It can be clearly seen a modification of morphology between the machined surface (M) and acid-etching (Ac) for those that were treated with an alkaline solution (AcAk), in which presented an effective osteoconductive behaviour.
A previous study [74] achieved a nanotopography by electrochemical anodization in Ti-6Al-7Nb alloy. The contact angle was dramatically decreased from 61.4° (machined surface) to 14.8° (treated surface). On the other hand, a nanoporous surface (pores <15 nm) created by the anodization of Ti-25Nb-25Zr alloy did not changed its roughness (121 nm) and contact angle (56°) when compared with the machined surface (123 nm; 59°) [81]. Despite that, the morphology alteration and surface composition was able to ensure the adsorption of albumin and fibronectin, as well as significantly improving cellular responses such as adhesion, migration, proliferation and mineralization of mesenchymal cells. Rough or porous surfaces present reinforced biomechanical properties due to the larger area of contact between recently formed bone and the implant surface, stabilizing the connection between the implant and the surrounding tissue through microscale mechanical interlocking [67].
A hydrothermal treatment using an alkaline Ca-containing solution combined with a simple post-heat treatment, seems to be an effective way to enhance cell viability and differentiation on Ti-13Nb-13Zr surfaces via increased surface hydrophilicity (contact angle = 21°), possibly by increasing the surface area at the nanoscale and the formation of anatase structure, without markedly altering surface morphology [82]. The presence of anatase phase proved to be important in increasing the surface energy and the surface reactivity [78].
It has already been observed that a cpTi surface that combines a submicron roughness with high hydrophilicity promotes cell differentiation and maturation with osteogenic capacity, besides ensuring greater cell adherence [83,84]. Moreover, few studies have compared the roles of surface characteristics of different alloys in the material properties and tissue response. Most research studies the surface properties only when it is related to surface treatments. From this perspective, further studies should deeply investigate the influence of
