Surface Modification of Titanium Dental Implants 3031215648, 9783031215643

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Surface Modification of Titanium Dental Implants

Karan Gulati Editor

Surface Modification of Titanium Dental Implants

Editor Karan Gulati School of Dentistry The University of Queensland Herston, QLD, Australia

ISBN 978-3-031-21564-3    ISBN 978-3-031-21565-0 (eBook) https://doi.org/10.1007/978-3-031-21565-0 © The Editor(s) (if applicable) and The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 This work is subject to copyright. All rights are solely and exclusively licensed by the Publisher, whether the whole or part of the material is concerned, specifically the rights of translation, reprinting, reuse of illustrations, recitation, broadcasting, reproduction on microfilms or in any other physical way, and transmission or information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed. The use of general descriptive names, registered names, trademarks, service marks, etc. in this publication does not imply, even in the absence of a specific statement, that such names are exempt from the relevant protective laws and regulations and therefore free for general use. The publisher, the authors, and the editors are safe to assume that the advice and information in this book are believed to be true and accurate at the date of publication. Neither the publisher nor the authors or the editors give a warranty, expressed or implied, with respect to the material contained herein or for any errors or omissions that may have been made. The publisher remains neutral with regard to jurisdictional claims in published maps and institutional affiliations. This Springer imprint is published by the registered company Springer Nature Switzerland AG The registered company address is: Gewerbestrasse 11, 6330 Cham, Switzerland

Preface

This book aims to present advances in the surface modification of titanium dental implants, from the macro and micro to nanoscale surface modifications, focusing on advanced bioactive and nano-engineered dental implants. Through eight chapters, the book covers a wide array of topics that provide an improved understanding of the fabrication, bioactivity, therapy, and stability of modified titanium dental implants. Overall, the book significantly contributes to the ever-changing field of dental implants. From the basics of why the surface modification is needed to the advanced state-of-the-art electrochemically anodized nanostructures fabricated on implants, the book covers the domain of dental implants from a clinical, materials science, and nano-engineering perspective. The first chapter, “Titanium: The Ideal Dental Implant Material Choice”, details the ideal characteristics of titanium that make it the most popular dental implant material choice. While modern titanium-based dental implants provide optimum treatment outcomes in healthy conditions, enhanced bioactivity and therapy are needed to ensure long-term success in compromised patient conditions. The need to modify the implant surface (especially in compromised conditions that present a significant therapeutic challenge) is thoroughly reviewed in the chapter “Titanium Dental Implants in Compromised Conditions: Need for Enhanced Bioactivity and Therapy”. Advances in dental implants have evolved from macro- to micro- to nanoscales. The chapter “Macro to Micro: Surface Modification of Titanium Dental Implants” is devoted to various macro and microscale modifications performed on titanium-based dental implants. The next generation of dental implants has controlled nanotopography that augments the bioactivity and therapy toward achieving timely integration and long-term success. The fourth chapter, “Nano-scale Surface Modification of Dental Implants: Fabrication”, compiles various nano-engineering tools and techniques that enable effective nanoscale surface modification of titanium dental implants, focusing on easy, scalable and cost-effective electrochemical anodization that fabricates controlled nanotopographies on titanium implants. Titanium dioxide (or titania) nanotubes (like nanoscale test tubes) can be fabricated on dental implants via anodization with excellent control over their dimensions. The nanotube-modified implants offer v

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Preface

various functionalities, including enhanced bioactivity and local therapy. The fifth chapter, “From Micro to Nano: Surface Modification for Enhanced Bioactivity of Titanium Dental Implants”, and the sixth chapter, “Local Therapy from Nano-­ engineered Titanium Dental Implants”, categorically explain the strategies employed to orchestrate implant integration and achieve tailored local therapy from anodized nanotubular dental implants, respectively. The seventh chapter, “Mechanical Stability of Anodized Nano-engineered Titanium Dental Implants”, focuses on the mechanical stability considerations of anodized dental implants. Finally, the eighth chapter, “Cytotoxicity, Corrosion and Electrochemical Stability of Titanium Dental Implants”, presents the advances and challenges associated with the cytotoxicity and corrosion of modified and nano-engineered dental implants. All chapters present clinical translation challenges and recommend future directions to advance the domain, ensuring long-term success, even in compromised patient conditions. The book is interdisciplinary and will profoundly interest a broad audience, including dentists, undergraduate/postgraduate/research students, academics, and material/biomaterial scientists. Since the book describes cutting-­edge nanotechnology advances in dental implants, it will be valuable to entrepreneurs aiming to understand the next generation of nano-engineered implants. Herston, QLD, Australia

Karan Gulati

Contents

 Titanium: The Ideal Dental Implant Material Choice ��������������������������������    1 Himanshu Arora Titanium Dental Implants in Compromised Conditions: Need for Enhanced Bioactivity and Therapy������������������������������������������������   23 Necla Asli Kocak-Oztug and Ece Irem Ravali Macro to Micro: Surface Modification of Titanium Dental Implants ��������   61 Yifan Zhang, Shuai Li, Ye Lin, Ping Di, and Yan Liu  Nano-scale Surface Modification of Dental Implants: Fabrication������������   83 Ruben del Olmo, Mateusz Czerwiński, Ana Santos-Coquillat, Vikas Dubey, Sanjay J. Dhoble, and Marta Michalska-Domańska From Micro to Nano: Surface Modification for Enhanced Bioactivity of Titanium Dental Implants ������������������������������������������������������  117 Tianqi Guo, Sašo Ivanovski, and Karan Gulati  Local Therapy from Nano-engineered Titanium Dental Implants��������������  153 Anjana Jayasree, Sašo Ivanovski, and Karan Gulati Mechanical Stability of Anodized Nano-­engineered Titanium Dental Implants ������������������������������������������������������������������������������  199 Divya Chopra and Karan Gulati Cytotoxicity, Corrosion and Electrochemical Stability of Titanium Dental Implants��������������������������������������������������������������������������  219 Tianqi Guo, Jean-Claude Scimeca, Sašo Ivanovski, Elise Verron, and Karan Gulati Index������������������������������������������������������������������������������������������������������������������  255

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Titanium: The Ideal Dental Implant Material Choice Himanshu Arora

Abbreviations Å Angstrom cpTi Commercially pure titanium GPa Gigapascal HA Hydroxyapatite MPa Megapascal PEEK Polyether ether ketone Ti-6Al-4V Titanium aluminium vanadium alloy TiZr Titanium zirconium alloy ZrO2 Zirconium oxide

1 Introduction The relationship between edentulism and dentistry is as long as dentistry itself. Since then, dentists worldwide have been busy finding novel ways to limit or restore edentulism. Edentulism, whether partial or complete, has seen an increasing trend in the last few decades, with reports estimating around 120 million Americans are missing at least one tooth and approximately 35 million are completely edentulous (American College of Prosthodontists, 2022). Consequences of partial or complete edentulism range from functional, esthetic, physical, and psychological limitations affecting the overall oral health related quality of life. Various treatment options have evolved to solve this health crisis over the past few centuries with oral H. Arora (*) School of Dentistry, The University of Queensland, Herston, QLD, Australia e-mail: [email protected] © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 K. Gulati (ed.), Surface Modification of Titanium Dental Implants, https://doi.org/10.1007/978-3-031-21565-0_1

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implantology the latest addition to this list of options. Till date no treatment option is complete. Each treatment option must compete with the natural dentition in performance and long-­term success. This drives the current research and advances in oral implantology to find the best performing dental implant. Dental implant is defined as a prosthetic device made of alloplastic material(s) implanted into the oral tissues beneath the mucosal and/or periosteal layer and on or within the bone to provide retention and support for a fixed or removable dental prosthesis; a substance that is placed into and/or on the jawbone to support a fixed or removable dental prosthesis (The Glossary of Prosthodontic Terms: Ninth Edition, 2017). Major advances have occurred over the last few decades in the clinical use of oral and maxillofacial implants. Latest statistics on the use of dental implants reveal that, in the United States alone, an estimated 5 million implants are placed annually, and a total of 15–20 million implants are placed worldwide (Misch & Misch, 2015). Dental implants are currently used to replace missing teeth, rebuild the craniofacial skeleton, provide anchorage during orthodontic treatments, and even aid in new bone formation in the process of distraction osteogenesis. In modern dentistry, the dental implant is the one of the best tooth replacement options for nearly all situations where a tooth is missing or is failing. The primary reason for this is the extremely high success rate achieved with dental implants. Saving teeth at all costs is no longer the norm because of the unpredictability of the longevity of heroic dentistry. In other words, preserving bone and tissue regeneration are now considered to be more important than trying to prolong tooth retention (Massa & Von Fraunhofer, 2021). One of the main reasons for the high success rate of dental implants is their ability to integrate with bone in the oral environment (Misch, 2008). The goal of placement of endosseous dental implants is to achieve osseointegration of the bone with the implant in order to support a prosthesis (Brånemark et  al., 1983; Branemark et  al., 1977). The physical, chemical, and biological properties of dental implant materials along with their surface characteristics are key factors in their success (Binon, 2000; Buser et  al., 1991). A wide variety of materials has been used for these implants, but only a few promote osseointegration and biointegration (Weiss & Weiss, 2001). Titanium and titanium alloys have been the most widely used of these materials. This chapter will look into the historical aspect of dental implant materials, drawing comparisons with the modern-day contemporary materials in an attempt to arrive to a conclusion ‘why titanium is the most suitable dental implant material?’.

2 Osseointegration The godfather of modern implants was a Swedish physician and anatomical and experimental biologist named Per-Ingvar Branemark. He studied bone healing response and regeneration in the 1950s and in order to observe the functioning of bone marrow in vivo, he used titanium to make a chamber that could be inserted into

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Fig. 1 Radiographic image of the original titanium screw placed in rabbit tibial bone by P. I. Branemark showing the integration of the implant with bone that led to discovery of osseointegration. (Albrektsson et al., 2017)

rabbit legs to allow microscopic visualization of vital processes (Fig. 1). After a few months-long series of investigations, he sought to retrieve the chamber for reuse and found to his annoyance that it could not be removed from the rabbit bone (Branemark, 1983). Branemark reportedly was not struck by the significance of this turn of events until sometime after 1960, when he accepted a professorship in the Department of Anatomy at Gothenburg University. There, using an adaptation of the titanium chamber placed in the upper arms of human volunteers, he and his team investigated the workings and structure of human blood cells under a number of conditions. This work yielded a great deal of information about the nature of blood, and it showed the researchers that the titanium serving as lens casings appeared uniquely compatible with the human soft tissue and skin, provoking no adverse immunological reactions. At this point, Branemark began to contemplate using titanium for medical applications (Albrektsson et al., 2017). As this understanding advanced, Branemark believed it necessary to coin a new term to refer to the in-growth of the bone into the threads and crevices of titanium. He finally settled upon “osseointegration,” derived from the Latin words os (bone) and integro (to renew) (Branemark et al., 1977). The first and the most important event that occurs when an implant is placed in host tissue is surface adsorption of proteins. The amount, composition, and conformational changes of the adsorbed proteins influence the entire biological response to the material, including antigenic response, attachment, and growth of cells. The host response to implants placed in bone involves a series of cell and matrix events, ideally culminating in tissue healing that is as normal as possible and that ultimately leads to intimate apposition of bone to the biomaterial, i.e. an operative definition of osseointegration. For this intimate contact to occur, gaps that initially exist between bone and implant at surgery must be filled initially by a blood clot, and bone damaged during preparation of the implant site must be repaired (Szmukler-­ Moncler et al., 1998). The material used to construct oral implants plays a major role in the host response and has been one of the most researched topics in the field of oral implantology.

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3 Materials for Endosseous Dental Implants Today, the goal of the placement of endosseous implants is to achieve osseointegration at the surface of the implant. Osseointegration, as defined by Branemark, is the direct contact of the loaded implant material with living bone (Brånemark et al., 1983) (Fig. 2). The concept of osseointegration has been developed largely from the work of Branemark and was introduced in 1982 after several decades of animal work and at least a decade of work in humans (Fenton, 1992; Albrektsson et al., 2017). This concept represented a fundamental shift away from the prevailing dogma of the time. Previously, implant materials were sought which would act as inert substances, usually eliciting a fibrous encapsulation around them (Lemons, 1990). However, by definition, osseointegration demands the absence of a fibrous layer (Meffert et al., 1992), and implies that the biological response of the bone is not one of inertness toward a foreign material but rather one of integration of the material with the bone as if it were part of the body. By today’s standards, the presence of a fibrous layer between bone and implant indicates failure of the implant (Albrektsson et al., 2017; Buser et al., 2017). In spite of these definitions, there is still controversy about what osseointegration really represents. For some, osseointegration does not represent an

Fig. 2  Histological photomicrograph showing a direct contact between a titanium screw/implant (black) and bone tissue using hematoxylin-eosin staining technique. (Albrektsson et al., 2017)

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‘advantageous’ response of the body to the material, but simply the lack of a negative response (Stanford & Keller, 1991). One of the key requisites for osseointegration is compatibility between bone tissue and the implantable material/device. Biocompatibility has traditionally been concerned with implantable devices that have been intended to remain within an individual for a long time. The selection criteria for implantable biomaterials involves as a list of events that has to be avoided, on the basis that they would be non-toxic, non-immunogenic, non-thrombogenic, non-carcinogenic, non-irritant and so on, such a list of negatives becoming, by default, the definition of biocompatibility (Williams, 2008). A wide variety of materials has been used for endosseous implants. These could be divided into: • Materials of historical interest • Currently used materials

3.1 Materials of Historical Interest These include various ceramics, polymers, and metals which have been used clinically in the past but are not being used currently due to their disadvantages/complications or the advent of new and better materials. Ceramics Carbon is a ceramic which was introduced as an endosseous implant in the 1970s (Lemons, 1990). Most of the vitreous (amorphous) carbons were used as coatings on stainless steel cores since the bulk form was too brittle (Albrektsson et al., 1986). To increase the mechanical properties of the carbons, silicon was added to form CSi ceramics. In addition, forms with isotropic (crystalline) properties, such as low temperature isotropic (LTI) and ultra-low temperature isotropic (ULTI), were developed and had higher strengths, moduli and better toughness (Kent & Bokros, 1980). The low corrosion and lack of toxic elements of these implants were viewed as advantageous initially, but the biological response was far inferior to today’s standards, and many of these implants were exfoliated. Five-year survival rates for these implants, even under the best conditions, were 24–65% (Mah, 1990; Albrektsson et al., 1986). Other ceramics have also been used as endosseous implants. Alumina, hydroxyapatite and tricalcium phosphate were introduced in the 1960s and 1970s (Mah, 1990), although the brittleness of the pure forms was not acceptable for most implants (Williams, 1981b). The single crystal sapphire is a form of alumina which has sufficient strength to be used as a bulk material in some clinical situations. Sapphire dental implants (single crystals of aluminium oxide) have been used since late 1970s, but there are few well-documented long-term follow-up studies. Such materials lack the strength and abrasion requirements of an implant material (Kawahara et al., 1980). The biological response to sapphire can be quite favourable

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and osseointegration probably occurs. The clinical success rates for sapphire implants range from 69% to 91% after 5 years (Albrektsson et al., 1986). Fartash et al. (1990) in their experimental study placed nine single crystal sapphire dental implants bilaterally into pre-extracted areas in the lower jaw of two beagle dogs. Implants were analysed after 180  days in situ. Eight implants were stable, and radiographs disclosed complete bone healing. The ninth implant was mobile and surrounded by a non-mineralized connective tissue capsule containing bundles of collagen. Histometric analysis of the alveolar bone surrounding the stable implants revealed that the value of the bone contact surface ranged from 37.1% to 86.9% at the light microscopic level. Single-crystal alumina, which has good mechanical properties and superb biocompatibility, has been used successfully for screw types of implants but cannot reproduce the physiologic function of natural teeth in the free-standing form. Therefore, a porous alumina dental implant was fabricated for free-standing applications such that bone ingrowth would provide additional stability (Mah, 1990; Williams, 1981b). Yamagami et al. (1988) placed porous alumina dental implants in free-standing form and bearing occlusal stress in the jaw of rhesus monkeys for 4, 6, and 8 months. The porous alumina dental implant was designed with a polished cylindrical core of single-crystal alumina, an outer porous root layer 1 mm thick, and a smooth apex of poly-crystalline alumina 3–5  mm. The 20  mm long implant used in this animal experiment had a porous root portion 4  mm in diameter and 7  mm in length. Implantations were observed from 4 to 8 months. The implants were free-standing throughout the examination while bearing occlusal stress. Fourteen of the 15 implants were considered successful. Radiographs showed prolific new calcified bone growth at the sites of the porous alumina root portions. These data demonstrated that secure bone fixation had been achieved and that a good biologic seal was provided at the gingival interface. Ceramics can make up the entire implant, or they can be applied in the form of a coating onto a metallic core. Low flexural strength and various degrees of dissolution/solubility of an all-ceramic implant make coating the application of choice in the field of implant dentistry (Wataha, 1996; Piconi et al., 2003). Coatings can be dense or porous, depending on the chemical composition of the parent material and the coating method employed. The goal is to achieve strong adherence between the coating and the metallic core that withstands functional loading and avoid fragmentation. Hydroxyapatite (Ca10(PO4)6(OH)2), tricalcium phosphate (Ca3(PO4)2), and bioglasses are some of the more commonly used bioactive ceramics, which possibly develop a chemical bond of a cohesive nature with bone (Wataha, 1996; Lacefield, 1998). Hydroxyapatite and tricalcium phosphate are still used today as coatings on metallic cores with good biological response. Bioglass is a complex glass which has been researched for used as an implant material. Its bio integration with bone seems to result from dissolution of the ceramic surface to give a silica-rich gel layer covered by a layer rich in calcium and phosphorous. These ceramic layers seem to merge chemically with the bone (Hench & Wilson, 1984).

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Some of the concerns associated with HA-coated implants were reviewed by Biesbrock and Edgerton (Biesbrock & Edgerton, 1995) and included microbial adhesion, osseous breakdown, and coating failure. However, the authors suggested that in cases where more rapid and enhanced bone implant contact is needed, such as in type IV bone (low quality porous bone), grafted bone sites, or when short implants are indicated, HA-coated implants may be preferable. Caulier and co-­ workers (Caulier et al., 1997) found improved performance with threaded calcium phosphate–coated implants placed in less mineralized trabecular bone, although the thickness of the coating decreased over time. While the clinical success of HA-coated implants has been reported to be comparable to non-coated implants (Alsabeeha et al., 2012), various concerns have been reported with their use. The degradation of ceramic coatings has been a point of controversy (Lozada et  al., 1993) and concerns have been expressed about their long-term stability and success (Zablotsky, 1992). The enhanced bacterial susceptibility of the HA coating due to the surface roughness has been a concern when compared to titanium implants (Ong & Chan, 2017). Polymers A variety of polymers has been used as endosseous implants, but their inferior mechanical properties or poor biological response have limited their use (Lemons, 1990). The advantage of polymers is their ability to be easily fabricated into the desired shape. In general, the polymer implants promoted a fibrous response even when coated with carbon (Williams, 1981b). Metals Metals are probably the oldest form of material used for dental implants and are still by far the most common type of materials used today. A diverse number of metals and alloys have been used as endosseous implants. The gold-based alloys were among the first alloys to be used for implants, probably because these alloys were available in dentistry and the technology existed to cast them (Lemons, 1990). As endosseous implants, they promoted a fibrous interface with bone, and were therefore supplanted by stainless steel and tantalum in the 1940s and 1950s (Mah, 1990). Cobalt chromium alloys were also developed and used as endosseous implants during this time. In hindsight, however, the fundamental problem with all these metals and alloys was the fibrous response which they promoted with bone. By today’s standards, none of these materials appropriately osseointegrate, probably because of their inferior corrosion in the body and release of elements into the tissues. All these metals have been largely replaced today by titanium or titanium alloys.

3.2 Currently Used Materials The materials being used, almost exclusively currently, for dental implants are titanium and its alloys which are discussed in detail in the next section. Various other metal alloy combinations, polymers, and ceramics have been proposed and/or

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potentially used as dental implant materials. Two of them warrant a discussion: Polyetheretherketone (PEEK) and Zirconia. PEEK is an organic synthetic polymeric material which is biocompatible and has good chemical resistance. Young’s modulus of PEEK material in pure form is 3.6 GPa, carbon-fiber-reinforce PEEK (CFR-PEEK) is around 18 GPa which is very close to bone (14 GPa) (Mishra & Chowdhary, 2019). When compared to titanium and other metal alloys, PEEK has been reported to exhibit less stress shielding when used as an implant in load bearing situations (Lee et  al., 2012). Although initial reports have been promising, there is very limited clinical research available on the use of PEEK as a dental implant material. More research is needed in this interesting material before it could be used as a potential alternative to titanium dental implants (Mishra & Chowdhary, 2019). Zirconia has gained considerable interest in implant dentistry over the last decade (Chopra et al., 2022). Zirconium dioxide (zirconia) ceramics with improved properties have been introduced as an alternative material to aluminium oxide implants which were withdrawn from the market in the early 1990s. Currently, tetragonal zirconia polycrystal, particularly 3 mol% yttrium oxide (yttria) -stabilized zirconia, is the ceramic of choice for dental implants (Kelly & Denry, 2008). The white, opaque color of zirconia, along with early reports of good biocompatibility and low affinity to bacterial plaque, make it a material of interest in biomedical sciences (Cionca et al., 2017). Zirconia also exhibits several promising physical and mechanical properties, including low thermal conductivity, high flexural strength (900–1200 MPa), favourable fracture resistance, as well as wear and corrosion resistance. A phenomenon termed phase transformation toughening gives zirconia its excellent properties (Sanon et al., 2015). It stops crack propagation resulting from the transformation of zirconia from the tetragonal phase into the monoclinic phase and the consequent 4% volume expansion and induction of compressive stresses. However, one of zirconia’s negative properties is its low-temperature degradation or aging. In the presence of water or water vapor, slow transformation from the tetragonal phase into the monoclinic phase leads to slow development of roughness, thus producing progressive deterioration of the material (Kelly & Denry, 2008). Zirconia implants have several advantages over the gold standard titanium implants. Their opaque colour is beneficial in the aesthetic regions of the mouth. They have a reduced affinity to bacterial plaque and have been reported to have more favourable soft-tissue integration as compared to titanium implants (Roehling et al., 2019). It has been established that micro-rough ZrO2 implants are equivalent to the ‘gold standard’ Ti micro-rough implants in terms of osseointegration capacity (Janner et al., 2018). Although zirconia implants have been reported to have outcomes comparable to titanium implants in in-vitro studies, clinical reports have failed to replicate these findings. A recent systematic review reporting on zirconia implants estimated an overall survival rate of zirconia one- and two-piece implants was 92% (95% CI: 87–95) after 1 year of function (Hashim et al., 2016). Interestingly, early failure of one-piece zirconia implants ranged between 1.8% and 100%, with the overall early

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failure rate calculated at 77% (95% CI: 56–90). Despite the progress made, significant research gaps remain, including mechanical stability and local cytotoxicity concerns. The next generation of zirconia implants will be nano-engineered with controlled bioactivity to accelerate implant integration, even in compromised patient conditions (Chopra et al., 2022).

4 Titanium and Its Alloys Titanium and its alloys are the most common materials used for endosseous implants and are the materials of choice according to some researchers (Massa & Von Fraunhofer, 2021). The two forms of titanium used for endosseous dental implants are commercially pure titanium (cpTi) and the titanium alloy Ti6A14V. These alloys are basically dilute alloys of oxygen and titanium, with other elements added.

4.1 Titanium in Its Elemental Form Titanium (Ti) exists as a pure element listed in the periodic table with an atomic number of 22 and anatomic weight of 47.9. It is the ninth most abundant element and the fourth most abundant structural metallic element in the earth’s crust, following aluminium, iron, and magnesium. Pure titanium is a rather soft nonmagnetic material. The principal titanium ore reserves, rutile and ilmenite, are found in abundance in the United States, Canada, and Australia. Though the bulk of titanium ore is mined for use in the pigment industry, 5–10% of titanium ore is used to produce cp titanium and titanium alloys (Bannon & Mild, 1983). The element was discovered by Wilheim Gregor, a clergyman, who found the metal in a “black magnetic sand” in Cornwall in 1791. Three years later, Klaproth found a rutile that was the oxide of a new metal he named titanium, after the Greek Titans. He recognized that this metal was identical to the material Gregor had discovered (Williams, 1981a). The commercial production of titanium was not viable until the 1930s when the refining process was mastered. In 1925, van Arkel refined the ore using titanium tetraiodide, producing a metal with acceptable properties and ductility. In the 1930s, Krol developed commercial extraction procedures that are still used today (Williams, 1981a). Titanium is produced by heating titanium ore (rutile, ilmenite) in the presence of carbon and chlorine and then reducing the resultant TiCl4, with molten sodium to produce a titanium sponge. This sponge is then fused under vacuum or in an argon atmosphere into ingots composed of the familiar metal (Cotton & Wilkinson, 1971). Titanium will burn in air and is the only metal that will burn in the presence of nitrogen. Pure titanium undergoes a crystallographic change on heating to 882 °C. This type of transformation occurs in many materials and produces properties significantly different from those of the original state.

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4.2 Titanium in Alloyed Form Titanium is a dimorphic alloy with two phases: α and ß phase. α-Titanium is hexagonal close-packed (hcp) crystal lattice, and ß-Titanium is body-centered cubic (bcc) lattice. Titanium alloys of interest to dentistry exist in three forms: alpha, beta, and alpha-beta. These types originate when pure titanium is heated, mixed with elements such as aluminium and vanadium in certain concentrations, and then cooled. This treatment produces true solid solutions. These added elements are said to act as phase-condition stabilizers (Noort & Barbour, 2013). Aluminium has been called an alpha-phase condition stabilizer. Aluminium also serves to increase the strength and decrease the weight of the alloy. Vanadium has been called a beta-phase stabilizer. As aluminium or vanadium is added to Ti the temperature at which the alpha-to-beta transformation occurs changes to a range of temperatures. In these ranges, both the alpha and beta forms may exist. The alloy form desired is maintained at room temperature by quenching the alloy from the temperature at which the desired form exists. These combination alloys, especially alpha-beta, may be heat treated to increase their strength. One of such alloys is Ti-6Al-4V, also known as Grade V titanium alloy. It is composed of 6% and 4% of aluminium and vanadium, respectively, together with addition of maximum 0.25% of iron and 0.2% of oxygen. The remaining of the alloy is titanium (Liu et al., 2017). Another currently used titanium-based alloy for dental implants is an alloy of Titanium and Zirconium (Ti-Zr). Zirconium belongs to Group 4 (according to new IUPAC name) in the periodic table, which is the same as titanium and hafnium, have similar chemical structure and properties. Thus, they have been recognized as non-­ toxic and non-allergic. Zirconium is usually used in dentistry in its ceramic form (ZrO2). Binary Ti-Zr alloys have been developed to improve bioactivity, biocompatibility, and mechanics of titanium for biomedical application. Currently these alloys are marketed under the name Roxolid (Straumann, Basel, Switzerland) and have been shown to significantly improve osteoblast adhesion (Sista et al., 2013). A recent systematic review reported that TiZr implants exhibited similar soft tissue behaviour when compared with Titanium and Zirconia implants (Fernandes et  al., 2022). Various currently used materials for dental implants have been summarised in Table 1.

4.3 Physical Properties of Titanium and Its Alloys The atomic structure of titanium is 1s2, 2s2, 2p6, 3s2, 3p6, 3d2, 4s2. The lightly held 3d2 and 4s2 electrons are highly reactive and rapidly form a tenacious oxide that is responsible for the metal’s biocompatibility. At temperatures up to 882  °C, pure titanium exists as a hexagonal close-packed atomic structure (alpha phase). Above that temperature, the structure is body-centred cubic (beta phase) with the metal finally melting at l665 °C (Park & Lakes, 1992). The element titanium dissolves several other elements to form alloys. Among these are silver, aluminium, arsenic, copper, iron, gallium, uranium, vanadium, and

Yttria stabilized tetragonal zirconia

Polyetheretherketone

ZrO2

PEEK

Ti-Zr

Alloy with 6% Aluminium and 4% Vanadium Titanium 85% Zirconium 13–15%

Composition Oxygen content (0.4%)

Ti-6Al-4V

Dental implant material Grade IV cp-Ti

3–18

200

98

85–115

Elastic modulus (GPa) 110 Disadvantages Aesthetic issues, Corrosion Possible hypersensitivity to released Ti Improved strength Tissue toxicity due to Al and V leakage Biocompatibility, Aesthetic issues, Improved strength Corrosion, Possible hypersensitivity to released Ti Biocompatibility, Low temperature Aesthetics, degradation, Reduced affinity to Limited long-term plaque, clinical data High flexural strength Biocompatibility, Very limited clinical Aesthetics data

Advantages Biocompatibility, Machinability, Soft tissue integration

Dental implants and abutments

Dental implants and abutments

Implant abutments Dental implants

Clinical use Commercial dental implants

Table 1  Currently used materials for dental implants and their advantages/disadvantages and clinical uses

Mishra and Chowdhary (2019)

Chopra et al. (2022), Roehling et al. (2019), Kelly and Denry (2008) and Janner et al. (2018)

Liu et al. (2017) and Sista et al. (2013)

Liu et al. (2017)

References Noort and Barbour (2013), Guo et al. (2021a) and Darvell (2018)

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zinc. The addition of trace amounts of carbon, oxygen, nitrogen, and iron will markedly improve the mechanical properties of pure titanium (Weast & Astle, 1981). Most commercially or surgically pure titanium products have some of the trace elements present. Cp-Ti are categorized into four grades depending on impurity content (e.g., carbon and oxygen) under the International Organization for Standardization (ISO) standards 5832-2. The different grades vary mostly in the oxygen content and have various corrosion resistance ability, ductility, and strength. • • • •

cp Grade I titanium cp Grade II titanium cp Grade III titanium cp Grade IV titanium

Grade 4 cpTi has the most oxygen at 0.4% (Fig. 3). Nitrogen, carbon, hydrogen and iron are also present, but do not vary much between grades while iron is added for corrosion resistance. The mechanical and corrosion properties of these alloys may change significantly with relatively small changes in the concentrations of the minor elements. Alloying of titanium helps to enhance some of its properties like strength, corrosion resistance, machinability, as well as lower the modulus of elasticity (Liu et al., 2017).

Fig. 3  The strength and oxygen contents variation of various grades of commercially pure titanium. (Darvell, 2018)

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4.4 Mechanical Properties of Titanium and Its Alloys In general, titanium is a good choice for intraosseous applications not only due to the biocompatibility, but also mechanically. Titanium could be processed and machined in a rapid manner such that the shapes and sizes could be easily controlled. The elastic modulus of cpTi is 110 GPa, which is half that of stainless steel or cobalt chromium alloy. Tensile properties of cpTi depend significantly on the oxygen content and, although the ultimate tensile, proof stress and hardness increase with increased oxygen concentration, this is at the expense of the ductility. The alloys most commonly used for dental implants are of the alpha-beta variety. Of these, the most common contains 6% aluminium and 4% vanadium (Ti-6Al-4V). After heat treatment these alloys possess many favourable physical and mechanical properties that make them excellent implant materials. They are light, strong, and highly resistant to fatigue and corrosion. Although they are stiffer than bone, their modulus of elasticity (stiffness) is closer to bone than any other important implant metal; the only exception is cpTi. This property leads to a more even distribution of stress at the critical bone-implant interface because the bone and implant will flex in a more similar fashion (Liu et al., 2017). Titanium alloys are largely used in industrial applications such as jet engines, air frames, and the aerospace industry, which require high strength-to-weight ratios and good corrosion resistance. Other applications include chemical processing, nuclear waste containment, heat exchange units, seawater desalinization, marine equipment, deep-well drilling, and food processing situations that require resistance to corrosion. When compared with cpTi, Ti-6Al-4V has an excellent yield strength and fatigue properties, excellent corrosion resistance ability and lower elastic modulus (Liu et al., 2017). However, Ti-6Al-4V alloy has the disadvantage of low wear resistance and low shear strength (Kong et al., 2011) that could impair the usage as implant and as in screw form. Such a phenomenon is termed as ‘stress shielding effect’ (Niinomi & Nakai, 2011), which is a due to the stiffness mismatch between implant material and surrounding bone. Suitable surface treatments have been recommended to improve this situation.

4.5 Biological Properties of Titanium and Its Alloys 4.5.1 Oxide Coating Most metals form oxide layers when exposed to the atmosphere. The nature of this oxide depends on the metal and the conditions under which it was oxidized. Anything that comes in contact with the implant surface has the potential to change it. Assuming that the physiologic conditions of the body remain fairly constant, the behaviour of a metal in the body depends on the character of the oxide layer.

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Pure titanium, theoretically, may form several oxides. Among these are TiO, TiO2, and Ti2O3. Of these, TiO2 is the most stable and therefore the most commonly used under physiological conditions. These oxides form spontaneously on exposure of Ti to air. Within a millisecond of exposure to air, a 10  Å oxide layer will be formed on the surface of pure titanium (Kasemo, 1983). Within a minute, this layer can become 100 Å thick. The metal may be passivated in this way, although the U.S.  Food and Drug Administration (FDA) requires manufacturers of titanium implants to passivate their products with a nitric acid bath prior to sale. Theoretically, breakdown of this oxide layer should not occur under physiological conditions. Many of the titanium alloys, in which titanium is present in concentrations of 85–95%, maintain the passivity of pure titanium. When an implant is introduced into the body, complex reactions begin to take place at the oxide/bioenvironment interface. The oxide film grows as ions diffuse outward from the metal and inward from the environment. The oxide that forms in the body may, therefore, be somewhat different than that which forms in air. The rate of formation and composition of this film is important. Although there is no universally accepted definition of the term “passivity,” for our purposes, if an implant metal is oxidized and the oxide does not break down under physiological conditions, the metal is said to be passive or passivated. Few metals display such a high degree of passivity under physiologic conditions as does titanium. Titanium, both as a pure metal and as an alloy, is easily passivated, forming a stable TiO2 (titania) surface oxide that makes the metal corrosion resistant. This oxide will repair itself instantaneously on damage such as might occur during insertion of an implant (Guo et al., 2021a). In the passive state, the rate of dissolution of TiO2 is extremely low. With time, little change can be seen on the surface of the metal implant but an accumulation of titanium in tissue can be observed. The normal level of titanium in human tissue is 50 ppm (Williams, 1981a). Values of 100–300 ppm are frequently observed in soft tissues surrounding titanium implants. At these levels, tissue discoloration with titanium can be seen. This rate of dissolution is one of the lowest of all passivated implant metals and seems to be well tolerated by the body. The clinical significance of this data is substantiated by more than 20 years of clinical experience with cpTi and Ti-6A1-4V alloys (Mombelli et al., 2018). The surface properties of implants are increasingly emphasized as important to the biological response that the materials will elicit from the body (Guo et  al., 2021b). Thus, the surface oxide which forms on the titanium alloys is of paramount importance to its favourable biological properties (Mombelli et al., 2018). In air, the oxide begins to form in nanoseconds and reaches 20–100 Å thickness in 1 second. The thickness of the oxide depends upon factors such as the type of machining which created the metallic surface, roughness of the surface, coolants used during the machining, and treatments to passivate or sterilize the surface (Donley & Gillette, 1991). The oxide layer can be mechanically disrupted, and such damages can result in the release of titanium particles. Mechanical wear of implant surfaces can occur at different instances: during implant placement, during the fitting of a dental

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prosthesis, due to mechanical cleaning in the context of prevention and therapy of peri-­implant infections, and as a result of micromovements of parts of the implant and the suprastructure during function (Mombelli et al., 2018). An in vitro microstructural analysis of dental implants subjected to insertion torque and pullout tests suggested that inserting and removing implants reduced the oxide layer (Valente et al., 2014). These findings were also confirmed by another in vitro study (Deppe et al., 2018) reporting that the insertion forces could provoke release of particles from the implant surfaces by stripping them off. Some chemical agents like acidic products or fluorides used in dental prophylaxis agents could decrease the protection of the oxide layer and initiate a corrosion process. Different patterns of corrosion were observed when titanium grade II or IV implants were in contact with saliva containing fluoride ions suggesting that the fluoride ions were incorporated in the oxide layer decreasing its protective properties (Souza et  al., 2015). Chlorhexidine, a chemical commonly used in mouthwashes, has also been implicated to affect the oxide layer. Although a 0.12% concentration of chlorhexidine digluconate did not affect the corrosion resistance (Faverani et  al., 2014), a 0.2% chlorhexidine digluconate might induce pitting (Quaranta et al., 2010). Dental implants are different from other implantable devices regarding the way they interact with environment. Dental implants are permanently exposed to the oral microflora composed of various bacterial species. Bacteria play a prominent role in the initiation of corrosion by mainly lowering the pH and release of lipopolysaccharide (LPS) (Barão et al., 2012). LPS has been reported to negatively affect the resistance to corrosion and increase the surface roughness of titanium (commercially pure and grade IV) (Mathew et al., 2012). Interestingly, it has also been speculated that bacterial biofilm might lubricate the implant surface, thereby lowering the frictional forces and, in turn, decreasing corrosive wear (Souza et al., 2010). 4.5.2 Metal Ion Leakage The pioneers of cpTi use for implants occasionally noticed blackening in the tissue surrounding the implant. This reaction is an indication of titanium leakage from the implant, which has been described by other authors (Emneus, 1967). When titanium alloys are implanted, higher levels of the component elements can be detected in tissues locally and systemically. In a clinical evaluation on patients with peri-implant disease, high contents of particulate and submicron titanium were present in peri-implantitis tissue. The authors concluded that these high titanium contents in peri-implant mucosa can potentially aggravate inflammation, which might reduce the prognosis of treatment interventions (Pettersson et  al., 2019). A systematic review evaluating titanium release from dental implants reported that titanium particles surrounding periimplant tissues are a common finding. Periimplantitis sites presented a higher number of particles compared to healthy implants. The particles were mostly around the implants and inside epithelial cells, connective tissue, macrophages, and bone.

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Various mechanisms were described as causes of titanium release, including friction during implant insertion, corrosion of the implant surface, friction at the implant– abutment interface, implantoplasty, and several methods used for implant surface detoxification (Suárez-López del Amo et al., 2018). Although Ti-6Al-4V alloy has been widespread in use as an implant biomaterial, it has been reported that the alloy could release of aluminium and vanadium ions (Smith et al., 1997). In particular, vanadium exhibits a high cytotoxicity and aluminum may even induce senile dementia. This said, these leachable metal ions might cause various health issues such as allergic, cytotoxic effect and even neurological disorders. The released titanium ions from a dental implant could potentially lead to allergic or hypersensitivity reactions. Hypersensitivity reactions to metals may arise in predisposed patients chronically exposed to metallic materials, including dental implants made of titanium alloys. Although the evidence is weak, and titanium allergy is rare, hypersensitivity reactions should not be underestimated (Poli et al., 2021). This hypersensitivity in susceptible patients could lead to implant failure, and the need for long-term clinical and radiographic follow-up of all implant patients who are sensitive to metals has been emphasized in the literature (Siddiqi et al., 2011).

5 Conclusions and Future Directions The advent of dental implants has revolutionised the field of oral rehabilitation. From the variety of materials that have been used for manufacturing dental implants over the last 50 years, none has been as successful as titanium and its alloys. The unique physical and biochemical properties of titanium such as the presence of surface oxide layer which is responsible for its inherent inertness and biocompatibility as well as appropriate physical and mechanical properties like strength and elastic modulus make it an ideal material choice for dental implants. The fact that titanium can integrate with bone and its surface can be tailored to enhance the process of osseointegration, make its use possible for supporting prosthetic restorations in a challenging oral environment. The favourable soft tissue response around titanium and its alloys helps in maintaining a healthy peri-implant mucosa, thereby promoting the longevity of implant restorations. Over the last decade zirconia has been developed and marketed as an alternative to titanium especially in the anterior region of the mouth where greyish hue of titanium can pose certain aesthetic challenges. Zirconia has the inherent inertness and biocompatibility, as well as successful osseointegration comparable to titanium implants. The high elastic modulus and potential for low temperature degradation can pose challenges with this material in the harsh oral environment. Future research should explore the possibility of using newer materials with elastic modulus comparable to bone to distribute the stresses more evenly around the implant. The use of polymeric materials like PEEK seems promising in this regard.

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More long-term clinical research is needed to better understand the outcomes and complications with zirconia implants. Surface modification of titanium has been the focus of research over the last 2 decades and continues to do so. More research is needed to tailor the titanium surface to integrate at a cellular level with the soft tissue to help prevent the incidence of peri-implantitis and improve the longevity of dental implants. Acknowledgements  The author declares no conflicts of interest in relation to this manuscript.

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Piconi, C., Maccauro, G., Muratori, F., & Brach Del Prever, E. (2003). Alumina and zirconia ceramics in joint replacements. Journal of Applied Biomaterials & Biomechanics: JABB, 1(1), 19–32. Poli, P. P., de Miranda, F. V., Polo, T. O. B., Santiago Júnior, J. F., Lima Neto, T. J., Rios, B. R., Assunção, W.  G., Ervolino, E., Maiorana, C., & Faverani, L.  P. (2021). Titanium allergy caused by dental implants: A systematic literature review and case report. Materials (Basel, Switzerland), 14(18). https://doi.org/10.3390/ma14185239 Quaranta, A., Ronconi, L.  F., Di Carlo, F., Vozza, I., & Quaranta, M. (2010). Electrochemical behaviour of titanium in ammine and stannous fluoride and chlorhexidine 0.2 percent mouthwashes. International Journal of Immunopathology and Pharmacology, 23(1), 335–343. Roehling, S., Schlegel, K. A., Woelfler, H., & Gahlert, M. (2019). Zirconia compared to titanium dental implants in preclinical studies—A systematic review and meta-analysis. Clinical Oral Implants Research, 30(5), 365–395. https://doi.org/10.1111/clr.13425 Sanon, C., Chevalier, J., Douillard, T., Cattani-Lorente, M., Scherrer, S. S., & Gremillard, L. (2015). A new testing protocol for zirconia dental implants. Dental Materials: Official Publication of the Academy of Dental Materials, 31(1), 15–25. https://doi.org/10.1016/j.dental.2014.09.002 Siddiqi, A., Payne, A. G. T., De Silva, R. K., & Duncan, W. J. (2011). Titanium allergy: Could it affect dental implant integration? Clinical Oral Implants Research, 22(7), 673–680. https://doi. org/10.1111/j.1600-­0501.2010.02081.x Sista, S., Nouri, A., Li, Y., Wen, C., Hodgson, P. D., & Pande, G. (2013). Cell biological responses of osteoblasts on anodized nanotubular surface of a titanium-zirconium alloy. Journal of Biomedical Materials Research, 101(12), 3416–3430. https://doi.org/10.1002/jbm.a.34638 Smith, D. C., Lugowski, S., McHugh, A., Deporter, D., Watson, P. A., & Chipman, M. (1997). Systemic metal ion levels in dental implant patients. The International Journal of Oral & Maxillofacial Implants, 12(6), 828–834. Souza, J. C. M., Henriques, M., Oliveira, R., Teughels, W., Celis, J. P., & Rocha, L. A. (2010). Do oral biofilms influence the wear and corrosion behavior of titanium? Biofouling, 26(4), 471–478. https://doi.org/10.1080/08927011003767985 Souza, J. C. M., Barbosa, S. L., Ariza, E. A., Henriques, M., Teughels, W., Ponthiaux, P., Celis, J.-P., & Rocha, L. A. (2015). How do titanium and Ti6Al4V corrode in fluoridated medium as found in the oral cavity? An in vitro study. Materials Science & Engineering. C, Materials for Biological Applications, 47, 384–393. https://doi.org/10.1016/j.msec.2014.11.055 Stanford, C. M., & Keller, J. C. (1991). The concept of osseointegration and bone matrix expression. Critical Reviews in Oral Biology and Medicine: An Official Publication of the American Association of Oral Biologists, 2(1), 83–101. https://doi.org/10.1177/10454411910020010601 Suárez-López del Amo, F., Garaicoa-Pazmiño, C., Fretwurst, T., Castilho, R.  M., & Squarize, C.  H. (2018). Dental implants-associated release of titanium particles: A systematic review. Clinical Oral Implants Research, 29(11), 1085–1100. https://doi.org/10.1111/clr.13372 Szmukler-Moncler, S., Salama, H., Reingewirtz, Y., & Dubruille, J.  H. (1998). Timing of loading and effect of micromotion on bone-dental implant interface: Review of experimental ­literature. Journal of Biomedical Materials Research, 43(2), 192–203. https://doi. org/10.1002/(sici)1097-­4636(199822)43:23.0.co;2-­k The Glossary of Prosthodontic Terms: Ninth Edition. (2017). The Journal of Prosthetic Dentistry, 117(5s), e1–e105. https://doi.org/10.1016/j.prosdent.2016.12.001 Valente, M. L., Lepri, C. P., & dos Reis, A. C. (2014). In vitro microstructural analysis of dental implants subjected to insertion torque and pullout test. Brazilian Dental Journal, 25(4), 343–345. https://doi.org/10.1590/0103-­6440201302402 Wataha, J. C. (1996). Materials for endosseous dental implants. Journal of Oral Rehabilitation, 23(2), 79–90. https://doi.org/10.1111/j.1365-­2842.1996.tb01214.x Weast, R. C., & Astle, M. (1981). Handbook of chemistry and physics. CRC Press. Weiss, C., & Weiss, A. (2001). Principles and practice of implant dentistry (1st ed.). Mosby. Williams, D. (1981a). Titanium and titanium alloys. In D. F. Williams (Ed.), Biocompatibility of clinical implant materials (Vol. 1, 1st ed., pp. 9–44). CRC Press.

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Williams, D. F. (1981b). Implants in dental and maxillofacial surgery. Biomaterials, 2(3), 133–146. https://doi.org/10.1016/0142-­9612(81)90039-­9 Williams, D. F. (2008). On the mechanisms of biocompatibility. Biomaterials, 29(20), 2941–2953. https://doi.org/10.1016/j.biomaterials.2008.04.023 Yamagami, A., Kotera, S., Ehara, Y., & Nishio, Y. (1988). Porous alumina for free-standing implants. Part I: Implant design and in vivo animal studies. The Journal of Prosthetic Dentistry, 59(6), 689–695. https://doi.org/10.1016/0022-­3913(88)90384-­8 Zablotsky, M. H. (1992). Hydroxyapatite coatings in implant dentistry. Implant Dentistry, 1(4), 253–257. https://doi.org/10.1097/00008505-­199200140-­00004

Titanium Dental Implants in Compromised Conditions: Need for Enhanced Bioactivity and Therapy Necla Asli Kocak-Oztug

and Ece Irem Ravali

Abbreviations 5FU Fluorouacil AIDS Acquired immune deficiency syndrome AP Antiplatelet APTT Activated partial thromboplastin time BMPs Bone morphogenetic proteins COL1 Collagen 1 CVD Cardiovascular disease ECT Ecarin clotting time GNAS1 Guanine nucleotide binding protein 1 HAART Highly active antiretroviral treatment HbA1c Glycohemoglobin HIV Human immunodeficiency virus hs-CRP High-sensitivity C-reactive protein IHD Ischemic heart disease INR International normalized ratio IL-6 Interleukin 6 MRONJ Medication-induced osteonecrosis of the jaw NSAIDs Non-steroidal anti-inflammatory drugs PEEK Polyetheretherketone PRP Platelet-rich plasma N. A. Kocak-Oztug (*) Faculty of Dentistry, Department of Periodontology, Istanbul University, Fatih/Istanbul, Turkey School of Dentistry, The University of Queensland, Herston/Brisbane, Australia e-mail: [email protected]; [email protected] E. I. Ravali Faculty of Dentistry, Department of Oral and Maxillofacial Surgery, Istanbul Aydın University, Kucukcekmece/Istanbul, Turkey e-mail: [email protected] © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 K. Gulati (ed.), Surface Modification of Titanium Dental Implants, https://doi.org/10.1007/978-3-031-21565-0_2

23

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N. A. Kocak-Oztug and E. I. Ravali

PT Prothrombin time RA Rheumatoid arthritis RANKL Receptor activator of nuclear factor kappa-Β ligand RUNX2 Runt-related transcription factor 2 SERMs Selective estrogen receptor modulators SLA Sandblasted, large grit, acid-etched SLE Systemic lupus erythematosus SS Sjögren’s Syndrome TCT Thrombin clotting time Ti Titanium TiO2 Titanium oxide VEGF Vascular endothelial growth factor WHO World Health Organization

1 Introduction Brånemark defined the osseointegration as, “the direct functional and structural adhesion between bone and supporting device surface” (Brånemark et al., 1969). Today, this term is being used to explain the working mechanism of dental implants. For a successful osseointegration, the implants must be placed carefully into the prepared area in the jaw, which is structurally and anatomically sufficient. In fact, the most important factors for a sustainable osseointegration are biocompatible materials and a healthy bone (Albrektsson et al., 1981). Early establishment and long-term maintenance of osseointegration after implant surgery is the key for a long-term success in implant dentistry. Osseointegration is prompted through diverse factors, which depend on the host bone and the implant. This mutual interaction between the host and the implant reveals two important factors necessary for implant success (Fig.  1). The first factor is patient associated, including the quality/quantity of the bone in the implant surgery area and the patient’s immune response. The second factor is dental implant characteristics and whether it will establish rapid osseointegration in the region (Velasco-Ortega et al., 2019). While patient factors may require therapeutic intervention, the dental implant allows for ease of modification to allow for successful implant therapy, even in compromised conditions (Do et al., 2020). In the literature, failure of the implants is divided either as early implant failure or late implant failure. The failure prior to the prosthesis loading is defined as early implant failure whereas late implant failure represents the unsuccessful implants up to 2  years after prosthesis loading. This implies that loading protocol is also a vital motive for implant success (DeLuca et al., 2006). Numerous titanium (Ti) alloys have been produced and used to reinforce implants as mentioned in the first chapter. According to literature, Ti alloy dental implants have been successful compared to their pure Ti ancestors with fewer fracture problems (Ngeow et al., 2020). In addition to strengthening these implants to resist fractures, many attempts have been made to improve the interface between biomaterials

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Fig. 1  Factors affecting the long-term success of dental implants. (Modified after Elias et  al. (2012), Ngeow et al. (2020) and Albrektsson et al. (1981))

and bone. In addition to looking for alternative materials for strength and/or aesthetic reasons, recent research focuses on improving the interaction between biomaterials and bone to achieve faster and improved osseointegration. In the past decade the trend for the implant surfaces has shifted to rough surfaces (Al-Zubaidi et al., 2020). Until so far, various types of surface modification have been applied to optimize the roughness and the morphology of the implant surface. Some of these modifications can be listed as electrochemical anodization, calcium phosphate coating, acid etching, and various combinations of these processes (Rupp et al., 2018; Gulati et al., 2021b; Guo et al., 2021b). In the case of successful osseointegration after implant placement surgery, long-­ term success is strongly dependent on the bone remodelling rates (Diz et al., 2013; Ngeow et al., 2020). Bone remodelling is defined as the bone’s physiological reaction to implant loading in the first year of function (Kocak-Oztug et  al., 2022). Physiological changes in ageing and pathological changes can affect bone health during this period (Ngeow et  al., 2020). Several bone conditions can impair the healing in alveolar bone such as Paget’s disease, osteoporosis, osteogenesis imperfecta, etc. Plus, antiresorptive therapy protocols or corticosteroids used to treat these conditions also affect the bone quality. Numerous autoimmune diseases, diabetes,

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N. A. Kocak-Oztug and E. I. Ravali

smoking, chronic kidney diseases, radiation therapy, some cardiovascular conditions, oral hygiene, and presence of preoperative periodontal disease are the other risk factors affecting the outcome of implant surgery (Diz et  al., 2013; Anner et al., 2010). Except for bone changes caused by antiresorptive drugs and radiation treatment, the effect of other systemic conditions on bone are still being investigated in the compromised patient groups. Therefore, recent knowledge about implications of osseointegration for these risk factors is more hypothetical than evidence based. The current literature lacks sufficient clinical evidence for comparing the long-term survival of dental implants in healthy patients and compromised patients (Bornstein et al., 2009; Dutta et al., 2020; Duttenhoefer et al., 2019). Only a few conditions that may cause pre- or post-operative medical side effects and implant failure have been reported as absolute contraindications for implant application. Until now, recent open-heart surgery or aortic surgery, severe bleeding disorders, some psychiatric illnesses, drug abuse, active treatment phase for cancer and intravenous bisphosphonates therapy have been listed as absolute contraindications for dental implant placement (Diz et al., 2013; Hwang & Wang, 2006). Again, lack of clinical evidence exists to understand how these contraindications may influence implant success rates. However, for patients combating serious medical conditions, performing implant surgery can be considered as an important problem in terms of medical ethics. On the other hand, as improvements are made in the treatment of patients’ general health, focusing on oral health can also benefit the general health of individuals in the long term (Nickenig et al., 2016). In addition to the view that implant applications will increase the quality of life of patients in the long term, it is also an accepted fact that selective treatment should be applied in patients whose general health has been severely affected. For instance, dental implant placement can always be rescheduled until the patient’s systemic health is in a more stable condition. Additionally, in such conditions, personalized surgical procedures should be applied with modified implants with active surface to augment healing and reduce failure rates (Vissink et al., 2018; Al-Zubaidi et al., 2020).

2 Ageing With age and requirements, the oral cavity’s characteristics and shape will change. Likewise dental problems and tooth loss will become more common. Partial or total edentulism caused by periodontal disease or tooth decay can be treated with traditional fixed or removable dentures. However, implant-supported dentures can significantly improve the comfort of patients at this stage of their life by preventing further bone loss, providing stability while speaking and eating (Do et al., 2020; Reissmann et al., 2017; Chan et al., 2021). Biological ageing alters the immune response, inflammation, regeneration, and the stages of the wound healing. Further, ageing slows down the immune response

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27

and prolongs the inflammation period by boosting the release of inflammatory biomarkers. Additionally, ageing affects the regeneration process by reducing the number of mesenchymal stem cells and altering the angiogenesis (Gündoğar et  al., 2021). The reduced quantity of mesenchymal stem cells also negatively influences the bone tissue wound healing (Maxson et al., 2012). Attributed to changes in cell activity, reduced collagen production, decreased matrix metalloproteinase levels and increased apoptosis, ageing can lead to an imbalance in bone healing. From a dental implant perspective, ageing might have an adverse effect on osseointegration after implant surgery (Bartold et al., 2016). Related studies in the literature represents the cumulative survival rate of dental implants to be around 94% (Hoeksema et  al., 2016). However, considering the increase in the bone/soft tissue pathology around the implants and the changes in the marginal bone level by age, it is appropriate to state that more clinical trials are needed to show how ageing effects the implant survival (Srinivasan et al., 2017). Oral hygiene practices may also be adversely affected by the slowdown of muscle activities in old age and the increase in the incidence of diseases such as dementia and Parkinson’s disease (Chan et al., 2021). These reasons are responsible for the possibility of increased incidence of implant loss in old age patients (Bartold et al., 2016). Detailed clinical and radiological examinations are necessary to track complications, minimize loss of implants, and identify risk factors, particularly in elderly patients. In addition, these measures can improve the early acceptance and long term success of dental implants (Gündoğar et al., 2021).

3 Periodontal Disease It is known that the human oral cavity is home to over 600 different types of bacteria (Dewhirst et al., 2010). At least 400 types of bacteria are located in the subgingival area, making the periodontal pocket a reservoir of periodontal pathogens (Paster et al., 2001; Taba et al., 2005). Several studies have reported that periodontal pathogens spread from the remaining dentition to the implant surface (Lasserre et  al., 2018; Dabdoub et al., 2013; Casado et al., 2011). Therefore, patients who have a history of periodontal disease with many periodontal pathogens have an increased potential to contaminate peri-implant area (Casado et  al., 2011, 2013; Zhang et al., 2015). Various studies have proven that the periodontal pathogens including Porphyromonas gingivalis, Prevotella intermedia, Aggregatibacter actinomycetemcomitans, Treponema denticola and Tannerella forsythia can be found in the peri-­ implant sulcus (Casado et al., 2011; Cortelli et al., 2013). Around 1 month after the second stage surgery, these bacteria can be detected in the peri-implant area (Aoki et al., 2012). Furthermore, it was stated that patients would have the same kind of periodontal bacteria in their peri-implant sulcus as in their remaining periodontal

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pockets (Casado et al., 2011; Zhang et al., 2015). Not only the presence of periodontal pathogens, but also the local immune response driven by the reaction between bacteria and the host can provide susceptibility to various inflammations and periimplant diseases (Lasserre et  al., 2018). During periodontal disease, the local inflammatory response to the bacterial pathogens initiates the host’s immune response. This response will activate a high volume of biomarkers and spread of the inflammation through the gingival tissues (Kim & Amar, 2006). In this stage, the gingival inflammation is reversible, but if this inflammation expands to adjacent alveolar bone, resorption may occur (Casado et al., 2013). In addition, several risk factors such as genetic factors, can alter the host’s response. Recent studies have shown that chronic periodontal disease and peri-implant disease have a genetic background. Patients who lose their teeth due to periodontal disease are more likely to develop peri- implant bone loss (Fig. 2) (Zhang & Finkelstein, 2019; Dirschnabel et al., 2011). A long-term study showed a significant increase in bone loss around implants, in patients with a history of periodontitis (Levin et al., 2011). In a review, Schou et al. analysed studies up to 10-year follow-up (Schou et  al., 2006). According to this review, the number of patients suffering from peri-implantitis increased significantly and the number of implants with a bone loss around implants increased for the patients with a periodontitis history (Schou et al., 2006). Wang et al. stated that with a precise long-term periodontal follow-up, patients that lost their teeth due to

Fig. 2  Clinical representation of peri-implantitis. (Courtesy of Prof. Dr. Aslan Yasar Gokbuget from Istanbul University/Turkey)

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chronic periodontitis can present relatively high implant survival rate (Wang et al., 2021). However, bone augmentation with implant surgery, and positioned implants in the anterior region showed lower implant survivals in patients with chronic periodontitis (Wang et al., 2021). Ong et al. explored bone loss outcomes around the implants in patients undergoing periodontal disease treatment (Ong et al., 2008). This review stated that, in comparison to healthy patients, higher amounts of periimplant complications and implant loss can be seen in periodontal disease patients (Ong et al., 2008). Briefly, short-term survival for dental implants seems acceptable for patients with a history of periodontitis. On the other hand, there is not enough long-term clinical trials especially for individuals with a history of aggressive periodontitis (Theodoridis et al., 2017; Wang et al., 2021). Due to common risk factors such as diabetes, smoking and different treatment options for periodontitis optimising a randomised clinical trial for those patients are quite challenging (Liddelow & Klineberg, 2011). Another problem to identify periodontal disease as a risk factor is the application of the new classification for periodontal and peri-implant diseases to clinics. Most of the studies still classify periodontal disease based on the 1999 periodontal disease classification system. The new classification has been recently started to be in use to diagnose and classify the periodontal diseases. For example, a recent study by Adler et al. using the new periodontal disease classification reported in the 2017 World Workshop, showed implant loss related with treated severe periodontitis (Stage III–IV) (Adler et al., 2020).

4 Smoking Smoking is still a common habit that affects bone loss and long-term success of dental implants (Naseri et al., 2020). Several studies stated that smoking has a dose-­ dependent negative impact on osseointegration. Smoking disrupts the osseointegration by inhibiting the growth of progenitor cells that are important for bone healing, thereby slowing the healing process of healthy bone tissue (D’Haese & De Bruyn, 2013; Naseri et al., 2020; Bazli et al., 2020). According to clinical studies, smokers are more common than non-smokers with loss of attachment, gingival recession, and severe periodontal disease, which indicates poor periodontal health for these individuals (Windael et al., 2020; Ong et al., 2008; Meyle et al., 2019). Kasat and Ladda reported that heavy smoking will accelerate the loss of marginal bone and subsequent formation of pockets which will jeopardize the survival rate of implants (Kasat & Ladda, 2012). According to a systematic review by Strietzel et al., smoking is a significant risk factor for dental implant surgery (Strietzel et  al., 2007). Authors also stated that augmentation procedures for implant treatments in smoking patients contain increased risks for complications (Strietzel et al., 2007). Tobacco products consist numerous harmful ingredients, however, nicotine remains to be the most damaging component. Nicotine is the main chemical component leading to tobacco addiction. Moreover, it is stated to be related to negative

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N. A. Kocak-Oztug and E. I. Ravali

prognosis of most diseases (Bazli et al., 2020). Tobacco products also reported to contain benzene and aldehydes that can impair bone healing (Bezerra Ferreira et al., 2016; Bazli et al., 2020). In studies observing the very early stages of bone healing around dental implants, bone healing was found to be severely impaired in smokers, as compared to non-smokers (Levin & Schwartz-Arad, 2005; Bezerra Ferreira et al., 2016; Anner et al., 2010). To sum up, smoking has been shown to impair both osseo- and soft-tissue integration and healing, and therefore may cause early implant loss (Windael et  al., 2020). The frequency of smoking also plays an important role in the failure of the implant. In addition, smoking is associated with increased bone loss around implants and decreased bone density in the jaw, which is related to late implant failure. Therefore, even after successful osseointegration, smoking may shorten the lifetime of the implant, and smokers should be aware of the negative effects of smoking on implant survival before implant surgery (Levin & Schwartz-Arad, 2005; DeLuca et al., 2006).

5 Diabetes The WHO states that prevalence of diabetes have increased over the past few decades (Lin et al., 2020). Diabetes has two types and both types impair the general well-being of human physiology. First version is Type 1 diabetes named as juvenile or insulin-dependent diabetes which is a consequence of the autoimmune destruction of pancreatic beta cells in the earlier periods of life. In all cases of diabetes only around 10% is Type 1. Second version is Type 2 diabetes named as non-insulindependent or adult diabetes which is more common (King, 2008; Ngeow et  al., 2020). Studies on humans and animals have shown that diabetes changes bone properties, decreases bone mineral density, and can result in fracture healing disorders (Romero-Díaz et al., 2021; Yaturu et al., 2007). The leading cause of impaired healing of bone in diabetes is related to its direct effect on osteoblasts. Diabetes slows the formation of new bone by inhibiting the development of bone cells. This leads to a decrease in bone tissue density and for Type 1 diabetes patients it means a risk of osteoporosis in older ages (Romero-Díaz et al., 2021; Yaturu et al., 2007; Diz et al., 2013). Diabetes is also a burden for cardiovascular system. Due to microvascular problems and circulatory disturbances, diabetes also has a negative impact on the healing of soft-tissue, and hence periodontal problems are more common in these patients (Ngeow et al., 2020). Some authors stated that in diabetic patients with well-controlled metabolism, the success rate of routine dental implants is similar to corresponding healthy control group (Sghaireen et al., 2020). Also in recent studies, implant failure was associated with inadequate or non-existent glycaemic control (Dubey et  al., 2013). Attributed to the known impact of hyperglycaemia on recovery before and after the dental implant therapy; strict blood glucose, HbA1c controls and medical consultation are recommended (Table 1). Since diabetes, smoking and periodontal disease

Contraindication No, beware of other systemic conditions

No, not enough data for aggressive periodontitis

No, beware of other systemic risk factors

No, it depends on how controlled the disease is

Acute MI, recent cardiovascular surgery

Consultation required

No

Systemic conditions Ageing

Periodontal disease

Smoking

Diabetes

Cardiovascular disease

Bleeding disorders/congenital and acquired conditions

Bleeding disorders/drug-related Similar

Similar, high risk of peri-implant health issues Similar

Reduced

Reduced

Reduced

Survival rate Similar

Strict blood glucose, HbA1c controls and medical consultation are recommended Prophylaxis, consultation, oral health maintenance avoid major surgeries Coagulation tests, factor transfusion, consultation, avoid major surgeries Consultation, coagulation tests, avoid major surgeries

Early detection of negative changes in peri-implant tissues with controls

Precautions/recommendations Consultation, oral health maintenance, avoid major surgeries Oral health maintenance

Table 1  Dental implant treatment options and precautions in medically compromised patients

(continued)

Frequent control, avoid early contacts and mucosal pressure, avoid immediate implantation

Frequent control, avoid early contacts and mucosal pressure

Aseptic surgeries minimize the risk of peri-implant infections

A strict oral care regime, prefer implants with antimicrobial surface modifications A strict recall regime, prefer implants with osteogenic and antimicrobial surface modifications Frequent control, avoid immediate implantation

Dental implant modifications Prefer implants with osteogenic surface modifications

Titanium Dental Implants in Compromised Conditions: Need for Enhanced Bioactivity… 31

No (beware of cytotoxic drugs causing bone marrow depression and bisphosphonates/ antiresorptive drugs) No (beware of bisphosphonates and antiresorptive drugs)

Head and neck cancer/ chemotherapy

Relative (high accumulation of drugs in oncology patients)

No (beware of bisphosphonates and antiresorptive drugs)

Bone diseases/bisphosphonates, antiresorptives

Bone diseases/Paget’s disease

Bone diseases/osteoporosis

Contraindication Yes (during RT, 9 months post-RT, if the cumulative dose of RT is >50 Gy)

Systemic conditions Head and neck cancer/ radiotherapy

Table 1 (continued) Precautions/recommendations Implant placement before or during ablative surgery or 9 months after RT, consultation, prophylaxis. If RT dose >40 Gy, hyperbaric oxygen therapy recommended, avoid major surgeries Consultation, blood tests, oral health maintenance, prophylaxis

Dental implant modifications Late loading, prefer rough-­ surfaced implants, SLA surfaced implants avoid mucosal pressure, prefer implants with osteogenic and antimicrobial surface modifications, PRP application may be beneficial Late loading, prefer implants with osteogenic and antimicrobial surface modifications

Prefer rough-surfaced implants, Similar survival, high Minimally invasive, avoid marginal bone loss excessive force, implant placement late loading, undersized drilling, with osseodensification technique avoid immediate implantation, prefer implants with osteogenic surface modifications Reduced Consultation especially if drug use Frequent control, avoid early contacts and mucosal pressure, >2 years or in the presence of prefer implants with osteogenic predisposing factors, avoid surface modifications excessive force, use antiseptic agents Prefer high surface energy Reduced If jaws are not affected by the implants, late loading, disease, implant placement with undersized drilling, avoid osseodensification technique immediate implantation, prefer implants with osteogenic surface modifications

Reduced

Survival rate Reduced in the irradiated bone, first 9 months after RT, implantation in the maxilla or augmented bone

32 N. A. Kocak-Oztug and E. I. Ravali

Yes (successful implant cases limited) Yes (in severe cases)

No (if the disease is under control and bone regenerative potential is adequate)

Yes (if the general health is at risk or total white cell count below 1500–3000 cells/mm3)

Yes (in severe cases)

Yes

Bone diseases/Cementoosseous dysplasia Bone diseases/osteogenesis imperfecta

Immunocompromised patients/ SLE, scleroderma, Sjögren’s syndrome, rheumatoid arthritis, Crohn’s disease, mucosal diseases

Immunosuppressive therapy/ organ transplantation

HIV and AIDS

Ti allergy

Consultation, oral health maintenance, prophylaxis, use antiseptic agents, minimally invasive surgery

If necessary: apply unaffected areas and maxilla Minimally invasive, avoid using force

Precautions/recommendations If jaws are not affected by the disease, in the remission phase

Reduced

Similar

Consultation, oral health maintenance, prophylaxis, use antiseptic agents, check CD4+ T lymphocyte level Patch testing in patients with a predisposition to allergies or allergies to other metals

Reduced, high risk of Consultation, oral health peri-implant bone maintenance, prophylaxis, use loss antiseptic agents

Higher risk of peri-implant health issues, increased marginal bone resorption

Reduced

Reduced

Survival rate Similar

Alternative implant materials such as zirconia should be preferred

Prefer late loading, avoid immediate implantation, prefer implants with osteogenic surface modifications, anodic oxidation and, SLA active surface Frequent control, guided, flapless, septic surgeries minimize the risk of peri-­ implant infections, prefer implants with antimicrobial surface modifications Guided, flapless, aseptic surgeries minimize the risk of peri-implant infections, prefer rough surface implants, prefer implants with antimicrobial surface modifications Prefer implants with antimicrobial surface modifications

Dental implant modifications Prefer long implants, late loading, avoid immediate implantation Avoid excessive surgery

Modified after Gündoğar et al. (2021), Levin et al. (2011), Strietzel et al. (2007), Naujokat et al. (2016), Hwang and Wang (2006), Diz et al. (2013), Koudougou et al. (2020), Shugaa-Addin et al. (2016), Schiegnitz et al. (2021), Granström et al. (1999), Buddula et al. (2011), Heberer et al. (2011), Ocaña et al. (2017), de Medeiros et al. (2018), Wagner et al. (2017), Ngeow et al. (2020), Guo and Yuan (2020) and Vissink et al. (2018)

Contraindication No (in the remission phase)

Systemic conditions Bone diseases/fibrous dysplasia

Titanium Dental Implants in Compromised Conditions: Need for Enhanced Bioactivity… 33

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are common risk factors, it is also recommended that individuals should be informed about these conditions before the implant application (Strietzel et al., 2007). Patients with uncontrolled diabetes are generally considered unsuitable for implant therapy. According to reports, blood glucose levels affect the stability of early implants and therefore an increase in the implant failure is expected (Naujokat et  al., 2016). In a recent study evaluating type 2 diabetes patients operated with implants located in the anterior maxillary ridge, the resorption of the alveolar ridges was proportional with the glycaemic blood levels (Gómez-Moreno et  al., 2015). Compared with other patients, diabetic patients with increased HbA1c levels showed more marginal bone loss (Gómez-Moreno et al., 2015). Elevated amount of pro-inflammatory cytokines in the tissue is reported to be another relevant factor that cause tissue inflammation in diabetic patients. This leads to an increase in osteoclasts and a decrease in bone mass (King, 2008). Until now, many studies have focused on judging the success of osseointegration in various levels of diabetic diseases with different blood sugar levels. Currently there are limited investigations to determine whether the osseointegration can be improved with different macro designs, surface coatings and surface modifications in this compromised population (Dubey et al., 2013; Diz et al., 2013).

6 Cardiovascular Disease The cardiovascular system is the organ system that transports nutrients and oxygen through the blood vessels to the tissues and removes carbon dioxide and metabolic waste. Cardiovascular disease (CVD) is a general term for diseases that affect organs such as the heart and blood vessels. Risk factors for cardiovascular disease include obesity, diabetes, high blood pressure, stress, smoking, alcohol and drug use, inadequate physical activity, and unhealthy diet. At least 70% of patients in the high-risk group of CVD have more than one risk factor (Khan et al., 2020). Ischemic heart disease (IHD) manifests as myocardial infarction or ischemic cardiomyopathy. IHD in its non-fatal forms can leave permanent damage to the physical health and reduces the quality of life of patients. The patient’s suitability for the surgical procedure can be assessed by the stability of the existing disease and evidence of recent CVD. However, as mentioned before, CVD such as recent acute myocardial infarction, stroke and cardiovascular surgery have been reported as absolute contraindications for dental implants (Hwang & Wang, 2006). CVD impairs blood flow and tissue nutrition. Hypercholesterolemia, hyperglycaemia, and hypertension, which are risk factors for cardiovascular disease, increase blood viscosity and prolong blood flow time. Since inadequate blood supply causes chondrogenic differentiation of mesenchymal cells and adequate blood supply is one of the most important factors for the success of osseointegration, CVD might affect the further success of the implant by slowing the blood supply to the tissue (Staedt et al., 2020; Alsaadi et al., 2007).

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Apart from the strict contraindications mentioned above, CVD has been reported not to be a contraindication for dental implants. Preoperative antibiotic prophylaxis is necessary because CVD increases the risk of infective endocarditis. In addition, considering factors such as acute cardiovascular conditions and bleeding that may occur due to stress, medical consultation should be requested following a detailed anamnesis (Table 1). Also, in the postoperative period, it is recommended to administer antibiotics and anti-inflammatory drugs approved by the cardiologist to keep the allotted time short for a fast recovery (Farbod et al., 2009). CVD and periodontal disease share common risk factors, including diabetes and smoking. Further, oral pathogens such as Porphyromonas gingivalis, Fusobacterium nucleatum, Tannerella forsythia, and Aggregatibacter actinomycetemcomitans have been detected in atheroma plaques due to periodontal inflammation. In the light of this information, it can be said that periodontal disease and CVD pose a risk for each other (Figuero et al., 2011). In conclusion, implant success is directly related to oral hygiene, regular periodontal treatment, informing the patient about the ­hazards of smoking, and postoperative care in CVD patients.

7 Bleeding Disorders Bleeding disorders are conditions that occur because of hereditary diseases, syndromes, or medications and may manifest as epistaxis, postoperative, or spontaneous bleeding. Decreased platelet production (e.g., aplastic anemia, myelodysplastic syndrome, bone marrow suppression due to chemotherapy and radiation therapy, viral infections such as HIV and rubella, medications) and platelet destruction (e.g., idiopathic thrombocytopenic purpura, hemangiomas, vasculitis, hemolytic uremic syndrome) can lead to uncontrolled bleeding (Dutta et  al., 2020). If the platelet counts in the blood falls below 50,000/mm3, there is a risk of spontaneous bleeding. Postoperative bleeding can lead to fatal outcomes by obstructing the airway through the neck fascia (Dutta et al., 2020; Diz et al., 2013). Use of anticoagulants and thrombolytics, chronic renal failure, liver disease, vitamin K deficiency, Von Willebrand disease and deficiency of coagulation factors, multiple myeloma, and hemophilia can cause coagulation disorders. Antiplatelet (AP) drugs such as aspirin, dipyridamole, and thienopyridines as well as direct thrombin inhibitors and factor Xa, are direct oral anticoagulants that have been widely used in recent years because of their clinical benefits. Anticoagulant drugs such as coumarin, warfarin, and heparin are used to prevent thromboembolic events. Warfarin is a vitamin K antagonist and is commonly used to prevent thromboembolism. It should be remembered that drugs such as metronidazole, erythromycin, and clarithromycin, which we use in daily dental practice, enhance the effect of warfarin. Although there are not enough studies showing the success of dental implants in

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patients taking anticoagulants, it has been reported that minor surgical procedures can generally be performed under proper haemostatic measures (Bajkin et al., 2020; Kalpidis & Setayesh, 2004; Zeevi et al., 2017). Anemia is defined as a fall in hemoglobin level below normal. It may occur with blood loss, heavy menstruation, iron deficiency or as a symptom of a disease. While mild anemia may be asymptomatic, symptoms occur in severe anemia because the oxygen-carrying capacity of the blood decreases. As mentioned in the previous section, inadequate oxygen supply can negatively affect osseointegration and bone healing. Conditions such as factor deficiency, platelet disorders, and the use of anticoagulant medications can lead to uncontrolled bleeding in surgical practice. In such cases, it is recommended to consult with the appropriate physicians after a detailed anamnesis. To minimize the risk of uncontrolled bleeding, it is recommended to perform the necessary blood tests before the procedure, to regulate any medications used, and to use replacement factors in case of factor deficiency. Since standard blood tests are not sufficient to detect bleeding disorders, it has been found useful to check renal functions and markers such as coagulation factors, international normalized ratio (INR), prothrombin time (PT), activated partial thromboplastin time (APTT), thrombin clotting time (TCT), ecarin clotting time (ECT) and factor Xa (Diz et al., 2013; Lupi & Rodriguez, 2020). During surgery, the use of anesthetics containing vasoconstrictors, the use of local haemostatic and antifibrinolytic agents such as tranexamic acid and desmopressin, the use of antiseptic mouthwashes and oral care to minimize the risk of local infections, and the preference for minimally invasive surgical methods while avoiding major surgery are recommended (Table 1). The use of non-steroidal anti-­ inflammatory drugs (NSAIDs) should be avoided unless otherwise recommended. Since oral procedures are usually minor surgeries, they are among the procedures with a low risk of bleeding. However, more than 3–4 dental implants at a time in a patient are considered as a risk factor because high number of implants lead to an increase in the surgical area. Major procedures such as sinus lifts, augmentation, and regional osteotomies should be avoided as much as possible in these patients (Lupi & Rodriguez, 2020). Marković et  al. investigated peri-implant bone healing, implant survival, and success rates on small-diameter implants that can be used in place of augmentation in patients on anticoagulant therapy (Marković et al., 2017). A 100% survival and success rate at the end of a one-year follow-up period and a decrease in implant stability score at 3 months compared to healthy samples were reported, which may be attributed to the effects of oral anticoagulants on bone healing (Marković et al., 2017). In conclusion, although no adverse effects of bleeding disorders on the success of dental implants have been reported, consultation is required because of the risk of prolonged haemorrhage and blood loss in this patient population. In cases where healthy haemostasis cannot be achieved, reducing the risk of major surgical procedures by dividing them into multiple operations is vital (Table 1).

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8 Head and Neck Cancer In cases of malignant tumours of the head and neck, patients often undergo radical surgical procedures that result in the loss of large amounts of bone, followed by therapeutic phases such as radiotherapy and chemotherapy. After such radical surgeries, it may be necessary to reconstruct the tissues with bone grafts and tissue flaps to restore aesthetics and function (Koudougou et al., 2020). Radiotherapy and chemotherapy are used to destroy rapidly proliferating cancer cells but can also interfere with wound healing and tissue blood supply by suppressing the immune system’s response. As radiation therapy causes a decrease in osteocytes along with the malignant cells, osteoclastic and non-osteoclastic resorption of bone occurs (Anesi et  al., 2020). Dental implant treatments are contraindicated because the effects of radiotherapy persist for 6 months after treatment. Minimal trauma and infection can cause osteoradionecrosis, as hypoxic and hypovascular healing is observed in the injured bone. Further, osteoradionecrosis affects the mandible more than the maxilla due to its proximity to the radiation field and less vascularized structure (Hwang & Wang, 2006). It has been reported that the risk of osteoradionecrosis decreases when the total radiation dose is less than 66 grays, and the probability of osteointegration increases when it is less than 50 grays (Diz et al., 2013). It has been also reported that high-­ dose radiation (cumulative dose >50 gray) can lead to osseointegration deficiencies and consequent dental implant failure due to its adverse effects on bone/soft tissue damage and vascularization (Table 1) (Yu et al., 2021). Successful implant treatment should be performed at least 21 days before the start of radiotherapy or at the earliest 9 months after the end of treatment (Diz et al., 2013). According to the current literature review, the success rate of implant-supported prosthesis in the irradiated area was 67.4%, while the survival rate of implants placed 1 year after completion of radiotherapy was reported to be 93.1% (Koudougou et al., 2020). Treatments should be performed under aseptic conditions after prophylactic measures and under appropriate premedication with antibiotics with good bone penetration, such as clindamycin. Any procedures that compromise osseointegration and increase the risk of osteoradionecrosis, such as early implant loading, should be avoided (Table 1). A study by Schiegnitz et al. reported the survival rate of 711 implants in 164 patients with a history of oral cancer up to 10 years, and a significant difference was found between the survival rates of implants placed immediately after surgical treatment (92.5%) and implants placed after completion of oncological treatment (89.5%) (Schiegnitz et  al., 2021). While no effect of irradiated bone alone on implant survival was seen, significantly lower survival rates were reported for implants placed in augmented bone after radiation therapy. During the follow-up period, a total of 70 implants were lost, including 6 implants due to primary loss,

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42 implants due to peri-implant disease, 17 implants due to tumour recurrence, and 3 implants due to osteoradionecrosis (Schiegnitz et al., 2021). In a study conducted with 1405 implants in 466 oncologic and non-oncologic patients, the overall implant survival rate was 96.65%, while this rate was 93.02% in oncologic patients and 97.16% in non-oncologic patients (Silva et  al., 2020). However, the result was limited because the number of implants in oncologic patients was only 172 (Silva et al., 2020). Also, in a study of 435 implants in 93 patients, implant losses in irradiated bone tissue generally occurred in the short-­ term, and long-term implant losses were comparable to those in patients who had not received radiotherapy [Irradiated (0.81%), Non-irradiated (1.29%)] (Nelson et al., 2007). Another study investigated the effects of dental implant manufacturing methods on survival with 271 implants and a 5-year follow-up in 48 patients with a history of oral cancer (who received radiation of at least 50 gray to the neck and head region) (Buddula et  al., 2011). In the maxilla, the survival rate was 72.6% for implants with a turned surface, and 87.5% for implants with a roughened surface. Whereas in the mandible, 91.7% of implants with a turned surface, and 100% with a roughened surface survived. Implant loss occurred more frequently in the 2 years after radiotherapy than in the late phase (Buddula et  al., 2011). Heberer et  al. reported that chemically modified and conventional SLA surfaces had a high success rate in the irradiated bone (Heberer et al., 2011). Although there is no study showing clear effects of chemotherapy on dental implant success, it is suspected that conditions that negatively affect the immune system may also affect implant success. A 2019 systematic review and meta-­analysis about implants in immunocompromised patients examined four retrospective studies and two prospective studies, of which only one of the retrospective study was controlled and found no effect of chemotherapy on implant success rate (Duttenhoefer et al., 2019). In another retrospective study, 106 mandibular dental implants were placed in 30 patients with postoperative oral cavity cancer treated with either adjuvant cisplatin or carboplatin plus fluorouacil (5FU). It was reported that there was no significant effect on implant survival compared to control groups at 10  years (Kovács, 2001). However, some cytotoxic anticancer drugs can cause immunosuppression by inducing bone marrow depression and making tissues susceptible to infection and bleeding. Therefore, elective procedures such as implant treatments are contraindicated in patients taking this group of medication (Hwang & Wang, 2006). In summary, bone metabolism and immune response are damaged by the exposure of radiotherapy, chemotherapy, and various cytotoxic anticancer drugs, and hence it would be advantageous to use implants with osteogenic and antimicrobial surface modifications (Table 1).

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9 Bone Diseases 9.1 Osteoporosis Osteoporosis is a disease that results from destruction in the bone remodelling mechanism. Although the bone retains its shape, spontaneous fractures may occur as it loses its density. Since osteoporosis is a non-localized bone disease, resorption also occurs in the alveolar bone, mainly in the maxilla and edentulous regions. It usually occurs in women over 50 years of age and affects 30% of postmenopausal women and around 200 million people worldwide (Guo & Yuan, 2020). According to studies, the risk of developing osteoporosis decreases in women who receive hormone therapy during the postmenopausal period (Rozenberg et  al., 2020). Although hormone therapy has been attributed to have positive effects on the success of dental implants, the number of studies on this topic are insufficient, and there are even studies stating that it increases the risk of implant loss (Wagner et al., 2017; Rozenberg et al., 2020). In addition to hormone therapy, vitamin D used in the treatment of osteoporosis has been reported to decrease marginal bone loss, while bisphosphonates have been reported to have various adverse effects (Wagner et al., 2017). A comprehensive meta-analysis conducted in 2018 found that osteoporosis had no significant effect on implant survival, but marginal bone loss around the implant was significantly higher in osteoporotic patients than in healthy patients (de Medeiros et al., 2018). It has also been reported that osteoporosis is not a contraindication for the implant treatment, but some precautions should be taken for successful outcomes. In osteoporotic patients, it is recommended to prefer minimally invasive procedures to avoid excessive force applied to the jaw and to avoid augmentation if possible. Early loading of all implants should be reviewed, as osteoporotic bone tissue may be subject to stress fractures due to masticatory forces (Guo & Yuan, 2020).

9.2 Paget’s Disease Paget’s disease, defined by deformity and progressive enlargement of the long bones, replaces bone with more vascular soft bone. This disease, which is usually asymptomatic, can manifest itself with malocclusions in the jaws, avulsed teeth, root resorption, or in some cases, excessive bleeding after an extraction. Removable dentures are not considered a good treatment option for Paget’s patients, as they require constant adjustment due to the continuous expansion of the bones. In cases where bone density is good and the jaws are not compromised according to radiological analysis, dental implant treatments are suitable. Monitoring serum alkaline phosphatase, calcium and phosphorus activity can provide information on the course of the disease (Guo & Yuan, 2020). In studies of implant therapy in Paget’s

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disease patients, it has been suggested to omit the final drill, use implants with high surface energy, apply late loading as in osteoporotic patients, and avoid immediate procedures (Ngeow et al., 2020).

9.3 Cementoosseous Dysplasia Cementoosseous dysplasia is a non-neoplastic, slow-growing bone disease usually affecting the mandible in which the normal trabecular structure of the bone is replaced by dense acellular cementum and osseous tissue. Cementoosseous dysplasia is highly susceptible to infection due to inadequate vascularization. Therefore, even the most minimally invasive procedures such as elective periodontal scaling should be avoided in these patients. In mandatory cases, tooth extractions and implant treatments should be performed under aseptic conditions, especially in the maxilla and in areas not affected by this disease (Esfahanizadeh & Yousefi, 2018; Merlini et al., 2016).

9.4 Fibrous Dysplasia Fibrous dysplasia is a disease characterized by abnormal, non-malignant growths in long bones. It results from the growth of disorganized fibrous bone due to GNAS1 gene mutation with increased proliferation of osteoblastic cells. It accounts for 10% of benign bone tumours. Since primary stabilization cannot be achieved in fibrous bone, it is recommended that dental implants should be used during the remission phase when bone growth stops. In addition, it is recommended to take measures such as increasing the surface area by choosing long implants, avoiding immediate surgery and premature loading (Ngeow et al., 2020).

9.5 Osteogenesis Imperfecta Osteogenesis imperfecta is a hereditary disease in which brittle bones occur due to abnormal type I collagen synthesis. It is classified according to the severity of the disease and can lead to spontaneous fractures of the jaw. In addition, dentinogenesis imperfecta may also occur along with conditions such as osteopenia, growth and hearing retardation, blue sclera, lung, and heart malformations in these patients. Although no significant effect of implant surface characteristics on implant survival has been found in this group of patients, it has been reported that late loading at the end of a prolonged osteointegration period may be beneficial, as in other similar bone diseases (Table 1) (Prabhu et al., 2007, 2018; Wannfors et al., 2009).

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9.6 Medication-Induced Osteonecrosis of the Jaw (MRONJ) Medication-induced osteonecrosis of the jaw (MRONJ) involves exposed and necrotic areas of bone in the jaw caused by certain drugs that prevent rapid bone resorption. Osteonecrosis can remain asymptomatic for a long time before reaching areas of surrounding tissue, such as nerves and soft tissue (Ruggiero et al., 2014). MRONJ can occur spontaneously but is usually due to tooth extractions and trauma. Because of the trabecular structure and vascularity, the mandible is more likely to be affected than the maxilla. Bisphosphonates cause suppression of bone resorption and are used as a treatment option for osteoporosis, hypercalcemia, Paget’s disease, and malignant bone disease. Although the mechanisms of action are not fully known, they are thought to suppress osteoclast precursor cells and promote osteoclast apoptosis (Hwang & Wang, 2006). Many studies show that the use of bisphosphonates causes MRONJ. Since their intravenous forms are more effective, they are prescribed more frequently, but the risk of osteonecrosis increases to the same extent. The half-lives of bone-bound bisphosphonates range from months to years. Although not as common as their intravenous forms, studies have reported MRONJ in dental implants where long-term oral bisphosphonates have been used (Bedogni et  al., 2010; Gelazius et al., 2018; Rawal & Hilal, 2020). Denosumab is a monoclonal antibody that has been widely used in recent years to suppress bone resorption in the treatment of osteoporosis, giant cell tumours, hypercalcemia, and bone metastases. Denosumab does not bind to bone and has a half-life of 1  month. However, according to a 2020 systematic review and metaanalysis, denosumab has a higher risk of developing drug-induced osteonecrosis of the jaw than zoledronic acid (Limones et al., 2020). The recommended dose of bisphosphonates for osteoporosis patients is much lower than for cancer patients. A 2018 systematic meta-analysis reported that low-­ dose bisphosphonate treatments have no significant effect on marginal bone loss around implant and implant survival (Stavropoulos et al., 2018). As studies on this topic are insufficient, it has been reported that minor dentoalveolar surgical procedures such as implants may cause drug-induced osteonecrosis of the jaw, as they are comparable to tooth extractions in terms of trauma (Ruggiero et al., 2014). Although there is insufficient evidence to support MRONJ after dental implant surgery, it would be appropriate to perform elective procedures in high-risk patients (Table 1). In cases where bone metabolism is impaired, coating the implant surfaces with various osteogenic agents has a positive effect on osseointegration. For this purpose, VEGF, BMPs, extracellular matrix proteins, magnesium, various drugs, graphene, hydroxyapatite, and various metals have been used in implant surface coatings (Zhang et al., 2021). Side effects such as MRONJ do not occur with the local application of bisphosphonates, as the effect increases at the target site and the toxicity decreases in the non-target sites. According to studies, coating dental implant surfaces with osteoporosis-inhibiting drugs such as RANKL antibodies, selective

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estrogen receptor modulators (SERMs), bisphosphonates, parathyroid hormone, zoledronic acid, ibandronate, and alendronate might promote osseointegration at the cellular level and suppresses bone destruction (Zhang et al., 2021). Although bioactive molecules are not very popular due to ethical debates about their use and high production costs, many studies have been conducted in this field. Bone morphogenic protein, insulin-like growth factor, platelet-derived growth factor, fibroblast growth factor and vascular endothelial growth factor, collagen I and various genes are the most used biomolecules. Studies report that growth factors promotes angiogenesis, induces osteogenic differentiation of stem cells, and regulates bone regeneration (Zhang et al., 2021; Kocak Oztug et al., 2021). In addition to the above mentioned methods, the use of inorganic elements on dental implant surfaces such as silicon, calcium, magnesium, strontium, and zinc, which are less expensive, promotes osteogenesis when modifying the implant surface (Gulati et  al., 2021a; Gulati, 2022). In osteoporotic animal models, studies report that choosing implants with micro- and nanostructured, hydrophilic and SLA active surfaces and coating the surfaces with materials such as bisphosphonates, hydroxyapatite, strontium, collagen I, fibroblast growth factor, simvastatin, and calcium phosphate increase bone-surface contact, bone density, and pull-out force (Günes et  al., 2016; Lin et  al., 2019; Lotz et  al., 2020; Wermelin et  al., 2007). Another study reported that the use of hydrophilic surfaces plays an important role in activating BMP signalling (Siqueira et al., 2021). In 235 osteoporotic female patients treated with bisphosphonates, an implant survival rate of 98.7% was observed when a plasma-rich growth factor was administered with the implant (Mozzati et al., 2015). Similarly, growth factors have been reported to increase implant survival when calcium ion-loaded implant surfaces are supported with plasma-rich growth factor preparations in osteoporotic patients (Mozzati et al., 2021). It has been also reported that treatment of PRP on TiO2 surfaces promotes early osteogenesis by increasing RUNX2 and COL1 gene expression and suppresses osteoclastogenesis by increasing OPG expression (Jiang et al., 2016). When PRP was applied to the nanomodified TiO2 implant surface, an increase in the bone volume surrounding the implant and implant stability was observed (Jiang et al., 2016). When the micro-roughness of dental implant surfaces is increased by methods such as grit-blasting and acid- etching, bone-implant contact and osseointegration also increase. Mechanical polishing, abrasive blasting, grinding, polishing, laser texturing, micro-arc oxidation, hydrothermal treatment, magnetron sputtering, ultraviolet radiation, and selective laser melting are among the osteogenic surface roughening techniques commonly used (Stich et al., 2022; Li et al., 2021). Studies have reported that surface energy can be increased by nanoscale modification of implant surfaces, and bone cell migration and proliferation increase due to adsorption of matrix proteins (Alghamdi, 2018). Osteosupportive implant surface modifications and coatings, such as those mentioned above and others, can be used to enhance osseointegration. Although there are many in vitro and in vivo studies on osteogenic modifications, clinical trials are very limited due to ethical discussions.

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As more studies are conducted in this area, modified implants could be used more frequently in medically compromised patients (Table 1).

10 Autoimmune Diseases Autoimmune diseases are caused by the inflammation of the immune system by the production of antibodies to its antigens or by the activation of lymphocyte-like cells. Autoimmune response is responsible for more than 80 diseases (Zeher & Szegedi, 2007).

10.1 Rheumatoid Arthritis (RA) Rheumatoid arthritis (RA) is an autoimmune disease characterized by chronic inflammation, oedema, and pain in the joints, leading to joint destruction over time. Anti-inflammatory drugs, corticosteroids and immunosuppressants are used to treat it. It often occurs in conjunction with osteoporosis and other connective tissue diseases. A 2010 study reported that in the presence of RA with concomitant diseases, an increase in marginal bone resorption and bleeding should be expected (Krennmair et al., 2010).

10.2 Systemic Lupus Erythematosus (SLE) Systemic lupus erythematosus (SLE) is an autoimmune disease with a wide spectrum of symptoms affecting almost all organ systems. Although the aetiology is not precisely known, the diagnosis is made based on the common occurrence of symptoms. The oral mucosa is affected in most SLE patients, with discoid lesions and ulceration being the most common lesions. Long-term use of corticosteroids is one of the most used options in the treatment of SLE. In addition, drugs such as cyclophosphamide, mycophenolate mofetil, and azathioprine are also used (Theofilou et al., 2021).

10.3 Scleroderma Scleroderma is a connective tissue disease which causes involvement of the internal organs and the multisystems in its progressive forms, resulting in thickening of the skin caused by fibrosis. Uncontrolled collagen deposition is one of the main features

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Fig. 3  Clinical representation of implant and free gingival graft treatment in a scleroderma patient. (Courtesy of Prof. Dr. Aslan Yasar Gokbuget and Asst. Prof. Dr. Necla Asli Kocak-Oztug from Istanbul University/Turkey)

of this disease. The oral symptoms resemble those of the whole body: tense mucous membranes, shrunken and lost commissures, a hardened tongue (Fig. 3). In cases where the masticatory muscles are involved, resorptions of the mandible may occur. In these cases, pathological fractures may also occur during minor surgery and tooth extraction (Theofilou et al., 2021).

10.4 Sjögren’s Syndrome (SS) Sjögren’s Syndrome (SS) affects periodontal health by decreasing the quality and flow of saliva and increasing the susceptibility of teeth to caries. For this reason, patients often experience tooth loss and the associated need for prosthetic rehabilitation. SLE may co-occur with other autoimmune diseases such as scleroderma. A 2017 systematic analysis reviewed 6 studies with a mean follow up period of 3.97 year. This review found high survival and low complication rates for implants in Sjögren’s patients (Almeida et al., 2017).

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10.5 Crohn’s Disease Crohn’s disease is a chronic inflammatory bowel disease that usually affects the digestive tract, but the oral mucosa can also be affected. Immunosuppressive and anti-inflammatory drugs are often used in their treatment. Studies on implant survival have shown an association between Crohn’s disease and early implant loss (Alsaadi et al., 2007). Autoimmune diseases affecting the oral mucosa also include oral lichen planus, pemphigus vulgaris, bullous pemphigoid, epidermolysis bullosa, and systemic lupus erythema. These diseases are manifested by the formation of painful bullae, vesicles, erosions, and papules in the mouth, as well as xerostomia in Sjögren’s syndrome (de Mendonça Invernici et al., 2014; Schifter et al., 2010). Since these are painful lesions, they interfere with routine activities such as brushing teeth and eating. As a result, poor oral hygiene leads to loss of periodontal health and teeth over time. Hence the use of removable dentures is very difficult in these patients, fixed restorations on teeth or implants are preferred. However, studies report that some medications used in the treatment of these diseases may affect bone quality and thus osseointegration of implants (Mustafa et al., 2015). Autoimmune diseases, except for diseases with joint involvement such as rheumatoid arthritis, generally do not affect bone and bone metabolism. Studies report that implant survival in patients with rheumatoid arthritis who received an abrasive-­ blasted, acid-etched surface implant is 94.6% and decreases to 92.3% when connective tissue disease is concomitant (Krennmair et al., 2010). When planning dental implant treatment for these patients, the complex disease symptoms and the effects of the drugs used for treatment on the immune system, skeletal system, and oral environment should be known. For example, in SLE with renal involvement, bleeding disorders due to dialysis dependency, susceptibility to infections, and decreased renal clearance should be considered (Theofilou et  al., 2021). Although very rare, studies have reported that peri-implant carcinomas may occur in patients with oral lichen planus (Moergel et al., 2014).

11 Organ Transplantation Organ transplantation is currently a common procedure to restore the functions of organs. With the knowledge of immune response occurring after these surgeries, immunosuppressive drugs have been developed to minimize the side effects of long-term chronic immunosuppression and to prevent organ rejection. Glucocorticoids and immunosuppressive drugs such as cyclosporine and nifedipine impair bone healing (Duarte et  al., 2001). Corticosteroids suppress osteoblasts and stimulate differentiation of bone marrow cells into adipocytes (Fu et al., 2012). Although drugs may affect osseointegration by this mechanism, studies on this topic are inadequate (Ouanounou et  al., 2016; Smith et  al., 1992). Animal

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studies show negative effects of cyclosporine use on osseointegration and periimplant bone healing (Sakakura et al., 2007). Immunosuppressive agents have been reported to cause severe and rapid bone loss in animal models (Movsowitz et al., 1988). Despite these findings, many studies have reported successful implant and augmentation procedures under long-term immunosuppressive therapy (Heckmann et al., 2004; Montebugnoli et al., 2015; Paredes et al., 2018). Generally, oral surgical procedures should be completed before transplantation (Duttenhoefer et  al., 2019). Considering that this patient group is immunosuppressed, all surgical procedures are associated with risks. Elective procedures, such as dental implants after transplantation, should be performed at times when general health permits the surgical procedure, in consultation with the appropriate branch physician and with adherence to a prophylactic medication regimen. Oral surgery is contraindicated when the total white blood cell counts falls below 1500–3000 cells/ mm3 because of increased susceptibility to infection (Hwang & Wang, 2006). Patients are susceptible to infections because the host immune response is suppressed by drugs after organ transplantation. In these cases, it may be beneficial to use implants with antimicrobial surface modifications. In cases where drugs are used that suppress bone metabolism, the use of modified osteogenic implants may be considered. Implants embedded in bone tissue without oral mucosal contact with guided flapless aseptic procedures in transplant patients may minimize the risk of peri-implant infections (Guo & Yuan, 2020).

12 HIV and AIDS Human immunodeficiency virus (HIV) is a virus that can be transmitted through unprotected sexual intercourse, blood contact, and maternal transmission and can cause Acquired Immune Deficiency Syndrome (AIDS) in humans. According to the report published by the United Nations and the World Health Organization in 2020, 38 million people are living with HIV, of whom 25.4 million are receiving the necessary treatment (Asfaw & Adamu, 2020). Thanks to antiviral treatment protocols developed in recent years, the quality of life and the life expectancy of HIV+ patients is similar to that of healthy people. Highly active antiretroviral treatment (HAART) is recommended in the presence of AIDS or a CD4+ cell count of 2.0 μm (“rough”)

Applications Abutments “Machined” experimental implants Turned implants, most implants used before 1995 SLActive TiUnite® Most implants of today Plasam-sprayed Hydroxyapatite-­ coated implants

Advantages Reduced bacterial adhesion

Drawbacks Too smooth for proper osseointegration and soft tissue integration Long clinical Inferior osseointegration, documentation of all less forgiving for implants untrained surgeons Stronger bone response, Bacterial plaque adhesion better clinical results than turned implants Enhanced micro retention, increased corrosion resistance

Severe marginal bone resorption due to the delamination of layer

Sandblasting and acid etching are the most clinically utilized techniques by manufacturers of titanium dental implants (Souza et al., 2019). Compared with a smooth implant surface, a rough implant surface could not only enhance bone anchorage but also promote mesenchymal cell differentiation toward osteoblastic phenotype, leading to the augmented osseointegration (Nagasawa et  al., 2016; Khandelwal et al., 2013). In the 1980s, the majority of marketed implants had turned or machined surfaces, with an estimated average roughness (Sa) of 0.5–0.8  μm. Later, a much rougher surface namely titanium plasma sprayed surface (TPS) and surfaces coated with hydroxyapatites (HAp) and other calcium phosphates (CaPs) emerged, with the Sa value >2 μm (Wennerberg et al., 2018). However, these TPS implants coated with HAp soon disappeared from the market, owing to the delamination of the HAp-­ coating, which could cause severe marginal bone resorption, even implant failure. Next, moderately rough surfaces manufactured via blasting, etching, and oxidation techniques were introduced to the market during the 1990s and early 2000s. One of the most successful surfaces in current clinical implant dentistry is the sandblasted, large-grit, and acid-etched (or SLA) surface. The smooth titanium implant surface is converted into a roughened surface with cavities of about 200 μm by sand blasting technology, and then cleaned by acid etching to generate a secondary cavity of 20  μm, resulting in a multi-level rough implant surface that is favorable to bone bonding. Another comparable surface is produced by an anodic oxidation (or anodization) technique, which uses titanium as anode to form a thickened and roughened TiO2 layer. This surface is characterized as isotropic with Sa value between 1 and 1.5 μm (Wennerberg & Albrektsson, 2010). Apart from moderate micro-roughness to increase surface area and oxide thickness, anodized implants could also improve their surface by increasing OH-groups and adhesion points for proteins and cells, which enables augmented osseointegration (Karl & Albrektsson, 2017).

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3 Macro-scale Design of Dental Implant Surfaces Usually, macro-scale alterations provide primary stability and mechanical interlock between the implants and the tissues, avoiding micro movements that can be deleterious to osseointegration (Shalabi et  al., 2006; Hyo-Sook et  al., 2014). Adopting geometric design to increase implant surface area and avoid highly focused stress for bone is critical for osseointegration and survival (Barfeie et al., 2015). Implant body design (shape, length of implant, outer and inner diameters), thread pattern as well as pitch distances are mechanical implant features, which are related to implant macro-design (Dagorne et al., 2015; Abuhussein et al., 2010).

3.1 Implant Body Shape A number of in vitro and in vivo studies have proven that the implant body shape plays a vital role not only for the primary stability of an implant, but also for the long-term success (Kohn, 1992; Steigenga et al., 2003; Kong et al., 2008). In 1906, Greenfield first implanted the cylindrical hollow circular implants made of iridium platinum alloy into the jaws, and followed up the clinical results up to 7 years, which was recognized by the American Philadelphia society of Stomatology (Bell, 1992). This implant body shape is considered to be the predecessor of hollow cylindrical implants. In 1937, Müller used iridium platinum alloy to make a mesh implant that can be placed between the periosteum and the alveolar bone (Bell, 1992). This subperiosteal implant included four protrusions exposed to the oral cavity, which was intended to prevent bone structure damage compared to the intraosseous implant. However, further studies confirmed such implants can be easily infected, leading to severe bone loss. Attributed to these reasons and a complex manufacturing process resulted in their elimination from the implant market. In 1938, in a pioneering attempt, Adams implanted a screw implant with a healing cap for a patient, which is considered to be the pioneer of modern two-stage implant technology (Bell, 1992). In 1947, Formiggini first introduced the threaded implant made with tantalum wire. After implantation, good healing and successful repair were achieved (Bell, 1992). In 1968, in order to increase the contact area between the intraosseous implant and the bone tissue, Linkow designed a sheet structure on both sides of the implant, called leaf implant.’ This kind of implant was widely recognized and used in the 1970s. However, a large number of clinical practices had exposed many disadvantages of this body shape design. Due to the lack of standard technology for preparing implant bone bed (including tools and surgical technology), there was inevitably a large gap between the implant and the alveolar bone after implantation, which mostly were fibrous healing, and not osseointegration. In addition, most of these implants were one-stage implants, which were connected with the oral cavity directly and hence prone to infection, which can result in implant failure. By the late

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Fig. 1  Implants of various shapes and thread types: (a) simple cylindrical, such as IMZ®; (b) conical thread, such as Camlog® screw line; (c) root thread, such as NobelReplace®; (d) cylindrical thread, such as Straumann®

1980s, the design concept of this leaf implant had been gradually abandoned (Linkow & Wagner, 1993). Since Bränemark established the theory of osseointegration, a large number of clinical practices have confirmed that it is difficult for supra-periosteal and periosteal implants to achieve satisfactory osseointegration and long-term favorable outcomes. The cylindrical and root implant design of pure titanium or titanium alloy have achieved appropriate bone bonding and long-term stability, and became the mainstream implant widely used in clinic. At present, there are more than 200 implant systems registered with the FDA in the United States and CE certified in Europe. There are a wide variety of different body shapes, but the basic design is mainly cylindrical and root (Fig. 1).

3.2 Various Geometric Thread Patterns At present, cylindrical and root shape with special thread patterns have become the mainstream design, substituting the simple cylindrical one. The implant’s macro thread structure improves stability and facilitates mounting. At the same time, it increases the attachment area of bone cells, and provides a favorable environment for osseointegration in the later stage. It can also optimize the stress distribution, influence the conduction of bite force and improve the long-term stability of implant. Common thread shapes include standard V-shape, square shape, sawtooth shape, anti-sawtooth shape, circular shape, spiral shape, etc. (Fig. 2). An implant may have

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Fig. 2  Threads of different shapes: (a) standard V-shape; (b) square shape; (c) sawtooth shape; (d) anti-sawtooth shape; (e) spiral shape

only one single thread shape, but also could have different thread shapes at neck, middle section or root tip with different thread depth, width, pitch, thread angle and root plane angle. For example, micro thread with small pitch can be designed in the neck of implant, wide thread or double thread can be designed in the middle, and self-tapping thread is often designed in the 1/3rd of root tip.

3.3 Different Connection Between the Implant and the Abutment The connection between the implant and the abutment has important functions such as connecting the abutment, transmitting dispersed bite force and anti-rotation, which are directly related to the long-term performance of the implant. Therefore, the design of this connection is regarded as one of the major changes in modern implant. Implant connection can be divided into the external and the internal type (Fig. 3). External connection has a certain mechanical structure on the implant shoulder that is used for abutment connection. For example, the classical Brånemark implant system has an outer hexagonal connection structure with a height of 0.7 mm on the implant shoulder, and the base of the repair abutment is fixed on the outer hexagonal structure through a central bolt. By contrast, internal connection has no structure above the implant shoulder, but extends into the implant body through the extension under the repair abutment for connection and fixation. Such connection warrants functional anti-rotation, which helps to resist the clockwise or counterclockwise rotation when the prosthesis faces

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Fig. 3  Structure schematic of implant connection (a) External connection; (b) Internal connection

with lateral force. There are several different internal connection structures such as tube-in-tube connection (e.g., Camlog®), Platform abutment (e.g., NobelSpeedy®), and Taper connection (e.g., Ankylos®). Most of the tube-in-tube and platform connections have anti-rotation structures. The common anti-rotation designs include inner triangle and inner hexagon. Taper connection is considered to have good stress conductivity, but anti-rotation relies on its mechanical embedment. In 2006, a new internal connection called platform switching emerged, where the abutment edge ends at the inner side of the implant top platform rather than flush with the edge. During the last 10–15 years, the internal connection combined with platform switching design has been widely used in contemporary implant systems owing to its good mechanical properties, better abutment connection stability, anti-­ rotation function and stress conduction. Further, it has been gradually considered to influence the formation of biological width, preserve and reduce the absorption of neck bone tissue, and improve the long-term stability of implant neck soft and hard tissue.

3.4 Surface Modifications on the Neck of Dental Implants Dental implants, as an open system connected with oral cavity, are different from implants in other parts of the human body, such as orthopedic implants. The long-­ term clinical effect of dental implants depends not only on its appropriate bone bonding after implantation into the jaw (osseointegration), but also on the sealing effect of healing soft tissue (soft-tissue integration). Contemporary dental implant systems can be roughly divided into one-stage and two-stage implants. Because of their large clinical tolerance and convenience, two-stage implants are the mainstream design of contemporary implants. The bone implant and abutment of two-­ stage implants are two distinct parts that need to be connected with specific structures. This special neck structure has an important impact on the reconstruction

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of soft and hard tissue around the implant, the generation of stress, the conduction of bite force and the long-term stability of implant prosthesis. Within the transmucosal area of implants, epithelial cells proliferate more quickly on micro-machined roughened surfaces (Sa = 2.972 μm ± 0.126 μm), while their initial adherence and activation is boosted on polished surfaces (Sa = 0.012 μm ± 0.002 μm) (Guo et al., 2021; Cao et al., 2018). In the connective tissue layer, bundles of collagen fibers from periosteum and subepithelial connective tissue are found to grow parallel to the long axis of the machined implant without surface treatment (Shioya et al., 2009). However, on roughened surfaces with Sa around 70–100 nm and laser-modified microgroove surfaces, fibroblasts and fibers are observed inserted into the surfaces with an oblique direction, leading to a more robust and stable soft tissue integration (STI) (Zhao et al., 2013; Nothdurft et al., 2015). Embedment of collagen fibers provides the basis for the soft-tissue integration around the implant that prevents the epithelial tissue in the upper part from further growing into the root of the implant. However, compared with epithelial sealing around natural teeth, STI around implants is very fragile due to its fewer hemi-desmosomes and prolonged establishment time (Ivanovski & Lee, 2018). The formation of STI is related to the material properties and surface morphology of the implant, as well as the position of the micro-gap between implant and prosthetic suprastructure (Roehling et  al., 2019). For better aesthetics, white coloured zirconia abutments and implants are gaining attention. According to some studies, soft tissue adherence to zirconia is equivalent to titanium (Hanawa, 2020). There was no significant difference in the soft tissue response between zirconia and titanium abutments (van Brakel et al., 2012). In terms of reaction to bacteria, more remodeling and/or inflammatory phenomena around titanium abutments than those around zirconia abutments (Nascimento et al., 2014). However, titanium tended to show a faster initial osseointegration process compared to zirconia (Roehling et al., 2019). Macroscopically, the implant neck can be designed as cylindrical, dished or reverse dished (Fig. 4). Some studies have shown that the design of dish or reverse dish changes the stress conduction and distribution of the traditional cylindrical neck, and can reduce the bone resorption at the neck to a certain extent (Messias et al., 2019; Rokn et al., 2015). Micro threads can also be designed to disperse neck

Fig. 4  Different implant neck morphology

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stress and increase the area of bone cell attachment (Bateli et  al., 2011; Messias et al., 2019). On the micro level, the implant neck surface can be designed as a machined smooth or rough  surface. However, a consensus on a specific surface roughness value range for enhanced STI is still controversial. Currently, to enhance the longevity of implants, the trans-mucosal regions on commercial implants (either implant neck or abutment) are mainly fabricated with a smooth surface that is easy to decontaminate and inhibits bacterial attachment, which also results in poor STI (Guo et  al., 2021). An ideal implant surface modification strategy would enhance the function of epithelial and fibroblast cells for improved attachment to the implant surface, modulate the inflammatory response to promote more rapid healing, while reducing bacterial attachment and colonization.

4 Micro-scale Design of Dental Implant Surfaces The classic Brånemark system is a machined surface at the beginning of its design, with simple processing techniques and low cost. However, it requires a more extended bone healing period and is less commonly used in the clinic (Buser et al., 2017). The surface modifications of current micro-rough implant systems have a variety of improved technologies or biological modifications, which are accomplished by different manufacturing techniques, including acid-etching, anodization, sandblasting, grit-blasting, or other coating procedures, significantly enhancing the surface area and augmenting osseointegration due to the formation of pits, grooves, and protrusions (Fig. 5). Next, we introduce several standard techniques to achieve Sa of 1–10 μm implant surface.

4.1 Strategies of Micro-scale Surface Modifications 4.1.1 Sandblasted, Large-Grit and Acid-Etched (SLA) This technique involves bombardment of particles (such as silicon, aluminum, titanium dioxide and absorbable bio-ceramic) with 110–500 μm onto the implant surface at high speeds, followed cleaning via acid etching (HCl and H2SO4). Besides, in order to improve hydrophilicity, Ti implants are immersed in isotonic solution at low pH to produce a super-hydrophilic titanium surface. This procedure creates a new hydrophilic and chemically active surface, called SLAactive (Agroya et  al., 2020). Compared with acid-etching surfaces, the super-hydrophilic surface can increase BIC in 2–4 weeks (Lang et al., 2011). The average BIC on SLA surfaces showed to be in the range of 67–81% for 6 months (Bornstein et al., 2009). By utilization of an SLA procedure in isotonic solutions, some spike-like nanofeatures can be produced on the surface. This moderately rough (Sa of 1–2 μm)

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Fig. 5  Micro-morphology of different implant surfaces by scanning electron microscopy (Jarmar et al., 2008): (a) machined surface; (b) the hydroxylapatite coating of Ti plasma sprayed surface; (c) sand blasted and acid etched surface; (d) anodic oxidized surface

surface provides a favorable interface conducive to the generation of bone bonding. Experimental investigations have shown that SLA treated implants creates a greater bone contact and stability at early healing phase (Cochran et al., 2010; Taba Júnior et al., 2003). Moreover, other clinical studies (Nelson et al., 2016; Roccuzzo et al., 2014) of immediate provisional restorations on implant have reported a clinical success rate of about 100% on this super-hydrophilic implant surfaces with positive aesthetic outcomes. At present, many mainstream implants have adopted SLA technology to treat the surface, including Straumann®, Ankylos®, Camlog®, Astra®, Osstem®, etc. 4.1.2 Plasma Spraying Deposition Plasma spraying deposition can be mainly divided into hydroxyapatite sprayed surface (HAS) such as Zimmer® spline reliance implant and Ti plasma sprayed surface (TPS) such as early Straumann® implant. In order to increase the surface areas, TPS, for instance, uses special titanium slurry flame jet coating technology. It adds titanium particles and hydrides into the pressurized inert gas (argon), which quickly (3000 m/s) passes through a high-temperature arc (15,000 ~ 20,000 °C). Next, tiny titanium droplets are sprayed to the implant surface at a distance of 10 ~ 20 cm to form a 30 ~ 50 μm thick layer, with chemical compatibility and biocompatibility remaining unchanged (Ong et al., 2004). The ideal film thickness is around 50 μm

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(Buser et al., 1991), and the average roughness of the coating is around 7 μm, which enhances the implant surface area (Buser et  al., 1991). This rough surface (Sa  >  2.0  μm) showed comparable, even superior bone responses compared to machined surface in some animal studies (Ong et  al., 2004; Ong et  al., 2002). Studies on the function of HA crystallinity in bone formation and bone bonding have been conducted, but no consensus on the ideal characteristics has been achieved thus far (Lo et al., 2000; Mohammadi et al., 2004). However, it is suspected that the coated titanium particles may fall off during implant implantation or after implant stress, which can cause severe marginal bone resorption and even implant failure (Arcos & Vallet-Regi, 2020). Therefore, the implants with sprayed surface have been gradually replaced by the implants with medium roughness (1.0–2.0  μm) after being widely used in clinic for more than 20 years. 4.1.3 Anodic Oxidation The aim of anodic oxidation is to increase the thickness of TiO2 layer in order to improve the surface characteristics of dental implants (Anil et al., 2011). The anodic oxide film is generated by the charging of the double electric layer at the metal-­ electrolyte interface. The process involves dissolving oxide layer supported by the electric field and it is accelerated by temperature, including the production of a soluble salt comprising the metal cation and an anion in the electrolytic bath. This method enables the growth of 10 nm to 40  μm of TiO2 oxide layer and can also allow the adsorption and incorporation of ions from the electrolyte. Oxidation duration, oxidation voltage, electrolyte solution type, electrolyte solution concentration, and the subsequent heat treatment process are the influencing variables (Wang et al., 2020). Through anodic oxidation, controlled topographies can be fabricated on implants, which also offers corrosion resistance and augmented bioactivity. Alternatively, in electrochemical anodization (EA), fluoride and water in electrolytes drive the self-ordering of controlled metal oxide nanostructures when the implant (anode) and counter electrode (cathode) are immersed, and appropriate current/voltage is supplied (Gulati et al., 2015). In comparison to machined surfaces, anodized surfaces result in a substantial strengthening of the bone response, with higher results for biomechanical and histomorphometric testing (Rocci et al., 2013). When compared to turned titanium surfaces of identical forms, anodized titanium implants had a greater clinical success rate (Jungner et al., 2005). According to a recently published meta-analysis, in addition to a moderate microroughness that increases surface area and oxide thickness, anodized implants also provide additional adhesion points for proteins and cells, which contributes to the augmentation of osseointegration (Karl & Albrektsson, 2017). There are two mechanisms to explain the osseointegration: mechanical interlocking and biochemical interaction found between implant material and bone (Sul et al., 2005).

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4.1.4 Laser Surface Processing Laser treatment of implants is emerging as a surface modification strategy. It is also capable to form micro-roughened, as well as uniformly or randomly distributed grooves/holes on Ti surfaces, which provide various advantages over mechanical treatments, including precise crafting, accurate controllability, and generating fewer metal streaks and particles (Chen et al., 2017; Blázquez-Hinarejos et al., 2017). It can not only control its surface roughness, but also treat it according to a predetermined angle (towards the crown or root or perpendicular to the implant surface). Besides, different from aforementioned techniques, laser processing focuses more on improving the integration of dental implants in the surrounding soft tissue. The neck surface of the implant which has been treated in a laser micromachining stage could form a pattern of micro- and nanoscale channels. These microchannels have been postulated to operate as a biologic seal by inducing the adhesion of connective tissues and bone and limiting epithelial downgrowth (Nevins et  al., 2010). Furthermore, it has been identified that laser modified implants upregulate the expression of keratinized proteins from junctional epithelial cells and enhance the formation of collagen fibers, therefore resulting in improved STI in both the epithelium and connective tissue layers (Leong et al., 2018). In a dog model, Nevins et al. (2010) demonstrated histologically that connective tissue formation around laser-­ processing abutments was organized in a perpendicular manner. In a clinical study, peri- implant soft-tissues were retrieved from patients at 15 months after surgery and histological staining showed significantly enhanced gingiva-implant contact area (98.8%  ±  3.78%) on the laser modified (Laser-Lok®) Ti implant system as compared to a smooth surface (24.1% ± 16.63%) (Blázquez-Hinarejos et al., 2017). 4.1.5 Other Modifications Physical Vapor Deposition (PVD).  After thermal oxidation treatment of the implant in pure oxygen at 800 °C by atmospheric heating method, a dense and thick oxide film is formed, which increases the corrosion resistance of the implant and improves the bone bonding ability (Mendonça et al., 2008). Prachar et al. examined the characteristics of TiN with ZrN on pureTi, Ti-6Al-4V, and Ti35Nb6Ta titanium alloys. It was proven that TiN had stronger cell colonization than ZrN (Prachar et al., 2015). Furthermore, their color overcomes the problem of aesthetics in oral implantology since the color of these coatings prevents Ti visibility through the gingiva (Prachar et al., 2015). The advantages of this technique include short processing time and simple equipment; while the shortcomings are the reduced bond and wear resistance between surface and deposition coatings (Xue et al., 2020). Micro-Arc Oxidation (MAO).  MAO also known as micro plasma oxidation or anodic spark deposition. Through this technique, a corrosion and wear resistance ceramic film can be formed directly on the surface of non-ferrous metals, which is porous and conducive to the formation of new bone (Yin et al., 2012). As a hot spot

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technique for surface modification, MAO was employed in plenty of study schemes, including the preparation of titanium dioxide and HA layers (He et  al., 2018; Shimabukuro, 2020). The improved surface hydrophilicity of the porous coating generated by the MAO method may promote the interaction between the implant and the surrounding biological environment. It also brings excellent antibacterial capabilities owing to the presence of metal ions. It offers benefits such as a simple procedure, a compact footprint, high processing capacity, high production ­efficiency, suitability for large-scale industrial production, and environmental protection (Xue et al., 2020). Chemical Modifications.  Chemical modifications mainly promote early bone integration through hydroxyapatite deposition, an important bone biomimetic material (Zweymüller, 2012). For instance, the sol-gel method was applied to the implant surface with the appropriate colloidal calcium and phosphorus ratio, and then heated to form a solid hydroxyapatite film, which significantly improved the osseointegration ability of the implant (Abrishamchian et al., 2013). Further, ion beam assisted deposition technology synthesizes the coating by bombarding the growing surface with a specific energy, type and current ion beam, during electron beam evaporation deposition or sputtering deposition (Coelho & Lemons, 2010; Granato et al., 2010). The hydroxyapatite coating prepared by this method has strong adhesion with titanium matrix and can be combined at low temperature, which overcomes the disadvantage of delamination.

4.2 Biological Response to Micro-rough Implant Surfaces: Cellular Responses, Gene Expression and In Vivo Tests Over the last few decades, the influence of implant surface characteristics on the biological response has been extensively investigated. The reactions include induction of angiogenesis  and osteogenesis by cellular responses (cell adhesion, morphology, proliferation and differentiation) (Bosshardt et al., 2016; Liviu et al., 2015). Compared to macro-rough, the micro-rough surface maximizes interlocking between the mineralized bone and the implant surface, in addition to enhancing mechanical stability (Ralf et al., 2016; Junker et al., 2010). One possible way that topography may impact cellular differentiation is by forced changes in cell shape (Dike et al., 1999). From a microcosmic point of view, the cytoskeleton senses the surface texture by actin protuberance of lamellipodia. Interestingly, the micro-­ roughness surface presents a strong influence on the lamellipodia direction, which imposes cytoskeleton mechano-transduction determining cell shape and differentiation fates (Schönichen & Geyer, 2010; Dominguez & Holmes, 2011). In osteogenesis aspect, one function that microscale surface roughness may play in better osseointegration is the stabilization of fibrin clots by the implant surface (Park et al., 2001). The described physical interlocking of fibrin fibers with surface

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features facilitates the directed ongrowth of bone forming cells directly at the implant/bone contact. Topographic improvement may help in stability of extracellular matrix scaffolds for conduction of cells toward and onto the implant surface (contact guiding) (Ricci et  al., 2008). Several authors have reported surface topography-­ specific impacts on titanium-adherent osteoblastic cell activity (Schneider et al., 2003; Ogawa & Nishimura, 2006). These studies show that the surface adhesion-mediated modulation of cell activity favours bone formation. Investigations have established that the micro-level surface topography improves the adhering osteoblasts’ development and extracellular matrix formation/mineralization (Abron et al., 2001). Micro-roughness induces platelets to secrete biological mediators that attract differentiated osteogenic cells and promote adhesion, together with the formation of the fibrin matrix for stabilization of the blood clot (Feller et al., 2014). Together these experiments have demonstrated that enhanced surface topography significantly promotes extracellular matrix formation of adherent cells and produces a quicker and more reliable osseointegration response. The micro-­ topography alters the growth, metabolism, and migration of these osteogenic cells. The alteration allows for the induction and regulation of the expression of specific osteoblastic integrin subunits that are in contact with the implant. In turn, bone matrix proteins interact with these integrins-mediating osteoblast activity (Vlacic-­ Zischke et  al., 2011; Zhao et  al., 2007). Besides, the micro-level topography enhances the secretion of VEGF-A, TGF-β1, FGF-2, osteoprotegerin, and angiopoietin-­1 by osteoblast-like MG63 cells (Olivares-Navarrete et al., 2013; Saghiri et al., 2016); and increases the production of pro-angiogenic factors such as VEGF-A, fibroblast growth factor (FGF)-2, and epidermal growth factor (EGF) in primary human osteoblasts (HOB) through α2β1 signaling pathway (Raines et al., 2010). In immunological aspect, it is demonstrated that the Ti implant surface topography and roughness created by SLA treatments stimulated the macrophages to secrete proinflammatory cytokine including tumor necrosis factor (TNF)-α, as well as down-regulated the production of chemokines like the monocyte chemoattractant protein (MCP)-1 and macrophage inflammatory protein (MIP)-1α (Refai et  al., 2004). However, when the macrophages were stimulated by lipopolysaccharide (LPS), higher level expressions of these cytokines (TNF-α, IL-1β, IL-6) and chemokines (MCP-1, MIP-1α) were observed (Refai et al., 2004). Additional to cellular responses examined in vitro, the in vivo tests are carried out providing information on tissue level around surface materials (Ernst et  al., 2014). Parameters related to osseointegration phenomena include: bone-to-implant contact, bone mineralization, removal torque, histomorphometry and quantification analysis, all of which can illustrate the osseointegration efficacy of a given implant material (Bagherzadeh et  al., 2013; Ernst et  al., 2014). In animal experiments, a moderately rough surface with a Sa of about 1.5 μm and a Sdr of about 50% leads to favorable bone remodeling, in contrast the most common implant surfaces provides a Sa of 1.1 μm and a Sdr of 37% for an anodized surface (TiUnite™, Nobel Biocare® AB, Gothenburg, Sweden) and a Sa of 1.75 μm and a Sdr of 143% for a hydrophilic, sandblasted, large grit and acid etched surface (SLActive™, Straumann® AG, Basel, Switzerland) (Gottlow et  al., 2012). Zhang et  al.

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demonstrated the osteogenic performance of SLA and 3DA (three-dimensional printing and acid-etching) implants in the femoral condyle of SD rats for 3 and 6 weeks (Zhang et al., 2020). Yet, micron-scale topographic alteration of the cpTi surface is acceptable in dental implant industry (Albrektsson & Wennerberg, 2004a; Albrektsson & Wennerberg, 2004b).

5 Contemporary Implant Surface: Clinical Application and Evidence Nowadays, surface modification methods (for instance, grit-blasting, acid-etching, and anodization) have proved clinical efficacy. There is no doubt that osteogenic cells prefer to recognize and response to the micro-rough Ti surface, compared to the machined one. In a systematic review which evaluated 7711 implants from well-­ documented implant systems, a mean success rate was 89.7% (34.4–100%) over a mean follow up time of 13.4 years (10–20 years). Cumulative mean values for the survival rates were 94.6% and marginal bone resorption values were reported to be 1.3 mm (Moraschini et al., 2015). It was concluded that current dental implants are safe and present a high survival rate with minimal marginal bone resorption in the long term. The 10-year survival rate of SLA Ti implants was reported to be 95–97% (Buser et al., 2012; Roccuzzo et al., 2014; Rossi et al., 2018). As one of the mainstream treatment technologies of implant surface, SLA surface has been tested in clinics for the longest period. A recently published meta-analysis comparing 10-year clinical outcome of different dental implant surfaces (machined, blasted, acid-etched, sandblasted and acid-etched, anodized, Ti-plasma-sprayed, sintered porous and micro-textured) demonstrated that the anodized implants had the lowest failure rate (1.3%, 0.2–2.4%) and minor peri-implantitis rate (1–2%) (Wennerberg et al., 2018). It is well established that osteogenic cells prefer and respond to micro-rough Ti surfaces, as compared to the machined surfaces (Buser et  al., 1991; Klokkevold et al., 2001). However, additional investigations are needed to find the most optimized implant surface topography (SLA or anodized) that enhances bioactivity and osteogenesis (Yeo, 2019). Currently, both SLA and anodized implants present a suitable topography for clinical use. However, SLA surfaces remain the preferred choice in clinical dentistry, with many manufacturers opting for SLA over anodized implants. Despite the favorable clinical results, there are still implant-related mechanical, biological and functional complications (Wennerberg et  al., 2018; Albrektsson et  al., 2016). One major complication is peri-implantitis, which can cause bone loss around the implant, eventually leading to implant failure. Although there are a great range of surface treatment technologies commercially available for generating Ti implant surfaces, effect of one surface treatment on the durability and performance of dental implants over the other has not been fully researched. It’s worth emphasizing that there are currently no clear rules and recognized standards for implant surface morphology design (Wang et al., 2020). Besides, the high cost

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is another barrier that causes many difficulties during the clinical validation stage of implant design. In order to address these problems, the future dental implants should meet the following characteristics: biocompatibility and antimicrobial properties, biomimetic and standardized qualities, biological safety and inexpensive cost. Furthermore, a significant pre-clinical and clinical tests needs to be performed to assure the security and dependability of implants employing innovative technology.

6 Future Directions While micro-roughness is regarded as the ‘gold standard’ towards establishment of appropriate implant-bone bonding, nano-engineering is emerging as a new platform for further enhancement of the dental implant bioactivity. This new trend in titanium surface engineering aims to create biologically inspired surfaces that can imitate natural bone architecture and stimulate osteoblast adhesion, differentiation, proliferation, and migration, resulting in improved bone formation and osseointegration. To reduce the risk of periimplantitis-induced implant failure, antibacterial and anti-­ inflammatory therapeutics can also be physically adsorbed on such nano-scale surfaces in order to limit primary bacterial adherence and biofilm formation. In particularly, synthesis of TiO2 nanotubes using anodization approach on surface of titanium has shown remarkable potential to promote cellular behavior such as adhesion, proliferation and differentiation (Gulati et  al., 2018). In addition, hydroxyapatite mineralization is enhanced and bacterial adherence is lowered on nanotubular surfaces compared with normal smooth surfaces (Mei et  al., 2014). Recent attempts have confirmed that nano-engineered implants promote osseointegration, and holds great promise as the next generation of dental implants (Hamlekhan et al., n.d.; Gulati et al., 2021; Zhang et al., 2021). Acknowledgements  This work was supported by the National Natural Science Foundations of China 81871492 (Yan Liu), Ten-Thousand Talents Program QNBJ2019-2 (Yan Liu), ITI Research Grant 1544-2020 (Yan Liu), Key R & D Plan of Ningxia Hui Autonomous Region 2020BCG01001 (Yan Liu), Innovative Research Team of High-level Local Universities in Shanghai (SHSMU-­ ZLCX20212402, Yan Liu), Key Research Program of Central Health Commission 2022ZD18 (Ye Lin).

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Nano-scale Surface Modification of Dental Implants: Fabrication Ruben del Olmo, Mateusz Czerwiński, Ana Santos-Coquillat, Vikas Dubey, Sanjay J. Dhoble, and Marta Michalska-Domańska

Abbreviations A Anatase AC Alternating current bFGF Fibroblast growth factor CAD Computer-assisted design CaP Calcium phosphate DC sputtering Direct current sputtering DCD Discrete crystalline deposition DDS Drug delivery systems DLIP Direct laser interface pattering EA Electrochemical anodization GO Graphene oxide HA Hydroxyapatite LBL Layer-by-layer LPD Laser pulse deposition MAPLE Matrix-assisted pulsed laser evaporation NR Nanorod R. del Olmo · M. Czerwiński · M. Michalska-Domańska (*) Institute of Optoelectronics, Military University of Technology, Warszawa, Poland e-mail: [email protected]; [email protected] A. Santos-Coquillat Experimental Medicine and Surgery Unit, Inst. De Investigación Sanitaria Gregorio Marañón, Madrid, Spain V. Dubey Department of Physics, Bhilai Institute of Technology Raipur, Raipur, India S. J. Dhoble Department of Physics, R.T.M.Nagpur University, Nagpur, India © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 K. Gulati (ed.), Surface Modification of Titanium Dental Implants, https://doi.org/10.1007/978-3-031-21565-0_4

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PDA Polydopamine PTL Phase transition lysozyme PVD Physical vapor deposition PVDMS Physical vapor deposition magnetron sputtering R Rutile Ra Profile roughness average RF sputtering Radio frequency sputtering RGDC Arginine-glycine-aspartic acid-cysteine Sa Area roughness average SA Supramolecular assembly TA Tannic acid Ti Titanium TNT TiO2-based nanotubes TNWs Titanium nanowires TPS Ti plasma-sprayed

1 Introduction 1.1 Titanium: The Gold Standard in Dentistry Metallic biomaterials/implants are widely used in dentistry and dental surgery. These biomaterials are applied in clinical practice for dental restoration, endodontic implantations, or orthodontics applications. Dental implants from commercially pure titanium  or titanium alloys have an extensive and successful history of clinical application of more than 40 years and can be considered the gold standard in dentistry (Chen & Thouas, 2015; Qu et al., 2007). Ti CP grade 4 (ASTM F67) is the most common Ti-based alloy investigated in dentistry. It presents outstanding corrosion resistance, biocompatibility, and osseointegration. Ti CP has a single-phase alpha microstructure and is available in four grades where the oxygen (O) content varies between 1.8 and 0.40 wt.%, and the iron (Fe) content between 0.20 and 0.50  wt.%. Ti CP grade 4 contains up to 0.40 wt.% of oxygen, that influences the physical and mechanical properties (highest tensile and yield strengths) (Geetha et al., 2009; Lyndon et al., 2014; Mathieu et al., 2014). The main characteristics of Ti are excellent biocompatibility, high strength, stiffness, and relatively low density. More notably, due to surface passivation Ti implants can osseointegrate with bone tissue (Guo et  al., 2012). The passivation process occurs when pure titanium or its alloys are exposed to the air, producing a ~2–7 nm thick TiO2 layer in a few seconds. The TiO2 layer provides biocompatibility, chemical inertness, and high corrosion resistance (Guo et al., 2012; Navarro et al., 2008; Wang et al., 2016).

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As established in previous chapters, to increase implant bioactivity and improve the long-term success (especially in compromised patient conditions such as osteopenia and diabetes), their surface has been modified in the macro, micro, and nanoscales. This chapter focuses on the advanced nano-engineering of Ti-based dental implants.

1.2 Nano-scale Surface Modification of Ti Dental Implants The long-term success of an implant is dependent on the surface characteristics of the implant (i.e., the porosity, roughness and chemical composition). The surface topography of dental implants is crucial for the adhesion and differentiation of osteoblasts during the initial phase of osseointegration and in the long-term bone remodeling. However, the implant surface also dictates bacterial adhesion and biofilm formation. The usual topographic parameters to describe the surface roughness are the 2-dimensional Ra (profile roughness average) and the 3-dimensional Sa (area roughness average). Most dental implants on the market have a Ra of 1–2  μm, because this range provides an optimal degree of roughness to promote osseointegration (Dohan Ehrenfest et al., 2010; Guo et al., 2021a, b). Even though micro-­ roughness is considered the gold standard, new nano-engineered implants are being designed to enhance bioactivity and minimize bacteria adhesion to prevent the loss of implant due to peri-implantitis (Zhang et al., 2021). This nano-scale surface modification can help shield implant structures from bacterial attack and biofilm formation (Navarro et al., 2008; Zhang et al., 2021). Alternate strategies to promote osseointegration are focused on surface charge and wettability modulation. After the implantation surgery, proteins are absorbed onto the implant surface, paving the way for cell-implant interactions. Some cell types, including osteoblasts and fibroblasts, have a preferential union to the absorbed proteins rather than the implanted material. The orientation of the adsorbed molecules after implantation can change as a result of the surface energy of the biomaterial and after cell adhesion (Guo et  al., 2012). Moreover, surface wettability is highly dependent on surface energy, and enhanced wettability improves the implant surface biological interaction (Qu et al., 2007). Titanium implant surfaces present contact angle measurements between 0° (hydrophilic) and 140° (hydrophobic) (Le Guehennec et al., 2007; Zhao et al., 2005). Considering these implant surface characteristics, to enable long-term implant success, current research aims to optimize implant bioactivity to promote early osseointegration, maintain it in the long-term, and at the same time favor the soft-­ tissue integration to prevent bacterial infection (Smeets et al., 2016).

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1.3 Current Nanoscale Surface Modification Methods of Ti Dental Implants Nowadays, there are more than 1300 implant systems with different shapes, dimensions, bulk and surface material, thread design, implant-abutment connection, surface topography, surface chemistry, wettability, and surface modifications (Shimizu et al., 2009). Therefore, numerous surface modifications by different subtractive and additive methods have been applied to improve implant integration (Table 1). Examples of conventional subtractive processes include machined, electropolishing, mechanical polishing, sand-blasting, acid-etching, and electrochemical anodic oxidation. Examples of conventional additive processes are hydroxyapatite (HA) and calcium phosphate (CaP) coatings, Ti plasma-sprayed (TPS) surfaces, and ion deposition (Rupp et al., 2018; Wennerberg & Albrektsson, 2009). For the past few decades, Brånemark Standard implants were considered the gold standard for implant surfaces (Table 1). These implants were machined with a turning process, and the imperfections along these machined surfaces allowed osteogenic cells to adhere and deposit bone, thus creating a bone-implant interface (Abraham, 2014; Wennerberg & Albrektsson, 2009). These features promoted alternative designs to achieve microrough titanium surfaces (e.g., sandblasting and/or acid etching) with bioactive properties. However, each is associated with pros (augmented bioactivity) and cons (as described next). For example, surface roughening methods can lead to increased soft tissue growth at the bone-implant interface, thus reducing the osseointegration between the implant site and the bone Table 1  Commercial surface treatments on dental implants Treatment Mechanical

Methodology/composition Machined Ti

Surface plasma spray

Welding Ti powder in an inert atmosphere (e.g., Ar). Plasma spraying of Ti alloy powder onto the dental implant. Electrophoretic deposition, sol-gel processing, hot isostatic pressing, flame spraying, plasma spraying, and laser pulse deposition. Mixtures of acids (HCl + H2SO4, HF + HNO3) Electrochemical anodic oxidation

Hydroxyapatite (HA)

Double-etched (DE) Anodized

Product/company Brånemark Standard Implants (Nobel Biocare), Restore Machined Implants (Lifecore Dental) IMZ TPS (Densply Friadent), Bonefit (Straumann Institute), Restore TPS (Lifecore Dental), Steri-Oss TPS (Nobel Biocare) IMZ HA (Densply Friadent), Restore HA (Lifecore Dental), Steri-Oss HA (Nobel Biocare) Osseotite (Zimmer Biomet), Steri-Oss Etched (Nobel Biocare) Xeal and TiUltra (Nobel Biocare)

Discrete Crystalline Calcium phosphate (CaP) particles Nanotite/T3 (Zimmer Biomet) Deposition (DCD) are deposited on a double acid-etched surface by the sol-gel process.

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(Santos-­Coquillat et al., 2018). For plasma sprayed HA coating, precise control of the chemical composition, crystallographic structure, and crystallinity of the coating is not achieved, and the resulting HA layer is mechanically and chemically unstable (Qu et  al., 2007). Therefore, HA tends to disintegrate after the coating formation, causing cracks in the implant surface. Surface roughness can determine the fate of the implant as it modulates the osseointegration and can promote bacteria adhesion and proliferation on the implant. Commercial implants have been functionalized by different surface treatments in the past decades (Table 2), obtaining diverse roughness values. Dental implant surfaces are classified into different groups considering their surface roughness values (Ra or Sa). Where smooth implant surfaces present less than 0.5 μm, minimal rough surfaces are between 0.5 and 1 μm, moderately rough are between 1 and 2 μm, and rough surfaces are above 2 μm. We can find different smooth surfaces evaluated in the literature, however with limited or no use in clinical practice (De Bruyn et al., 2017). Taken together, the use of moderately rough or rough surfaces can promote bacteria colonization in the implant area, making nano-scale (Ra 10 years) survival rates of dental implants inserted by adequately trained physicians could be up to 98% (Albrektsson et  al., 2017). Under constant load bearing conditions, the long-term survival and functioning of dental implants is dictated by the integration with the surrounding tissues, including osseointegration (OI) at the implant screw surface and gingival/ T. Guo · S. Ivanovski · K. Gulati (*) School of Dentistry, The University of Queensland, Herston, QLD, Australia e-mail: [email protected]; [email protected]; [email protected] © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 K. Gulati (ed.), Surface Modification of Titanium Dental Implants, https://doi.org/10.1007/978-3-031-21565-0_5

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mucosal integration at the transmucosal level (Berglundh et al., 2018). Pioneered by Branemark et al. in 1969, osseointegration involves the direct contact between living bone and the implant at the microscopic level, which is a basis for the success of a dental implant (Albrektsson & Wennerberg, 2019; Chrcanovic et  al., 2014). Successful OI requires the absence of micro-motion at the bone-implant interface under constant masticatory loading. Further, a robust and functional OI at the implant-bone interface is a key determinant of the long-term functioning of the associated implant prosthodontics (Albrektsson & Wennerberg 2019; Chrcanovic et al., 2014). Timely bone healing, intramembranous osteogenesis, limited foreign body reaction (FBR) and absence of chronic inflammation are essential conditions for obtaining successful OI (Berglundh et al., 2003; Trindade et al., 2018). Starting from the initial angiogenesis induced by the blood clot at the implant-bone interface, micro vascularisations are observed at the margin of the host bone within 24 h (Berglundh et al., 2003). The blood clots are then infiltrated by the mesenchymal cells from the bone marrow, which migrates to the wounded sites with various newly formed blood vessels (Berglundh et al., 2003; Trindade et al., 2018). The migrated mesenchymal cells are modulated by the growth factors (GFs) from the blood to differentiate into osteoblasts, finally establishing osteogenesis around implants (Berglundh et al., 2003). The earliest new bone formation could be observed at around 5–7 days post-implantation, which is regarded as the woven bone structures stretched from the host bone, with calcified tissue and collagen matrix deposited (Berglundh et al., 2003). Similar to fracture healing in an orthopedic implant setting, continuous bone remodelling occurs during the bone formation process upon dental implant placement, transforming the initial woven bone into lamellar bone (Wang et al., 2016). Further, the connection of newly formed bone with the implant surface could be observed at approximately 4 weeks after implantation (initial OI) (Davies, 2003). Finally, at approximately 8–12 weeks, the transformation and maturation of lamellar bone is completed, which is regarded as established OI. Under appropriate conditions and maintenance, modern implants should obtain stable osseointegration with surrounding bones without ongoing bone loss, after the initial bone remodelling within the initial 1 ~ 2 years after implant placement (Albrektsson et al., 2022). Alongside establishing appropriate OI, obtaining soft-tissue integration (STI) at the transmucosal region is also critical for the success of dental implants. Appropriate STI leads to formation of a soft-tissue sealing or barrier that protects the implant structures from bacteria ingress (Guo et al., 2021a, b). Unlike the OI at the implant-­ bone interface, the STI around dental implants is significantly weaker than those around natural teeth (Atsuta et al., 2016). Such results are attributed to its histological characteristics, composed of the superimposed peri-implant epithelial (PIE) sealing and the connective tissue adaption (Atsuta et al., 2016). The PIE initiated from the oral epithelium (OE) horizontally migrates to the surgical site within 3–4 days after implant surgery (Atsuta et al., 2005b). Attachment structures such as hemidesmosomes (HDs) and internal basal lamina (IBL) are generated from the epithelial cells and are recruited at the epithelium-implant interface within 1 week, representing the initial epithelium attachment formation. Such initial

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attachment matures until 2–4 weeks when a dense epithelium barrier is established (Atsuta et al., 2005b). However, compared with the junctional epithelium around natural teeth, significantly fewer adhesive structures (HDs and IBL) are observed at the PIE-implant interface, which significantly limits the adhesion strength of PIE sealing (Atsuta et al., 2005b; Tomasi et al., 2014). Further, such adhesive structures are only present at the lower 1/3rd of the PIE-implant interface, reducing the adhesion region of PIE sealing and compromising its strength (Atsuta et al., 2005a). Fibroblasts are responsible for STI in the underlying connective tissue layer, which induces connective tissue regeneration and wound healing by secreting and remodelling collagen and extracellular matrix (ECM) (Gulati et al., 2020). Unlike the epithelium barrier that establishes within 1 week, the formation and maturation of peri-implant connective tissue is a prolonged process. Typically, the formation of collagen fibres requires 4–6  weeks and another 2–6  weeks for their maturation (Ericsson & Lindhe, 1993; Fujii et al., 1998). Thus, the delayed connective tissue healing and regeneration around dental implants significantly restricts the STI formation. Further, compared to the periodontal ligament that firmly connects a tooth to alveolar bone, the peri-implant collagens run parallel to the implant surface yielding only a physical “adaption” without biological integration (Fujii et al., 1998). Immune cells such as polymorphonuclear leukocytes (PMNLs) and macrophages also influenced the formation of OI and STI around dental implants. Immediately after implant placement, a universal FBR will occur within a few seconds, starting from protein adhesion on implant surfaces to forming a transient surface matrix (Brown & Badylak, 2013). Such FBR will initiate acute inflammation response, with the immediate recruitment of PMNLs that release enzymes and reactive oxygen species (ROS) at the surgical sites (Brown & Badylak, 2013). PNMLs normally undergo apoptosis within 48  h, followed by  the recruitment of macrophages, which ends the acute inflammatory responses and initiates chronic inflammatory responses. Such acute inflammatory responses should be alleviated and reduced within 1 week; and any delay beyond 3 weeks will significantly increase the risk of implant failure (Chen et  al., 2016). Chemoattractants and cytokines released by the PMNLs involve the post-surgical macrophage infiltration at the implant site, which tailors the inflammation and the host responses (Guihard et al., 2012). Based on the activation pathway, macrophages could be categorized as classical (M1) and alternative (M2) activated types (Guihard et  al., 2012). Although M1-activated macrophages were also reported to induce osteogenesis in mesenchymal stem cells (MSCs), the various proinflammatory cytokines may significantly aggravate the inflammatory response and prolonged the chronic inflammatory process after surgery. Driven by interleukin-4 (IL-4) and interleukin-13 (IL-13), macrophages continuously form foreign body giant cells (FBGC), degrading the surrounding tissue and resulting in implant failure (Freytes et  al., 2013). On the other hand, macrophages with M2-activation promote tissue repair by secreting osteogenic cytokines such as bone morphogenetic protein-2 (BMP-2) and vascular endothelial growth factor (VEGF), which regulates excessive inflammatory responses and finally result in a homeostasis around the implant (Champagne et al., 2002; Freytes et al., 2013; Ivanovski et al., 2022).

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The ingress of bacteria and biofilm formation is a critical factor that negatively influences the OI and STI around dental implants. Some specific pathogens, including P. Gingivalis, F. Nucleatum and A. Actinomycetemcomitans are closely related with inflammation in the peri-implant mucosa and the progressive bone loss (peri-­ implantitis) (Shibli et al., 2008). The endotoxin of P. Gingivalis (PgLPS) has been reported to initiate excessive inflammatory responses, which is the keystone mechanism for aggravated tissue inflammation, swelling and attachment loss (Irshad et al., 2013). Further, the PgLPS could also induce ECM degradation around implant sites via upregulating the expression of monocyte chemotactic protein-1 (MCP-1) and matrix metalloproteinase (MMP), thereby damaging peri-implant tissue and causing bone resorption (Irshad et al., 2013). F. Nucleatum could promote the superoxide anion production from fibroblasts, which are favourable for the proliferation of P.  Gingivalis, thus aggravating the peri-implant tissue damage (Metzger et  al., 2009). Moreover, it has been reported that A. Actinomycetemcomitans damages the intracellular connections between fibroblasts/osteoblasts, compromising their functions and the related tissue integration (Gutiérrez-Venegas et al., 2007). Apart from these specific bacteria types, it is accepted that general biofilm is substantially destructive to peri-implant tissues (Berglundh et al., 1992; Mombelli & Décaillet, 2011). It is noteworthy that biofilm protects the embedded pathogens and impairs host immunity, resulting in the progressive formation and maturation of pathogenic biofilms (Mombelli & Décaillet, 2011). Further, the antibiotic resistance gene could be horizontally transferred into other bacteria within the biofilm, improving their resistance against antibiotics and yielding uncontrolled biofilm accumulation (Mombelli & Décaillet, 2011). Hence, maintaining healthy oral conditions, and periodically disrupting biofilm formation especially against the pathogenic bacteria is critical for establishing and maintaining tissue integration around the implant surfaces. The bio-inertness of non-modified Ti implants also influences tissue integration and wound healing, especially with respect to limiting STI at the transmucosal region (Guo et al., 2021b). Typically, 2–5 nm thick TiO2 film readily forms on Ti upon exposure to air/moisture, which is amorphous and provides biocompatibility (Lausmaa, 1996). Such a naturally formed oxide layer is responsible for the bio-­ inertness of non-modified smooth Ti. Additionally, the native oxide layer could be corroded within the human body, exposing the underlying Ti to leach ions, which raises toxicological concerns. Thus to improve the bioactivity of Ti dental implants, studies have been performed to modify their surfaces, aiming at obtaining improved OI and STI on the implant-tissue interface, as detailed in the following sections. In summary, this chapter provides an overview of the formation and characteristics of OI and STI around Ti dental implants. Next, the progress on the various microscale modifications of Ti implants for enhancing their bioactivity is detailed, including topographical, chemical and bioactive coatings. Further, the progress towards novel nano-engineered Ti implants with controlled nanotopographies towards augmenting implant bioactivity is reviewed. This chapter compares and contrasts the surface modification of Ti implants towards bioactivity enhancements, evolving from micro to nano, aimed at ensuring long-term implant success.

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2 Microscale Surface Modification 2.1 Enhancing Osseointegration In 1991, Buser et al. reported that microscale topography and roughness could significantly influence osseointegration on a dental implant surface (Buser et al., 1991). Compared to polished implants, roughened implants demonstrated an increased bone apposition in vivo, with further enhancements possible via hydroxyapatite (HA) coatings (Buser et  al., 1991). Additionally, acid etched and grit blasted Ti enhances osteogenesis (Gotfredsen et  al., 1990). These observations in the early 1990s led to the evolution of Ti implants from minimally rough microgrooved surfaces to moderately rough microscale surfaces (Gulati et  al., 2018a). Compared with the relatively smooth early Ti implants, microscale implants promote proliferation and migration of osteoblasts/osteoprogenitor cells, which accelerate wound healing and tissue integration (Albrektsson & Wennerberg, 2019; Buser et  al., 1991). Another critical characteristic for dental implants is surface hydrophilicity, which significantly affects the protein/plasma adsorption and the adhesion/recruitment of osteoblasts (Devgan & Sidhu, 2019). To date, numerous microscale approaches have been utilized or studied on Ti implants, targeted at modifying different surface characteristics to enhance osseointegration and bone regeneration on their surfaces. 2.1.1 Physical and Chemical Modifications One solution for obtaining microscale rough implants is grit blasting of ceramic particles such as alumina (Al2O3) and titania (TiO2), driven by compressed air to collide Ti surfaces (Aparicio et al., 2003). However, the blasted particles can get embedded on the Ti surface, thus an additional acid-washing is recommended to remove those embedded particles (Aparicio et al., 2003). The embedded Al2O3 is resistant to acid-washing and hence the surface chemistry of Ti is significantly altered. Further, the Al2O3 gritted implants can potentially release Al ions and particles that can be detrimental to the surrounding bone (Ivanoff et al., 2001). Thus, less toxic and more acid-soluble TiO2 has been regarded as a preferred option for blasting Ti implants. It has been reported that 25 μm diameter TiO2 particles could generate a roughened Ti surface with a consistent roughness value of 1  ~  2  μm, which is favourable for the adhesion and proliferation of osteoblasts towards promoting osseointegration (Rasmusson et  al., 2001). Compared to the smooth and machined Ti, a significantly enhanced bone-implant contact (BIC) area was obtained on the TiO2 blasted Ti implant surface (Ra = 1 ~ 2 μm, gritted by 25 μm diameter TiO2) (Rasmusson et  al., 2001). Further, the clinical reliability of TiO2 blasted implants was also validated by showing a 96.9% cumulative implant survival rate over 10 years (Rasmusson et al., 2005). Finally, since washing the embedded particles is recommended after grit-blasting, the fabrication of commercial implants

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always combines sandblasting and acid etching (SLA) to obtain a reliable and consistent implant surface. Implant surface chemistry also influences the performance of adhered cells, and as a result, numerous chemical modifications have been utilized, including acid-­ etching, plasma spraying, sputter deposition, sol-gel coating and electrophoretic deposition, to either augment osteoblasts activity or promote calcification/mineralization on implant surfaces (Le Guéhennec et al., 2007). Acid etching via HCl, H2SO4 and HF are practical options for improving the roughness of Ti implants, which facilitate tiny pits and poles on Ti surface that range around 0.5  ~  2  μm in diameter. Cho and Park reported that micro-roughened implants achieved by acid etching significantly increased removal torque at 3 months after implantation in rabbit tibia in vivo, indicating their influence on early-stage osseointegration (Cho & Park, 2003). In addition, acid-etched Ti implants have been reported with enhanced osteoconductive potential by promoting the attachment of osteogenic cells (Park & Davies, 2000). Compared with the micromachined or plasma sprayed Ti, the acid-etched implants could establish a larger bone-implant contact area with reduced bone resorption, which is both mechanically and biologically stable (Cochran et al., 2002). Hydroxyapatite (HA) is extensively utilized on dental implants, attributed to its similarities with bone minerals in the structural, chemical and mechanical properties to promote the biological apatite precipitation on the implant surfaces (Davies, 2003). HA precipitation on implant structures upregulates bone healing as the deposited biological apatite layer serves as a matrix for the osteoblast attachment and proliferation (Davies, 2003). Further, the release of Ca and P from HA-coated implants improves the new bone formation during the bone remodelling process (Ciobanu & Harja, 2019; Ciobanu et al., 2012; Daculsi et al., 2003; Davies, 2003). Ciobanu and Harja reported significantly accelerated osteogenesis around HA-coated implants, with enhanced implant stability and wound healing (Ciobanu & Harja, 2019). It is noteworthy that HA-coated implants not only promote biological apatite formation but also enhance osseointegration by directly forming an osteoid layer with osteoblasts to improve their proliferation (Goodman et al., 2013). 2.1.2 Incorporation of Bioactive Agents Numerous biomolecules have been incorporated on Ti implants, including growth factors, peptides and proteins. Besides, bioactive polymers like chitosan have also been utilized to modify implants to promote bone healing and osseointegration (Stadlinger et al., 2008). Stadlinger et al. reported that the coating of Ti implants with collagen and chondroitin sulphate (CS) resulted in significantly enhanced in vivo bone-to-implant-contact (BIC) at 5  weeks after implant surgery (Stadlinger et al., 2008). Similarly, another study reported an increased BIC and bone-volume density (BVD) after 4 and 8 weeks within the minipigs’ maxilla on the glycosaminoglycans coated Ti implants (Stadlinger et al., 2012). Functional proteins such as growth factors effectively modulate the differentiation of osteoblasts and stem cells,

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induce angiogenesis, and attract osteoprogenitor cells via chemotaxis effects. However, loading these biofunctional macromolecules onto implant surfaces is a complex process that requires incorporating polymeric coatings as the frame to incorporate those bioactive molecules. For instance, Al-Jarsha et al. spin-coated a layer of poly-ethyl acrylate (PEA) and fibronectin (FN) complex onto Ti implants to establish a thin film to load bone morphogenetic protein-7 (BMP-7) (Al-Jarsha et al., 2018). Compared with the Ti implants directly coated with BMP-7, implants incorporated with PEA-FN-BMP-7 complex could significantly enhance the adhesion/proliferation of human mesenchymal stem cells (hMSCs) and promote the expression of osteogenic marker at day 28 (Fig. 1) (Al-Jarsha et al., 2018). Incorporation of bioactive metal ions is an alternative option to enhance bioactivity on Ti implants. This could be achieved via immersion in varied ionic solutions with controlled pH and concentrations. It is known that Sr2+ and Mg2+ ions promote new bone formation (Song et al., 2018), and soaking Ti implants in a mixture of 50 mM CaCl2 + 50 mM SrCl2 enabled a Sr-containing calcium hydrogen titanate surface that slowly releases Sr2+ ions (Yamaguchi et al., 2014). Similarly, on the Sr/ Mg loaded Ti implants, the in vitro expression of integrin β1, ALP and β-catenin were enhanced from MC3T3-E1 osteoprogenitor cells, indicating an enhanced proliferation and osteogenic differentiation potential (Okuzu et al., 2017). Another modification utilized on Ti implants is coating with bisphosphonates that promotes the apoptosis of osteoclasts to inhibit bone resorption (Costa & Major, 2009). Biphosphonates modified implants accelerated the bone healing by reducing the bone resorption around Ti implants during bone remodelling, and reported a 41% enhanced  pull-out force after 4  weeks within rats tibiae in vivo (Agholme et al., 2012). Thus, biophosphonates coated implants may be favourable for patients

Fig. 1  The osteogenic differentiation of human mesenchymal stem cells (hMSCs) on modified Ti implants. (a) The fluorescence images of hMSCs with osteogenic marker osteocalcin (OCN) and osteopontin (OPN), which were stained green in both top and bottom column; (b) Significantly promoted OCN and OPN expression from the hMSCs on the Ti/PEA/FN/BMP-7 implants, showing a promoted osteogenic potential. Ti Ctrl: Non-treated Ti (Control); Ti/BMP-7: Ti coated with bone morphogenic protein-7 (BMP-7); Ti/PEA/FN: Ti coated with poly ethyl acrylate (PEA) and fibronectin (FN); Ti/PEA/FN/BMP-7: Ti/PEA/FN infiltrated with an additional BMP-7 layer. (Reproduced with permission from Al-Jarsha et al. (2018))

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with compromised bone conditions such as diabetes and osteoporosis, to improve wound healing and osseointegration (Adler et al., 2016). In another study, Shim et al. utilized poly(lactide-co-glycolide) particles to load fibroblast growth factor-2 (FGF-2) onto Ti implants (Shim et  al., 2014). FGF-2 loaded particles could be gradually released, promoting the in vitro alkaline phosphate activity from osteoblasts. Further, compared with the bare Ti, FGF-2 coated Ti implants established significant higher BIC percentage and obtained enhanced new bone formation at 12 weeks within rabbit tibiae (Shim et al., 2014). Another option to coat bioactive molecules is to apply chemical linkers to immobilize biomacromolecules. Zheng et al. reported use of dopamine (DA) and polydopamine (PDA) to immobilize the chitosan particles with BMP-2 loadings on Ti implants (Zheng et  al., 2013). Further, the layer-by-layer self-assembling technique could also be utilized for biomacromolecule coatings. For example, BMP-2 encapsulated BSA (bovine serum albumin) particles were functionalized with chitosan and finally loaded on Ti implants (Wang et al., 2015), and the modified implants significantly increased the proliferation and spreading area of BMSCs, and upregulated their in vitro ALP activity (Wang et al., 2015). In summary, biofunctionalized Ti implants are a favorable option to augment implant integration via incorporation of potent proteins and growth factors, however sustained release for months/years of sensitive proteins still needs further investigation, specially in compromised conditions in vivo.

2.2 Microscale Approaches to Enhance Soft Tissue Integration (STI) Robust STI forms a transmucosal barrier against the ingress of oral microbes; breach of such barrier results in biological complications such as mucosal inflammation, peri-implant bone loss and implant failure (Guo et al., 2021b). However, unlike the osseointegration with robust bone anchorage on the implant structure, the STI at the transmucosal region of the implant is only a ‘physical adaption’ with significantly weakened sealing strength compared to the soft tissue attachment formed by inserting collagen fibers at teeth (Guo et al., 2021b). Various attempts have been made to improve the STI around dental implants by modifying implant topography, chemistry and utilizing bioactivity coatings (Guo et al., 2021a). 2.2.1 Surface Topography Modification The topography of dental implants is a critical factor that influences the morphology, proliferation and activity of adhered epithelial cells/fibroblasts. Compared with the irregular roughened surfaces, Ti implants with aligned microgrooves significantly improved the function of fibroblasts, including their spreading morphologies,

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extracellular matrix (ECM) and collagen secretion (Chou et al., 1995). It has been reported that the human gingival fibroblasts (hGFs) on wider grooves (width 25 ~ 50 μm) were aligned parallel to the grooves and obtained a dense ECM-like structure, translating into the peri-implant wound healing (Yoshinari et al., 2003). Guillem-Marti et  al. reported augmented early-stage adhesion and activation of fibroblasts on microgrooves with width 50 μm) were more effective in maintaining the long-term activity of fibroblasts to secrete more fibres (Guillem-Marti et al., 2013). Laser treatment has been used to create microtopography on Ti with precise dimensions (evenly distributed holes/grooves). For instance, Weiner et  al. inserted various Ti implants into dog mandibular for 6  months and found that laser modified Ti implants with microgrooves (width 12 ~ 24 μm) established stable STI around its surface, with significantly reduced tissue recessions than the smooth counterparts (Weiner et  al., 2008). Similarly, another in vivo study reported augmented adhesion strength of peri-implant epithelium around laser-treated implants with microgrooves within 3 months after implant placement (Nevins et  al., 2010). Moreover, in the connective tissue layer, some fibres were perpendicularly aligned to the laser-treated implant surfaces, indicating the direct connection of fibres with implant surfaces (Nevins et al., 2010). To date, laser-modified dental implants have already been clinically applied (e.g. Laser-­ Lok®). Compared with the conventional (smooth) implants, which only acquired a limited gingiva-implant contact area (24.1%  ±  16.63%), the laser-modified Ti implants enabled significantly enhanced gingiva integration (98.8%  ±  3.78%) at 15 months after surgery. Such enhanced STI establishment and maintenance could be attributed to the augmented junctional epithelium-specific proteins from epithelial layer that enhanced the epithelial attachment, contributing to improved STI strength and stability (Leong et al., 2018). To summarize, compared with smooth surfaces, Ti implants with aligned microgrooves augments the activity of peri-implant epithelial cells and fibroblasts. 2.2.2 Chemical Approaches The chemistry of a material surface is also critical in determining the cell behaviour, including cell adhesion and cytokine expression, which influences the STI formation around implants. Calcium is a critical element related to cell adhesion and cell-­ substrate interaction. As reported in an in vivo study, Ca doped Ti implants (via hydrothermal treatment) could obtain stable peri-implant epithelium attachment throughout the implant-epithelium interface at 6 weeks in the transmucosal region of Wistar rats, significantly surpassing the untreated Ti surface that only presented PIE attachment at the apical 1/3rd of implant-epithelium interface (Oshiro et al., 2015). Meanwhile, the adhesion strength of PIE on modified implants was also enhanced by showing increased resistance against the horseradish peroxidase penetration (Oshiro et al., 2015). Further, phosphate (PO43−) modified Ti implants by ion beam-assisted deposition (IBAD) promoted connective tissue integration on Ti implants, enhancing STI formation and stability (Bao Hong et  al., 2007;

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Zhao et al., 2005). Compared with hydroxyapatite (HA) coated Ti implants, a hybrid coating of tricalcium phosphate/hydroxyapatite (TCP/HA) on Ti significantly enhanced the fibroblast proliferation and adhesion (Zhao et al., 2005). Moreover, similar conclusions were obtained by another in vivo study, in which the Ti implants were coated with calcium phosphate (CaP) and inserted in dog mandibles and healed for 3 months (Bao Hong et al., 2007). Compared with the uncoated Ti, the peri-implant connective tissue around CaP-Ti surface obtained more granular tissue and fibre bundles. Further, some oblique/perpendicular collagen fibres were obtained around CaP-Ti implant surface, suggesting the direct implant-gingiva connection, which was absent from uncoated implants (Bao Hong et al., 2007). 2.2.3 Coating with Proteins Compared with the abovementioned topographical and chemical modifications, the application of biological coatings on the implant surface can further enhance the adhesion and proliferation of surrounding cells. Since cells are attached to the implant surface via adhesive structures, including hemidesmosomes (HDs) and internal basal lamina (IBL), several studies utilized bioactive coatings to promote the secretion of these adhesive structures from cells, to enhance STI formation. As the critical components of HDs and IBL in the epithelial layer, laminin-1 and laminin-5 and their related proteins have been utilized as coatings on Ti implants. Initial attempts included physical deposition, however the binding strength of such coatings on implants was suboptimal (El-Ghannam et al., 1998). Later, chemical functionalizations were utilized to pretreat the Ti implants to reinforce the binding strength of the bioactive molecules. Werner et al. chemically cross-linked laminin-5 on Ti implants that enhanced the in vitro proliferation and adhesion of epithelial cells, translating into enhanced epithelial sealing (Werner et al., 2009). Further, Liu et al. loaded a bioactive domain of laminin-5 (LNA3G3P protein) via the chimeric peptide linking onto Ti implants that yielded a stable and firmly adhered PIE layer within 3  weeks after implantation into the transmucosal region of minipigs (Liu et  al., 2019). Further, facilitating a multilayered porous frame by layer-by-layer self-assembly also contributed to an effective protein/peptide coating. Briefly, LBL self-assembly of a multi-layer hydroxyapatite/collagen (HA/Col) was achieved on Ti implants, which was utilized as a frame to load recombinant adenovirus containing LAMA3 gene (critical gene for synthesizing laminin) (Zhang et  al., 2018). Compared with the uncoated and HA/Col coated Ti, the HA/Col/LAMA3 coated implants obtained larger PIE adhesion area, with stronger adhesive structures presented at the PIE-implant interface at the transmucosal region of Wistar rats at 4 weeks after implant surgery (Fig. 2) (Zhang et al., 2018). Platelet aggregation on Ti implants also play a crucial role in influencing the epithelial seal, since adhered platelet could release cytokines and chemotaxis attractants, which induce cell proliferation and migration. Peptide-activated receptor 4-activating peptides (PAR4-AP, known to influence the initial platelet adsorption), were incorporated on Ti implants to accelerate the aggregations of platelets (Maeno

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Fig. 2  Enhanced epithelium attachments on the Ti implants modified with the LAMA-3 recombinant gene. In vivo immunohistochemical staining of Wistar rats’ periodontal and peri-implant soft tissue at 4 weeks showed laminin α3 (component of adhesive structures, black arrow) expression at the tooth/implant-gingiva interface. Compared with non-modified smooth Ti and Ti coated with chitosan-hydroxyapatite-collagen composite CS/(HA/COL)5, significantly higher laminin α3 expression was present around the Ti implants coated with a combination of LAMA3 gene and HA/Col film (HA/Col/AdLAMA3). The expression strength of laminin α3 around the HA/Col/ AdLAMA3 implant was comparable to that around natural teeth. (Reproduced with permission from Zhang et al. (2018))

et al., 2017; Sugawara et al., 2016). Compared with the bare Ti, PAR4-AP modified implants improved the expression of laminin-5 from epithelial cells, and augmented their adhesion (Maeno et al., 2017). Further, the expressions of collagen IV, which is a critical component of IBL, was also upregulated from epithelial cells on PAR4-AP coated implants (Maeno et al., 2017). In another study, Kihara et al. synthesized a synthetic peptide A10 to coat Ti implants (Kihara et al., 2018), which enhanced the adhesion of the epithelial cells, and established a dense epithelial barrier layer with pericellular junctions on the surface (Kihara et al., 2018).

3 Nano-engineered Implants for Enhanced Osseointegration Compared with the microscale modified Ti implants, nano-engineered implants with customized nanostructures exhibit the following advantages: • Cell functions are significantly influenced by the nanoscale topography via the mechanotransduction effect on cell filopodia, thereby influencing cell spreading and alignment • Nanoscale roughness promotes extracellular matrix secretion from cells to enhance their functions (Gulati et al., 2018a, 2020);

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• Hollow nanostructures (nanotubes/nanopores) could be utilized as drug loading/ releasing reservoirs, to achieve local therapy, including antibacterial and anti-­ inflammatory functions leading to enhanced tissue regeneration (Jayasree et al., 2021) • Nano-engineering modifications can simultaneously alter different surface characteristics of modified implants (such as surface topography and chemistry), which enhances the surface bioactivity (Chopra et al., 2021b; Guo et al., 2021d); • The nanostructured surfaces have increased surface roughness, which enhances hydrophilicity and cellular adhesion and proliferation (Chopra et al., 2021b; Guo et al., 2021d). Various characteristics, including surface roughness, topography, chemistry and wettability have been reported to influence osseointegration and bone healing (Cho & Park, 2003; Ivanoff et  al., 2001; Mendonça et  al., 2008). The commercial microscale implants require 3 ~ 6 months for complete bone healing and osseointegration, and nanoscale implants targets to achieve osseointegration faster (Mendonça et al., 2008; Nagasawa et al., 2016; Wennerberg et al., 2014). Mimicking the morphology of natural bone structures, nanoscale surfaces promote bone regeneration at both cellular and molecular levels (Albrektsson & Wennerberg, 2019; Ivanoff et al., 2001; Rasmusson et al., 2001; Wang et al., 2013).

3.1 Laser Treatment Laser modification is an option to create nanostructured surface on Ti implants, that offers  the capability to control the surface texture by adjusting laser parameters (Ercan et al., 2014). Various laser-related modifications such as laser ablation and laser-induced periodic surface melting can precisely tailor the geometries of modified Ti implant surfaces (Cunha et  al., 2013; Valle et  al., 2015). As reported by Hallgren et al., a nanopatterned surface with uniformed hemispherical pits can be fabricated on implants by Nd:YAG laser beam (Hallgren et al., 2003). Further, when implanted for 12 weeks in rabbits tibiae in vivo, the bone-implant contact (BIC) on such laser-treated implants was significantly enhanced compared to the non-­ modified Ti implants (Hallgren et  al., 2003). Similarly, Faeda et  al. modified Ti implants by sequential Nd:YAG laser ablation and hydroxyapatite (HA) coating, achieving a nanoscale roughened Ti implants that accelerated the bone healing at 4 ~ 12 months with dramatically increased BIC percentages (Faeda et al., 2012). Further, such laser modified surface not only accelerated the bone healing speed, but also improved the osseointegration strength by showing increased removal torque at 12 weeks within rabbit tibiae (Faeda et al., 2012). Additionally, the heat generated by Nd:Yag laser could thicken the TiO2 barrier layer on Ti surface while generating nanoscale rods, which enhanced the stability of nanostructured TiO2 layer (Brånemark et al., 2011). The layer with nanorods obtained significantly increased BIC and showed increased implant removal torque at 8 weeks within rabbit tibiae

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(Brånemark et  al., 2011). It is noteworthy that laser nano-engineering is mainly restricted to several hundred nanometers, and mechanotransduction of osteocytes and osteoblasts (guiding the filopodia extension and stretching) requires smaller nanostructures (diameter 90%

300 days

30 h

7 days

Initial Burst Total Release (IBR) Release

–

S. aureus

–

Bacterial Cell Studied

Table 3  Various approaches for localized delivery of antibiotics using titania nanotubes (TNTs) for antibacterial efficacy

–

Only HAP coatings ensure an excellent antibacterial activity, but faster release. All coatings displayed antibacterial functions.

–

Antibacterial Functions

Release was significantly delayed

Drug was released faster (aqueous electrolyte)

Co-delivery of ibuprofen (anti-inflammatory)

Other Findings

(continued)

Moseke et al. (2012)

Ionita et al. (2017)

Pawlik et al. (2017)

Ref

Cecropin B

Gentamicin

Cefuroxime

4

5

6

No. Drug Loaded

Table 3 (continued)

Nano-tubular

Nano-rugged

Nano-smooth

BMP2-loaded TNTs via LbL [TNT-BMP2LbLg]

Hyaluronidase sensitive multilayers of chitosan/ sodium hyaluroniccecropin B [(Chi/ SH–CecB)5] via LbL technique

Loading Method/ Substrate

D: 70 nm

–

153.2 μg

Maximum release in 1–2 min

24 h: pH 7.4: 17.0 μg pH 5.8: 38.1 μg

90 min

10 days pH 7.4: 44 μg pH 5.8: 130.3 μg

72 h

Initial Burst Total Release (IBR) Release

200 μg/ 6 h: 37% implant A: 10 × 10 mm. LC:2ug/mm2

Amount Loaded (A: area, LC: loading capacity)

L: 300-400 nm 25 mg/mL D: 70–90 nm and 150 mg/mL

pH-responsive D: 70 nm multilayer film: alginate dialdehyde-­ gentamicin (ADA-Gen) and chitosan (Chi)

Chitosan coating

Additional Feature

Nanotube Dimensions (Diameter: D Length: L)

–

S. aureus E. coli

S. aureus S. epidermidis

Bacterial Cell Studied

–

TNT-BMP2LbLg had excellent antibacterial capacity both in early (6 h) and in long-term (72 h)

TNT–CecB– LbLc substrates had good early (4 h) and long term (72 h) antibactericidal capacity against both bacteria

Antibacterial Functions Other Findings

Nano-smooth samples released the least cefuroxime, and the nano-tubular samples released the most

Acidic environment could trigger the release of Gen from the multilayer films and BMP2 from TNTs. In vitro: TNT-BMP2-­ LbLg promoted osteoblast functions.

Good cytocompatibility for osteoblasts, even co-culture with S. aureus

Ref

Chennell et al. (2013)

Tao et al. (2019)

Shen et al. (2016)

Vancomycin (NT-V)

9

Circular Ti substrates with presynthesized TiO2 nanoparticles (Ti-NPs).

Simplified lyophilisation method

–

Immersion – method (I) or Electrophoresis method (E)

Mesoporous thin films composed of TiO2 nanoparticles on anodized Ti for loading drugs at high doses.

500 μg/cm2 In vitro 15 min: 600 μg/mL In vivo Unable to measure

75 V 24 h: >0.14 mg (I); 0.18 mg (E)

75 V D: 115 ± 5 nm L 3.4 ± 0.1 μm

In vitro D: 80 nm L: 800 nm In vivo Rod, D: 1 mm H 20 mm

60 V 24 h: >0.2 mg (I); 0.16 mg (E)

60 V L 3.2 ± 0.3 μm D: 93 ± 3 nm

–

4.4 μg/day during the first 2 days

IBR@ 9 h 1.12 μg/h

IBR@ 9 h 1.06 μg/h

IBR@ 9 h 1.28 μg/h

50V 24 h: >0.1 mg (I); 0.2 mg (E)

2.90 μg

5.70 μg

3.01 μg

50V – L 2.5 ± 0.1 μm D: 78 ± 3 nm

Ti substrate D: 12 mm Ti-NPs ∼20 nm

210 min

28 days

–

16 days

119 h

119 h

119 h

S. aureus

S. aureus

S. aureus P. aeruginosa A. actinomycetemcomitans P. intermedia P. gingivalis

Adapted with permission from Chopra et al. (2021) TNTs titania nanotubes, LbL layer by layer, LC loading capacity, IBR initial burst release, HA/HAP hydroxyapatite

Vancomycin

Ag + Minocycline

Ag NPs

Amoxicillin

Minocycline

Cephalothin

8

7

Good antibacterial effect both in vitro and in vivo.

TNTs anodized at 60–75 V showed strong antibacterial behaviours against S. aureus due to high IBR

Minocyclineloaded samples showed maximum efficacy Minocycline combined with Ag NPs showed effectiveness against all tested bacteria.

Good biocompatibility Zhang et al. (2013)

Most uniform Mansoorianfar morphology, et al. (2019) appropriate drug release, cell viability behaviour achieved by TNTs [60–75 V]. Drug loading efficiency increases up to 60% via electrophoresis method (for 75 V TNTs).

Park et al. (2014)

While the localised delivery of antibiotics helps prevent infection at the implant site, these treatments might not be effective against resistant strains of bacteria like methicillin-resistant Staphylococcus aureus and polymicrobial systems (Godoy-­ Gallardo et al., 2021; Lin et al., 2021). Delivery of metallic ions or NPs like Ag, Cu, F, Zn have been explored recently to address the shortcomings of localised delivery of antibiotics. Chen et  al. (2013) demonstrated a dual action system (loading of AgNPs into TNTs followed by immobilisation of quaternary ammonium salt (QAS)) where high positive charge of QAS attracts negatively charged bacteria and induces contact killing, while AgNPs released into the surrounding environment eliminates remaining bacteria. They observed approximately 90% killing efficiency for TNT-QAS against Escherichia coli over 30 days, while TNT-Ag-QAS showed an enhanced killing efficiency of 99.9%. Though Ag-loaded TNTs showed minor cytotoxicity towards osteoblasts, the presence of QAS enhanced osteoblast activity. Jia et al. proposed a triple action mechanism of preventing infections at implant surfaces using AgNPs (Jia et al., 2016). Initially a mico-nano surface was fabricated on Ti using micro-arc oxidation (MAO) followed by coating self-polymerising polydopamine and immobilization of AgNPs. The release of Ag+ ions repel the attachment of planktonic bacteria onto the implant surface, although the action is not 100% effective and some bacteria still make it to the implant surface. However, upon landing, these bacteria come in contact with the AgNPs present on the surface and undergo apoptosis. Finally, negatively charged bacteria are attracted to the surface micropores where they undergo ‘trap-killing’ attributed to either collision with pore walls or Ag NPs-initiated membrane destruction (Fig.  5). Additionally, an effect of the modified implants to enhance osteoblast adhesion, pseudopodal extension and ALP activity was demonstrated in vitro, while a minimal immune response to their subcutaneous implantation in a rabbit model supported clinical applicability. Table 4 summarizes the key advances relating the loading and release of metal-­ based ions and NPs from TNTs towards local therapy. Even though the incorporation of metallic nanoparticles and ions have shown great potential to eliminate infection in in vitro and in vivo studies, detailed investigations regarding their long-term effects on therapeutic efficacy and toxicity are required (Gulati et al., 2021).

6 Strategies of Regulating Drug Release To ensure clinical translation of localised therapy from nano-engineered Ti implants the drug release pattern should be tailorable, matching specific therapeutic requirements. The release of drugs from TNTs usually follows Fick’s first law of diffusion and depends primarily on: (a) drug- size, molecular weight, charge; (b) TNT- diameter, length, charge, and surface chemistry; (c) interaction between drug and implant surface; and (d) other factors like pH and temperature (Losic et al., 2015).

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Fig. 5  Bioinspired AgNPs loaded titanium implants for trap killing of bacteria. (a) Schematic representation of fabrication of AgNPs immobilized micro-arc oxidation (MAO) system. (b) Schematic representation showing the three proposed killing mechanisms. (c) SEM images demonstrating ‘trap killing’. The colours represent: Yellow: bacteria, blue: nanosilver and pink: nanosilver bound to bacteria. (Adapted with permission from Jia et al. (2016))

Soon after implant placement, high concentration of drug is released consequent to rapid diffusion of the drug present on the surface of the implant (initial burst release or IBR) (Gulati et al., 2015a,b). Regulating and minimising IBR is critical as these abrupt high doses of drugs can be toxic to nearby cells thereby affecting osseointegration and STI. Even though the most desirable strategy is to obtain a zero-order kinetics where the drug is released at uniform rate irrespective of time and concentration, in certain situations a lower or higher dose of drug might be preferred. Several strategies have been utilized to tailor the release pattern of drugs from the implant surface and will be discussed in the following sections.

Ag NPs

Ag2O- NT 01.27 Magnetron Ag2O- NT 04.67 sputtering and anodization Ag2O- NT 07.43

4

5

Ag2O- NT 014.63

Ag NPs

3

D: 20 nm

Photo-reduction

D: 5–20 nm

D: 8 nm

Spin coating and D: 40 nm annealing

Visible-light irradiation

Au NPs

2

Nanotube Dimensions

Length decreases with increase of Ag content

D: 100 nm L: 15 μm

D: 110 nm L: 900 nm

D: 150 ± 10 nm L: 1.5 μm

7 days: 50%

42 ppb 7 days: 50%

7 days: 25%

28 ppb

7 days: 50%

50 ppb

–

–

–

–

28 days

24 h

24 h

24 h

24 h

Initial Burst Release [time + amount Total released (%)] Release

45 ppb

–

–

–

Chemical D: D: 3.35– reduction using 102 ± 21 nm 116.2 ± 6.4 nm 14.6 ppm δ-gluconolactone (GL)

Ag NPs

1

No. Metal Loaded

Loading Method/ Size of Substrate Metal NPs

Amount Loaded (μg)

E. coli S. aureus

E. coli

E. coli

P. gingivalis F. nucleatum

S. aureus

Immunomodulatory functions from TNT-Au

–

Bioactivity/ Toxicity Evaluation (special features)

Xu et al. (2019)

Gunputh et al. (2018)

Ref.

Antibacterial rates higher than 97%

Bacterial inactivation

No cytotoxicity and supports cell proliferation

Gao et al. (2014)

Enhanced TNTs Hajjaji et al. crystallinity leads to (2018) reduced surface defects.

The dual action Displayed long-term Chen et al. antibacterial biocompatibility. (2013) efficacy achieved via contact and release killing.

UV irradiation increased antibacterial functions

Ag NPs exhibited antibacterial effect in both micron- and nano-sized clusters.

Bacterial Cell Antibacterial Studied Efficacy

Table 4  Localised delivery of metallic/semi-metallic ions and nanoparticles (NPs) from titania nanotubes (TNTs) and evaluation of their antibacterial efficacy

176 A. Jayasree et al.

Cu NPs-0.3 Cu

7

Ag NPs

Cu-Ti-O NTAs

8

9

Cu NPs-3.0 Cu

Ag-doped hydroxyapatite (Ag-HAp)

6

D: 100 nm

AMS Cu-Ti-O20.47

AMS Cu-Ti-O15.14

AMS Cu-Ti-O4.62

AMS Cu-Ti-O2.69

Anodizing magnetron-­ sputtering (AMS) Cu-Ti-O0.00

Micro-arc oxidation

–

D: 50 nm

Micro-arc D: 1–5 μm oxidation (MAO)

Electrophoretic Ag-HAp-0 Electrophoretic Ag-HAp-0.02 Electrophoretic Ag-HAp-0.05 Electrophoretic Ag-HAp-0.08 Electrophoretic Ag-HAp-0.1

–

–

T: 5–10 μm

T: 24.2 μm

–

8.57 μg/ cm2

140 ppt

1.6 ppm

–

6 h: 2.65 μg/ cm2

24 h ~ 135 ppt

24 h ~ 115 ppt

28 days

28 days

14 days

24 h: >0.3 ppm 14 days

S. aureus

S. aureus

S. aureus

No colony observed on Cu- NTAs during the first 21 days. Only few colonies can be seen after immersion for 28 days. Antibacterial rate remains >90% at the end of test duration (permanent bactericidal effect).

Nearly 70% decrease of viable bacteria and 300% increase of dead cells

Excellent antibacterial activity

S. aureus Ag-HAp-0.05 P. aeruginosa showed excellent antimicrobial efficacy (> 99% reduction in viable cells)

Mirzaee et al. (2016)

No cytotoxicity of Cu-Ti-O NTAs to endothelial cells (ECs) and upregulated cell proliferation. Enhanced in vitro angiogenesis activity of endothelial cells.

Antibacterial activity is reduced by over 64.2%

(continued)

Zong et al. (2017)

Jia et al. (2016)

0.3 Cu promotes Zhang et al. osteoblast functions. (2018) 3.0 Cu shows cytotoxicity.

The passive current densities of the HAp -TNTs are lower than those of Ag-HAp-TNTs, leading to a slightly lower corrosion resistance.

Local Therapy from Nano-engineered Titanium Dental Implants 177

D: 200 nm

D: 0.5 μm

D: 100 nm

Fluoride incorporated Ti–6Al–4V

Ti/Van HA Drop-casting (Hydroxyapatite)

Ti/Van HA-collagen

Ti/Van TiO2

12

T: 200 nm

EA with fluoride-free TiO2 BL

L: 3.5 μm

–

–

D: 20-40 nm T: 150 nm

T: 5.1 ± 1.6 μm

Nanotube Dimensions

EA with fluoride-TiO2 barrier layers (FBL)

–

11

B-CaP coating

Plasma electrolytic oxidation (PEO) CaP coating

B, P, Ca

Loading Method/ Size of Substrate Metal NPs

10

No. Metal Loaded

Table 4 (continued)

–

–

–

Amount Loaded (μg)

IBR @ 5 h Ti/Van TiO2 : >54%

IBR @ 5 h Ti/Van HA-collagen : >62%

IBR @ 5 h Ti/Van HA : >75%

–

B-CaP coating Ca ~ 1.8 ppm P ~ 0.6 ppm B ~ 0.4 ppm

IBR @ 24 h CaP coating Ca ~1 ppm P ~ 0.6 ppm

30 h

–

28 days

Initial Burst Release [time + amount Total released (%)] Release Bioactivity/ Toxicity Evaluation (special features) Ref.

S. aureus

S. aureus S. epidermidis

TNTs have long term release. Both coatings show antibacterial functions.

Decrease in the bacterial adhesion as compared to fluorine free barrier layers (BL).

–

FBL surface (increased roughness and surface energy) promotes increased protein adsorption. No change in surface topography

Ionita et al. (2017)

Arenas et al. (2013)

Sopchenski S. aureus Boron in B-CaP B incorporation et al. (2018) P. aeruginosa coating prevents does not change biofilm formation coating morphology and crystallinity in comparison with free B coating. B presence promoted ADSCs spread after 1 day of culture, with no cytotoxicity.

Bacterial Cell Antibacterial Studied Efficacy

178 A. Jayasree et al.

Gallium (III) Drop casting loading on 3D printed nanopiller

15

–

–

ZnO QDs D: 3–5 nm

L: 550 nm

D: 146 nm L: 7.12 μm

TNTs ID:~80 nm

Adapted with permission from Chopra et al. (2021)

Dip coating

(poly-DL-lactic acid) with gallium (III) dip coated on TNTs

14

Vancomycin loaded TNTs capped by ZnO-FA QDs

Folic acid conjugated ZnO quantum dots. TNTs-Van@ ZnO-FA QDs

13

683 μg

100 μL ZnO-FA solution of 1 mg/ mL

30% in 1 d

50 ppm in 1 d

5 h

60% in 5 d

400 ppm in 14 d

24 h

Ga-doped TNTs showed excellent anti-bacterial property (assessed in the spinal infection rat model in vivo)

Antibacterial action enhanced from pH 7.4 to 5.5: NTs-Van@ ZnO-FA QDs 60.8% to 98.8% TNTs-Van 85.2% to 95.1% Reduced inflammation and favourable compatibility with osteoblasts

Excellent biocompatibility of NTs-Van@ZnO-FA QDs with MC3T3-E1 cells in vitro

S. aureus 100% eradication – P. aeruginosa of both bacteria was observed

S. aureus E. coli

S. aureus

Maher et al. (2022)

Dong et al. (2019)

Xiang et al. (2018)

Local Therapy from Nano-engineered Titanium Dental Implants 179

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6.1 Altering TNTs Dimensions Based on Fick’s law of diffusion, the dimensions of TNTs play a major role in determining the rate of drug release and can be easily tailored by varying the parameters of anodization. TNTs of fixed diameter (110 nm) but varying lengths (25–100 μm) were fabricated and indomethacin drug loading and release was analysed to evaluate the effect of TNTs length on drug release kinetics (Aw et al., 2014). An initial drug loading capacity of 15% and 26%, and overall drug release of 6  days and 23 days, were observed for TNTs with length 25 μm and 100 μm, respectively. A reduced IBR was observed with an increase in the length of TNTs indicating that longer tubes ensure deeper loading of higher drug concentration that delays its diffusion and reduces IBR and total release (100% release). Similar experiments with nanoporous anodized alumina (NAA) studied the effect of varying nanopore diameter (65–160 nm) with constant length and observed that greater the diameter of the nanostructure, greater the contact with surrounding media leading to higher rate of diffusion (Aw et al., 2014). Hamlekhan et al. demonstrated that TNTs of diameter 60–80 nm with 1–5 μm length released the entire drug amount within 25–110 min, and TNTs of diameter 110–170 nm with length of 40–70 μm prolonged the release for 4–11  days (Hamlekhan et  al., 2015). Overall, varying the aspect ratio of the TNTs can help tune the drug release kinetics. To evaluate the correlation between the size of the loaded drug molecule and the dimension of TNTs, Peng et al. studied the release of albumin (large protein molecule), paclitaxel and sirolimus (small molecules) from TNTs of various dimensions (Peng et al., 2009). TNTs (100 nm diameter) of length 1 μm held less than half the amount of drug held in 5 μm, confirming that longer the tube, the greater the volume for drug entrapment. They also observed that TNTs of diameter 100 nm and 5 μm length could prolong the release of the larger sized albumin to 30 days, while small molecules were released within 7–14 days. To summarise, longer and wider TNTs can be loaded with higher drug amounts but can also demonstrate high IBR and faster release in comparison to short and narrow TNTs (that can be loaded with lower drug amounts but show lower IBR and prolonged drug release). Further optimization and techniques to effectively alter these parameters to obtain an improved control over the release pattern need to be developed.

6.2 Polymeric Modifications of TNTs A thin polymer coat on the open pores of drug-loaded TNTs can act as a barrier towards the diffusion of drug from the TNTs and thereby achieve controlled release. Further, polymer encapsulation of sensitive drugs prior to loading inside the TNTs can also be utilized as a strategy to control drug release. Vasilev et al. utilised plasma polymerisation (PP) to deposit a layer of poly(allylamine) onto TNTs to create a thin chemically reactive chemical coating containing amine functional (Vasilev et  al., 2010). The functional groups present in this coating

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was utilised to fabricate two types of TNT surfaces; (a) TNTs coated by LbL assembly of poly(sodium styrenesulfonate) (PSS), and (b) TNTs covalently coupled with poly(ethylene glycol) (PEG). They demonstrated that by altering the time of plasma deposition the diameter of the nanotubes can be changed from 20 to 140 nm. The proposed system is an interesting proof of concept for incorporation of polymeric coatings on the surface of nanotubes. However, PP requires costly equipment and which makes it highly unsuitable for translation purposes. As an alternative, scientists have developed a much simpler alternative of dip coating polymers onto surface of TNTs. Gulati et  al. dip coated indomethacin loaded TNTs with chitosan and PLGA and evaluated its influence on the drug release kinetics (Gulati et al., 2012). Chitosan modified TNT showed a 35% IBR and PLGA modified TNT a 12% IBR within 6 h in comparison to uncoated samples which showed an IBR of 77%. While the entire drug in uncoated samples was released within 4 days, the coated TNTs showed a prolonged release up to 30 days, thereby establishing dip coating as an efficient technique to prolong the drug release from TNTs. These biopolymeric coatings also demonstrated enhanced osteoblast cell adhesion at early time points indicating that these surfaces have great potential for improving osseointegration and control local therapy synergistically. A similar pattern was observed when BMP-2 loaded TNTs were coated with gelatin and chitosan (Hu et al., 2012) and LbL deposition of alginate dialdehyde-gentamicin and chitosan (Tao et  al., 2019). Briefly, the PLGA coating on BMP-2 loaded TNTs prolonged the release of BMP-2 to 28  days in comparison to the 80% IBR observed in non-coated TNTs within 1 day (Zhang et al., 2021b). In a pioneering attempt, Fathi et  al. developed a silk fibroin nanofiber coated vancomycin loaded TNTs (Fathi et al., 2019). The vancomycin was drop casted on the TNTs followed by electrospinning of silk fibres onto their surface. They demonstrated that the amount and duration of drug release can be modulated by altering the diameter of the silk nanofibers. By decreasing the diameter of nanofibers from 350 to 180 nm, the IBR within 6 h could be reduced from 73% to 30% and the total release can be prolonged from 1 to 28 days.

6.3 Encapsulation of Drug in Nano-carriers Ensuring the stability of sensitive therapeutics such as drugs, growth factors and proteins, and preventing their denaturation and precipitation within the physiological environment is one of the major challenges in LDD (Atanase, 2021; Son et al., 2021). Polymeric micelles with hydrophobic and hydrophilic cores have shown great potential in localised delivery of sensitive payloads (Ghosh & Biswas, 2021; Gigmes & Trimaille, 2021). Aw et  al. loaded TNPs with five types of polymeric micelles (i) d-α-tocopheryl polyethylene glycol 1000 (TPGS), (ii) Pluronic F127, (iii) PEO(260)–PPO(400)–PEO(260), (iv) 1,2-distearoyl-sn-glycero-3-­­ phosphoethanolamine-N-[methoxy (polyethylene glycol)-5000] (DGP 5000), and

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(v) 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy (polyethylene glycol)-2000] (DGP 2000) (Aw, 2011). All groups showed similar IBR pattern for 6 h and sustained release was observed in all micelles for 21 days irrespective of micelle type. The largest micelle, PEO-PPO-PEO (75  nm), showed a slightly quicker drug release. To further prolong the drug release the TNP loaded with micelles were further coated with allylamine using PP and it was observed that a 70  nm thick coating reduced the IBR by 50% and prolonged the drug release to 30 days.

6.4 Triggered Therapy Various strategies to control and regulate a sustained local release might address therapeutic demands in most cases, although they fall short in cases when a change in therapeutic dosage is needed to obtain a desired effect (Jayasree et al., 2021). A further drawback of sustained release is complete consumption of loaded drugs/ antibiotics that may allow re-infection in later stages. The ability to control drug release based on the needs of the local microenvironment is ideal. ‘On-demand’ triggered therapy can minimize IBR and achieve maximized therapeutic potential. Triggered therapy using an external or internal stimulus has been explored to overcome shortcomings of sustained drug release. TNTs are an ideal candidate for triggered therapy due to their enhanced surface properties and ease of functionalisation. The chemical/temperature changes in the implant microenvironment (onset of infection/inflammation) can be utilized as internal stimuli for localised therapy (Liu et al., 2015). While internal triggers cater to the microenvironment and cannot be regulated by a clinician/patient, external triggers such as magnetic/electric fields and electromagnetic waves provide more control as they can be managed externally (Timko et al., 2010). Some of these triggered therapy systems are discussed in the following sections (Fig. 6). 6.4.1 Enzyme Trigger At various stages of implant surface microbial infection, the pathogens secrete enzymes like hyaluronidase (HAase) and chymotrypsin that can also be utilised as triggers to initiate drug release from TNTs (Yu et al., 2020). Grafting antibiotics with polymers like hyaluronic acid that are easily degraded by HAase can be used to achieve enzyme triggered release. Yu et  al. loaded defroxamine (DFO) inside TNTs followed by spin coating gentamycin-grafted hyaluronic acid (Yu et  al., 2020). Observing a sustained release of DFO from the TNTs in absence of HAase and a burst release when HAase was added confirmed that the presence of HAase enzyme triggered drug release. Further, the presence of E. coli and S. aureus resulted in a release of gentamycin that reduced the microbial load. It is noteworthy that the

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Fig. 6  Triggered release of drugs from titania nanotubes (TNTs). Schematic representation: (a) uncontrolled drug release from TNTs, (b) triggers based on internal factors like pH and enzyme, (c–f) externally triggers: ultrasound, magnetic field, electrical stimulation and electro-magnetic radiation. (Adapted with permission from Jayasree et al. (2021))

microbes secrete large amounts of HAase as part of their metabolism, which degrades the polymers and triggers the release of gentamycin. A similar principle was employed by Yuan et al., in developing vancomycin-loaded TNTs coated with dopamine-modified hyaluronic acid and 3,4- dihydroxyhydrocinnamic acid-­ modified chitosan via LbL assembly (Yuan et al., 2018). The researchers reported a sustained release of vancomycin for 6 h followed by a triggered release upon introducing HAase. They further demonstrated an increase in the amount of drug release (80% of drug within 6 h) in the presence of S. aureus, compared with only 25% release in the absence of bacteria, confirming that bacterial load can act as an internal trigger to facilitate drug release. Further, the antibacterial efficacy of the trigger system was evaluated in vivo in a rat femur infection model, with significantly lower microbial load, lower neutrophil infiltration and higher bone formation observed, demonstrating the antibacterial efficacy of the system while simultaneously improving osseointegration. 6.4.2 pH Trigger Bacterial attachment and biofilm formation at an implant site can reduce the local pH from 7.4 to an acidic 5.5, and this change can be utilized as a trigger to facilitate an ‘on demand’ release of therapeutics (Dong et  al., 2017; Ribeiro et  al., 2012). Dong et al. utilised a pH sensitive acetal linker to immobilise AgNPs on TNTs and

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assessed its release pattern with changes in pH (Dong et al., 2017). Though a sustained release of AgNPs was observed at both neutral (7.4) and acidic pH (5.5), the amount of drug released in pH  5.5 was 2.5 times higher than pH  7.4. Further, a sustained release was observed for 28 days in pH 7.4, followed by a burst release of high dose of AgNP when the pH was dropped to 5.5. The proposed system showed great potential as an implant surface modification that limits drug release at physiological conditions for prolonged periods and provide a burst release upon infection. Wang et al. used the same principle to develop TNTs loaded with AgNPs and vancomycin and coated them with a pH sensitive coordination polymer 1,4-bis (imidazol-1-ylmethyl) benzene (BIX) (Wang et al., 2017). Capping of vancomycin loaded TNTs with folic acid modified ZnO quantum dots demonstrated a similar triggered release pattern, with release of 500 μg/mL within 15 days in pH 5.5 compared to the 200 μg/mL observed at pH 7.4 (Xiang et al., 2018). 6.4.3 Electrical Triggers/Electrical Stimulation Therapy (EST) EST or the modulation of cellular activity by providing an electrical stimuli has shown to augment differentiation of osteoblasts (De Giglio et al., 2000), fibroblasts (Shi et  al., 2008), neurons (Gomez & Schmidt, 2007), endothelial cells (Garner et al., 1999) and cardiomyocytes (Nishizawa et al., 2007) through signalling cascade activation. In pioneering studies, Sirivisoot et al. developed Polypyrrole (PPy) doped TNTs loaded with penicillin, streptomycin and dexamethasone via electrodeposition (Sirivisoot et al., 2011). The continuous redox reactions occurring within PPy upon electrical stimulation lead to rupture of bonds between the polymer and drug molecules resulting in an abrupt release of 80% of the drug within 5 cycles of electrical stimulation. Cyclic voltammetry measurements by applying voltages of −1 to 1 V showed an ON-OFF release pattern for drug release corresponding to electrical stimulation. Further, Shi et al. (2013) demonstrated that while chitosan coated vancomycin loaded TNTs showed sustained release of drug in the absence of trigger, application of a voltage of 3 V triggered an abrupt release of 50% of drug within 10 min. Upon application of voltage, a reduction in pH caused degradation of the chitosan layer and thereby enabling quick diffusion of vancomycin. Gulati et al. converted TiO2 nanotubes on Ti wires into Ti nanotubes via magnesiothermic reaction and demonstrated that such systems can be used to enable simultaneous EST and LDD (Gulati et al., 2016c). Interestingly, as the nanotubes were not modified, they reported that drug release was not dependent on the EST. 6.4.4 Magnetic Field Due to its ease in handling, high tissue penetration and minimal adverse reactions, magnetic field is the most versatile trigger mechanism widely investigated for targeted therapy and imaging (Bi et al., 2016). The success of magnetic NPs in the field

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of targeted drug delivery and imaging has inspired researchers to explore their potential as a localised triggered therapy mechanism (Bi et al., 2016; Hoare et al., 2009). In a pilot study, Gulati et al. demonstrated an interesting system of loading TNTs with magnetic iron oxide nanoparticles followed by indomethacin for triggered therapy (Gulati et al., 2010). The same group later evaluated the release of indomethacin loaded polymeric micelle from TNTs via magnetic trigger (Aw et al., 2012a). TNTs were initially loaded with dopamine-conjugated iron oxide NPs followed by incorporation of polymeric micelles loaded with indomethacin. Three types of micelles, d-a-tocopheryl polyethylene glycol 1000 (TPGS), Pluronic F127 and PEO–PPO–PEO were used, and a sustained release pattern was observed in the absence of a magnetic field, with an IBR of 50% within 5 h. However, upon introduction of a magnetic field an abrupt drug release of 95–100% was observed within 1.5 h. 6.4.5 Radiofrequency (RF) Increased skin permeation of electromagnetic waves in the range of 3 kHz–300 GHz makes RF highly suitable for non-invasive imaging and drug delivery applications. Bariana et  al. used a solenoid copper coil and a variable frequency generator to evaluate RF triggered release of indomethacin from gold nanoparticles (AuNPs)loaded TNTs upon application of a 1 GHz RF field (Bariana et al., 2014). TNTs were initially loaded with AuNPs followed by either bare indomethacin (Ind-TNT) or indomethacin preloaded in TPGS micelle (Ind-TPGS-TNT). Ind-TNT and Ind-­ TPGS-­TNT demonstrated a burst release of 100% of drug within 5 min when RF was applied, while in the absence of RF trigger a release of 25–30% was observed within 3 h. The release was attributed to the application of RF that generated eddy currents causing friction amidst the vibrating AuNPs, leading to temperature increase and a convective displacement of indomethacin from the TNPs. The study also showed that drug release can be further tailored by varying the time of RF exposure, for example, 2–5 min RF exposure corresponded with 70–90% of drug released, respectively. 6.4.6 Near Infra-Red (NIR) NIR waves in the range 650–900 nm can penetrate tissues and bone with minimal phototoxicity and minimal attenuation of blood and soft tissues (Cho et al., 2015). Several NIR based therapeutic drug delivery systems utilise the photolytic cleavage mechanism, where the NIR waves causes cleavage of polymeric chains and release payloads (Cao et al., 2013; Wu et al., 2008). Recently, upconversion nanoparticles (UCNPs) were investigated for triggered therapy due to their ability to convert NIR into different wavelengths such as visible and UV (Carling et al., 2010; Chen et al.,

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2014). Zhao et  al. combined UCNP and photolytic cleavage potential of NIR to develop amphiphilic TNTs with AuNP grafted hydrophobic cap and an ampicillin loaded hydrophilic bottom for triggered therapy (Zhao et al., 2020). Upon application of NIR, the UCNPs absorbed the photons and emitted waves of lower wavelength that generated reactive oxygen species (ROS) resulting in cleavage of the link between ampicillin and TNT, and thus its release. Upon triggering with NIR, a burst release of 75% of drug was observed within 120  min, compared to 20% release observed without the trigger. A sustained release of only 10% was observed for 10 days, with a subsequent burst release at day 10 upon NIR illumination. Further, bioactivity evaluation with human keratinocyte cell line (HaCaT) in vitro showed that NIR stimulation can cause production of ROS leading to minor cytotoxicity. 6.4.7 Visible Light Visible light of wavelength 380–700 nm has been used extensively for phototherapy in medicine since 1960. The photothermal, photocatalytic and photodynamic properties of light on interaction with polymeric and metallic surfaces make them highly suited for triggered therapy applications (Yun & Kwok, 2017). An amphiphilic TNTsbased LDD system has recently been developed where TNTs are loaded with ampicillin via a silane linker (3-glycidozypropyl) trimethoxysilane (GPMS) (hydrophilic part) and an AuNP-octadecylphosphonic acid (ODPA) hydrophobic cap as shown in Fig. 7 (Xu et al., 2016). Upon illumination, surface plasmon resonance of AuNPs triggers incision of the hydrophobic cap (followed by the cleavage of ODPA) resulting in ampicillin release. In this system, bare drug-loaded TNTs exhibited an IBR of 80% within 10  min, while the amphiphilic TNTs produced 60 nm, reduced stability was observed (Liu et al., 2011). Further attempts at augmenting corrosion/electrochemical stability involve the treatment of TNTs via thermal oxidation (Grotberg et al., 2016) and Ca/P/Zn biofunctionalization using reverse polarization (Alves et al., 2018). For biofunctionalized TNTs, the segmented tribo-electrochemical resistance was attributed to forming a thin oxide film at the interface between TNTs and Ti and forming a P-rich tribofilm (Alves et al., 2018). Chapter 8 focuses on understanding and augmenting the corrosion and electrochemical stability of modified and nano-engineered Ti dental implants.

3.2 Stability During Sterilization Effective cleaning and sterilization post-fabrication are essential steps for implants and their surface modifications. Interestingly, for anodized Ti-based dental implants with TNTs or TNPs, only a few studies have explored and optimized this aspect. Zhao et  al. reported, for the first time, the influence of varied sterilization techniques, including autoclaving, UV irradiation and ethanol treatment, on the bioactivity of TNTs (Zhao et al., 2010). The study reported that UV irradiation was the most favorable sterilization option attributed to observing higher proliferation and mineralization of osteoblasts in vitro and effective removal of organic impurities on UV-treated TNTs. Further investigations evaluated wet autoclaving (TNTs in water) and dried autoclaving (TNTs sealed, no water) and found that dry autoclaving was favorable towards osteoblast proliferation in vitro (Oh et al., 2011). Next, comparing various sterilizations, Kummer et  al. reported that UV and ethanol treatment reduced bacterial growth while autoclaving enhanced the bacteria on TNTs (Kummer et al., 2013). While the abovementioned studies investigated the influence of various sterilization methods on TNTs bioactivity, the mechanical stability of the sterilized TNTs was not examined. Junkar et al. performed mechanical testing during sterilization and reported that among autoclaving, UV, H2O2 plasma and O2 plasma, autoclaving was unsuitable as it caused delamination and mechanical failure of the TNTs anodic film (especially

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Fig. 6  Autoclaving of TNTs damages the nanostructures and alters surface morphology. AFM images of 100 nm TNTs sterilized using (a) autoclaving, (b) UV irradiation, and (c) plasma treatment. (Adapted with permission from Junkar et al. (2016))

for large diameter TNTs, confirmed via SEM and AFM imaging) (Junkar et  al., 2016). The damage to TNTs during autoclaving is a combined influence of moisture and high temperature/pressure, which induced TNTs crystallization and changed the surface morphology of TNTs. Figure 6 presents the AFM images of the various sterilized TNTs and clearly shows that autoclaving resulted in TNTs destruction and change in surface morphology. To perform disinfection of TNTs, Beltrán-Partida et al. used superoxide water (SOW with H2O2 and oxidizing radicals) to clean TNTs, which resulted in significantly enhanced osteoblast (pig periosteal osteoblasts) functions while reducing bacterial viability (Staphylococcus aureus) in vitro on cleaned TNTs (Beltrán-­ Partida et al., 2016). In another study, Radtke et al. reported that the structure and morphology of TNTs (amorphous, fabricated at low voltages) were not influenced by autoclaving (Radtke et al., 2019). The authors also found that TNTs fabricated at higher voltages must have absorbed water adequately removed prior to autoclaving to prevent structural damage. More recently, Guo et al. thoroughly investigated the influence of various sterilization techniques on topography, chemistry, bioactivity and stability of the titania nanopores (TNPs) (Guo et  al., 2021d). Briefly, using appropriately aged electrolyte, micro-rough Ti substrates were anodized to fabricate anisotropic TNPs, followed by sterilization using autoclaving (wet and dry), ethanol immersion, gamma irradiation and UV irradiation (various times). The findings revealed that autoclaving compromised the mechanical stability of the anodic film. Among other techniques, UV irradiation (irrespective of the time of exposure) resulted in favorable hydrophilicity, protein adhesion capacity and proliferation of gingival fibroblasts in vitro. Next, nanoindentation testing revealed that ethanol immersion reduced the TNP elasticity, while UV and gamma irradiation showed similar modulus and hardness values as the non-sterilized TNPs.

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4 Testing Stability Post-implantation Ti substrates and implants modified with TNTs and TNPs have been inserted in bones ex vivo and in vivo to observe integration, therapy and mechanical stability. While most attempts have been made on bones, they lack the use of a masticatory simulator to mimic an oral cavity.

4.1  Ex Vivo Implantation As discussed previously, the oral cavity represents a complex environment, and the implant micro-environment is further challenged with implantation surgery. Ex vivo models allow for mock clinical surgery and can be used to investigate if the implant can survive the handling and surgical placement. To assess the mechanical stability of bare and silver nanoparticles (Ag NPs) modified TNTs on Ti rods, Shivaram et al. implanted the TNTs into equine cadaver bone ex vivo (Shivaram et  al., 2016). Briefly, the implant placement involved drilling holes and hammered insertion, and the implants were retrieved after 14 days and analyzed using SEM. Both bare and Ag-TNTs showed no visible damage or delamination, and the implantation did not impact the release profile of therapeutic Ag NPs. To ease bone ex vivo maintenance and manipulation for complex experimental conditions, three-dimensional (3D) bone reactor-Zetos™ was used to investigate the therapeutic release and stability of TNTs modified Ti wires (Aw et al., 2012). The Zetos™ system maintains the bone viable for up to 3 weeks ex vivo via continuous media perfusion and can also exert load cycles to evaluate mechanical characteristics of the bone/implants (David et al., 2008). Rahman et al. inserted rhodamine B dye-loaded TNTs/Ti wires into bovine trabecular bone cores (three types: marrow removed; marrow intact: coagulated; and marrow intact: coagulation prevented via heparin) ex vivo using Zetos™ system (Rahman et al., 2016). Upon retrieval of the TNTs/Ti wire implants after 11 days, the implants were analyzed via SEM, and the data confirmed that the TNTs retained their structures without any visible deformation or delamination. It is known that anodization of curved substrates like wires, abutments or implants can result in surface cracks, which are a result of substrate curvature and roughness, internal and mechanical stresses in the anodic film, weak spots and collapse of nanotubes (Chopra et  al., 2021a; Gulati et  al., 2015). At the same time, these cracks may be reduced by tuning anodization voltage/time, water content and electrolyte aging, the cracking of anodic film on curved substrates must be tested for mechanical stability. To achieve this, TNTs on Ti wires with micro-scale cracks were inserted into bovine trabecular bone cores ex vivo using Zetos™ for 5 days, followed by retrieval and surface morphology analysis via SEM that confirmed the stability of the TNTs (Gulati et  al., 2016a). Additionally, ex vivo 3D cell culture models have also been used to investigate bioactivity and local drug release

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functions of drug/protein loaded TNTs (with micro-cracks on anodic film) on Ti wires, and the examination of TNTs confirmed that their surface integrity was maintained (Gulati et al., 2016a; Kaur et al., 2016).

4.2  In Vivo Implantation In one of the pioneering attempts relating to in vivo placement of TNTs, von Wilmowsky et al. inserted TNTs on Ti rods into the front skull of domestic pigs in vivo for up to 90  days and found that retrieved implants had intact TNTs (von Wilmowsky et  al., 2009). Next, Bjursten et  al. performed the pull-out testing of heat-treated TNTs on Ti implants via placement in rabbit tibia in vivo (Bjursten et al., 2010). The findings confirmed that a fracture force of up to 10.8 N was insufficient to compromise adhesion strength between TNTs and Ti. Advancing the domain and testing the anodized Ti implants in vivo, Choi et al. implanted bare and anodized orthodontic miniscrews in the mandible of beagle dogs in vivo for 12 weeks (Choi et al., 2012). Upon retrieval, SEM confirmed that the thread edge (close to tip) of the anodized screw was smoothened via smearing, and the thread edge (close to the top) of the machined implant became rough, as compared to the unused respective implants. This observation was attributed to the stress concentration in the tip area that made it prone to mechanical damage. Next, AFM measurements of the implants were performed, and the roughness of the anodized screws was significantly reduced compared to the unused anodized screws. It is noteworthy that while the roughness of the anodized screw was reduced due to the implantation procedure, their values were higher compared to non-anodized bare screws. Overall, the damage to the anodic film was attributed to the insertion shearing force or orthodontic tension, which are important parameters to consider while placing anodized nano-engineered implants. Although not directly applicable to a dental implant setting, the mechanical stability of TNTs modified Ti wires have also been confirmed in a mice tumour model in vivo (Kaur et al., 2016). Briefly, anti-cancer drug-loaded TNTs were implanted into mice tumour sites and retrieved after 6 days, and SEM imaging confirmed the stability of the structures. The abovementioned ex vivo and in vivo studies demonstrate that TNTs are a favorable bioactive and therapeutic modification for implantable applications and maintain their structural integrity while implanted; however, more dental implant-focused in  vivo investigations are needed, specially under mechanical loading conditions.

5 Future Directions and Conclusions Optimizing electrochemical anodization (EA) has enabled the fabrication of controlled and stable nanostructures, including nanotubes and nanopores on dental implants, to achieve superior bioactivity and mechanical stability. Various

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investigations have evaluated the mechanical characteristics of nano-engineered implants in corrosive environments, sterilization techniques and ex vivo and in vivo implantations. The stability of nano-engineered titanium dental implants is crucial to ensure ease of clinical translation; however, various research gap remains unaddressed, as described below: Dimensions vs Stability The mechanical investigations on anodized implants often reveal that nanotube dimensions dictate stability. For physically/chemically enhanced nanotubes, dimensions decide the mechanical performance and need for further modification. At the same time, it is well established that wettability, protein adhesion capacity, cell adhesion/proliferation, drug loading capacity and drug release kinetics are all influenced by nanotube dimensions. This means one or more functions must be sacrificed to accommodate appropriate mechanical performance. For example, very high drug loading amounts and prolonged release require longer and wider tubes/pores, which may not be as robust as a thin anodic film with smaller/shorter tubes. Cytotoxicity While titania nanoparticle cytotoxicity has been investigated, the findings cannot be adequately translated for anodized implants with nanotubes and nanopores. This is because anodic film is attached to the underlying substrate (implant), and if any mechanical failure happens, the anodic film may break in the form of aggregations of nanotubes or loose nanotubes. These will behave significantly differently compared to nanoparticles’ geometry and size. Further, the anodic film contains organic electrolyte and fluoride ions, and only limited studies have explored its complete removal. Hence, a thorough cytotoxic evaluation of nanotubular or nanoporous implants is needed. Stability Enhancements Alters Topography Various physical (crystallization, deposition of secondary material and heat-­ treatment) and chemical (carbonization and F sedimentation) strategies have shown promising outcomes towards achieving augmented mechanical performance of anodized Ti with TNTs or TNPs. However, such modifications can change surface chemistry, topography or the dimensions of the nanostructures. These alterations can further influence implant bioactivity and the ability to load and release therapeutics. Long-Term In Vivo Testing Anodization performed on clinical implants has been placed inside bones ex vivo, and in vivo. However, these investigations only partially achieve the dental implant-­ relevant physiological conditions. For instance, long-term in vivo investigations in the oral cavity with mechanical loading is needed to ensure the implant experiences physiological conditions with appropriate mechanics, chemical/saliva environment and possible implant placement trauma. Such investigations must be carried out for months to ensure the anodic film survives the implantation.

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Acknowledgements  Divya Chopra is supported by a UQ Graduate School Scholarship (UQGSS) funded by the University of Queensland. Karan Gulati is supported by National Health and Medical Research Council (NHMRC) Early Career Fellowship (APP1140699).

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Cytotoxicity, Corrosion and Electrochemical Stability of Titanium Dental Implants Tianqi Guo, Jean-Claude Scimeca, Sašo Ivanovski, Elise Verron, and Karan Gulati

Abbreviations ASTM American Society for Testing and Materials BL Barrier layer BMSCs Bone marrow mesenchymal stem cells COF Coefficient of friction EA Electrochemical anodisation EDXS Energy dispersive X-ray spectroscopy EIS Electrochemical impedance spectroscopy FBGC Foreign body giant cell GNPs Graphene nanoplatelets HA Hydroxyapatite ICPMS Inductively coupled plasma mass spectroscopy MAO Micro-arc oxidisation NPs Nanoparticles OCP Open circuit potential PIII Plasma immersion ion implantation PVD Physical vapour deposition PVP Polyvinylpyrrolidone ROS Reactive oxygen species T. Guo · S. Ivanovski (*) · K. Gulati (*) School of Dentistry, The University of Queensland, Herston, QLD, Australia e-mail: [email protected]; [email protected] J.-C. Scimeca Université Côte d’Azur, CNRS, Inserm, iBV, Nice, France E. Verron (*) Nantes Université, CNRS, CEISAM, UMR 6230, 44000, Nantes, France e-mail: [email protected] © The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 K. Gulati (ed.), Surface Modification of Titanium Dental Implants, https://doi.org/10.1007/978-3-031-21565-0_8

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SBF SLA SLM SMAT TiN TNPs TNTs UMCA

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Simulated body fluids Sandblasted and acid-etched Selective laser melting Surface mechanical attrition treatment Titanium nitride Titania nanopores Titania nanotubes Ultrasonic mechanical coating and armouring

1 Corrosion of Ti Implants According to the American Society for Testing and Materials (ASTM), five grades of Ti are used for implant biomaterials. Grades I–IV differ according to the purity grade, and the amount of various interstitial elements (carbon, oxygen, nitrogen, hydrogen, and iron), while Grade V refers to the Ti alloy Ti-6Al-4V, the most commonly used in the medical implant industry. Regardless of the grade, Ti dental implants have been widely used due to their favorable biocompatibility and mechanical properties such as high fatigue strength (140–1160 MPa) and fracture toughness (Guo et al., 2021a). It is estimated that around 5 million implants are placed in the USA per year, and 15–20 million worldwide (Pettersson et al., 2019). Despite this huge success, increasing reports of implant instability related to peri-­implantitis, a multifactorial disease caused by several factors impacting implant/bone tissue interactions, is a growing concern (Guo et al., 2021b). Dental implant surface characteristics, its composition and mechanical properties and especially the Ti oxide layer, are critical for osteointegration. A variety of nanomaterials (NM) can be used for the surface treatment of Ti-based dental implants (Zhang et  al., 2021). For example, NM of titanium nitride (TiN) has been shown to improve the chemical and wear resistance of Ti implants (Xuereb et al., 2015). Unfortunately, changes in chemical and topographic structures occur over time in response to mechanical stimuli and oral cavity environment, leading to the release of particles and ions from the coating layer or from the implant itself. All these changes, mostly irreversible, dramatically compromise implant osteointegration.

1.1 Reasons for Ti Implant Degradation Identifying the underlying mechanisms governing implant alterations could help (i) engineers to improve the design of implants, and (ii) clinicians to adapt their practices for optimizing long-term survival of implants. To this end, Romanos et al. have analyzed the processes leading to the release and distribution of Ti particles and ions

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within peri-implant hard and soft tissues (Romanos et al., 2021). They have highlighted that implant wear mechanisms occur at different stages including implant bed preparation, insertion step, and implant decontamination/maintenance. 1.1.1 Mechanical Corrosion Mechanical corrosion first appears during the implant bed preparation, which requires bone-cutting instruments that cause the release of Ti particles due to metal attrition, wear, and corrosion. Indeed, Ti particles and ions have been observed in irrigation liquid collected from the implant bed preparation by piezosurgery or drill procedures (Rashad et al., 2013). In fact, drills induce abrasive wear, coating damage and blunting (especially at the tip and flanks), and plastic deformation of the cutting point (Mishra & Chowdhary, 2014). In addition, substance loss and the condensation of particles detached from the tool is correlated to the number of drills used, suggesting that single-use drill may be an optimum choice. Moreover, sterilization of the cutting tools can generate particles by initiating corrosion (Allsobrook et al., 2011). Consequently, the number of sterilization procedures should be controlled, although there is no consensus on this point. During the insertion procedure, friction between the implant and bone tissue generate Ti particles of various size concomitantly with implant insertion in the mandible and maxilla, as observed by Romanos et  al. in periprosthetic tissues (Romanos et al., 2014). Abundant irrigation may prevent this accumulation of remnant particles within the surgery site. Implant-abutment fit strongly impacts on implant longevity, and as described by Stimmelmayr et al. (2012), material characteristics influence wear of the implant-­ abutment connection. Indeed, the association of an implant with an abutment generates the release of particles. These particles can remain inside the connection area, resulting in frictional wear, or can migrate to adjacent tissues and contribute to a foreign body reaction (Delgado-Ruiz & Romanos, 2018). The gap (i.e., the mismatch) between abutment and implant can grow in the presence of micromotion resulting from functional loading. During mastication, a larger gap favors micromovements and fretting at the interface which amplifies implant destabilization in a vicious cycle (Schwarz, 2000). Moreover, high loading forces observed during osteotomy and manual bone condensation promote microcracks and particles release at the implant-bone interface (Romanos, 2015). Regardless of their origin, micromovement of the abutment is deleterious through inducing Ti particles release and further metal corrosion (i.e. fretting). Notably, the gap can be colonized by microorganisms, glycoproteins, and fluids that can easily accumulate and form a stable biofilm responsible for implants microbiological corrosion (Apaza-Bedoya et al., 2017), although, as compared to butt-joint implant-abutment connections, conical implant-abutment connections have been shown to minimize the micro gap at the connection, and to reduce bacterial accumulation (Zipprich et al., 2018).

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1.1.2 Chemical and Electrochemical Degradations In addition to production of metal particles through mechanical degradation, the implant surface can undergo chemical and electrochemical degradations resulting in the release of soluble ions. A stable Ti oxide layer (1.5–10 nm thickness) due to the high affinity of Ti for oxygen is spontaneously formed at the surface of implant. This thin film provides sufficient protection of the bulk material from corrosion (i.e., electrochemical phenomenon) and the formation of reactive oxygen species (ROS) (Abey et  al., 2014). Although this film can be reformed through a re-­ passivation phenomena, it is exposed to functional stimuli (i.e., micromotions, micromovement of the implant-abutment interface) and various environmental conditions such as acidic pH and electrolytes that can degrade this protective layer over time and expose the bulk material to the oral environment (Delgado-Ruiz & Romanos, 2018). A wet corrosion of dental implants is predominantly observed given that oral cavity is a humid environment. Depending on its composition, pH, buffering capacity and surface tension, saliva can play the role of an electrolyte and contribute to the dissolution of the oxide layer. In addition, acidic metabolites (citric acid and lactic acid) are produced by biofilm microorganisms such as Streptococcus mutans and Candida albicans (Noronha Oliveira et al., 2018). The resulting low pH can block the re-passivation phenomena and favor corrosion of the thin oxide layer and implant instability/failure (Siddiqui et al., 2019; Souza et al., 2015). Moreover, very low levels (or absence) of oxygen below the gingival margin around dental implants stimulate the proliferation of anaerobic microbes (Porphyromonas gingivalis, Actinomyces) generating lactic acid (Noronha Oliveira et al., 2018). 1.1.3 Tribocorrosion Oral environment and mechanical constraints (loading  or velocity) impact dental implant integrity. Considering the different mechanisms involved in Ti implants degradation, the term “tribocorrosion” has been defined as a combination of tribological (i.e., wear and fretting) and corrosive (i.e., chemical or electrochemical reactions) phenomena (Apaza-Bedoya et  al., 2017; Revathi et  al., 2017). Among different types of corrosion characterized by Noumbissi et al., crevice, galvanic and pitting corrosions are mostly observed in patients (Noumbissi et  al., 2019) (see Table 1). Implant maintenance (cleaning, disinfection) constitutes a risk of tribocorrosion. Despite this risk, these procedures are essential to prevent the formation of biofilm involved in the pathogenesis of dental caries, gingivitis and peri-implantitis. Saliva contains multiple proteins and electrolytes that can adsorb onto the implant surface. The resulting organic layer favors adhesion of cells and bacteria through specific membrane protein- or glucan-binding sites (Busscher et al., 2010; Song et al., 2015). These interactions are also influenced by implant surface characteristics such as surface tension (Banas & Vickerman, 2003). Various bacteria populations progressively accumulate and grow, and the composition and the thickness of the resulting

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Table 1  The three main forms of corrosion observed in patients with dental implants Crevice Corrosion – Occurs within constricted spaces where there is no exchange of oxygen (e.g., at the implant-abutment interface) and is favored by an acidic environment due to chloride ion concentration increase. Galvanic Corrosion – Takes place when dissimilar alloys are in direct contact within the oral cavity. In such cases, the implant plays the role of an anode, and metal ions are released because of the resulting galvanic activity (i.e., ion exchange between implant and its prosthetic components). Pitting Corrosion – A localized form of corrosion arising on openly exposed metal surfaces in the absence of any apparent crevices. This occurs usually along with fluoride based solutions used during dental procedures and daily care.

biofilm can evolve in response to the continuous introduction and removal of microorganisms and nutrients within the oral cavity (Chin et al., 2006). In addition, hard structures such as teeth and implants provide non-shedding surfaces favorable to the formation of a stable biofilm, whereas soft tissues (including oral epithelia) compromise this stability due to rapid tissue turnover (Sanz-Martín et  al., 2018). Excessive accumulation of bacteria produces favorable conditions for the development of dental caries, gingivitis and peri-implantitis. Hence, to prevent these deleterious events and despite the risk of damage, decontamination of the implant surface is required. While no consensus has been reached in terms of the method, soft procedures are strongly recommended (Khan & Sharma, 2020; Sato et  al., 2021; Schwarz et al., 2011). Delgado-Ruiz and Romanos extensively analyzed the causes of particle and ion release during the maintenance phase (Delgado-Ruiz & Romanos, 2018). Among the different procedures, they mentioned that chemical decontamination methods can damage the Ti layer and induce corrosion because of the pH. This release can be amplified by friction leaving the bulk surface of implant exposed. Minimal degradations have been observed with saline, chlorhexidine, hydrogen peroxide and citric acid, while there is no information regarding the impact of tetracycline and doxycycline use (Noronha Oliveira et  al., 2018; Souza et  al., 2019). Mechanical methods using lasers can be used in combination with chemical methods to improve the efficiency of decontamination. However, surface alterations following laser treatment have been reported, depending on the delivered energy, the irradiation duration, and surface characteristics of the treated area (i.e., roughness) (Vayssette et al., 2018). Upon irradiation, temperature increase at the irradiated spot can induce surface modification. Thus, it is recommended to use pulsed modes to reduce the intensity/duration of irradiations, and to cool the area using devices implementing air-water flux delivery with proper ratios (Giannelli et al., 2015).

1.2 Factors Influencing Ti Corrosion/Degradation As described above, tribocorrosion can result in release of metal particles and ions results. The amount of released Ti is dependent on many factors, including the type of implant. Joseph et  al. (2009) evaluated the tribocorrosion of Ti implants in

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simulated biological fluids in vitro, and showed the rate was 23 to 448 μg/L/week for commercially pure Ti. The release rate of soluble Ti from Ti6Al7Nb alloy was lower and ranged from 17 to 50 μg/L/week. These amounts seem to be relatively small, but considering that implants are present for many years, Ti can gradually accumulate in the body (Golasik et al., 2016). Tribocorrosion depends also on the oral cavity environment, which it is known to fluctuate in terms of pH, salt concentrations, oral flora, temperature, oxygen content, food, beverages, tobacco consumption, and toothpaste use. All conditions that can lower the pH under 6 favor the corrosion process. These constant aggressions lead to the degradation of the Ti oxide thin layer, and the removal of metallic debris and ions from the surface (Mouhyi et al., 2012; Peixoto & Almas, 2016). Most toothpastes or gels contain fluoride at concentrations 0.1–2 wt% that can induce the dissolution of the thin oxide layer (Perinetti et al., 2012; Schiff et al., 2002). Addressing this issue, Kaneko et al. characterized compounds formed at the implant surface and documented the presence of Ti fluoride, Ti oxide fluoride and sodium Ti fluoride (Kaneko et al., 2003). By replacing the original Ti oxide layer, these Ti/fluoride compounds form a more soluble layer which undergoes chemical events leading to accelerated dissolution. Lastly, inflammatory episodes occurring within surrounding tissues can also alter the corrosion resistance of implants. Notably, the presence of bacteria-derived lipopolysaccharides (LPS) exacerbates peri-implant tissues inflammation (Yu et al., 2015).

1.3 Importance of Augmenting Anti-corrosion Capacity of Ti Implants (Cytotoxicity Concerns) 1.3.1 Molecular Interactions As previously mentioned, Ti displays a high reactivity when exposed to oxygen. After conversion into Ti dioxide/titania (TiO2), the resulting layer may interact and bind calcium, hydroxyapatite and organic compounds including serum proteins. This binding process leads to the formation of a dynamic protein corona, and its composition depends not only on the environment at a given time (i.e., nature and concentration of proteins present in the environment) but also on material characteristics (shape, structure, size) (Gulati et al., 2021). The presence of the corona enables the cellular internalization of TiO2 particles through a “Trojan horse”-like mechanism. Once inside the cytoplasm, particles can interact with various proteins, lipids and genetic materials. These interactions may alter cell homeostasis, create intracellular lesions and result in toxic effects (Mano et al., 2013; Ribeiro et al., 2016). Particles may be exposed to an acidic environment mediated by endosomes (pH 6.0) or lysosomes (pH 4.5), thus generating free metal ions. These metal ions can

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potentially affect cellular homeostasis and functions by influencing gene/protein expression (Sabella et al., 2014). Moreover, by disrupting electron transport in the mitochondria inner membrane, they may interfere with energy production and generate endogenous ROS. Low levels of ROS can activate transcription factors, leading to enhanced expression of inflammation-associated genes. Furthermore, a high and sustained production of ROS that are not detoxified by endogenous antioxidant defenses affects cellular membranes (lipid peroxidation, structural proteins alterations) as well as enzymes, ion pumps, or nuclear DNA (Bressan et al., 2019). Ti (ions, NPs, oxide) and ROS may diffuse via nuclear pores and translocate into the nucleus. TiO2 can act as a DNA intercalator and generate severe DNA damage (Pogribna et al., 2020). Additionally, interactions between nano-scale Ti particles and plasmatic/cellular proteins affect protein structure and/or functions. Interactions with the cytoskeleton may also induce conformational changes in tubulin and disrupt its polymerization, as shown with TiO2 NPs that considerably alter intracellular transport, cell division and cell migration (Hou et al., 2019). 1.3.2 Cellular Interactions Immune Cells Ti particles generate a pro-inflammatory response via interaction with immune cells (e.g., macrophages and T lymphocytes) and stimulating the release of multiple inflammatory cytokines such as IL-1β, IL-6, TNF-α and prostaglandins (Noronha Oliveira et al., 2018). Among the interactions with inflammatory cytokines, TiO2-NP seems to preferentially adsorb CXCL8 and INF-α, thus resulting in the disruption of neutrophil chemotaxis. This adsorption impacts local concentration of inflammatory mediators that can hamper physiological inflammatory responses (Batt et al., 2018). Pettersson et  al. also described a pro-inflammatory response in macrophages (Pettersson et  al., 2017). Nevertheless, they demonstrated that Ti particles alone displayed a limited impact on cytokines secretion, strongly increased only in the presence of bacteria. This suggests that Ti particles act as a secondary stimulus enhancing the production by macrophages of cytokines such as IL-1β. As expected, the presence of Ti nanoparticles (NPs) exacerbates this feature as compared to Ti microparticles (MPs). Indeed, compared to MPs, NPs offer a greater potential of reactivity and interaction with various biological macromolecules and cells (Gulati et al., 2021). However, it is worthwhile to note that NPs can agglomerate and increase their size, thus lowering their reactive surface area. Accordingly, NPs agglomeration may reduce the severity of the NP toxic effect, as explored by Bruno et al. (2014).

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Bone Cells Osteointegration of implants is crucial and strongly dependent on the activity of osteoblasts and bone marrow mesenchymal stem cells (BMSCs) (Takeda et  al., 2012). Unfortunately, Ti particles can disturb the cytoskeleton of BMSCs, thus affecting their migration and differentiation capabilities. These particles can also display cytotoxicity toward BMSCs (Meng et al., 2009), and change in osteoblastic viability has been demonstrated at high concentration of Ti particles (0.15–1%) (Pioletti et al., 1999). Similarly, Mine et al. demonstrated that the viability of murine osteoblastic cells (MC3T3-E1) was lowered by high Ti concentrations (e.g. 20 mg/L), while concentrations in the range of 1–9 mg/L were ineffective in altering cell viability (Mine et al., 2010). Saldaña et al. reported that osteoblast proliferation was reduced by more than 50% when cells were exposed to 48 mg/L of Ti during 3 days, or to 24 mg/L of Ti for 7 days (Saldaña et al., 2006). Further, the metabolic activity of osteoblasts was reduced by 20% for a 7 day exposure to 24 mg/L of Ti, and to very low levels when a 48 mg/L concentration was used. Other studies have shown that Ti in the form of powder is cytotoxic towards human osteoblast-like cells (SAOS-2) in a time and dose-dependent manner (at concentrations above 15.5 μg/L), while Ti in bulk form (disc made from metal powder) displayed no cytotoxicity (Li et al., 2010). Viability of MC3T3 cell line was less than 10% after 5–7 days incubation with Ti particles (5 and 10 g/L), with caspase-3 activity and apoptosis increased after 48 h of incubation with Ti particles at 2.5 g/L (Qiu et al., 2015). Ti particles can inhibit osteoblastic cell viability through the stimulation of excessive interleukin secretion (Il-6 and Il-8) by abnormally recruited neutrophils (Chen et al., 2020; Happe et al., 2019; Schulze et al., 2013). Expression of osteoblastic proteins such as matrix metalloproteinase-2 (MMP-2), membrane type 1 MMP (MT1-MMP), and p38 protein is also upregulated in the presence of Ti particles (0.1 g/L). Consequently, Ti particles are involved in periprosthetic osteolysis (Chen et al., 2014). Similarly, Ti particles inhibit the activity of periodontal ligament cells, and the osteogenic differentiation of alveolar bone cells (Zhou et al., 2021). Regarding bone resorbing cells, Ti particles tend to enhance the differentiation and resorption activity of osteoclasts by promoting the release of RANKL and pro-­ inflammatory cytokines (Il-1 and TNFβ) by macrophages and lymphocytes (Wachi et al., 2015). Osteoclasts are involved in the process of Ti surface corrosion, and a subsequent accumulation of metal ions in their cytoplasm and nuclear heterochromatin has been observed. Cadosch et  al. showed that in over 20% of 22 healthy donors, monocytes differentiated into mature osteoclasts after 10 days’ exposure to Ti at a concentration of 48  μg/L (Cadosch et  al., 2010). They also demonstrated increased expression of specific osteoclastic biomarkers such as cathepsin K and TRAP in monocytes after 10 days incubation with Ti at 48 μg/L, a level measured in blood and tissue from patients with loosened implants. A great concentration of mature osteoclasts surrounding Ti implants leads to a large resorption area.

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Consequently, the osteolysis generated around implants progressively contributes to the aseptic loosening (Zhou et  al., 2021), and is exacerbated by an alteration of osteoblastic cell viability (Park et al., 2013). Other Cells Ti particles also interact with peri-implant tissues including gingival epithelium. High concentrations of Ti (>13 ppm) can alter gingival epithelial cell viability, and induce cell necrosis (Makihira et al., 2010). This damage can be amplified when Ti particles surface contains phosphate-enriched TiO2 or fluoride-modified species (Wei et  al., 2005). Furthermore, low particle concentrations (0.001%) reduce the proteolytic and collagenolytic activities of fibroblasts, while their proliferation remains unaffected. By contrast, higher contents of particles are toxic for fibroblasts in terms of both viability and proliferation (Kheder et al., 2021). These deleterious effects impact the epithelial barrier integrity, which in turn favors bacterial colonization (Dini et al., 2020). It has been observed that Ti can stimulate the expression of Toll-like receptor-4 by fibroblasts, and also that the sensitivity of these cells to oral bacteria is modified (Wachi et al., 2015). With oral infection, Ti particles exacerbate the inflammatory response caused by oral biofilm (Berryman et al., 2020) resulting in chronic inflammation of peri-implant tissues (Wilson Jr. et al., 2015). In summary, inflammation of peri-implant tissues associated with an excessive osteoclastic activity contribute strongly to implants loss, and oral infections worsen this situation. As identified in chapter “Titanium Dental Implants in Compromised Conditions: Need for Enhanced Bioactivity and Therapy”, several risk factors such as smoking, premature loading and diabetes have been associated with weak osseointegration or excessive bone resorption (Berryman et al., 2020; Dini et al., 2020; Ferreira et al., 2018; Lee et al., 2017; Mombelli et al., 2018; Stacchi et al., 2016; Wilson Jr. et al., 2015). 1.3.3 Tissue Interactions Dental implants are considered a source of Ti particles in both intra-oral and extra-­ oral tissues. In fact, Ti particles have been observed in tissues surrounding Ti implants such as the submucosal plaque, peri-implant soft tissues and bone, but also in distant locations such as lymph nodes (Frisken et al., 2002). Various investigations have evaluated Ti levels in the peri-implant mucosa in patients with pure or alloy Ti implants (Frisken et  al., 2002; He et  al., 2016; Mercan et  al., 2013). In comparison with samples collected from patients with stainless steel brackets and tubes, Ti level in oral mucosa cells increased by a factor of 3 after a 30 days orthodontic treatment. Similarly, when compared with patients with healthy peri-­ implant tissues, submucosal biofilm of implants with peri-implantitis displayed high concentrations of dissolved Ti (He et al., 2016). Consequently, serious concerns regarding the long term safety of Ti implants have been raised.

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Mercan et al. investigated Ti release from cover screws of endosteal implants to adjacent gingival tissues (Mercan et  al., 2013). A 1.4-fold increase in gingiva Ti content was found after 3 months. It was also observed that metallic debris could migrate throughout the different layers of peri-implant soft tissues, and then be involved in inflammatory responses (Paknejad et al., 2015). Moreover, the inflammatory status was correlated with particle size. Histological analysis of tissues adjacent to dental implants harvested from patients (e.g., oral mucosa in contact with Ti implant, mucosa of overlying Ti cover screws during submerged healing of 2-piece implants, gingiva) revealed peri-­ implantitis lesions as well as mucosa displaying various grades of inflammation (Mombelli et  al., 2018). Ti particles have been detected in tissue, and a gradual decrease in their density observed between connective tissue and epithelium 6 months-post Ti implant placement (Flatebø et al., 2006). Further, an inflammatory infiltrate was evidenced within the connective tissue facing the cover screw, and it was shown that the inflammatory infiltrate/fibroblast ratio decreased with time. Based on energy dispersive X-ray spectroscopy (EDXS) and inductively coupled plasma mass spectroscopy (ICPMS) analysis, Ti particles have been localized and quantified in tissues surrounding dental implants in patients (Table 2). For example, Ti particles phagocytosed by macrophages in soft tissue has been reported and tissue inflammation intensity was generally correlated with the presence of the particles (Olmedo et al., 2013). Paknejad et al. noted that particles were not exclusively observed close to tissues adjacent to Ti implants (Paknejad et  al., 2015), and it seemed that keratinocytes might transport particles to more superficial tissue layers.

Table 2  Distribution and quantification of Ti particles in tissues surrounding dental implants in human Type of the biopsy analyzed Diseased implants (N = 20) Control: healthy implants (N = 20) Diseased implants (N = 12) Control: ceramic implants (N = 1) Diseased implants (N = 7) Control: healthy jawbones without implants (N = 6) Diseased implants (N = 15) Control: healthy implants (N = 15)

Distribution of particles Submucosal peri-­ implant plaque

Quantification of Ti level 0.07 ± 0.2 ng/μL

Ref Safioti et al. (2017)

Peri-implant tissue – mainly in soft tissue

High concentrations (7.53 × 10 5 count)

Fretwurst et al. (2016)

Upper half of bone crest Bone marrow (60–700 nm from the implant surface) Peri-implant tissue

1940 ± 469 μL/kg bone weight Presence confirmed (but not quantified)

He et al. (2016)

Based on Suárez-López del Amo et al. (2018)

2–2.44 ppb Olmedo Presence of particles (individual et al. (2013) or clusters) in epithelial cells and macrophages

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Due to the ability of TiO2 to bind biomacromolecules such as proteins, new antigenic determinants can be exposed and lead to hypersensitivity reactions in vulnerable patients (Gulati et al., 2021). Various signs and symptoms of Ti allergy have been described including swelling, redness, rash, vesicular lesions of the skin, and facial eczema reactions (Fage et  al., 2016). Fortunately, removing Ti implants relieves these clinical symptoms, thus supporting the causal relationship between tissue exposure to Ti and the incidence of these reactions. In addition to temporal association, this causal relationship is strongly suggested in several cases by the proximity between implants and oral lesions (Mombelli et al., 2018). For example, one patient developed facial eczema following the placement of a Ti implant for a mandibular overdenture (Egusa et al., 2008), and reactive lesions of the peri-implant mucosa (i.e., pyogenic granuloma and peripheral giant cell granuloma) have been observed (Olmedo et al., 2010). During a clinical study, a patient displayed a chronic inflammatory response (with concomitant fibrosis around all implants) associated with a foreign body giant cell (FBGC) reaction (du Preez et al., 2007). In addition, persistent proliferation of the peri-implant soft tissue has been described after a mandibular vestibuloplasty and the placement of a split-thickness skin graft (Mitchell et  al., 1990). In other cases, patch or blood tests performed in patients with Ti implants evidenced an allergy/hypersensitivity to Ti (Müller & Valentine-­ Thon, 2006; Sicilia et al., 2008). To complete the analysis of hypersensitivity reactions to Ti, Mombelli et al. compared alleged cases of Ti hypersensitivity in dental and non-dental implants, and reported few differences (Mombelli et al., 2018). For example, all dental implants were contaminated by bacteria during surgery, whereas it was not the case for indwelling devices. Moreover, clinical symptoms in the non-dental field are not restricted to tissues surrounding the implant. By contrast, clinical signs are predominantly observed in tissues in direct contact with dental implants. However, given the few reported cases in the literature, there are still controversies around the existence of Ti allergy/hypersensitivity in patients receiving Ti implants. Finally, a “yellow nail syndrome” has been described in patients with Ti implants (Kim et  al., 2019). This syndrome results in a high content of Ti in the nails of patients. It is characterized by a change in nails, postnasal drip, cough-associated sinusitis, and bronchial obstruction (Berglund & Carlmark, 2011). Lymphedema can also be associated in chronic pathologies. Pleural effusion was the most common lung change, and chronic sinusitis with an early onset was reported to occur. Several studies have established a correlation between Ti and this yellow nail syndrome (Cheslock & Harrington, 2022; D’Alessandro et  al., 2001; David-Vaudey et al., 2004; Decker et al., 2015). Ti ions could be released by galvanic corrosion due to electro-chemical coupling between Ti implants and gold elements present in the teeth. Indeed, in few patients, disease symptoms disappeared after several months when gold elements were removed. Pathogenesis is still under investigation.

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2 Physical Modification and Utilizing Ti Alloys Chemical and corrosion resistance are important criteria for the success of Ti dental implants, as they regularly interact with fluids having various pH and temperature. Suboptimal corrosion resistance may leach Ti ions and cause toxicity, while the corrosion can challenge mechanical stability. To augment the corrosion resistance of Ti-based dental implants, as described below, physical modifications with altered topography and alloying Ti have been explored.

2.1 SLA Implants and Corrosion Sandblasting involves the collision of compressed air-driven alumina (Al2O3) or titania (TiO2) particles on Ti implants, enabling the fabrication of micro-roughened Ti implants and augmented osseointegration (Burnat et al., 2013; Jiang et al., 2006; Li et al., 2001; Wang et al., 2017). Sandblasted implants possess varied corrosion resistance as compared with bare Ti. Jiang et al. reported blasting of Ti implants with 200 ~ 300 μm diameter TiO2 particles to achieve a surface with an average roughness of 1.1 μm (Jiang et al., 2006). Regardless of the microscale cracks by blasting, blasted Ti achieved lower corrosion current density (Icorr) within 3.5% NaCl solution compared to non-modified counterparts, indicating a decreased passive corrosion rate (Jiang et al., 2006). Such enhanced anti-corrosion capacity could be attributed to the newly generated layer with microscale pits and peaks with thickened oxide layer, which reduces chemical reactivity (Jiang et al., 2006). Interestingly, alternate studies revealed increased corrosion of Ti implants post-sandblasting. Burnat et al. blasted Ti6Al4V and Ti6Al7Nb alloys with 110 μm diameter Al2O3 grains at a pressure of 4.5 bar at an angle of 45°, resulting in microscale pits and holes (Burnat et al., 2013). Blasted alloy surfaces showed increased corrosion rates within PBS, attributed to the enlarged surface area and the embedded Al2O3 grits (Burnat et al., 2013). Similarly, Li et al. reported increased Ti ion release (soaking in simulated body fluids (SBF) for 3  months) from the Al2O3 blasted Ti (0.65 ± 0.014 μg/mL) compared to non-blasted polished Ti (0.23 ± 0.020 μg/mL) (Li et al., 2001). Also reported were the significantly reduced open circuit potential (OCP) on blasted substrates, indicating their aggravated passive corrosion (Wang et al., 2017). It is noteworthy that the sharp edges and contours on blasted Ti implant surfaces can increase the tribocorrosion and the contact area, making them prone to corrosion. Alternatively, some studies utilised the acid etching of blasted surfaces to reduce surface protrusions to augment corrosion resistance of blasted implants (sandblasting and acid-etching, SLA). Li et al. reported utilising 40% nitric acid for 30 min on sandblasted Ti substrates to significantly reduce the Ti ion release within SBF (Li et al., 2001). Acid etching reduced the sharp protrusions from the blasted surface and dissolved the embedded sand particles. A similar conclusion was drawn by

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Ogawa et al., demonstrating that HCl treatment significantly improved the corrosion resistance of sandblasted implants by showing a marked decreased corrosion current (Icorr) value (Ogawa et al., 2016). Interestingly, some studies reported that the chemical resistance of Ti implants was not significantly improved after SLA treatment, as Xu et al. indicated, the Icorr values of SLA implants were comparable to smooth-Ti within SBF solution (Xu et al., 2020). Further, it is noteworthy that SLA-­ treated implants with micro-rough surface could inevitably increase bacteria aggregation and proliferation, among which the S. mutans (generates lactic acids) and P. gingivalis can compromise TiO2 layer leading to chemical corrosions (Guo et al., 2021b).

2.2 Cryogenic Treatment Cryogenic treatment involves freezing Ti at approximately −185  °C with liquid nitrogen to significantly reduce the residual stress within the implant structures to improve resistance against wear. As reported by Bhaskar et al., cryogenic treatment could refine the grain size of Ti implants significantly (from around 5 μm to 17 nm) to reduces its dislocation density (number of dislocations within a specific volume of a crystalline material) and augment wear resistance (Bhaskar et  al., 2014). Similarly, it was shown that the grain refinement by cryogenic treatment reduced the composition of β-phased Ti, contributing to enhanced wear resistance (Revathi et al., 2017). The scalability and cost-effectiveness of cryogenic treatment has seen its application to treat and increase mechanical strength in other prosthodontic materials such as cobalt-chromium. Compared with non-treated or annealed Ti, the significantly improved corrosion resistance of cryogenic treated Ti is attributed to the thickened oxide film on the treated surface (Bhaskar et al., 2014). This is supported by Zhu et al., who observed a significantly reduced passive corrosion current value for cryogenic Ti (0.536 μA/ cm2) compared to non-treated pure Ti (0.917 μA/cm2) (Zhu et al., 2014). Additionally, increasing the cryogenic treatment time could increase corrosion resistance. Compared with the untreated Ti alloys with a corrosion current density (Icorr) of 153.1  nA/cm2, reductions in values up to 86.3  nA/cm2 and 43.3  nA/cm2 were obtained after immersion in liquid nitrogen for 24 and 48 h, respectively (Fig. 1) (Gu et al., 2018). Such enhancements are attributed to the resistance value changes of the outer porous layer (Rp) on Ti surfaces (Gu et  al., 2018). Compared with implants at room temperature, the cryogenic-treated implants showed significantly reduced surface roughness due to the dislocation of grain size and increased corrosion resistance (Tang et  al., 2017). Further, the cryogenic treatment divides and separates the grains on Ti, releasing internal strains and yielding a submicron-scale crystalline structure, thereby enhancing the mechanical resistance of the implant (Tang et  al., 2017). The refined grain size could reduce the Gibbs free energy between the boundary of grains (anodic area) and their interior region (cathodic area), which in turn increases corrosion resistance (Tang et  al., 2017). The rapid

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Fig. 1  Electrochemical potential dynamic polarisation curves demonstrating augmentation of the chemical resistance of Ti implants with cryogenic treatment. Cryogenic treatment of Ti-6Al-4V implants over 24 (DCT-24) and 48 (DCT-48) hours exhibited lower corrosion current values than untreated implants (UT) within Hanks solution. (Reproduced with permission from Gu et al. (2018))

establishment of passive film over the refined nanocrystalline surface with more boundaries further increases its corrosion resistance (Tang et al., 2017). In summary, cryogenic treatment remodels the crystalline structure of Ti implants by separating grains and refining their size. The grain refining process significantly reduces internal strain to improve chemical corrosion resistance and wear of Ti implants.

2.3 Alloying of Ti Apart from pure titanium (cp-Ti), Ti alloys such as Ti-Zr and Ti-Nb are suitable for fabricating dental implants (Chopra et al., 2022). Studies to date have reported that these alloys are substantially more resistant to chemical corrosions than cp-Ti (Akimoto et al., 2018; Han et al., 2014). This may be attributed to the composition of passive oxide films on Ti alloy surfaces, composed of ZrO2, Nb2O5 and NbO2 with TiO2, which are more electrochemically stable and compact than the thin, amorphous TiO2 from cpTi surface (Akimoto et al., 2018; Han et al., 2014). One example of such alloy is the Ti-Zr, with a dual surface oxide layer of TiO2 and ZrO2 that is more chemically stable and resistant to acid challenges than the TiO2 film on pure Ti (Akimoto et al., 2018). The composition of Zr within the Ti-Zr alloys was shown to influence their corrosion resistance (Akimoto et al., 2018). As reported by Akimoto et al., Ti-Zr alloys containing 30–50 wt% Zr showed significantly reduced pitting corrosions in artificial saliva (pH 4.9) than pure Ti

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Fig. 2  Galvanic passivation curves showing TiZr alloy with 5 ~ 15% Zr presents appropriate electrochemical stability. The data indicates that electrons flows from pure Ti (cp-Ti) to Ti-5/10/15Zr (TiZr with 5 ~ 15% Zr), but the direction is reversed from Ti-20Zr (TiZr with 20% Zr) to cpTi. (Reproduced with permission from Han et al. (2014))

counterparts, attributed to the stable Zr-rich protective layer on those Ti-Zr alloy surfaces (Akimoto et al., 2018). Another study investigating the chemical resistance of Ti-Zr revealed that corrosion potential (Ecorr) values increases with the Zr composition (Zr ranged 5 ~ 15 wt%) (Han et al., 2014). The further galvanic passivation experiment revealed that such changes attributed to the electron flow from cp-Ti to the Ti-Zr alloy when Zr composition ranged at 5 ~ 15 wt%, but reversed from Ti-Zr to cp-Ti when Zr composition increased to 20 wt% (Fig. 2) (Han et al., 2014). With constant oral cavity pH changes and fluorine incorporation by toothpaste, the optimised Zr concentration of Ti-Zr implants ranges between 10 ~ 15%, as the clinically utilized Zr concentration within the Straumann® Roxolid implant system (15% Zr). Niobium (Nb) incorporation can also influence the stability of the Ti implants. Ti-Nb alloys are fabricated using an arc melting furnace. Han et al. reported that the corrosion resistance of Ti-Nb alloys increased with an Nb content of 5–10  wt%, although continuously increasing Nb composition to 20 wt% reduced the resistance (Han et al., 2015). This effect could be attributed to the phase of Ti-Nb alloys, maintaining the α phase until 10 wt%, followed by a shift to the less stable β and ω phases with the increase in Nb content (Han et al., 2015). Similarly, Caha et al. reported an enhanced corrosion resistance of Ti with 15 wt% Nb (combination of α + β phase) compared to 40 wt% Nb (mainly β phase) (Çaha et al., 2020). An additional disadvantage for the Ti-40Nb (40 wt% Nb) alloy is its low resistance to the tribocorrosion, with higher weight loss and a higher coefficient of friction (COF) value (Çaha et al., 2020). While Ti-40Nb showed favourable Young’s modulus (51 Gpa) that is closer to natural bone, the optimal Nb content for fabricating Ti-Nb implants should be 10–15 wt%, which yields favourable resistance to chemical corrosion and tribocorrosion, towards long-term stability (Çaha et al., 2020).

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In summary, incorporating metal elements such as Zr or Nb could significantly improve the corrosion resistance of conventional Ti implants. It is noteworthy that the Ti-Zr implants are already commercially utilised.

3 Surface Chemical Modification to Augment Corrosion Resistance It has been reported that the surface TiO2 layer of conventional Ti implants is thin, amorphous and inconsistent, which is insufficient against chemical and electrochemical corrosion. Chemical modification of the TiO2 layer using nitriding, plasma spraying, coating and micro-arc oxidisation can enable the protection of the Ti implants against chemical corrosions.

3.1 Nitriding Nitriding diffuses nitrogen onto the metal surface via a hydrothermal process to fabricate a hardened surface shielding layer. Since the metal-nitrides such as TiN and TiAlN are ceramic coatings with exceptional chemical stability in corrosive medium, nitriding treatment could effectively improve the corrosion resistance of the modified Ti implants (Chung et al., 2004). It has been reported that the TiN layer possess favourable mechanical hardness and chemical stability, attributed to the covalent bonding between Ti and N; hence, TiN-coated Ti has been utilized across various industries, such as anoxic casting metals, precursor for wear-resistant and biomedical implants (Kazemi et  al., 2020). Compared with non-treated Ti, TiN-­ coated implants showed significantly reduced corrosion current within simulated body fluids, indicating their decreased degradation speed within the human body (Fig. 3) (Kazemi et al., 2020). Additionally, traditional physical vapour deposition (PVD) could enable TiN and other ceramic coatings on Ti substrates, although corrosion may occur through defects in the coating layer such as delamination, indicating the need for improvements in layer consistency and mechanical stability. The isothermal exposure technique involves gradually heating Ti in N2 (10−3Pa to 850 °С for 12 h) followed by cooling in N2 (0.028 °С/s till 500°С) in vacuum. This method results in TiN coating with improved stability and adherence, demonstrating augmented corrosion resistance in Ringer’s solution at both 36 and 40 °С (Pohrelyuk et  al., 2013). Alternatively, selective laser melting (SLM) with acid etching (HF  +  HNO3) has been used to fabricate firmly adhered TiN coating on Ti with evenly distributed spherical TiN particles (Zhou et al., 2020). Such TiN composite is composed of a combination of δ-TiN and ε-Ti2N that is strongly incorporated with the underlying hexagonal-Ti, reducing the passivation current density and corrosion current density (Icorr) of Ti in acidic solution (Zhou et al., 2020).

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Fig. 3  Polarisation curves of bare and coated Ti-6Al-4V alloys within the SBF solution. Results confirm reduced corrosion density values on TiN and hydroxyapatite (HA)-coated implants as compared to non-coated counterparts, indicating that both TiN and HA coatings could protect Ti implants against chemical corrosion. (Adapted with permission from Kazemi et al. (2020))

Magnetron sputtering and glow discharge have also been used to perform nitriding of Ti. Ananthakumar et al. deposited a 2 μm thick TiN coating via DC magnetron sputtering that significantly reduced the degradation speed of Ti as shown by the reduced corrosion current density (Icorr) within 3.5% NaCl solution (Ananthakumar et al., 2012). Glow-discharge ion nitriding also resulted in a consistent TiN coating layer on Ti implants, which showed exceptional corrosion resistance in 5  wt% HCl, with the corrosion potentials significantly increased from −150 mV (non-coated Ti alloy) to 300 mV (TiN coated alloy) (Rossi et al., 2003). While TiN coated Ti exhibits exceptional corrosion resistance, the main application focus was fabricating aeroplane/marine equipment and electrochemical cells (e.g. EV batteries), and not dental implants. Thus biocompatibility and bioactivity studies are needed for TiN application in dental implants.

3.2 Coating with Calcium Phosphate (CaP) CaP is a bioceramic with desirable corrosion resistance that could be utilised as a coating to protect Ti implants from the electrochemical corrosion and damage. It has been reported that the CaP coated Ti has significantly reduced corrosion current value and increased polarisation resistance in vitro, indicating corrosion resistance (Qadir et al., 2019). Similarly, Coelho et al. reported that CaP coating on Ti alloys

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Fig. 4  Potentiodynamic polarisation curves for non-modified Ti and hydroxyapatite (HA)-coated Ti implants via ultrasonic mechanical coating and armouring (UMCA). Results show that the corrosion density values (Icorr) of UMCA-Ti was significantly lower than non-modified counterparts, indicating a reduced corrosion speed of HA-coated implant surfaces. (Adapted with permission from Lin et al. (2022))

via sputtering significantly increased the polarisation potential of alloys within a PBS solution, demonstrating its enhanced corrosion protection (Coelho et al., 2009). Hydroxyapatite (HA), one of the most applied forms of CaP, is similar to human bones in both morphology and composition and has been coated on Ti implants to improve osseointegration (Gulati et al. 2022a). Lin et al. fabricated a porous HA coating on Ti by ultrasonic mechanical coating and armouring (UMCA) (Fig.  4) (Lin et al., 2022). Compared with the unmodified Ti, UMCA-modified Ti significantly reduced the corrosion current density (Icorr) within Hank’s solution. This could be attributed to the HA coating layer behaving as a resistant barrier between Ti implants and corrosive solution (Lin et al., 2022). Saeed et al. also utilised microarc oxidisation (MAO) on Ti alloys to fabricate a porous and roughened HA coating that significantly increased the corrosion potential and reduced the corrosion current of Ti alloy, improving corrosion resistance (Saeed et al., 2021). Incorporating the MAO treatment, a composite HA/TiO2 coating could be fabricated on Ti, with dual microscale texture (microgrooves) and superimposed nanorods (Khalid Naji et al., 2021). Such modification not only increased the corrosion resistance of Ti significantly, but also enhanced its hydrophilicity due to the generated nanoscale rods (Khalid Naji et al., 2021). Compared with CaP with similar chemical compositions, HA-coated implants were more resistant against chemical corrosion by presenting increased corrosion potential (Ecorr) and reduced corrosion current (Icorr) values

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within SBF (Liu et al., 2021). Moreover, as the main inorganic component of human bones, HA coatings can also accelerate bone regeneration on implant surfaces, thus the utilisation of HA-coated implants could potentially benefit both the osseointegration and long-term functioning of dental implants.

3.3 Micro-arc Oxidisation (MAO) Modifying the surface chemistry via MAO to induce ions on Ti implant surface is an alternative approach to improve corrosion resistance (Prando et al., 2018). Such ion implantation is achieved at high voltages, where the high current density initiates a dielectric breakdown of TiO2 surface to expose the underlying Ti to the electrolyte containing different ions (Prando et al., 2018). Next, the various ions and functional groups (i.e., -OH, Ca ions, P ions) could be implanted onto Ti from the electrolyte to modify the chemistry of the passive TiO2 layer (Prando et al., 2018). Compared with the non-treated Ti that only enables a thin and amorphous oxide layer (2 ~ 5 nm), MAO establishes a porous structure on Ti via crystallisation of surrounding electrolytes. Deposition of Ca ions enabled a uniformed Ca-Ti layer that promoted cellular adhesion and migration to improve surface bioactivity (Krupa et  al., 2004). However, several defects on Ca implanted Ti were observed after immersion in SBF (Krupa et al., 2004). Thus, Krupa et al. incorporated phosphorus (P) on Ca-induced implants by MAO, which resulted in significantly reduced pits and defects after SBF immersion (Krupa et  al., 2004). Besides incorporation of ions, another benefit of P-doped Ti implants by MAO against chemical corrosion is increase in thickness and modified topography of the implant surface. It has been reported that doping P on Ti implants generates TiP-containing crystalline structures on the TiO2 oxide layer which significantly increases the stability of amorphous TiO2 and fills the cracks to reduce defects (Prando et al., 2018). Further, the TiP-crystalline is more resistant against chemical corrosion than the native TiO2 film. This observation is supported by electrochemical impedance spectroscopy (EIS) results, where the resistance value against the polarisation of P-incorporated Ti implants was 2.5-fold higher than the non-treated counterparts (Diamanti et al., 2013). In summary, MAO enables the implantation of P onto the TiO2 surface and alters the chemistry of Ti implants, yielding a corrosion-resistance surface.

3.4 Plasma Spraying Plasma spraying is a scalable surface modification approach that sprays the plasma of melted metal under high-temperature to cover the material surface, with the thickness of deposited coatings controlled from microscale to nanoscale (Geetha et al., 2009). The deposited nanoscale particles/crystals from plasma spraying could

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improve the resistance against mechanical friction and chemical corrosion (Geetha et al., 2009). Hydroxyapatite (HA) coating could also be achieved on Ti surfaces via plasma spraying, exhibiting increased corrosion potential and reduced corrosion current during electrochemical corrosion tests (Kwok et al., 2009). However, some gaps inevitably existed between the deposited HA crystals/particles, which was inefficient to shield the underlying Ti. To address this, Singh et al. mixed 1 ~ 2 wt% graphene nanoplatelets (GNPs) in HA powder to form a composite HA-GNP coating. Such HA-GNP coated Ti implants showed higher corrosion resistance (Rcorr) values than the HA-coated Ti (Singh et al., 2020). This is attributed to the embedding of the wrinkled GNPs into the defects of HA matrix, which blocks the electrolyte from penetrating the HA matrix and causing Ti corrosion (Singh et al., 2020). Alternatively, a composite of Al2O3-TiO2 (AT) nanoparticles was coated on Ti implants via plasma spraying (Palanivelu et al., 2014). Both AT and AT-HA sprayed Ti implants had nanoscale roughened surfaces with numerous nanoparticles, with significantly decreased corrosion current density (Icorr) compared to non-modified implants (Palanivelu et  al., 2014). Moreover, the wear-resistance of such AT-HA coated Ti implant was significantly increased (with reduced weight loss and cracks formation throughout the wear experiment). Such results supported the long-term mechanical and chemical stability of dual AT-HA sprayed implants, indicating their favourable stability for clinical application as endosseous implants (Palanivelu & Ruban Kumar, 2014). Similar results were reported by Richard et  al., where AT could be agglomerated into nanoscale dots onto Ti via plasma spraying  (diameter 30 ~ 50 nm) (Richard et al., 2010). Such AT-nanodots were shown to significantly reduce the friction coefficient and the weight loss from modified surface during the tribocorrosion tests (Richard et  al., 2010). Further, SEM images indicated that the AT-sprayed Ti was intact, dense and significantly reduced the corrosion current density (Icorr) throughout electrochemical corrosion tests by shielding the underlying Ti (Richard et al., 2010). In summary, plasma spraying could effectively melt and spray the coating materials onto Ti implant, modifying the implant surface with a nanoscale layer containing nanorods/nanoparticles. Such technique alters both the topography and the chemistry of implant surface, contributing to their mechanical, chemical and electrochemical stability. Moreover, such modified implants exhibits enhanced bioactivity and osseointegration attributed to HA incorporation within the surface layer (Palanivelu & Ruban Kumar, 2014; Richard et al., 2010).

3.5 Plasma Immersion Ion Implantation (PIII) PIII enables superimposition of metallic/non-metallic ions in a nanoscale film on Ti that thickens the TiO2 layer and alters the chemistry. Metal ions such as Ca or Ag implanted on Ti via PIII can result in a thin composite coating with numerous nanostructured dots/peaks (Cao et al., 2016; Harrasser et al., 2015). It has been reported that the Ag-implanted Ti alloys have significantly reduced friction coefficient values

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(reduced from 0.78 to 0.20) indicating their enhanced protection against tribological corrosions (Hongxi et al., 2012). However, most studies on metal ions implanted Ti mainly showed their bioactivity and antibacterial enhancements. The incorporation of non-metal ions (e.g., N, O, C) via PIII can enhance corrosion resistance of Ti implants significantly (Cao et al., 2016; Harrasser et al., 2015). 3.5.1 Nitrogen Treatment Silva et al. treated Ti-6Al-4V alloy by PIII with different combinations of H2 and N2 gases for varied times, which resulted in N-enriched layer on Ti alloys (da Silva et al., 2007). After 90 min N2/H2 treatment, an 88 nm-thickened N-enriched layer could be obtained on Ti alloys with approximately 33% N concentration. Interestingly, the thickness of the N-enriched layer was significantly higher on the N2/H2 treated alloys, compared with the pure N2 treated counterparts. Thicker N layer on N2/H2 treated alloys is attributed to H2 plasma that removes the protective oxide layer from metal surface, thereby augmenting the nitrogen penetration for a thicker metal-nitride layer (da Silva et  al., 2007). Further, N2/H2-coated alloy showed a larger passive region in the potentiodynamic polarisation curve, indicating their improved resistance against passive surface dissolution due to the superimposed N layer that shielded the alloy surface (da Silva et al., 2007). 3.5.2 Oxygen Treatment Yang et al. utilised oxygen plasma with a dose of 1 ~ 4*1016/cm2 on Ti discs to augment corrosion resistance (Yang et  al., 2011). Unlike N2 implantation, oxygen plasma implantation does not change the topography of modified Ti implants, and only alters the surface chemistry by thickening the native TiO2 layer via forming Ti2+(TiO) and Ti3+(Ti2O3) (Yang et al., 2011). Compared with the non-treated Ti, the corrosion rate (Icorr) and passive current (Ipass) of oxygen modified Ti implants were significantly reduced (Fig. 5) (Yang et al., 2011). Further, fewer etching holes were observed on the oxygen implanted Ti after the electrochemical corrosion tests, attributed to the thickened and compact TiO2 layer of oxygen implanted Ti (Mohan & Anandan, 2013). This compact TiO2 layer behaves as a protective shield against the release of Ti ions from the underlying implant surfaces, ensuring their stability under the potential chemical challenges within the oral cavity (Mohan & Anandan, 2013). 3.5.3 Carbon Treatment PIII technique has also been utilized to modify Ti-based implants to deposit a nanoscale diamond-like carbon-rich layer. Shanaghi and Chu reported the fabrication of a 50 nm-thick carbon-rich layer after C-PIII for 2 h (Shanaghi & Chu, 2019).

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Fig. 5  Oxygen plasma modified Ti implants demonstrate superior corrosion resistance by the potentiodynamic polarisation curves. The corrosion rate (Icorr) and passive current (Ipass) of oxygen treated Ti implants (b & c, concentrations of implanted oxygen: b = 1*1016/cm2, c = 4*1016/cm2) were significantly lower than non-treated counterparts (a). (Reproduced with permission from Yang et al. (2011))

Compared with the non-treated Ti alloys, C-PIII modified surfaces showed significantly increased corrosion potentials (Shanaghi & Chu, 2019). Further, the Icorr and Ipass values of Ti alloys were reduced after the C-PIII treatment. This is attributed to the formation of chemically inert carbide bonds between Ti and carbon, which prevented the penetration of corrosive ions during the electrochemical corrosion tests (Shanaghi & Chu, 2019). Similarly, utilising C2H2 to enable C-PIII on Ti alloys also established a C-rich coating layer that improved both the chemical corrosion and the mechanical properties of Ti (Young’s modulus and hardness), attributed to the titanium carbide layer (Poon et al., 2005). In summary, PIII treatment is a scalable and tailorable technique that can implant numerous ions on Ti implants, aiming at specific enhancements, including mechanical stability and corrosion resistance. Doping Ti implants with N, C and O can improve the corrosion resistance via formation of a nanoscale chemically inert protective layer. However, whether such modifications influence the bioactivity of the Ti implants needs further investigations.

4 Ti Implant Nano-engineering To date, numerous modification techniques have been utilised to fabricate nanotopography on Ti implants, including physical (e.g., laser-texturing, deposition), chemical (e.g., chemical etching, plasma spraying) and electrochemical approaches (e.g.,

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electrochemical anodisation) (Guo et al., 2021a; Martinez-Marquez et al., 2022). Nano-engineering techniques (especially via the electrochemical approaches) can create TiO2 nanostructures and thicken the TiO2 layer on Ti implants, thereby forming a protective shield against the chemical corrosion.

4.1 Nano-crystallinzation Nanocrystal layers are defined as a layer of nanoscale grains in different shapes that can be obtained on Ti implants by refining the grain size via surface mechanical attrition treatment (SMAT), a typical severe plastic deformation (SPD) technique (Gleiter, 1989). Grain refinement by SMAT significantly reduces the residual compressive stress within the implant surface, thereby augmenting its mechanical strength (Lin et al., 2006). Attributed to the distortions and strain during treatment that splits the initial grains into smaller nanoscale grains, SMAT reduces the grain size from 10 μm to 50 nm (Jelliti et al., 2013). The chemical stability of SMAT-­ treated Ti implants was significantly enhanced, as evident by the reduced corrosion density and increased potential values from electrochemical impedance spectroscopy (EIS) tests (Jelliti et al., 2013). Besides the EIS results, the wear rate of SMAT-­ treated Ti was reduced 3- to 10-fold  than that of the non-treated surfaces (Jelliti et al., 2013). Further, surface cracks and delamination were significantly reduced on SMAT-Ti after the wearing tests (attributed to the thickened passive layer) (Lin et al., 2006). As reported by Huang and Han, the passive film resistance (Rp) of Ti alloys (7.5*105  Ω.cm2) was significantly augmented after the SMAT treatment (20.5*105  Ω.cm2), indicating a compact and thicker passive film formed on the SMAT-treated Ti alloys with refined grains (Fig.  6) (Huang & Han, 2013). The SMAT treatment was more significant on rough Ti for corrosion resistance enhancements, for which their grain size changed dramatically and yielded a thickened TiO2 protective layer (Skowron et al., 2021). In summary, by refining the grain size of Ti to nanoscale, nanocrystallinity achieved via SMAT could significantly increase the chemical stability and the corrosion resistance of Ti implants. SMAT enhances the mechanical strength of Ti implants by refining the grain size to release their internal stresses. Additionally, SMAT treatment does not alter the chemistry of Ti implants, preserving their favourable biocompatibility.

4.2 Nanowires Fabrication of nanowires via alkali-heat treatment or electrospinning can establish a mesh structure with a distinctive surface roughness on Ti implants (Alali et  al., 2021) to promote surface bioactivity (Stevens & George, 2005). Ti implants with nanowires/nanofibres exhibit enhanced corrosion resistance (Manole et al., 2018;

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Fig. 6  Enhanced corrosion resistance of SMAT-treated Ti implants as shown by the electrochemical impedance spectroscopy (EIS). The Nyquist plots indicate an increase in the passive film resistance (Rp) of SMAT-treated Ti implants in both (a) physiological saline (PS) and (b) simulated body fluids (SBF). (Reproduced with permission from Huang and Han (2013))

Zhu et al., 2019). For example, Zhu et al. fabricated interconnected nanowires via hydrofluoric acid etching and alkali-heat treatment on Ti implants that generated a mesh structure that shielded the implant against the chemical corrosion (Zhu et al., 2019). It is noteworthy that nanowires-interconnected mesh structures could be mechanically delaminated and require further optimization (Zhu et  al., 2019). Alternatively, electrospinning can be utilised to obtain interconnected nanowires.

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Manole et al. mixed titanium butoxide with polyvinylpyrrolidone (PVP) and spun them on Ti implants to distribute nanowires and form an interconnected mesh structure (Manole et al., 2018). Compared with the untreated Ti, PVP-TiO2 nanowires on Ti implants exhibited significantly reduced corrosion current and increased corrosion potential within artificial saliva (Manole et al., 2018). Bioactivity assessments revealed that PVP-TiO2 nanowires could significantly enhance fibroblast viability and inhibit their secretion of inflammatory cytokines. Together, these observations suggest PVP-TiO2 nanowires, in promoting both corrosion-resistance and bioactivity, are suited as dental implant surface modification (Manole et al., 2018).

4.3 Anodised Nanostructures Electrochemical anodisation (EA) is a scalable and cost-effective strategy to fabricate hollow TiO2 nanostructures (nanotubes or nanopores) on Ti implants (Chopra et al., 2023). Several studies have evaluated the influence of nanotubes/nanopores on enhancing the corrosion resistance of Ti implants (Demetrescu et al., 2010; Liu et al., 2011; Man et al., 2008). Man et al. reported that TiO2 nanotubes (TNTs) could significantly reduce the corrosion current density (ICorr) and increase the polarization resistance of modified Ti implants (Man et al., 2008). Similarly, Demetrescu et al. reported a significant higher polarisation resistance of 120 nm-diameter TNTs (20 nm thickened wall) than the non-treated Ti implants (Demetrescu et al., 2010). The data from electrochemical impedance spectroscopy (EIS) showed that the ­corrosion rate of TNTs (0.0076  mm/year) was significantly lower than bare Ti (0.27 mm/year) (Demetrescu et al., 2010). To evaluate the corrosion resistance of varied dimensions of nanotubes, Liu et al. anodised Ti foils at 5, 10, 15 and 20 V for 30 min to fabricate TNTs with the diameter of 22 nm, 39 nm, 59 nm and 86 nm, respectively (Liu et al., 2011). Interestingly, the testing revealed that the corrosion resistance of TNTs increased with the nanotube diameter until 59 nm, but then significantly reduced for the 86 nm-diameter TNTs (Fig.  7) (Liu et  al., 2011). This observation is attributed to the underlying barrier layer of 5 ~ 15 V fabricated TNTs (diameter 22 ~ 59 nm) that increased with voltage (and hence increased the corrosion resistance). However, the TNTs fabricated at 20 V had a significantly larger diameter, which increased the surface area for contact with electrolytes and hence reduced their corrosion-resistantance (Liu et al., 2011). Alternatively, Saji and Choe reported that TNTs anodised by H3PO4-­ based electrolyte had reduced corrosion resistance than the non-anodised counterparts, since the acidic electrolytes significantly reduced the thickness of TiO2 BL beneath the nanostructures to reduce its corrosion resistance (Saji & Choe, 2009). Apart from the TNTs diameter, length, and thickness of the barrier layer, the crystalline phase also influences their corrosion resistance. Compared with the amorphous TNTs, the corrosion resistance of anatase TNTs was reduced at the BL but increased at the nanotubes (Fatichi et al., 2022). Heat treatment during anatase transformation

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Fig. 7  The anodisation polarisation curves of the polished Ti (MPT) and various anodised TNTs (5 V = NT05, 10 V = NT10, 15 V = NT15, 20 V = NT20) in artificial saliva. It is notable that the corrosion resistance could be described as NT15 > NT10 > NT05 > NT20 > MPT. (Reduced with permission from Liu et al. (2011))

compromises BL integrity, reducing the corrosion resistance. Further, the annealing treatment collapses (closes the tops) the anatase nanotubes and reduces the electrolyte-­nanotube contact, thereby increasing the corrosion resistance (Fatichi et al., 2022). Increasing the annealing temperature to 650 °C to create dual anatase-­ rutile phase TNTs exhibited an increased corrosion resistance value at the underlying BL, mainly attributed to the increased thickness during rutile transformation (Fatichi et al., 2022). To further enhance the resistance against tribocorrosion, an additional reverse-­ polarisation was evaluated by Alves et al., which formed a protective P-rich layer with Ca and Zn deposition on the fabricated TNTs (Alves et al., 2018). Such protective film could significantly increase the open circuit potential (OCP) values of TNTs and reduce the formation of the crack during the tribocorrosion tests (Alves et al., 2018). Further, the reverse polarisation thickened the nanoscale oxide film at the Ti-TNTs interface, improving the adhesion of the TNTs layer and their mechanical stability (Alves et al., 2017). In summary, electrochemically anodised Ti implants with nanotubes demonstrate enhanced corrosion resistance attributed to the thickened TiO2 BL and the superimposed nanostructures. Rutile phase nanotubes, with a thickened barrier layer, smaller tube/pore diameters, and longer tube length, are optimal against chemical corrosions.

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5 Research Gaps and Future Directions Dental implants are subjected to the chemical challenges from the ever-changing oral microenvironment at the site of implantation. Modifying implant surfaces from micro- to nanoscale to enhance their surface corrosion resistance can contribute to long-term implant success. However, there remains unanswered questions with respect to the role of surface modification in limiting corrosion of Ti dental implants that will inform future research in this field: • Sandblasting and acid-etching (SLA) is widely utilised for fabricating commercial dental implants, however there are conflicting reports on their corrosion resistance. Additional research is needed to thoroughly evaluate the corrosion resistance of SLA implants by incorporating other anti-corrosion modifications (Guo et al., 2021b; Li et al., 2001; Ogawa et al., 2016). • While refining the grain size of Ti implants by cryogenic treatment could significantly release their internal strains and enhance their wear/corrosion resistance, investigations into bioactivity performance are needed (Gu et  al., 2018; Tang et al., 2017). It is noteworthy that corrosion protection modifications can alter surface topography and chemistry, which can influence surface bioactivity (Guo et al., 2021). For instance, nitriding forms a protective TiN layer on Ti implants (Rossi et  al., 2003), however the effect on bioactivity performance remains underexplored. • Utilising sputtering and plasma spraying to create CaP and HA coating could generate a protective layer on Ti implants against chemical corrosions, while simultaneously enhancing their bioactivity. However, the mechanical stability of such coatings needs thorough testing, for instance in a long-term in vivo study under loading (Khalid Naji et al., 2021; Richard et al., 2010). Ideally, all corrosion protection modifications should survive the constant mechanical forces encountered in an oral setting. For example, bioactive and corrosion protecting nanowire coatings can mechanically delaminate under loading and require further optimization (Manole et al., 2018). • The next generation of anodised nano-engineered Ti implants with nanotubes offer multiple functionalities including corrosion protection, ease of further modification, enhanced bioactivity and tailored drug release; however, their performance must be investigated in vivo for long term functionality (Gulati et al., 2022b). Whether nanotubes will survive mechanical handling associated with surgical insertion into the oral cavity needs examination.

6 Conclusions Physical/chemical treatments (SLA, plasma, nitriding), alloying (Zr and Nb) and nano-engineering (nanowires, nanocrystals and nanotubes) have been employed to modify Ti dental implants to augment their long-term implant survival in a corrosive

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oral environment. However, corrosion protection coating alters the topography, chemistry, mechanical stability and bioactivity performance of the implants, which to-date remains underexplored. Clearly, the next generation of dental implants will employ advanced nanotechnology that offers exceptional corrosion protection, while maintaining favorable stability and bioactivity, tested in a long-term and ‘under load’ in vivo setting. Acknowledgements  Tianqi Guo is supported by a UQ Graduate School Scholarship (UQGSS) funded by the University of Queensland. Karan Gulati is supported by National Health and Medical Research Council (NHMRC) Early Career Fellowship (APP1140699).

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Index

A Alloplastic materials, 2 Anodization, 25, 69, 71, 75, 76, 96, 99–106, 133, 138, 154, 156, 158, 176, 180, 202, 205–209, 211, 214–216 B Bacterial infections, 85, 155, 188, 203, 204 Bioactivity, 9, 10, 24–49, 71, 75, 76, 85, 86, 95, 96, 104, 117–143, 159, 166, 176, 186, 188, 201–203, 206, 209, 212–216, 237, 239–243, 245, 247, 248 Biocompatibility, 5, 6, 8, 10, 11, 13, 16, 70, 76, 84, 93, 95, 102, 106, 120, 131, 164, 173, 176, 179, 201, 202, 222, 237, 243 C Chemical corrosion, 233–240, 242–244, 246, 247 Chemical methods, 225 Corrosion, 5, 7, 8, 11–16, 47, 63, 71, 72, 84, 94, 142, 177, 201, 202, 207, 212, 222–248 Cytotoxicity, 9, 16, 174, 176–178, 186, 188, 203, 216, 228 D Dental implants, 1–17, 24–49, 62–76, 84–109, 117–143, 153–189, 201–216, 222–248 Diabetes, 25, 29–35, 48, 85, 124, 142, 229

E Electrochemical methods, 108 Electrochemical stability, 211–212, 222–248 L Limitations, 1, 88, 94, 104, 105, 187 Local drug delivery, 164, 188 M Mechanical stability, 9, 73, 93, 103, 142, 143, 188, 201–216, 232, 236, 242, 246–248 N Nano-engineering, 76, 85, 88, 128, 129, 133, 135, 188, 202, 242–247 Nanopores, 89, 101, 102, 104, 106, 128, 129, 131, 133, 135, 137–139, 141, 154, 180, 203, 204, 206, 207, 213, 215, 216, 245 Nanoscale surface modification, 86–107, 163 Nanotopography, 88, 104, 108, 120, 129, 131, 133, 137, 143, 158, 164, 208, 242 Nanotubes, 76, 87, 95, 98–101, 128, 133–135, 137–139, 141, 154, 156–159, 171, 176, 181, 184, 203, 204, 206, 207, 214–216, 245–247

© The Editor(s) (if applicable) and The Author(s), under exclusive license to Springer Nature Switzerland AG 2023 K. Gulati (ed.), Surface Modification of Titanium Dental Implants, https://doi.org/10.1007/978-3-031-21565-0

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256 O Osseointegration, 2–6, 8, 16, 24–27, 29, 30, 34, 36, 37, 41, 42, 45, 46, 48, 49, 62–65, 67–69, 71, 73, 74, 76, 84–87, 103, 105, 106, 117, 118, 121–124, 127–135, 142, 154, 155, 158–164, 166–169, 175, 181, 183, 201, 205, 229, 232, 238–240 Oxide layer, 13–16, 71, 96, 97, 100–102, 120, 157, 201, 222, 224, 226, 232, 234, 239, 241 P Physical methods, 89, 94 R Roughness, 7, 8, 14, 15, 25, 62, 63, 69, 71–74, 85, 87–89, 91, 93, 94, 96, 100, 106, 121, 122, 127, 128, 135–137, 156, 178, 214, 215, 232, 233, 243

Index S Sand-blasted and acid-etched surface, 86, 122, 129, 134, 232, 247 Smoking, 26, 29–31, 34, 35, 46–49, 154, 229 Surface modification, 17, 25, 31–34, 38, 42, 46, 48, 49, 62–76, 84–108, 117–143, 153, 154, 184, 188, 201–203, 212, 225, 239, 245, 247 Systemic diseases, 48 T Therapy, 15, 24–49, 105, 107, 128, 143, 153–189, 202, 214 Titania nanotubes, 96, 133, 156, 159–162, 165–168, 173, 176–179, 183, 187 Titanium, 1–17, 24–49, 62–76, 84–86, 94, 99–104, 106, 117–143, 153–189, 201–216, 222–248 Titanium alloys, 2, 7, 9, 10, 13–16, 65, 72, 84, 234