Calcium Phosphate Nanocoatings for Bone Regeneration (Tissue Repair and Reconstruction) 981995505X, 9789819955053

Table of contents :
Preface
Contents
1 Introduction
References
2 Why Surface Modification
References
3 Calcium Phosphate
References
4 Mechanical Integrity of Thin Films and Coatings and Their Clinical Significance
4.1 Determining Stresses in Thin Films and Coatings
4.2 Interfacial Adhesion Between a Coating and a Substrate
4.3 Coating Adhesion and Mechanical Interlocking
4.4 Enhancing the Adhesion of Nanocoatings to Titanium via Anodization Process
4.5 Adhesion Theories Based on Chemical Bond, Electrostatic, and Diffusion
4.6 Characterizing Coatings and Nanocoatings Mechanically Prior to Implantation
4.6.1 In Situ Microtensile Test
4.6.2 Shear and Tensile Pull-Off Testing
4.6.3 Nanoindentation Testing
4.6.4 Bend Delamination Test
4.6.5 Bulge and Blister Testing
4.6.6 Scratch Testing
4.6.7 Pin-on-Disk
References
5 Coating Deposition Techniques
5.1 Coatings Produced by Plasma Spraying
5.2 Comparison Between Coatings Produced by Plasma Spraying and Sol–Gel Technique
5.3 The Sol–Gel Deposition Approach
5.3.1 The Alkoxide Route
5.4 Other Coating Deposition Techniques
5.4.1 Pulsed Laser Deposition
5.4.2 Physical and Chemical Vapor Depositions
5.4.3 Electrodeposition
References
6 Cellular Responses
References
7 Enhancing Implant Osseointegration Through Nanocomposite Coatings
7.1 Collagen
7.2 Chitosan
7.3 Bone Morphogenetic Proteins (BMPs)
7.4 Peptides
7.5 Stem Cells
7.6 Functionalization Using Osteopontin
7.7 Simvastatin
References
8 Calcium Phosphate Nanocoated Coralline Apatite
References

Citation preview

SpringerBriefs in Tissue Repair and Reconstruction Andy H. Choi Besim Ben-Nissan

Calcium Phosphate Nanocoatings for Bone Regeneration

Tissue Repair and Reconstruction Series Editors Andy H. Choi, Carlingford, NSW, Australia Besim Ben-Nissan, Sydney, NSW, Australia

SpringerBriefs in Tissue Repair and Reconstruction provides a unique perspective and in-depth insights into the latest advances and innovations contributing to improved and better treatments for patients with damaged soft and hard tissues as a result of diseases, trauma, and implantations. The book series consists of volumes that offer biomedical researchers better insights into the advancements of biomaterials science and their translation from the laboratory to a clinical setting. Similarly, the series provides information to surgeons and medical practitioners on novel ideas in biomedical science and engineering on top of disseminating new ideas and know-hows in diagnostics and treatment options for patients from head to toe. The series will cover a number of key topics: Fundamental Concepts and Surface Modifications: The topic will provide detailed information on the discovery and advancements of biomaterials surface modification approaches and their use within the human body in a safe manner and without provoking any negative tissue response. Computational Simulations and Biomechanics: Anatomically accurate computational models are being in all fields of medicine particularly in orthopedics and dentistry to reveal the biomechanical functions and behaviors of bones and joints when damaged, diseased, and in the health state. They also contribute to our understanding during the design and applications of implants and prosthetics subjected to functional loadings and movements. Surgical Advances and Treatment Options: Discusses how surgical techniques are revolutionized by our deeper understanding into biomaterials science and tissue engineering. The section also focuses on the latest innovations and surgical advancements currently being used to treat patients with damaged tissues. Post-Operative Treatment and Rehabilitation Engineering: Expands the independence and functionality of the patient after surgery while at the same time reducing the chance of complications such as wound infections and dislocations. Advances in technologies are creating new opportunities in how physiotherapy rehabilitations are delivered.

Andy H. Choi · Besim Ben-Nissan

Calcium Phosphate Nanocoatings for Bone Regeneration

Andy H. Choi School of Life Sciences University of Technology Sydney Ultimo, NSW, Australia

Besim Ben-Nissan School of Life Sciences University of Technology Sydney Ultimo, NSW, Australia

ISSN 2731-9180 ISSN 2731-9199 (electronic) Tissue Repair and Reconstruction ISBN 978-981-99-5505-3 ISBN 978-981-99-5506-0 (eBook) https://doi.org/10.1007/978-981-99-5506-0 © The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 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 Singapore Pte Ltd. The registered company address is: 152 Beach Road, #21-01/04 Gateway East, Singapore 189721, Singapore Paper in this product is recyclable.

Preface

One of the primary concerns in biomedical materials research is the relationship between the biological responses and surface properties of materials. At present, orthopedic and dental implants and prosthetics are manufactured from metallic metals such as titanium and its alloy such as Ti-6Al-4V as well as bioinert cobalt-chromium alloys. Their major disadvantage is failure to adapt to the local tissue environment and does not chemically bond to bone unless modified. Currently, two different methods are used to insert orthopedic implants surgically. The first uses bone cement (predominately poly (methyl methacrylate) or PMMA) for strong adhesion. The second approach utilizes bioactive ceramic coating to coat porous or microtextured implants for chemical bonding and mechanical interlocking. This technique is extensively applied to dental and maxillofacial implants in addition to orthopedic prostheses. While metals such as Ti-6Al-4V have been successfully utilized for more than half a century, its relationship with aseptic inflammation particularly in orthopedic joint replacements believed to be the result of titanium particles releasing from the surfaces of implants and prostheses into the surrounding microenvironment has been a concern. In dental implantology, it has also been revealed that particles could be released in a surface type-dependent fashion after ultrasonic scaling of titanium implants that may aggravate peri-implantitis. As a result, surface modification and the utilization of bioceramic coatings (both nanocoatings and nanocomposite coatings) on these materials are intended to offer protection against the release of metal ions which might trigger a negative host response. Ultimately, this provides an improved environment and structure for the formation of new bone tissues. The ability to generate an ideal environment for bone growth has been used to determine whether an implant or prosthesis coated with a biocompatible material is successful from a clinical perspective. Even though many biomaterials have been introduced in dentistry, tissue engineering, and orthopedics, bioactive calcium phosphates are extensively utilized due to their well-documented biocompatibility and safety profile in addition to their similarities to bone. Studies based on human trials and animal models have demonstrated that a thin hydroxyapatite and other calcium phosphate coating deposited on the surfaces of implants accelerated early v

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bone formation as well as an increase in bond strength between bone and implant. It has been proposed that an increase in the concentration of calcium and phosphate due to the partial dissolution of the apatite into the microenvironment, followed by the formation of carbonate apatite microcrystals and their amalgamation with the organic matrix of bone, causes biological growth of bone tissue. Consequently, it is important to gain an understanding into the mechanisms behind the osseointegration of calcium phosphate-coated implants using both human trials and animal models. I would like to give very special thanks to my mentor, series co-editor, and coauthor Prof. Besim Ben-Nissan for his friendship, support, and advice for over two decades. Finally, I would like to acknowledge the people at Springer, especially Dr. Ramesh Premnath, Mr. Ramesh Kumaran, and the team at Springer Publishing for their help and for making this book possible. Sydney, Australia

Andy H. Choi

Contents

1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1 4

2 Why Surface Modification . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

5 7

3 Calcium Phosphate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 9 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11 4 Mechanical Integrity of Thin Films and Coatings and Their Clinical Significance . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1 Determining Stresses in Thin Films and Coatings . . . . . . . . . . . . . . . . 4.2 Interfacial Adhesion Between a Coating and a Substrate . . . . . . . . . . 4.3 Coating Adhesion and Mechanical Interlocking . . . . . . . . . . . . . . . . . . 4.4 Enhancing the Adhesion of Nanocoatings to Titanium via Anodization Process . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.5 Adhesion Theories Based on Chemical Bond, Electrostatic, and Diffusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.6 Characterizing Coatings and Nanocoatings Mechanically Prior to Implantation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.6.1 In Situ Microtensile Test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.6.2 Shear and Tensile Pull-Off Testing . . . . . . . . . . . . . . . . . . . . . . . 4.6.3 Nanoindentation Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.6.4 Bend Delamination Test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.6.5 Bulge and Blister Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.6.6 Scratch Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.6.7 Pin-on-Disk . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

13 15 16 17 17 20 20 21 23 23 25 27 28 29 29

5 Coating Deposition Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 33 5.1 Coatings Produced by Plasma Spraying . . . . . . . . . . . . . . . . . . . . . . . . . 34 5.2 Comparison Between Coatings Produced by Plasma Spraying and Sol–Gel Technique . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 35 vii

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5.3 The Sol–Gel Deposition Approach . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.3.1 The Alkoxide Route . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.4 Other Coating Deposition Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . 5.4.1 Pulsed Laser Deposition . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.4.2 Physical and Chemical Vapor Depositions . . . . . . . . . . . . . . . . 5.4.3 Electrodeposition . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

36 40 43 43 44 45 46

6 Cellular Responses . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 51 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 54 7 Enhancing Implant Osseointegration Through Nanocomposite Coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.1 Collagen . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.2 Chitosan . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.3 Bone Morphogenetic Proteins (BMPs) . . . . . . . . . . . . . . . . . . . . . . . . . . 7.4 Peptides . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.5 Stem Cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.6 Functionalization Using Osteopontin . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.7 Simvastatin . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

59 60 61 62 65 66 68 70 72

8 Calcium Phosphate Nanocoated Coralline Apatite . . . . . . . . . . . . . . . . . . 79 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 82

Chapter 1

Introduction

The research and utilization of bioceramics as coating materials in dentistry and orthopedics have become a state of the art necessity to enhance the integration of the implant to the surrounding bone tissue instead of simply being a curiosity. Biological fixation is a term used to define whether an implant or prosthetic device can be securely attached to the host tissue without the need for any adhesive but instead through bone ingrowth. In addition, this fixation can be conducted with or without the use of mechanical fixation [1–3]. From a clinical point of view, biofunctionality in addition to biocompatibility of any biomaterial are the primary factors governing their rate of success and both of which are directly correlated to how the biomaterial interacts with the human tissue at the implant interface [1]. The concept of biocompatibility as described in the review by Williams refers to the capacity of a biomaterial to carry out its intended function relating to a medical therapy without eliciting any unwanted systemic or local effects to the patient, while at the same time creating the most suitable and favorable tissue or cellular response in that specific situation, and optimizing the clinically relevant performance of that therapy [4]. Biomaterials and bioceramics were employed during the early 1970s to perform singular and biologically inert roles such as implants. Furthermore, these materials were capable of interacting with bodily fluids and tissues without causing any adverse reactions for an extended period. The inadequacies with these synthetic materials as tissue replacements were pointed out by our increasing perception that cells and tissues of the human body perform many other crucial metabolic and regulatory functions. Since then, the requirements centered on the clinical utilization of biomaterials and bioceramics have reformed to provide a more positive interaction with the host instead of simply preserving the basic function without provoking any undesirable response. This concept has been complemented by the ever-growing demands on medical devices to prolong the duration of life in addition to improving its quality. Bioceramics, when used as body-interactive materials,

© The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2023 A. H. Choi and B. Ben-Nissan, Calcium Phosphate Nanocoatings for Bone Regeneration, Tissue Repair and Reconstruction, https://doi.org/10.1007/978-981-99-5506-0_1

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possess the capacity to aid the human body to heal by promoting the regeneration of tissues and thereby restoring physiological functions. Clinically, the most widely used materials at the present moment are those selected from a handful of well-examined and available biocompatible ceramics, polymers, metals, and their combinations as composites or hybrids. Despite the fact that bioceramics are used extensively as implants in dentistry, maxillofacial surgery, and orthopedics, further developments are progressing to increase their function and achieve improvements in their reliability as well as performance. Promoting biological fixation and osseointegration as well as preventing osteolysis and inflammatory responses are vital prerequisites for any implants and prostheses regardless if they are manufactured using metals, polymers, ceramics, or composites. In modern medical applications, biomaterials are crucial and these materials have been improved and refined in the past several decades for applications seen in implantology, drug delivery, and tissue engineering scaffolds. There has been an increase in interests in nanostructured materials for advanced technologies and the focus now is on the production of nanocoatings and nanobioceramics, which is applicable in dentistry and orthopedics. Nanostructured materials by definition refer to a variety of materials that comprise of delicate structures and sizes that fall between 1 and 100 nm. A huge development of nanotechnology has been recognized, but nonetheless, such developments have not come as a surprise when it is valued that these nanostructured materials demonstrate the capability to be integrated and adapted into biomedical devices. This is possible since a majority of biological systems such as protein complexes, membranes, and viruses display natural nanostructures. Nanocoatings, nanolaminates, and their composites offer the possibility of altering the surface properties of surgical-grade materials to achieve improvements in in vivo performance and reliability as well as increased protection from the release of unnecessary or harmful metal ions through appropriate selection of coating materials. They can also be synthesized to deliver exceptional bioactivity and faster tissue bonding properties through their increased surface area and their nanocrystalline structure, which are interrelated. Nanocoatings can be described as thin films or coatings whose thicknesses are below the range of micron-sized coatings or less than 1 μm. In general, the thickness is below 100 nm for a single coating and they can consist of isotropic and homogeneous compounds. Multiple layers with suitable biological, mechanical, physical, and chemical properties can be synthesized with relative ease. More importantly, they can be applied as multilayered gradient coatings or nanolaminates with different compositions. A number of advantages are associated with the applications of nanocoatings including the ability to be amalgamated with other compounds or nanoparticles, purity as a result of raw materials selection, low cost due to small quantities of materials needed for thin coating processes, ease of synthesis, and a variety of application methods can be used [3, 5–7]. It should be mentioned that scientific notation is usually mixed regarding the definitions of “nanocoatings” and “thin-film” coatings, and it has been quite controversial as to which term is more appropriate with no known and widely accepted explanation and both are used interchangeably.

1 Introduction

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Nanocomposites can be defined as an amalgam of two or more materials in which at least one of those materials should be on a nanometer scale. Nanocomposites can be created by either mixing physically or through the addition of a new element into an existing nanosized material. This introduction alters the existing nanomaterials’ properties and this may in turn offer new function for the material. Furthermore, it is conceivable through the use of the composite approach as well as secondary substitution phases to synthesize nanocomposites with mechanical properties such as Young’s modulus similar to those of human cortical and cancellous bone [3, 5–7]. The gel system is an alternative form of nanocomposite and its development is ideal for biomedical applications such as bone tissue repair [8–10]. Matching the specific requirements of biomedical devices is made possible through the gel approach as the properties of nanomaterials can be enhanced and modified. Essentially, it is a three-dimensional network immersed in a fluid, and subsequently, it enables the encapsulation of nanostructured materials within the system. Based on observations using in vivo animal models, it has been suggested that the gel composite approach could be used to deliver growth factors, stem cells, or pharmaceutics to improve implant osseointegration [11], as well as enhance bone tissue regeneration [12–18]. Nanogel, which is a nanosized flexible hydrophilic polymer gel, is another example of a gel system that can be utilized in applications such as guided bone regeneration and as drug-delivery carriers [19, 20]. Cholesterol-bearing pullulan nanogel is a synthetic degradable biomaterial, and it has been postulated that such nanogel is capable of entrapping proteins or hydrophobic drugs; subsequently, it was utilized in the study by Kato et al. to examine the possible effect of prostaglandin E2 (PGE2), a lipid-signaling molecule, on bone formation when injected on to the calvariae of mice [19]. The authors claimed that low dosage of PGE2 delivered by the nanogels could induce new bone formation and the bone formation activities of PGE2 were enhanced by the nanogels only at the site of injection. Later, a nanogel-based scaffold in the shape of a disk was synthesized to examine the feasibility of combined delivery of a selective EP4 receptor agonist and low-dose bone morphogenetic protein-2 in an effort to repair critical-size circle-shaped bone defects in calvariae that otherwise did not heal spontaneously [20]. In vivo observations showed such a delivery system was able to efficiently activate bone cells to regenerate calvarial bone through the formation of both outer and inner cortical plates and bone marrow tissue to recreate a structure similar to that of intact calvaria. The ways in which nanostructured materials are synthesized will strongly determine its properties and microstructures. Consequently, nanomaterials should be manufactured using the most suitable approach so that the desired property and/ or a combination of properties can be attained. Despite the fact that nanomaterials and nanobioceramics are applied extensively in orthopedics, spinal and craniomaxillofacial surgeries and in dentistry as dental implants, progress has been ongoing to extend its applications and achieving improvements in their performance and reliability.

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References 1. Choi AH (2022) Biomaterials and bioceramics—part 1: traditional, natural, and nano. In: Choi AH, Ben-Nissan B (eds) Innovative bioceramics in translational medicine I. Springer series in biomaterials science and engineering, vol 17. Springer, Singapore, pp 1–45 2. Choi AH (2022) Biomaterials and bioceramics—part 2: nanocomposites in osseointegration and hard tissue regeneration. In: Choi AH, Ben-Nissan B (eds) Innovative bioceramics in translational medicine I. Springer series in biomaterials science and engineering, vol 17. Springer, Singapore, pp 47–88 3. Choi AH, Ben-Nissan B (2018) Anatomy, modeling and biomaterial fabrication for dental and maxillofacial applications. Bentham Science Publishers, United Arab Emirates 4. Williams DF (2008) On the mechanisms of biocompatibility. Biomaterials 29:2941–2953 5. Choi AH, Ben-Nissan B (2015) Calcium phosphate nanocoatings and nanocomposites, part I: recent developments and advancements in tissue engineering and bioimaging. Nanomedicine 10:2249–2261 6. Choi AH, Ben-Nissan B (2007) Sol-gel production of bioactive nanocoatings for medical applications: part II: current research and development. Nanomedicine 2:51–61 7. Ben-Nissan B, Choi AH (2006) Sol-gel production of bioactive nanocoatings for medical applications: part I: an introduction. Nanomedicine 1:311–319 8. Dhivya S, Saravanan S, Sastry TP et al (2015) Nanohydroxyapatite-reinforced chitosan composite hydrogel for bone tissue repair in vitro and in vivo. J Nanobiotechnol 13:40. https:// doi.org/10.1186/s12951-015-0099-z 9. Kim BS, Kim HJ, Choi JG et al (2015) The effects of fibrinogen concentration on fibrin/ atelocollagen composite gel: an in vitro and in vivo study in rabbit calvarial bone defect. Clin Oral Implants Res 26:1302–1308 10. Heo DN, Ko WK, Bae MS et al (2014) Enhanced bone regeneration with a gold nanoparticlehydrogel complex. J Mater Chem B 2:1584–1593 11. Lee JH, Kim J, Baek HR et al (2014) Fabrication of an rhBMP-2 loaded porous βTCP microsphere-hyaluronic acid-based powder gel composite and evaluation of implant osseointegration. J Mater Sci Mater Med 25:2141–2151 12. Zhang Y, Dou X, Zhang L et al (2021) Facile fabrication of a biocompatible composite gel with sustained release of aspirin for bone regeneration. Bioact Mater 11:130–139 13. Han SH, Jung SH, Lee JH (2019) Preparation of beta-tricalcium phosphate microspherehyaluronic acid-based powder gel composite as a carrier for rhBMP-2 injection and evaluation using long bone segmental defect model. J Biomater Sci Polym Ed 30:679–693 14. Kim S, Kim J, Gajendiran M et al (2018) Enhanced skull bone regeneration by sustained release of BMP-2 in interpenetrating composite hydrogels. Biomacromolecules 19:4239–4249 15. Liao HT, Tsai MJ, Brahmayya M et al (2018) Bone regeneration using adipose-derived stem cells in injectable thermo-gelling hydrogel scaffold containing platelet-rich plasma and biphasic calcium phosphate. Int J Mol Sci 19:2537. https://doi.org/10.3390/ijms19092537 16. Ma D, An G, Liang M et al (2016) A composited PEG-silk hydrogel combining with polymeric particles delivering rhBMP-2 for bone regeneration. Mater Sci Eng C Mater Biol Appl 65:221– 231 17. Kim K, Lam J, Lu S et al (2013) Osteochondral tissue regeneration using a bilayered composite hydrogel with modulating dual growth factor release kinetics in a rabbit model. J Control Release 168:166–178 18. Suzawa Y, Funaki T, Watanabe J et al (2010) Regenerative behavior of biomineral/agarose composite gels as bone grafting materials in rat cranial defects. J Biomed Mater Res A 93:965– 975 19. Kato N, Hasegawa U, Morimoto N et al (2007) Nanogel-based delivery system enhances PGE2 effects on bone formation. J Cell Biochem 101:1063–1070 20. Kamolratanakul P, Hayata T, Ezura Y et al (2011) Nanogel-based scaffold delivery of prostaglandin E2 receptor-specific agonist in combination with a low dose of growth factor heals critical-size bone defects in mice. Arthritis Rheum 63:1021–1033

Chapter 2

Why Surface Modification

The prevention of inflammatory response is a critical prerequisite for any implants and prosthetics used in the biomedical arena regardless of its design or materials used in its manufacture. In addition, orthopedic and dental implants and prostheses are also required to encourage osseointegration and provide sufficient biological fixation as described earlier. The intention of depositing bioceramic micro- and nanoscale coatings on these implants and prostheses is to enhance its bioactivity in addition to speeding up the healing process by altering the surface properties, and if possible safeguard the implant material against biodegradation and to provide protection against the release of metal ions into the surrounding tissue environment that could trigger a negative host response. Ultimately, this results in an improved structure and atmosphere for new bone growth [1–6]. Achieving the ideal environment for bone growth has been hypothesized to be the condition in which any implant or prosthesis that has been coated with a biocompatible material will provide a successful outcome from a clinical perspective. This can be accomplished for instance if adequate mechanical interlocking is generated along with the availability of a bioactive surface for biological and chemical bonding to take place. Reducing the amount of metal ion release is another approach to achieve clinical success. A number of tissue responses can occur at the interface between the biomaterial and hard or soft tissues after an implantation. Several factors can influence the success rate of a biomedical implant or prosthesis, and these include the design and materials selection of the implant in addition to the properties and structure of the material used. The surgical procedure or technique used as well as the health and medical condition of the patient will also influence its success. During the last several decades, issues and concerns related to adhesion and biological interactions and the way in which reliability and longevity of implants and prostheses can be improved have motivated the exploration of surface modification toward early osseointegration, bone-implant adaptability, and rapid healing. Surface enhancements through increasing bioactivity via biological and chemical © The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2023 A. H. Choi and B. Ben-Nissan, Calcium Phosphate Nanocoatings for Bone Regeneration, Tissue Repair and Reconstruction, https://doi.org/10.1007/978-981-99-5506-0_2

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2 Why Surface Modification

means as well as macro- and microtexturing have been the primary focus for a large number of research groups in the biomedical and surgical arena. Presently, dental and orthopedic implants and prostheses are manufactured using metals such as bioinert cobalt-chromium alloys and titanium along with its alloys. Unless modified, these metals do not chemically bond to bone. Consequently, they are surgically inserted and bonded either via mechanical fixation involving microtexturing or with bone cement. Surface-modified or microtextured structures have an added benefit that centers on mechanical interlocking, and in certain circumstances, chemical bonding with the surrounding bone tissue. This can be accomplished by fixing porous surfaces produced in such a manner that they possess suitable surface pores. Another approach is to use wires or beads to generate micro- or macrotexturing. The surface area increases as a result of these structures and hence the amount of fixation. Chemical bonding can also be achieved by coating these surfaces with bioceramics such as calcium phosphate. As mentioned earlier, even though dental and orthopedic implants and prostheses have been produced successfully from titanium and its ternary alloys such as Ti6Al-4V for a number of decades, they have also been linked to aseptic inflammation in orthopedic joint replacements believed to be the result of titanium particles releasing into the surrounding micro-environment from the implant surface. Furthermore, a study has demonstrated that ultrasonic scaling of titanium dental implants could release particles in a surface type-dependent fashion that may provoke periimplantitis [7]. Above all, a major drawback of synthetic implants is their failure to adapt to the local tissue environment, and in particular, cells do not adhere adequately or directly to a majority of metallic surfaces. The ideal mechanism for fixation is intimate tissue ingrowth. While the most vital factor on bone attachment is the dimensions of micropores, which regulates mechanical attachment and bonding, bioactivity in general can be improved using calcium phosphate bioceramics. It was demonstrated that the deposition of nanocoatings over meso- and nanoporous structures improves the mechanical properties and performance as a result of pore-filling effect of the nanocoatings. More importantly, covering the substrate with nanocoatings can result in a decrease in surface defects. The physical and mechanical properties of the substrate material can also be improved as the nanocoating penetrates into the meso- and nanopores. Similarly, the deposition of nanocoatings on titanium or other metallic substrates can have numerous advantages in comparison to polycrystalline large grain-coated materials. Due to their extremely low grain size, materials coated with nanocoatings have large surface areas and this in turn reduces their sintering or densification temperatures. As a result, they can be sintered at lower temperatures and at a lower cost. Furthermore, biological properties are also influenced by the small crystalline grain structures and surface morphologies of nanocoating, and it has been demonstrated that nanocoated materials can promote early osseointegration as well as accelerate the bonding of the implant to soft and hard tissues in clinical situations [1–6]. Coating deposition and surface modification on titanium are of paramount significance as they allow the application of an array of coatings while preserving the advantageous bulk properties of the titanium substrate. Several macro- and nanocoatings

References

7

including sol–gel-derived ceramic coatings demonstrate promise due to their relative ease of production, capacity to offer exceptional mechanical properties in part due to their nanocrystalline structure, and ability to form a physically and chemically uniform coverage over complex geometric shapes [1–6]. On the other hand, the mechanical properties of the titanium substrate can be reduced if thick coatings were deposited due to the necessity to sinter powder ceramics at elevated temperatures of 1000 °C and above. If the temperature is below 882.5 °C, which is also referred to as the β-phase transus temperature, the crystal structure of commercially pure titanium is a hexagonal close-packed structure. However, titanium will undergo transformation to a body-centered cubic structure if the temperature exceeds the β-phase transus temperature. A degradation of the bond strength between the titanium substrate and the bioceramic coating will be observed as a result of the strain produced by this phase transformation at that temperature [3, 8]. Different dental and orthopedic implants and prostheses will utilize coating materials with various designs, functions, and properties depending on their intended application. Consequently, ascertaining the mechanical properties of these coatings using accurate measurement techniques is vital as they can differ significantly from the bulk material. In addition, a better and more reliable approach is essential to measure quantitatively the adhesion strength and hardness of the coating as well as the fracture toughness at the coating-substrate interface [3, 9].

References 1. Choi AH (2022) Biomaterials and bioceramics—part 1: traditional, natural, and nano. In: Choi AH, Ben-Nissan B (eds) Innovative bioceramics in translational medicine I. Springer series in biomaterials science and engineering, vol 17. Springer, Singapore, pp 1–45 2. Choi AH (2022) Biomaterials and bioceramics—part 2: nanocomposites in osseointegration and hard tissue regeneration. In: Choi AH, Ben-Nissan B (eds) Innovative bioceramics in translational medicine I. Springer series in biomaterials science and engineering, vol 17. Springer, Singapore, pp 47–88 3. Choi AH, Ben-Nissan B (2018) Anatomy, modeling and biomaterial fabrication for dental and maxillofacial applications. Bentham Science Publishers, United Arab Emirates 4. Choi AH, Ben-Nissan B (2015) Calcium phosphate nanocoatings and nanocomposites, part I: recent developments and advancements in tissue engineering and bioimaging. Nanomedicine 10:2249–2261 5. Choi AH, Ben-Nissan B (2007) Sol-gel production of bioactive nanocoatings for medical applications: part II: current research and development. Nanomedicine 2:51–61 6. Ben-Nissan B, Choi AH (2006) Sol-gel production of bioactive nanocoatings for medical applications: part I: an introduction. Nanomedicine 1:311–319 7. Eger M, Sterer N, Liron T et al (2017) Scaling of titanium implants entrains inflammationinduced osteolysis. Sci Rep 7:39612. https://doi.org/10.1038/srep39612 8. Colling EW (1984) The physical metallurgy of titanium alloys. American Society for Metals, Cleveland 9. Ben-Nissan B, Choi AH, Bendavid A (2013) Mechanical properties of inorganic biomedical thin films and their corresponding testing methods. Surf Coat Technol 233:39–48

Chapter 3

Calcium Phosphate

Synthetic calcium phosphate would be the perfect choice when searching for an ideal biomaterial to imitate and substitute human bone tissue based on the notion that they can replicate the structure and composition of a bone mineral commonly known as natural hydroxyapatite (HAp). In addition, calcium phosphate has been widely accepted as a biocompatible material chemically resembling the mineral component of human teeth and bone [1–6]. Hydroxyapatite, according to the majority of published data, is classified as calcium phosphate to which it belongs. Subsequently, HAp will be considered from the chemical properties point of view as calcium phosphate despite the fact that the reactivities and solubility will differ when compared to other phosphates within the physiological environment. Due to their brittleness and inorganic nature, the mechanical properties of calcium phosphates are a long way away to those of human bone even though it possesses similar composition and chemistry. This places a limitation on the application of porous HAp in load-bearing scenarios without further modifications. Calcium phosphates are classified according to their specific solubilities such as their degradation over time after it is attached and bonded to bone tissue and slowly being replaced by advancing bone growth (Fig. 3.1). Once it is exposed to bodily fluids, an exchange in the surface ions of calcium phosphate or HAp can take place with those of the aqueous solution. On the other hand, different molecules and ions such as collagen and proteins can be adsorbed onto the surface [7]. Normally, it has been widely recognized that natural and synthetic calcium phosphate bioceramics are osteoconductive (exhibit the capacity to support tissue ingrowth and bone formation) but not osteoinductive (demonstrate the ability to generate bone as soon as it is implanted into non-osseous sites). Bone graft materials commonly utilize calcium phosphate with interconnecting pores between 100 and 500 μm in diameter. The structure and chemistry of calcium phosphate determine its rate of dissolution, and this in turn governs the in-situ strength and long-term stability. Other examples where calcium phosphate bioceramics are being applied in the dental, medical, and orthopedics arenas include artificial eye, ear, and ocular © The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2023 A. H. Choi and B. Ben-Nissan, Calcium Phosphate Nanocoatings for Bone Regeneration, Tissue Repair and Reconstruction, https://doi.org/10.1007/978-981-99-5506-0_3

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3 Calcium Phosphate

Fig. 3.1 Solubilities of various calcium phosphate compounds [6]

implants, bone cement additives, spinal fusion, maxillofacial reconstruction, bone space fillers, alveolar ridge augmentation, the repair of bone and periodontal defects, and composites and implant coatings [1–6]. Currently, commercially available synthetic calcium phosphate bioceramics are classified based on their composition (Fig. 3.1). These include HAp, α- and βtricalcium phosphate, and biphasic calcium phosphate which is a mixture of HAp and β-tricalcium phosphate with a variable ratio of β-tricalcium phosphate and HAp [1–6]. Other calcium phosphate biomaterials that are available commercially have been produced from biological materials such as those derived from bovine bone, marine algae, processed human bone, and hydrothermally converted corals [1, 2, 8]. A deeper understanding into the properties, structure, and compositions of biological apatites, and in particular human enamel apatites, were gained through early investigations on synthetic apatites and related calcium phosphates. Nevertheless, studies on synthetic apatites during the past thirty years or so had been focused on their synthesis and utilization in medicine and dentistry in addition to their relevance as scaffolds for bone and teeth regeneration. Nanotechnology has driven innovative approaches for the production of synthetic bone-like calcium phosphate nanomaterials and nanocoatings. The accessibility of calcium phosphate nanoparticles has certainly generated new possibilities in the research and development of superior biocompatible coatings for implants and prostheses and high-strength dental and orthopedic nanocomposites. Although the applications of bulk calcium phosphates began since the 1920s for surgical applications, the new macro-, micro-, and nanocoatings of calcium phosphate are comparably new and have been utilized as porous coating materials since the early 1980s [1–6]. As mentioned previously, due to a lack of sufficient bioactivity on the surface over time, metallic implants like those manufactured from titanium and its alloys suffer a long-term issue of loosening once it is surgically implanted [1–6]. The main intention of using calcium phosphate as a bioactive coating is to create a rapid and strong biological attachment to the soft or hard tissue. The advantage of using calcium phosphate coatings lies on the supply of calcium and phosphate ions that may facilitate and/or stimulate the growth of new bone tissue on and toward the surface of the implant or prostheses. Calcium phosphate-coated implants have been shown to demonstrate extensive bone apposition in both in vivo and in vitro studies. On the other hand, a good balance must be reached between the solubility of the coating and the growth rate of bone tissue in an effort to reach mechanical integrity under functional loading. This would enable sufficient mechanical properties and bonding at the interface between the implant or prosthesis and bone tissue with the purpose of achieving long-term survivability [1–6].

References

11

The development of good interfacial strength between the implant and bone tissue is the result of the biological interactions of released calcium and phosphate ions. Calcium phosphate-coated implants, if produced in an acceptable fashion, will heal faster and displayed improvements in bone attachment [1–6]. Variables governing the long-term performance and quality of an implant or prosthesis coated with calcium phosphate include the surface roughness, porosity, constituent phases, thickness, and crystallinity of the coating. Other factors include the overall design of the implant and/or prosthesis and the amount of biomechanical functional loading [1–7]. Furthermore, the chemistry and surface topography of calcium phosphate crystals deposited as coatings or thin films on implants exhibit an acceleration in early bone formation and an increase in bond strength at the implant–bone interface. Even though a number of production methodologies are available to produce bonelike calcium phosphate nanoplatelets and nanopowders, the sol–gel approach is one very promising technique that can be utilized to synthesize these materials and will be discussed in detail in Sect. 5.3. Consequently, sol–gel crystalline nanocoatings and nanocomposite coatings based on calcium phosphate were developed and deposited on a variety of substrates [4–6]. Furthermore, it has been demonstrated that utilizing calcium phosphate, which is similar from a chemistry perspective to the mineral component of natural bone, as a coating has the added benefit of providing a bioactive layer on bioinert implant materials such as cobalt-chromium alloys as well as alumina and zirconia in an effort to enhance their osseointegration potential. Observations from early studies suggested it is more challenging to produce monophasic calcium phosphate materials and coatings despite the fact that biphasic HAp products can be produced without too much hassle through the sol–gel approach. Excellent bioactivity for integration into bone tissue is provided by calcium phosphate nanoparticles and nanoplatelets as a consequence of their high surface areas [1–7].

References 1. Choi AH (2022) Biomaterials and bioceramics—part 1: traditional, natural, and nano. In: Choi AH, Ben-Nissan B (eds) Innovative bioceramics in translational medicine I. Springer series in biomaterials science and engineering, vol 17. Springer, Singapore, pp 1–45 2. Choi AH (2022) Biomaterials and bioceramics—part 2: nanocomposites in osseointegration and hard tissue regeneration. In: Choi AH, Ben-Nissan B (eds) Innovative bioceramics in translational medicine I. Springer series in biomaterials science and engineering, vol 17. Springer, Singapore, pp 47–88 3. Choi AH, Ben-Nissan B (2018) Anatomy, modeling and biomaterial fabrication for dental and maxillofacial applications. Bentham Science Publishers, United Arab Emirates 4. Choi AH, Ben-Nissan B (2015) Calcium phosphate nanocoatings and nanocomposites, part I: recent developments and advancements in tissue engineering and bioimaging. Nanomedicine 10:2249–2261 5. Choi AH, Ben-Nissan B (2007) Sol-gel production of bioactive nanocoatings for medical applications: part II: current research and development. Nanomedicine 2:51–61

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3 Calcium Phosphate

6. Ben-Nissan B, Choi AH (2006) Sol-gel production of bioactive nanocoatings for medical applications: part I: an introduction. Nanomedicine 1:311–319 7. Choi AH, Ben-Nissan B, Matinlinna JP et al (2013) Current perspectives: calcium phosphate nanocoatings and nanocomposite coatings in dentistry. J Dent Res 92:853–859 8. Choi AH, Ben-Nissan B (eds) (2019) Marine-derived biomaterials for tissue engineering applications. Springer, Singapore

Chapter 4

Mechanical Integrity of Thin Films and Coatings and Their Clinical Significance

Assessing the amount of stress generated within a thin film or coating due to the deposition technique and the subsequent heat treatments applied is vital in determining its mechanical stability. In addition, all coatings deposited on dental, craniomaxillofacial, and orthopedic implants will experience various amounts of stresses and loads that are applied externally. For that reason, gaining a deeper insight into the combined effects of the strength and fracture resistance of the coating in addition to the properties associated with the coating–substrate interface is vital. The successful applications of any biomaterial-coated implants are primarily determined by the prospect of the coating to crack or spall from inbuilt stresses (regardless if they are compressive or tensile) or due to external mechanical loading. Without question, de-adhesion and fracture of the coating can potentially expose the substrate to accelerated corrosion and wear. Furthermore, coating debris may provoke a negative response within human tissue. Accordingly, integrity must be preserved by these coatings even when subjected to repeated and concentrated applied loads. From a clinical perspective, it is crucial to gain an understanding into the vulnerability of any coating to deformation, fracture, and delamination. Experimental data obtained from mechanical testing such as nanoindentation and microscratch test are essential in depicting the properties of a coating at the submicron level. The information obtained from these tests related to the failure processes and mechanical behavior combined with other appropriate characterization methodologies will allow us to ascertain new knowledge on the materials used for thin films and coatings. The performance and reliability of any coated dental and medical implants and prostheses are governed by the role played by the interfacial adhesion between the substrate material and the thin film or coating. The magnitude of the adhesive force that takes place during the deposition of coating or thin film onto the substrate as well as during the drying and firing process will be regulated by the nature of the coating and substrate surfaces. Typically, these forces can be classified as either secondary

© The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2023 A. H. Choi and B. Ben-Nissan, Calcium Phosphate Nanocoatings for Bone Regeneration, Tissue Repair and Reconstruction, https://doi.org/10.1007/978-981-99-5506-0_4

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bonds (i.e., van der Waals bonding) or primary interatomic bonds (i.e., covalent and ionic bonds) [1–3]. A much higher adhesion is offered by primary interatomic bonds when compared to secondary bonds. This is based on the notion that secondary bonds are based on much weaker physical forces such as hydrogen bonds or dispersion forces. In general, hydrogen bonds occur on the surfaces of polar materials while interfacial dispersion forces are available on all surfaces. The forces holding the two bodies together may perhaps arise from mechanical interlocking, diffusion between the substrate and the coating, or chemical bonding through electrostatic attraction. A combination of one or more of these suggested mechanisms may be involved depending on the physical condition and chemistry of both the surface of the substrate or the material of the coating used [2]. As soon as intimate contact between two different materials is achieved, a new interface is formed at the cost of the two free surfaces. The bond strength created is related to the nature of the interactions observed at the interface between the two materials. A number of reactions can take place on the surface of the coating during any deposition process. Furthermore, a number of factors can influence the bonding and adhesion between the coating and substrate such as wettability and surface topography and interactions. The scales of these interactions will considerably be governed by the capacity of one phase to wet another. Wetting is a condition indispensable for adhesion. The contact angle will determine the capacity for a solid surface to be completely wetted by a liquid coating. Complete wetting will take place if a greater molecular attraction occurs between solid and liquid molecules than between similar liquid molecules. Under this scenario, the contact angle is zero and the liquid is able to spread without any restrictions over a surface. The wetting of a surface can be defined using thermodynamic conditions. The development of interfacial adhesion and bonding is affected by important factors such as the energetics of the surfaces of both the solid coating and the substrate and the surface tension of the coating while it is in the liquid state [2, 4]. In addition to the contact angle, the spreading coefficient will also govern the ability for a liquid to spread and wet a solid. Subsequently, it is imperative to ascertain the surface tension of both the solid surface and the liquid given the fact that a correlation exists between the surface tension and the spreading coefficient as well as determining whether the liquid coating will be able to wet the substrate. In situations where the coating is in the liquid state during deposition, its viscosity is also extremely important, and accordingly, wetting can be considered as an intimate relationship between the substrate and the coating. It is crucial that wetting and intimate bonding remained intact during initial wetting and after coating deposition to attain an adequate adhesion between the substrate and the coating. The mechanisms of adhesion can only become operational once effective wetting between the coating and the substrate is available.

4.1 Determining Stresses in Thin Films and Coatings

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4.1 Determining Stresses in Thin Films and Coatings Ascertaining the magnitudes of stresses developed in coatings due to the deposition process and the temperatures used during sintering are crucial with respect to mechanical stability. Three components contribute to the stresses generated within a thin film or coating [5]: a. Externally applied stresses (σ external ); b. Intrinsic stresses (σ intrinsic ), which are a result of factors such as deposition; and c. Thermal stresses (σ thermal ), which are determined based on the differences in the thermal expansion between the substrate and the coating. The stresses developed within a coating (σ coating ) can be calculated using Eq. (4.1) once the values of any externally applied, thermal, and intrinsic stresses are determined: σcoating = σexternal + σintrinsic + σthermal

(4.1)

Substrate curvature is a simple methodology to determine the amount of stress generated in coatings. One such approach is to use a phase-shifting interferometer. The utilization of optical techniques such as interferometry is advantageous as it is a straightforward technique to use and high accuracy of approximately 6 nm are offered and substrates with curvature of up to 3 µm are acceptable. Another approach is to use a stylus profilometer that is capable of scanning 10 mm in length or more. However, a relatively thin substrate is needed in order to appropriately measure the curvature. Furthermore, this method also requires careful and precise placement of the sample to allow for a cross-scan in both the x- and y-directions before and after coating deposition, preferably on both sides. The relationship concerning the variations in the curvature radius between an uncoated and a coated substrate as well as the resultant stress in the coated substrate can be described using a formula suggested by Stoney (Eq. 4.2) [6]. With this equation, the combined thermal and intrinsic stresses in the coating can be determined without any prior knowledge of the properties of the coating apart from information such as the radius of curvature before (Rbefore ) and after coating deposition (Rafter ), Young’s modulus (E substrate ) and Poisson’s ratio of the substrate (υ substrate ), and the thickness of both the coating (T coating ) and substrate (T substrate ). ] [ 2 E substrate Tsubstrate 1 1 1 σ = − 6 Rafter Rbefore (1 − υsubstrate ) Tcoating

(4.2)

The radius (R) can be determined using the equation below (Eq. 4.3) assuming that the final bow (B) is much less than the length of the substrate (L substrate ): R=

L 2substrate 8B

(4.3)

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Convex deflection is observed in coatings experiencing compressive stress. On the other hand, concave deflection is the consequence of coatings under tensile stress. During coating deposition, measurements can be obtained in situ via the utilization of the deflection technique. Similarly, a cantilever beam along with a low-power laser or interferometry can be used to obtain measurements once the coating has undergone thermal treatment after deposition. Likewise, X-ray diffraction can be employed to detect changes in crystal lattice d-spacing, which in turn provides information related to the elastic strains of coatings. Moreover, this knowledge can be used to calculate stresses in coatings. However, data such as strain-free lattice spacing acquired from X-ray diffraction and Young’s modulus of the coating may not be available in certain scenarios [7].

4.2 Interfacial Adhesion Between a Coating and a Substrate Adhesion, in the context of this book, will be defined as the mechanical bond strength between a substrate and a coating (both micro- and nanocoating). A force must be applied in order to separate the coating from the substrate and used to drive a crack propagating at the interface between the coating and the substrate. This driving force can be through internal residual stresses or from externally applied stresses. The energy required to fracture the coating–substrate interface is defined as the work of adhesion (ωadhesion ). It can be calculated based on the values of the surface energies (γ ) recorded at the coating–substrate interface and of the coating and substrate using the following equation: ωadhesion = γcoating + γsubstrate − γinterface

(4.4)

It is worthy to note that Eq. (4.4) is similar to the “Griffith fracture” where the work of adhesion is identical to the fracture resistance of the interface [8]. On the contrary, the work of adhesion calculated using Eq. (4.4) does not take into consideration variables such as surface roughness, bridging ligaments, asperity contacts, and plastic deformations even though they are extremely significant [9]. For this reason, a more appropriate determination of adhesion can be achieved using the equation based on the practical work of adhesion (ωprac. adhesion ) (Eq. 4.5). The practical work of adhesion is also called the strain energy release rate or interfacial toughness. It can be determined using the amount of energy consumed while plastically deforming the coating (Ωcoating ) and the substrate (Ωsubstrate ): ωprac. adhesion = ωadhesion + Ωcoating + Ωsubstrate

(4.5)

4.4 Enhancing the Adhesion of Nanocoatings to Titanium via Anodization …

17

4.3 Coating Adhesion and Mechanical Interlocking Mechanical keying or interlocking characteristics are possessed by coatings if it penetrates into the surfaces of substrate materials that contain scratches, pores, voids, and crevices during the deposition process. The surface roughness of the substrate material plays an influential role in this mechanism of coating adhesion. Based on observations using a number of surface analytical approaches, it has been demonstrated that the coating can successfully penetrate into undercut and complex tunnel-shaped cracks, which in turn provides a form of mechanical attachment once the coating is set and/or sintered [2]. As mentioned earlier, surface roughness will play a pivotal role in determining the area available at the substrate–coating interface in addition to the force magnitude the substrate is capable of producing to grip the coating within the actual contact area at the interface. Consequently, enlarging the surface area will increase the adhesion of a coating to the substrate. The process of increasing the surface area is only considered advantageous if the coating can completely infiltrate into all of the crevices generated during surface roughening. Conversely, if the coating fails to penetrate fully into the crevices, then fissures at the interface between the substrate and coating can be generated. This will result in the formation of a non-uniform coating, which has a smaller interfacial contact between the coating and substrate compared to the actual available geometric area.

4.4 Enhancing the Adhesion of Nanocoatings to Titanium via Anodization Process As discussed earlier, titanium and its alloys are frequently utilized for dental and orthopedic applications and a variety of approaches ranging from chemical to mechanical are currently used to modify its surface. One such method is anodization, and in essence, it converts the surfaces of titanium alloy to an oxide layer, which is intended to expedite the adhesion of deposited coatings such as nanocoatings onto this layer. As a biomaterial in general, titanium is a metal whose surface is always covered by an oxide layer generated naturally as soon as it is exposed to an environment that contains oxygen such as air and water. This is due to the fact that titanium is classed as an oxide film former based on its electrochemical characteristics and its position within the periodic table. Consequently, it is dependent on the formation of an oxide layer that is technically titanium oxide (TiO2 ). The thickness of this “natural” oxide layer can range between 5 and 70 nm depending on the composition of titanium as well as the production conditions used during its working such as the atmosphere and maximum temperature reached [10, 11]. Furthermore, the presence of alloying or trace elements during the manufacture of titanium can either improve or interfere with this hard adherent oxide layer. This in turn also governs whether reductions

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4 Mechanical Integrity of Thin Films and Coatings and Their Clinical …

or enhancements in corrosion resistance and biocompatibility are witnessed for that particular titanium alloy in comparison to commercially pure titanium [12, 13]. Anodization or anodic oxidation is a well-established approach that can be utilized to create a number of protective oxide layers on metals as well as producing thick oxide layers as porous coatings. It can also be applied to protect the metal substrate against corrosion and minimize metal ion release. The mechanism behind anodization is the diffusion of oxygen and metal ions driven by the combined efforts of electrode reactions and an electric field, which creates an oxide film on the surface of the anode [14]. There are procedures that regulate the anodic oxidation of titanium, and they are very similar to the rules that are applied to other “value” metals. The chemical reactions that can occur during the anodization of titanium are shown below [14]. The first stage involves the creation of an adsorbed layer of oxygen or certain oxygenated species on the surface of the metal to be anodized. To be precise, the oxidized layer is formed on top of a preexisting “natural” oxide layer [15, 16]. • Reaction at the titanium–TiO2 interface: Ti ↔ Ti2+ + 2e− • Reaction at the TiO2 –electrolyte interface: 2H2 O ↔ O2− + 4H+ (reaction between oxygen ions and titanium to form oxide) 2H2 O ↔ O2 (g) + 4H+ + 4e− (oxygen gas evolves at the surface) • Reaction at both interfaces: Ti2+ + 2O2− ↔ TiO2 + 2e− Processing factors such as electrolyte composition, temperature, current, and anode potential can alter the structural and chemical properties of anodic oxides on titanium. Diluted acids of H3 PO4 , H2 SO4 , and acetic acid are some of the examples of the frequently used electrolytes for the anodic oxidation of titanium [14]. The ions of titanium and oxide produced during these redox reactions are forced through the oxide layer once an electric field is applied externally, resulting in the growth of the oxide film. A significant decrease in the applied voltage will be detected across the oxide film at the anode due to the high resistivity behavior of the anodic titanium oxides in comparison with the metallic elements of the electrical circuitry and the electrolyte. The oxide film will continue to develop and the current will flow as long as the electric field is strong enough to drive the ions across the oxide [14]. This is based

4.4 Enhancing the Adhesion of Nanocoatings to Titanium via Anodization …

19

on the theory that an almost linear relationship exists between the thickness of the final oxide film (t oxide ) and the voltage applied (V appl ). α is a growth constant with a value typically between 1.5 and 3.0 nm V−1 (Eq. 4.6). toxide = αVappl

(4.6)

The linear relationship as described in Eq. (4.6) holds true only if the voltage applied is between 100 and 150 V (which is below the dielectric breakdown limit of the oxide) and subjected to the processing environment and the electrolyte applied [14, 17]. On the other hand, if the voltage applied is greater than the breakdown limit during the anodization process, then the oxide will not provide sufficient resistance to prevent current flow and further growth of the oxide. Furthermore, performing the anodization process under such high voltages will result in an increase in gas evolution and frequent sparking. This form of anodization is commonly referred to as spark anodization. More importantly, spark anodization will create oxide films that are less uniform and more porous compared to other anodization processes carried out at voltages below the dielectric breakdown limit [14]. An equation has also been proposed that describes the relationship between the electric field and the anodic current across the oxide film at low anodic potentials [18, 19]: i curr = A exp B · FS

(4.7)

in which A and B are constants, FS is the field strength, and icurr is the ionic current. Observations from a number of studies support the theory that the growth of the anodic film on titanium is due to the transfer of Ti2+ cations through the film, thereby the growth must take place at the oxide surface interface [20, 21]. However, studies have been carried out to dispute this fact and it was suggested that the growth of the oxide film is in fact due to the transfer of oxide ion [22, 23]. In addition, the possibility of ion transfers of both Ti2+ and O2− contributing simultaneously to the growth mechanism has been proposed and this theory is similar to that of oxidation in a gas [24]. It has also been suggested that an anodic layer can be created only if conditions such as the characteristics of the electrolyte support the development of Ti4+ instead of Ti2+ and Ti3+ ions during the creation of the final oxide film [25]. However, the function of the electrolyte during the formation of the anodic film has not been extensively examined. It has also been suggested that the nature of the anions will regulate both the initial passivation and the subsequent growth stages [26–28].

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4.5 Adhesion Theories Based on Chemical Bond, Electrostatic, and Diffusion Chemical bonding is expected to be the strongest and most durable, and it is often possible to form covalent bonds across the interface between the coating and the substrate. There is also the possibility that certain surfaces that contain various chemical functional groups such as previously coated surfaces, composites, and some plastics to generate chemical bonds with the substrate material under ideal conditions. For chemical bonding on the other hand will require the presence of mutually reactive chemical groups attached firmly to both surfaces of coatings and substrates. This process can be accelerated through the use of interfacial shims or additives [4]. It is possible to generate an electrostatic force at the coating–substrate interface in the shape of an electrical double layer. Preserved within the surfaces of both the coating and the substrate are residual electric charges that are dispersed all over the system. Some adhesions of coatings are the results of the interactions between these electric charges [2, 4]. As soon as molecular contact through wetting between the substrate and the coating is achieved, atoms will diffuse across the interface to a certain degree depending on the curing conditions and material properties. This phenomenon is a two-stage process that involves wetting followed by the formation of a chemical bond as a result of elements interdiffusion across the interface [2, 4].

4.6 Characterizing Coatings and Nanocoatings Mechanically Prior to Implantation Characterizing coatings and nanocoatings mechanically will require vigorous and dependable procedures as the demand imposed by their utilization in the biomedical arena, such as dentistry, craniomaxillofacial surgery, and orthopedics, continues to grow. Significant developments have been made in the approach and instrumentation as a consequence of the growing tendency toward the deposition of nanocoatings on tinier devices. This advancement has allowed for the extraction of adhesive and mechanical behavior of coatings prior to implantation. Ascertaining deeper know-hows into the long-term mechanical reliability of biomaterial micro- and nanocoatings is of paramount importance when it comes to their application within the clinical environment. As a result, numerous methodologies are necessary to determine quantitatively the adhesion and mechanical properties of coatings. There have been continuous advancements in the equipment proficient at extracting the adhesion properties of the coating to an underlying substrate [1, 29]. Based on the findings of Nix, the mechanical reliability of both thin films and coatings (macro, micro, and nano) are governed by a number of key factors and the most critical of them all are the substrate roughness, the thickness and geometry of

4.6 Characterizing Coatings and Nanocoatings Mechanically Prior …

21

the coating, the properties at the coating–substrate interface, and residual stresses [30]. Typically, quantifying the coating properties or adhesion strength is a technique frequently used to describe the performance of micro- and nanocoatings deposited onto a substrate material. It is important to determine the adhesion strength of a coating to confirm that it is properly adhered to the substrate material to which it is deposited. The adhesive strength of thin films and coatings can be measured using a number of approaches as indicated by the work of Mittal [31]. Testing methods such as pull-off testing, indentation scratching at increasing loads, and pin-on-disk are the most popular for establishing the bond strength between the coating and substrate. Other methodologies such as microtensile testing are also excellent in providing information concerning the adhesion integrity between coatings and substrates. It is also vital to assess the mechanical properties of thin films and coatings in addition to determining its adhesive behavior. As a result, a number of outstanding techniques were created for this intention and these include microtensile testing, nanoindentation, pull-out test, scratch testing, and bending and bulge testing.

4.6.1 In Situ Microtensile Test In situ microtensile test offers insight into the susceptibility of interfacial delamination between a coating and a substrate through the application of controlled external stresses. Moreover, it is also ideal in providing the properties of both thin and thick coatings on a variety of ductile substrates. Above all, performing tensile testing on coatings using this approach is beneficial as it uses relatively small test samples and the stress field generated is uniform along the gauge length. A tensile coupon is used as substrate during microtensile testing, and a coating is deposited on its surface. A universal testing machine or a specialized device that can be placed under the objective lens of an optical microscope or in a scanning electron microscope is used to pull the coated specimen. After the application of specific strains, damage evolution, cracking, and failure in the coating can be observed in situ or ex situ [29, 32–35]. Useful information into the mechanisms behind material failure can be obtained during loading by examining the damage in situ using optical or scanning electron microscopes [36, 37]. In addition, studying the evolution of cracking and debonding during loading can offer valuable qualitative understandings into the vulnerability of a coating to cracking and de-adhesion. It should be mentioned that the residual stress must be determined using substrate curvature measurements when investigating coating behavior. On the other hand, Young’s modulus of the coating can be verified based on the results obtained from nanoindentation testing [38]. Once it is stressed in the uniaxial direction, brittle coatings on ductile substrates such as calcium phosphate coatings on Ti-6Al-4V can generate parallel cracks normal to the tensile axis in most cases. A decrease in crack spacing is witnesses as the coated substrate elongates because of an increase in the quantity of these cracks developing.

22

4 Mechanical Integrity of Thin Films and Coatings and Their Clinical …

Moreover, the presence of these cracks can also lead to significant delamination of the coating. In contrast, cracking can be irregular and the possibility of the coating to debond from the substrate is drastically reduced for more compliant and softer coatings. The drawback with these semi-brittle and softer coatings is the difficulty centered on their quantitative analysis. Using fracture mechanics, the interface energy, interfacial shear stress, and the strength of the coating can be calculated [39]. Young’s modulus of the coating and the effect of the residual stress (σ residual ) are used to determine the critical stress for cracking (σ cracking ) via Eq. (4.8). The residual stress is the consequence of drying and firing of the coating in addition to the differences in the coefficient of thermal expansion between the coating and substrate. εcracking is the strain immediately recorded when cracking first appears in the coating as soon as the substrate is subjected to a tensile load. σcracking = εcracking E coating + σresidual

(4.8)

The interfacial shear stress (τ interfacial ) can be calculated based on the value of σ cracking determined in Eq. (4.8) as well as the thickness of the coating (t coating ) and the average crack spacing (λ) [32]: τinterfacial =

π tcoating σcracking 1.5λ

(4.9)

The fracture toughness of the coating (K IC ) can be determined using the yield stress of the substrate (σ yield-substrate ) [34], and the function of the elastic contrast determined experimentally between the coating and substrate (F(α D )) [40] as shown in Eq. (4.10): ( K IC =

[

2 σcracking tcoating

π F(αD ) + √

σcracking 3σyield−substrate

])1/2 (4.10)

The adhesion at the coating–substrate interface can be ascertained using the strain (εi ) recorded as soon as detachment of coating buckling is first observed. The apparent interfacial fracture energy (γ interfacial ) or the steady-state strain energy release rate for a phase angle of around 50° as soon as coating delamination occurs is determined using the following equation [41]: γinterfacial = 1/2 tcoating E coating εi2

(4.11)

4.6 Characterizing Coatings and Nanocoatings Mechanically Prior …

23

4.6.2 Shear and Tensile Pull-Off Testing A coated-test sample with a typical diameter of 25 mm is glued to an uncoated coupon using a structural adhesive. A tensile load applied perpendicular to the coating– substrate interface is used during tensile pull-off testing to estimate the adhesive or bonding strength between the coating and substrate [42]. Once the coating is separated from the substrate under a tensile load, the adhesive strength, which is essentially the maximum load over the coated area, can be calculated. In general, at least five coated samples are tested and the test results obtained are used to calculate the averaged adhesive strength. The shear test approach on the other hand is similar but different to the tensile pulloff test in that the load applied to the bonded coating layer is in a parallel direction to the interface. In this test, a relatively strong bonding adhesive is required when shear stress is applied to separate the coating from the substrate [43]. However, these testing techniques have a number of drawbacks and they are related to issues such as the strength of the adhesive used and its consistency during application and the pulling of the coating in a direction parallel or perpendicular to the interface without any misalignment issues. Furthermore, the diffusion of adhesives could occur if the coating is porous, leading to false or inaccurate strength measurements. In addition, the necessity to confirm visually after shear testing to ascertain adhesive failure over cohesive failure is also an issue.

4.6.3 Nanoindentation Testing Considered the starting point by numerous biomedical and dental researchers, nanoindentation testing is the method of choice when it comes to ascertaining the mechanical properties of coating and implants. In addition, it is a simple and effective approach that can be used to obtain vital information such as the hardness and Young’s modulus of thin films and micro- and nanocoatings. Moreover, this technique can also offer a thorough examination of the elastic–plastic response between different coating–substrate combinations such as hard/soft and compliant/rigid coating on hard/soft and compliant/rigid substrate from the loading and unloading graphs. During a normal microindentation testing, a diamond tip with a known geometry, typically Vickers, Knoop, or Rockwell, is used to transmit a load applied to the surface of the material to be investigated. As soon as the load is removed, the area of the residual impression is determined via optical means. The hardness of the material is estimated based on this residual impression. The size of the residual impression in nanoindentation testing on the other hand is often only a few microns in size, and it is extremely challenging to obtain a direct measurement using optical techniques. During nanoindentation testing, a set load in milli-newton is applied to the indenter that is situated on the surface of the test specimen. The depth of penetration in nanometers is recorded as the load is applied. The hardness or mean contact pressure

24

4 Mechanical Integrity of Thin Films and Coatings and Their Clinical …

is calculated by dividing the maximum load over the project contact area, which is determined using the contact depth and the indenter geometry [44]: Hardness =

Maximum Load Project Contact Area

(4.12)

The area of contact at maximum load is determined using the known angle or radius of the indenter in conjunction with the depth of impression. A load–displacement graph is normally used to display the result of a nanoindentation test (Fig. 4.1). Young’s modulus of the test subject can be determined using the gradient of the curve during unloading along with the type of the indenter used (e.g., Berkovich, Brinell, Knoop, or Vickers) and the software based on the model [44– 46]. Gaining accurate measurements of the mechanical properties of coatings such as Young’s modulus or hardness are vital as they can differ from the bulk material. In addition, information such as hardness or contact pressure can also be obtained from the graph. Moreover, different categories of loading and unloading methods can be employed to attain the required properties as a function of penetration depth. A number of factors will govern the accuracy of the results obtained during nanoindentation testing, and these include calibration of equipment, indenter tip shape, preparation of test sample, initial penetration, frame compliance, and corrections for thermal drift [44–46].

Fig. 4.1 Load (in milli-newton) versus displacement (in nm) graph obtained from a nanoindentation test

4.6 Characterizing Coatings and Nanocoatings Mechanically Prior …

25

Young’s modulus can be calculated using the gradient or the slope of the load–displacement curve during unloading and the maximum area (A) recorded at maximum load according to Eq. (4.13). It should be mentioned that this equation is valid for elastic contacts with axis-symmetric indenters such as cylindrical, conical, and spherical punches. Contact Stiffness =

ΔLoad 2 √ =√ E A ΔDisplacement π

(4.13)

Equation (4.14) can be used to determine the combined Young’s modulus of the test sample and the indenter (E combined ). ν indenter and ν sample are Poisson’s ratio for the indenter and the test sample, respectively. 1 E combined

( ) ) ( 2 2 1 − ν sample 1 − νindenter = + E indenter E sample

(4.14)

As previously described by Fischer-Cripps, nanoindentation can also be used to establish the magnitude of coating adhesion and residual stress via direct indentation or transverse scratching [46]. Nanoindenters in particular have also been applied to measure adhesions of coatings to substrates and residual stresses determined from the load that resulted in delamination, and this is taken from the “pop-in” that corresponds to a plateau or discontinuity in the load–displacement graph as shown in Fig. 4.1. Likewise, it is also possible to examine the viscoelastic and creep behavior of soft materials. This is particularly applicable to researchers in dental composites and resins as well as in biological and bone tissue studies.

4.6.4 Bend Delamination Test Three- and four-point bending tests have been used to study the interfacial fracture between coatings on substrates in addition to dissimilar materials. More importantly, it has been demonstrated that determining the interfacial fracture energy of thin films and coatings is possible through the use of three-point bending test [47]. Being a straightforward approach, the test samples used in three-point bending testing can be easily prepared, and for this test to work, it is dependent on the generation of cracks along the coating–substrate interface from the notch via a simple energy balance for the system prior to cracking and debonding. The method also relies on the crucial moment when the crack deflects into and along the interface. Conversely, factors such as the test sample, loading geometry, and strain rate will influence the accuracy of the results obtained. This technique can offer the flexural stress–strain response of the material being investigated. Using Eq. (4.15), the flexural stress (σ flexural ) can be determined using

26

4 Mechanical Integrity of Thin Films and Coatings and Their Clinical …

the load applied at any given point during the test (F), the thickness (t), and width (w) of the test specimen, and the length of the support span (L): σflexural =

3L F 2t 2 w

(4.15)

Four-point bending test on the other hand has become a dependable technique in obtaining the adhesion properties of a coating deposited onto a substrate. First described in the study by Charalambides et al., this technique requires a notch machined into the coating that is deposited on a bend bar [48]. A crack will initiate from the notch as the bending moment increases and propagates to the interface. The condition of constant moment is described as the period when the deflection and propagation of cracks alongside the interface take place at a critical load. For a phase angle of 41°, the strain energy release rate (γ ) can be determined using the following equation: ( )( ) 2 M 2 1 − υsubstrate 1 λ γ = − 2E substrate Isubstrate Icomposite

(4.16)

It should be mentioned that the strain energy release rate is a function related only to the specimen geometry and the critical load for delamination and not associated with the debond crack length. The value of M can be calculated using the width of the test sample (w), the plateau load (F), and the measured distance between the inner and outer loading rollers (L): M=

FL 2w

(4.17)

According to Eqs. (4.18), (4.19), and (4.20), the values of I substrate , I composite , and λ can be determined using the thicknesses (t), Poisson’s ratio, and Young’s modulus of both the coating and substrate: Isubstrate = Icomposite =

3 tcoating

12

3 tsubstrate 12

( )2 3 λtcoating tsubstrate tcoating + tsubstrate λtsubstrate ) ( + + , 12 4 tcoating + λtsubstrate

(4.18)

(4.19)

where λ can be calculated using the following equation: ( ) 2 E substrate 1 − υcoating ( ) λ= 2 E coating 1 − υsubstrate

(4.20)

4.6 Characterizing Coatings and Nanocoatings Mechanically Prior …

27

The interfacial fracture energy as well as determining the rate of strain energy release has been quantitatively carried out using four-point bending test. This is achieved using a technique centered on fracture mechanics to estimate the energy at the interface between dissimilar materials [48]. Typically, the thickness of these substrates is much smaller in comparison with their lateral dimensions. Consequently, their elastic response can be established using simple beam bending mechanics [49]. A crack is initiated from the notch as the bending moment increases and then the crack propagates to the interface. The crack will deflect and propagate along the interface if the interfacial bond is sufficiently weak. The benefit of determining the interfacial fracture energy using this approach rests in the notion that there is no correlation between the energy and the crack length as long as the crack tip is not in close proximity to the pre-crack or the loading points. The interfacial fracture energy can be calculated using the critical load for stable crack propagation (F critical ): ( ) 2 2 1 − υsubstrate 21L 2 Fcritical Interfacial fracture energy = 16w 2 t 3 E substrate

(4.21)

4.6.5 Bulge and Blister Testing For the bulge test to function, a gas or fluid is needed to generate pressure in a coating–substrate system, which is fed through an opening in the substrate. The height of the hemispherical bulge produced in the coating is measured using optical microscopy or an interferometer. Information such as plastic, elastic, and timedependent deformation can be obtained from the pressured used and the deflection height. The blister test on the other hand is somewhat different, in which the pressure is increased until the coating begins to debond from the substrate. Using the radius of the hole (r), the pressure applied (P), the thickness of the coating, and the mechanical properties of the coating such as Young’s modulus and Poisson’s ratio, the interfacial energy can be calculated from the critical pressure for debonding according to Eq. (4.22) [50]:

Interfacial energy =

( ) 2 P 2 r 4 3 − υcoating 3 16tcoating E coating

(4.22)

28

4 Mechanical Integrity of Thin Films and Coatings and Their Clinical …

4.6.6 Scratch Testing Considered as one of the most extensively used and popular approaches for examining the strength of adhesion between the coating and the substrate, scratch testing uses a 200 µm radius metal spherical tipped or a hard diamond indenter to apply continuously an increasing load on the surface of the coating while at the same time the sample is displaced at a constant velocity [41, 51–54]. Coating failure can be determined from the load and distance from which delamination takes place under optical microscopy. Apart from the critical load, information such as the penetration depth, the applied normal force, and the tangential or frictional force can also be obtained from the test. Additionally, failure in the coating can also be determined from the variations in friction or the use of an acoustic emission sensor. Scratching the coating surface will result in an increase in elastic and plastic deformation, and extensive spalling of the coating from the substrate will occur at some critical load, and normally this can be determined using friction force measurements, optical microscopy, or acoustic emission. Once the critical load (F critical ) is determined, the practical work of adhesion (W practical ) can be calculated using data such as the thickness and Young’s modulus of the coating, and the contact radius (a) according to Eq. (4.23) [55, 56]: ( Wpractical =

Fcritical π a2

)2

2tcoating E coating

(4.23)

The residual stress in the coating on the other hand is not taken into consideration in Eq. (4.23). Subsequently, a study has suggested that for purely elastic coatings deposited on stiff substrates, the practical work of adhesion can be described according to the following equation [53]: Wpractical =

σ 2 tcoating 2E coating

(4.24)

The value of σ in Eq. (4.24) is a function in the coating, and it can be ascertained based on the applied stress and the residual stress as shown below: σ = σapplied + σresidual

(4.25)

However, Eq. (4.25) is not completely valid for describing the stresses once a certain amount of plastic deformation occurs even though the equation takes into account the residual stress within the coating. These vital studies were further developed so that the distributions of both elastic stress and residual stress in the coating are taken into consideration. An improved equation was derived as a result that utilizes the average elastic shear (τ ij ) and normal (σ ij ) stresses in the delaminated coating and the shear modulus of the coating (Gcoating ) to calculate the strain energy release rate [57, 58]:

References

29

( )⎤ 2 2 tcoating 1 − υcoating σresidual ⎦ Strain energy release rate = ⎣ 2E coating ( ) ( )⎤ ⎡ 2 2 σi2j tcoating 1 − υcoating ∑ τi2j tcoating 1 − υcoating ⎣ ⎦ + + 2G coating 2E coating ⎡

(4.26)

Scratch testing in general requires great care as a consequence of the complex stress states involved in addition to the wide spectrum of damage process that can occur. Similarly, factors centered on the properties of the coating and the substrate such as thickness, hardness, and roughness of the coating can have an effect on the results of scratch testing. Furthermore, issues such as the tip shape, testing condition, the loading rate, and the scratching velocity will also need to be taken into consideration when performing the test.

4.6.7 Pin-on-Disk Used to determine the wear resistance and friction coefficient of surfaces, the pinon-disk tribometer tester composes of a pin typically a tungsten carbide, ruby, or metal sphere under a static load in contact with a rotating sample. The test can be conducted either in air or in a simulated body environment such as artificial saliva for an extended period to provide information on the degradation of the coating with time [59–68].

References 1. Choi AH, Ben-Nissan B, Bendavid A et al (2016) Mechanical behavior and properties of thin films for biomedical applications. In: Griesser HJ (ed) Thin film coatings for materials and biomedical applications. Woodhead Publishing/Elsevier, The Netherlands, pp 117–141 2. Kendall K (2001) Molecular adhesion and its applications. Kluwer Academic/Plenum Publishers, New York 3. Williams D, Callister J (1994) Materials science and engineering, 3rd edn. Wiley, New York 4. Lee LH (1991) Fundamentals of adhesion. Springer New York, New York 5. Lepienski CM, Pharr GM, Park YJ et al (2004) Factors limiting the measurement of residual stresses in thin films by nanoindentation. Thin Solid Films 447:215–257 6. Stoney GG (1909) The tension of metallic films deposited by electrolysis. Proc R Soc Lond 82:172–175 7. Tsui YC, Doyle C, Clyne TW (1998) Plasma sprayed hydroxyapatite coatings on titanium substrates. Part 1: mechanical properties and residual stress levels. Biomaterials 19:2015–2029 8. Lawn B (1993) Fracture of brittle solids, 2nd edn. Cambridge University Press, Cambridge 9. Lane M (2003) Interface fracture. Annu Rev Mater Res 33:29–54 10. Andreeva VV, Kazarin V (1966) Proceedings of the 3rd international congress on metallic corrosion 1966. Mir, Moscow 11. Andreeva VV, Shishakov NA (1958) Electron diffraction and optic data on the thickness of oxide films on metals. Zh Fiz Khim 32:1671–1672

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12. Davis JR (1998) Metals handbook desk edition, 2nd edn. ASM International, Ohio 13. Schutz RW, Thomas DE (1987) ASM metals handbook of corrosion, vol 13. ASM International, UK 14. Brunette DM, Tengvall P, Textor M et al (2001) Titanium in medicine. In: Engineering materials. Springer, Heidelberg 15. Kossyi G, Nivakoyskii V, Kolotyrkin YA (1969) Zashch Met 5:210 16. Tomashov N, Strukov N (1963) Dakl Akad Nauk SSSR 152:1177 17. Arsov LD (1985) Growth of anodic oxide films on titanium surfaces. In: Contemporary inorganic materials: process in ceramics, metals and composites. 7th German-Yugoslav meeting on engineering materials science and technology, Bad Herrenalb, 22–26 Apr 1985 18. Nakata N, Iida Y (1969) Denki Kagaku Oyobi Kogyo Bulsuri Kagaku 37:366 19. Güntherschulze A, Betz H (1934) The movement of the ion lattices of insulators at extreme electric field strengths. Z Phys 92:367–374 20. Krasilshchikov A (1966) Proceedings of the 3rd international congress on metallic corrosion 1968. Mir, Moscow 21. Hall C, Hackerman N (1953) Charging processes on anodic polarization of titanium. J Phys Chem 57:262–268 22. Dornelas W (1967) Dissertation, University of Paris 23. Tylecote RF (1965) Internal stresses in oxide films during growth. Mem Sci Rev Métall 62:241– 247 24. Aladjem A (1973) Anodic oxidation of titanium and its alloys. J Mater Sci 8:688–704 25. Cotton JB (1966) Proceedings of the 3rd international congress on metallic corrosion 1968. Mir, Moscow 26. Tomashov I, Matveeva T (1971) Zashch Met 7:272 27. Bogoyavlenskii A (1966) Tr Kaz Aviats Inst 90:3 28. Cheseldine DM (1964) Anodic oxidation of titanium in formic acid electrolytes. J Electrochem Soc 111:1005–1007 29. Ben-Nissan B, Choi AH, Bendavid A (2013) Mechanical properties of inorganic biomedical thin films and their corresponding testing methods. Surf Coat Technol 233:39–48 30. Nix DW (2006) Mechanical properties of thin films (class notes for a graduate class at Stanford University). iMechanica 31. Mittal KL (ed) (1995) Adhesion measurement of films and coatings. CRC Press, London 32. Agrawal DC, Raj R (1989) Measurement of the ultimate shear-strength of a metal ceramic interface. Acta Metall 37:1265–1270 33. Agrawal DC, Raj R (1990) Ultimate shear strengths of copper silica and nickel silica interfaces. Mater Sci Eng A 126:125–131 34. Ignat M (1996) Mechanical response of multilayers submitted to in-situ experiments. Key Eng Mater 116–117:279–290 35. Latella BA, Triani G, Zhang Z et al (2007) Enhanced adhesion of atomic layer deposited titania on polycarbonate substrates. Thin Solid Films 515:3138–3145 36. Ignat M, Marieb T, Fukimoto H et al (1999) Mechanical behaviour of submicron multilayers submitted to microtensile experiments. Thin Solid Films 353:201–207 37. Latella BA, Ignat M, Barbé CJ et al (2004) Cracking and decohesion of sol-gel hybrid coatings on metallic substrates. J Sol-Gel Sci Technol 31:143–149 38. Latella BA, Gan BK, Davies KE et al (2006) Titanium nitride/vanadium nitride alloy coatings: mechanical properties and adhesion characteristics. Surf Coat Technol 200:3605–3611 39. Hu MS, Evans AG (1989) The cracking and decohesion of thin films on ductile substrates. Acta Metall 37:917–925 40. Beuth JL, Klingbeil NW (1996) Cracking of thin films bonded to elastic-plastic substrates. J Mech Phys Solids 44:1411–1428 41. Bull S, Berasetegui E (2009) An overview of the potential of quantitative coating adhesion measurement by scratch testing. Tribol Int 39:99–114 42. Cheng K, Ren CB, Weng WJ et al (2009) Bonding strength of fluoridated hydroxyapatite coatings: a comparative study on pull-out and scratch analysis. Thin Solid Films 517:5361–5364

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43. Wei M, Ruys AJ, Swain MV et al (1999) Interfacial bond strength of electrophoretically deposited hydroxyapatite coatings on metals. J Mater Sci Mater Med 10:401–409 44. Field JS, Swain MV (1995) Determining the mechanical properties of small volumes of material from submicrometer spherical indentations. J Mater Res 10:101–112 45. Field JS, Swain MV (1993) A simple predictive model for spherical indentation. J Mater Res 8:297–306 46. Fischer-Cripps AC (2002) Introduction to nanoindentation. Springer, New York 47. Latella BA, Ignat M (2012) Interface fracture surface energy of sol-gel bonded silicon wafers by three-point bending. J Mater Sci Mater Electron 23:8–13 48. Charalambides PG, Lund J, Evans AG et al (1989) A test specimen for determining the fracture resistance of bimaterial interfaces. J Appl Mech 56:77–82 49. Saha R, Nix WD (2002) Effects of the substrate on the determination of thin film mechanical properties by nanoindentation. Acta Mater 50:23–38 50. Bennett SJ, Devries KL, Williams ML (1974) Adhesive fracture mechanics. Int J Fract 10:33–43 51. Benjamin P, Weaver C (1960) Measurement of adhesion of thin films. Proc R Soc Lond A Math Phys Sci 254:163–176 52. Laugier MT (1981) The development of the scratch test technique for the determination of the adhesion of coatings. Thin Solid Films 76:289–294 53. Laugier MT (1984) An energy approach to the adhesion of coatings using the scratch test. Thin Solid Films 117:243–249 54. Burnett PJ, Rickerby DS (1987) The relationship between hardness and scratch adhesion. Thin Solid Films 154:403–416 55. Bull SJ, Rickerby DS, Matthews A et al (1988) The use of scratch adhesion testing for the determination of interfacial adhesion—the importance of frictional drag. Surf Coat Technol 36:503–517 56. Burnett PJ, Rickerby DS (1988) The scratch adhesion test: an elastic-plastic indentation analysis. Thin Solid Films 157:233–254 57. Hutchinson JW, Suo Z (1991) Mixed-mode cracking in layered structures. Adv Appl Mech 29:64 58. Venkataraman SK, Kohlstedt DL, Gerberich WW (1993) Metal-ceramic interfacial fracture resistance using the continuous microscratch technique. Thin Solid Films 223:269–275 59. Santos A, Teixeira J, Fonzar C et al (2023) A tribological investigation of the titanium oxide and calcium phosphate coating electrochemical deposited on titanium. Metals 13:410. https:// doi.org/10.3390/met13020410 60. Osak P, Maszybrocka J, Kubisztal J et al (2022) Effect of amorphous calcium phosphate coatings on tribological properties of titanium grade 4 in protein-free artificial saliva. Biotribology 32:100219. https://doi.org/10.1016/j.biotri.2022.100219 61. Bandyopadhyay A, Shivaram A, Isik M et al (2019) Additively manufactured calcium phosphate reinforced CoCrMo alloy: bio-tribological and biocompatibility evaluation for load-bearing implants. Addit Manuf 28:312–324 62. Rojas C, Vera E, Aperador W (2019) Corrosion resistance hydroxyapatite assessment and tricalcium beta phosphate coating, deposited on stainless steel low carbon vacuum melted. J Phys Conf Ser. https://doi.org/10.1088/1742-6596/1386/1/012022 63. Elghazel A, Taktak R, Elleuch K et al (2018) Mechanical and tribological properties of tricalcium phosphate reinforced with fluorapatite as coating for orthopedic implant. Mater Lett 215:53–57 64. Bansal S, Singh VP (2017) Experimental investigation of plasma sprayed HA (hydroxyapatite)/Al2 O3 /Fe2 O3 composites coated metallic implants. In: 2017 international conference on advances in mechanical, industrial, automation and management systems (AMIAMS), Allahabad, India, 3–5 Feb 2017, pp 314–319 65. Johanna EA, Yesid AC, William AC et al (2015) Tribological behavior of bone against calcium titanate coating in simulated body fluid. Ing Investig Tecnol 16:279–286 66. Navarro CH, Moreno KJ, Arizmendi-Morquecho A et al (2012) Preparation and tribological properties of chitosan/hydroxyapatite composite coatings applied on ultra high molecular weight polyethylene substrate. J Plast Film 28:279–297

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67. Dittrick S, Balla VK, Bose S et al (2011) Wear performance of laser processed tantalum coatings. Mater Sci Eng C Mater Biol Appl 31:1832–1835 68. Gross KA, Babovic M (2002) Influence of abrasion on the surface characteristics of thermally sprayed hydroxyapatite coatings. Biomaterials 23:4731–4737

Chapter 5

Coating Deposition Techniques

For more than forty years, four general conventional industrial coating methodologies have been proposed for the synthesis of bioactive HAp or calcium phosphate coatings destined for use in the clinical arena. Ducheyne and coworkers developed the first technique and the approach utilizes relatively thick calcium phosphate coatings with thicknesses between 100 μm and 2 mm deposited via spray coating to support bone ingrowth [1]. Later, Hench and his colleagues developed the second method centered on the production of thick bioglass coatings with surface bioactivity [2]. Self-assembly through precipitation in a simulated body fluid (SBF) solution was the third approach that was developed by Kokubo et al., and this method was applied and utilized at a later stage to examine the biocompatibility of various new materials instead of being used as a coating [3]. Around the same period, similar technique was also created by De Groot and coworkers. The fourth approach developed by Ben-Nissan et al. involves dipping the substrate into a calcium phosphate solution derived from the sol–gel method to create strong single or multilayered nanocoatings with thicknesses between 70 and 100 nm [4–9]. At the moment, techniques such as electrophoretic deposition, thermal spraying, pulsed laser deposition, physical and chemical vapor deposition, and sol–gel have been used to deposit ceramic coatings such as calcium phosphate onto a variety of substrates including titanium and its alloys. However, each of these deposition techniques has its own benefits and drawbacks that prevent them from being the ideal coating system. The primary technique used today for the deposition of calcium phosphate-based coatings on dental and medical implants and prostheses is plasma or thermal spraying. In spite of its widespread utilization, this method has had serious shortcomings such as the coatings deposited are typically non-uniform, highly porous, comparatively thick and contain amorphous phases. Above all, coatings deposited onto metallic surfaces will develop inadequate bonding strength. Moreover, the dissociation of HAp to other phases such as β-tricalcium phosphate and calcium oxide as a consequence of the high-temperature processing method is well known. In comparison with HAp, these phases have much faster dissolution rates © The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2023 A. H. Choi and B. Ben-Nissan, Calcium Phosphate Nanocoatings for Bone Regeneration, Tissue Repair and Reconstruction, https://doi.org/10.1007/978-981-99-5506-0_5

33

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5 Coating Deposition Techniques

and create issues within the physiological environment. Through the application of post-treatments, modifications in the composition, and the utilization of new and improved coating techniques have enabled plasma spraying to be used successfully to coat both dental and orthopedic implants and prostheses.

5.1 Coatings Produced by Plasma Spraying As mentioned above, the most widely used technique for the deposition of calcium phosphate coatings is plasma spraying. This approach is capable of depositing coatings that are microns to millimeters thick. A gas plasma is generated using direct current arc or from other sources such as radiofrequency (RF). The plasma is used to heat the powder into a partially liquid form and it is propelled toward the substrate. The expansion of rapid gas induces speeds of up to 800 m/s. Plasma spraying can be conducted in an ambient or controlled atmosphere as well as under vacuum conditions. However, the high temperature used throughout plasma spraying can expose the substrate material to intense heat and this could lead to the creation of residual thermal stresses in the coatings. RF magnetron sputtering utilizes plasma generated by an RF field between the sputtering target and the substrate holder. A magnetic configuration is present in magnetron sputtering below the target trapping electrons in close proximity to the target [10]. Trapped electrons will lead to the ionization in the sputtering gas. Plasma sheath can be described as a drop in voltage that causes fraction of the ions to speed up toward the target and is present between the plasma and the target. The expulsion of target species and secondary electrons is due to the bombarding ions with the secondary electrons maintaining the discharge. The atoms that traveled from the target and bonded with the substrate material will determine the thickness of the coating or thin film. Depending on the desired performance of the coating and the materials used, various processing routes such as flame spraying, high-velocity oxyfuel spraying, detonation flame spraying, on top of plasma spraying can be employed. In recent times, suspension and solution thermal spraying have provided a more economical approach when it comes to the production of thin calcium phosphate coatings. Yet, the coatings produced by suspension plasma spraying are more porous in comparison to the coatings produced by powder plasma spraying and this porosity resulted in a decrease in Young’s modulus and hardness of the bulk coating. This fact was disputed in a study in which higher mechanical properties were recorded during site-specific indentations on dense areas in the coating produced by suspension plasma spraying, suggesting it might be due to factors such as finer grain size and crystal orientation [11]. One of the major concerns for cementless hip and knee and dental prostheses has been the adhesion of calcium phosphate coatings deposited by plasma spraying. Factors such as the integrity of the coating–substrate interface, residual stresses, crack population, and size and distribution of pores will govern the adhesion properties of

5.2 Comparison Between Coatings Produced by Plasma Spraying …

35

thermally sprayed coatings. Furthermore, the strength of adhesion can be affected by several issues and a number of them are centered on the powder characteristics, spray variables such as spray parameters, and substrate preparation [12]. A restriction was placed by the US Food and Drug Administration on the production of calcium phosphate-coated implants deposited via plasma spraying for applications in dentistry during the early years due to issues concerning the mechanical degeneration and dissolution of the splats. Nonetheless, the concerns regarding phase change and ultimately the dissolution were lessened through improvements in the coating deposition technique, application of appropriate heat treatment processes, and optimization and utilization of various chemistries such as fluoro-HAp. A mixture of crystalline and amorphous phases is created by thermal spraying, and this causes variable solubility that is governed by either the quantity of the amorphous phase or the dissolving phosphate phases. The thickness of calcium phosphate coatings deposited using plasma or thermal spraying on a commercial scale is between 30 and 100 μm. Adhesion and bone growth can be initiated easily due to the thickness and chemistry of the coating.

5.2 Comparison Between Coatings Produced by Plasma Spraying and Sol–Gel Technique As discussed above, plasma spraying is without question the most extensively used approach for the deposition of calcium phosphate macro- and microcoatings, and it has been comprehensively applied on a commercial level as the principal coating process in the production of dental and orthopedic implants and prostheses. Even though all the concerns centered on plasma spraying described earlier have been addressed, the adhesion of the splats produced during plasma spraying is still a key concern regarding the clinical applications of implants coated with calcium phosphate deposited using this approach. Another concern is the dissociation of calcium phosphate to calcium oxide as well as other phosphate phases during the deposition process. It is well known that the solubilities of these materials are more rapid within the physiological environment. The distribution and size of the pores created, defect population, residual stresses, and the integrity of the interface between the coating and the substrate are also vital and critical for long-term clinical stability. Moreover, a number of variables such as substrate preparation, the characteristics of the powders used, and the spray parameters can also have an influence on the strength of adhesion [1, 12–14]. Using mechanical examination techniques such as microtensile and push- and pull-off testing, the strength of adhesion of macro- and microcoatings deposited using plasma spraying have been analyzed quantitatively and the test results revealed that the processing conditions will regulate the adhesive bond strength, which can vary between 5 and 25 MPa [15, 16]. Based on the findings from several investigations, simply increasing the plate power can produce an increase in the bond strength.

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Conversely, a reduction in the bond strength is the consequence of increasing working distance. A maximum adhesive bone strength of 25 MPa was recorded when a plate power of 28 kW, and a working distance of 90 mm was used. Failures at the fracture surfaces were classified at the completion of a pull-off examination as either adhesive failure (delamination at the coating–substrate interface) or cohesive failure (delamination within the coating) [15]. In contrast, the ability for nanocoatings to exhibit superior hardness, bioactivity, and strength has been well-documented due to grain sizes fall within the nanometer range [4–9, 17]. The thickness of new generation nanocoatings produced by the sol– gel technique ranged from 70 to 200 nm, but more importantly, a permanent mechanical and chemical bonding can be created simply by depositing these nanocoatings onto macro- and microtextured surfaces. It should be mentioned that due to the thicknesses of these nanocoatings, bone mechanical interlocking could not be created compared to coatings deposited by plasma spraying. However, faster healing and osseointegration are the consequences of the increased surface area owing to nanostructured grains [4–9]. Extensive studies were conducted to establish exactly the ways in which the adhesion of coatings derived from sol–gel can be enhanced. The application of a TiO2 interlayer has been suggested as a possible option to increase the adhesion of calcium phosphate coatings to titanium substrates [18–20]. The selection of appropriate heat treatment temperature is another crucial factor that will influence the mechanical characteristic of calcium phosphate coatings [21]. Another approach that can be applied to enhance the bonding strength as well as to stabilize the coating is alkali treatment. It has been hypothesized that an improvement in the bond strength on the surfaces of titanium treated with sodium hydroxide could be the consequence of factors such as the formation of a thin sodium titanium oxide layer and high surface roughness [22].

5.3 The Sol–Gel Deposition Approach The production of bioceramic materials such as blocks and thin films can be achieved using a solution-based synthesis approach known as sol–gel. This chemical route can also be used to synthesize ceramic coatings such as calcium phosphate. Furthermore, objects with complex shapes can also be coated using this approach. The uniqueness of this methodology rests in its ability to generate powders, monoliths, fibers, coatings, and platelets of the same composition simply by changing the processing parameters such as chemistry and viscosity of a given solution [4–9]. Moreover, a variety of compositions can be synthesized such as single and mixed oxides and nonoxides such as borides, chlorides, and nitrides. As mentioned previously, coatings produced using this technique have been shown to display increased osseointegration and mechanical properties due to nanocrystalline grain structure. The history of the sol–gel approach dates back to the beginning of chemistry when it was first recognized in 1846 as an application technology after the hydrolysis and polycondensation of tetraethyl orthosilicate, also known as tetraethoxysilane, was

5.3 The Sol–Gel Deposition Approach

37

observed in the study by Ebelmen [23]. In 1939, the first patent on sol–gel was published covering the synthesis of SiO2 and TiO2 coatings [24]. The potential of the sol–gel approach was acknowledged during the fabrication of high-purity glasses as traditional ceramic processing techniques were not sufficient in the mid-1950s, and as a result, the notion of using sol–gel to produce multicomponent and homogeneous glasses was born [25]. By definition, a sol is essentially a suspension of colloidal particles within a liquid [26], and the main difference between a sol and a solution is that a sol is a two-phase solid–liquid system, whereas a solution is a single-phase system. The sizes of these colloidal particles can range between 1 and 1000 nm and, as a result, gravitational forces can be neglected. Instead, their interactions will be regulated by short-range forces such as van der Waals forces as well as surface charges. Diffusion of the colloids by Brownian motion contributed to the stability of the sol system, which results in a low-energy arrangement [27]. Decreasing the surface charges of the system further enhances the stability of the sol particles. A noticeable reduction in the surface charges will cause gelation, and the subsequent product can retain its shape without the assistance of a cast. A gel is composed of a network or solid skeleton that is able to encapsulate a liquid phase or excess of solvents, and subsequently, it can be thought of as a composite. The mechanical and physical behavior of the gel will be governed by its chemistry. Soft gels that have a low Young’s modulus can be synthesized simply by regulating the polymerization of the hydrolyzed starting compounds. The production of a high molecular weight polymeric gel is the outcome of the creation of a three-dimensional network, and it can be considered as a macroscopic molecule spreading all over the solution. The time required for the last bond to develop within this network is commonly referred to as the gelation point. Through the use of this gelation, nanosized coatings or nanostructured powders and monolith can be synthesized using the procedure applied [4–9]. The sol–gel approach can be categorized as aqueous- or alcohol-based. For aqueous-based approach, it is normally carried out in the presence of water. On the other hand, the presence of water is eradicated for the alcohol-based system until the hydrolysis stage. Moreover, there are non-hydrolytics, which are essentially sol–gel routes carried out in the absence of solvents. Likewise, the precursors utilized in sol–gel can be classified into alkoxides and non-alkoxides (Fig. 5.1). Due to their volatility, alkoxides are ideal as precursors for sol–gel synthesis. In addition, metal salts such as the chemical elements in Groups 1 and 2 of the periodic table can also be used as precursors. The alkoxides created from these elements are non-volatile solids with a low solubility in many cases [27]. The use of solvents, typically organic alcohols, is needed in the manufacture of sol solutions (Fig. 5.1). The main purpose of utilizing solvents is to dilute liquid precursors and to decrease the effect of concentration gradients. Above all, they are used to dissolve solid precursors. The type of solvents used will determine factors such as the crystallization temperature [28] and particle morphology [29]. A large amount of organic material is needed during the production of materials such as powders and coatings using the sol–gel approach. Cracking as a result can

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5 Coating Deposition Techniques

Fig. 5.1 Synthesis of a thin-film ceramic coating via the sol–gel approach. Reprint with permission from [4]

5.3 The Sol–Gel Deposition Approach

39

be an issue throughout the production stages caused by factors such as fast drying. Furthermore, the occurrence of shrinkage during drying is common when it comes to the synthesis of monoliths. On the other hand, cracking in thicker coatings is often the result of problems such as separation of phases, and inhomogeneities due to the thermal variations between the coating and substrate materials in addition to the drying process. Precursors are utilized to create a solution, which is then applied to produce a sol. Coatings are synthesized from these solutions using a number of deposition approaches. As soon as the coated substrate is exposed to water, then hydrolysis will take place and during this stage, a three-dimensional network is formed once the formation of hydrated oxides or hydroxides is accomplished and the gelation process concluded. Excellent examples are oxide coatings such as zirconia and alumina. The process of gelation can be defined as the growth of clusters through aggregation of particles up to a stage where the clusters collide or polymer condensation. The gel point is determined by the increase in viscosity given the fact that no latent heat is released. The next step in the sol–gel approach involves drying and firing the synthesized gel. Without question, the most crucial stage in the manufacture of monoliths via this technique is drying, and the process involves removing excess solvents from the pore network. A significant quantity of shrinkage will take place during the drying and firing (or sintering) of a monolith, and the magnitude is dependent on the moisture content [30]. The appearance of cracking, as discussed earlier, can be an issue that is caused by large capillary stresses created if the pores are small, say less than 20 nm. The issue of cracking can also stem from the loss of large quantities of materials and therefore the subsequent shrinkage that occurs [2]. The utilization of drying control chemical additives (DCCA), regulating the drying conditions, and increasing the amount of cross-linking can reduce the likelihood of shrinkage (Fig. 5.1). In contrast, drying of sol–gel-derived nanocoatings is not much of a problem in comparison with monoliths. This is based on the notion that all of that shrinkage within the nanocoatings will take place in the direction perpendicular to the substrate as the nanocoatings are constrained in the plane of the substrate [31]. In an effort to minimize the possibility for cracking, nanocoatings with thicknesses between 400 and 500 nm can be dried under controlled rates. On the other hand, avoiding cracking in nanocoatings with thicknesses greater than 500 nm will become difficult [21, 32–34]. Equation (5.1) can be used to calculate the critical thickness for cracking or debonding to occur in coatings or films (t coating ) [35]: tcoating =

2E coating γ , 2 σcoating

(5.1)

where γ is the fracture energy required to generate a new surface, E coating is Young’s modulus (plain strain) of the coating, and σ coating is the stress in the coating.

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5 Coating Deposition Techniques

The firing or sintering process is the final step in the synthesis of ceramic materials via the sol–gel technique. Typically, a temperature of around 600 °C is used to heat the dried gels and any remaining organic materials are combusted during this process. Sintering temperatures, due to the small grain sizes obtained using the sol– gel approach, are between 400 and 800 °C. Some specific chemistry permits the use of a much lower temperature, and it can be as low as 90 °C.

5.3.1 The Alkoxide Route Various sol–gel routes have been employed for the production of synthetic calcium phosphate coatings since the early 1990s. As discussed earlier, Ben-Nissan and coworkers introduced the first pure calcium phosphate nanocoatings in the late 1980s and the route was based on the work of Masuda et al. on the synthesis of HAp powders (Fig. 5.2) [36]. Changes in both the technique and chemistry have enabled this approach to synthesize a variety of nanocoatings with the highest homogeneity and purity. The precursor solution is produced through the amalgamation of calcium diethoxide (Ca(OEt)2 ) or calcium acetate monohydrate (Ca(OAc)2 ·H2 O) with ethylene glycol. As soon as the dissolution of the calcium precursors is complete, diethyl hydrogen-phosphonate ((C2 H5 O)2 P(O)H) is added until a stoichiometric ratio between calcium and phosphorus of 1.67 is achieved. Coatings are deposited using this synthesized solution, followed by the application of heat treatment at low temperatures [4–9]. Two coating techniques have been exploited to deposit calcium phosphate onto a variety of biomaterials such as ceramics, glass, 316L stainless steel, cobalt-chromium alloys, titanium alloys, as well as a number of calcium phosphates derived from marine structures such as foraminifera and coral for slow drug-delivery and bone graft applications [4–9].

5.3.1.1

Spin Coating

Used preferably on flat shapes such as disks and plates, spin coating is one of the deposition techniques that can be employed to coat surfaces with sol–gel-derived calcium phosphate materials. The process can be divided into four stages. The first stage involves the delivery of excess quantities of coating materials to the surface of the substrate material while it is stationary or spinning. As soon as the substrate is spinning at maximum speed, the rotational movement will cause the coating material to move radially outwards due to the presence of centrifugal force. The third stage involves separating excess material in the form of droplets once it drifts to the perimeter of the substrate. Similar to the dip coating process described later, drainage can take place if a hole is available at the center of the substrate. It is important to note that evaporation can occur at any given time throughout the entire course of the deposition process. The presence of edge effects as a consequence of surface tension

5.3 The Sol–Gel Deposition Approach

Fig. 5.2 Synthesis of calcium phosphate coatings using the alkoxide route

41

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5 Coating Deposition Techniques

causes the coating or film to thin down to a thickness that is relatively uniform except for the edges. Increases in drag forces lead to a reduction in the fluid flow, which in turn causes the thickness of the coating to decrease. Thinner coatings can be more affected by evaporation, and it becomes a critical step as soon as the spinning has ceased. This will also result in an increase in the viscosity of the solution by concentrating the quantity of non-volatiles available. Obtaining coatings with uniform thicknesses is possible if the viscosity of the coating material is insensitive to shear rate (Newtonian behavior). Once a balance between centrifugal forces and viscous forces is reached, then this uniformity can take place. According to the study by Scriven [37], the following equation can be used to calculate the final thickness of coatings deposited using spin coating technique:  tcoating =

3eη 2ωV0

1/3 

 V , V0

(5.2)

where e is the rate of evaporation, η is the viscosity of the solution, ω is the angular velocity, V 0 is the initial volume fraction of solvent, and V is the volume fraction of solvent in the coating. According to Eq. (5.2), the thickness of the coating deposited using spin coating can be regulated by fine-tuning the solids content and viscosity of the solution in addition to the spin speed used. Heterogeneities of the coatings can be induced under certain scenarios if the metallic substrate used is oxidized at different thicknesses. During sintering, these areas might cause cracking or spalling of the coating.

5.3.1.2

Dip Coating

Dip coating is another approach that can be used to deposit sol–gel-derived calcium phosphate materials. In general, dip coating is used as a batch process and is mainly applied to deposit coatings on flat glass substrates. However, it is also capable of coating complex-shaped objects such as fibers, pipes, rods, and tubes [37]. The overall process of dip coating can be divided into five stages: immersion, start-up, deposition, drainage, and evaporation. In most cases, the first two stages always take place in sequence. Stages three and four can occur in tandem during the entire coating process unless essential preventative measures are taken. For any given solution, the resultant thickness of a single-layered coating will be governed by the withdrawal speed. Conversely, the thickness of the coating produced can be altered by changing the number of layers deposited as well as modifying the viscosity of the solution [38–40]. Equation (5.3) can be used to calculate the thickness of coating or film deposited onto a substrate using the dip coating approach [37]: 

tcoating

Uη =c gρ

1/2 ,

(5.3)

5.4 Other Coating Deposition Techniques

43

where η is the viscosity of the sol solution, U is the speed of withdrawal, g is the acceleration due to gravity, ρ is the density of the coating material, and c is a constant and the value is approximately 0.8 for Newtonian liquids. There will be variations in the coating thickness between the center and edge of the substrate once it has been dip-coated, and this in turn will lead to the generation of cracks during the firing or sintering process. Thicker coatings can be produced using a quicker pullout rate. This is due to the fact that less drainage and evaporation can take place during pullout. Provided that all other variables are suitable, crackfree coatings with thicknesses ranging from 70 to 100 nm can be manufactured by controlling the pullout rate in an appropriate manner. As discussed earlier, another advantage possessed by the dip coating approach is its ability to deposit multilayered coatings with a high uniformity and thickness of up to 1000 nm. Furthermore, this approach is used on a commercial scale to deposit multilayered coatings with certain optical characteristics [34].

5.4 Other Coating Deposition Techniques 5.4.1 Pulsed Laser Deposition Pulsed laser deposition (PLD) is a versatile and simple deposition approach to fabricate calcium phosphate coatings. The process of PLD involves evaporating the surface of the target using a focused laser beam under vacuum or different gas atmosphere. The material vaporized is then deposited onto a parallel substrate [41]. A number of studies have been attempted to examine the effects of pulsed laserdeposited coatings on bone regeneration using both in vitro cell studies and in vivo implantations [42–49]. Furthermore, it has also been reported that PLD is more reliable and efficient due to its ability to generate adherent and crystalline films [50–53]. Like all coating processes, it is vital to examine factors that will influence the bioactivity of the deposited coating. In 2006, a study was attempted to show the correlation between the substrate temperature and the growth rate and composition of the deposited coating. Observations from their study revealed a reduction in the mass of the calcium phosphate coating with increasing substrate temperature to 700 °C from room temperature. More importantly, the Ca/P ratio was also found to increase gradually with increasing temperature. The authors postulated that this was due to the amount of phosphorus sublimated from the growing coating increasing more rapidly in comparison with the quantity of calcium that sublimated [54]. Observations from another study also suggested that the use of high substrate temperature could promote the oxidation of Ti-6Al-4V surfaces prior to the growth of calcium phosphate coatings despite the fact that the authors managed to obtain crystalline nanocoatings with thicknesses up to 1000 nm using substrate temperatures of 400 and 500 °C [55]. In this regard, studies have been carried out to determine

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5 Coating Deposition Techniques

the deposition parameters needed to obtain crystalline calcium phosphate coatings using PLD [54, 56]. A relationship between laser fluence and the crystallinity of the deposited calcium phosphate nanocoatings was postulated and the results suggested that the as-deposited coatings became more crystallized and denser once higher laser fluences were used. In addition, higher laser fluences also reduced the Ca/P ratio of the as-deposited coating [56].

5.4.2 Physical and Chemical Vapor Depositions Regarded as a versatile synthesis method, physical vapor deposition (PVD) is capable of producing thin-film materials with structural control at the nanometer scale and this can be accomplished by observing the processing parameters in a meticulous manner. The process of PVD involves the creation of vapor phase species through laser ablation, sputtering, ion beam, or evaporation. In many industrial sectors, PVD has been extensively used for the deposition of thin films. State-of-the-art magnetron sputtering processes permit the synthesis of metals, alloys, ceramics, and polymer thin films onto a wide range of substrate materials. As a result, PVD has found use in the biomedical arena for coating applications. There is a growing demand for coatings with enhanced and tailored properties such as corrosion resistant, high hardness, wear, and often, complex combinations of those properties are required. In spite of its versatility, coating non-planar substrates uniformly is not possible in a majority of PVD-based processing approaches due to the creation of vapor atoms in high vacuum resulting in nearly collisionless vapor transport to the substrate, and for that reason, only areas in the line-of-sight of the vapor source are coated. However, this issue can be resolved through the use of multiple, spatially distributed sources or sophisticated rotational and/or translational movement of the substrate. Chemical vapor deposition (CVD) is another form of vapor deposition, and this atomistic deposition approach can be used to coat complex-shaped biomedical prostheses as well as in the manufacture of composites and nanodevices due to its unique non-line-of-sight deposition capacity. More importantly, CVD can create highly pure materials with structural control at nanoscale or atomic levels [57], and for that reason, a number of studies have been carried out to examine the possibility of using CVD to deposit calcium phosphate coatings for biomedical applications [58–66]. CVD is also capable of synthesizing single-, multi-, functionally graded, nanostructured, and composite coating materials with unique structure and well-controlled dimensions at low processing temperatures. The procedure for fabricating coatings involves the chemical reactions of gaseous reactants on or in close proximity to a heated substrate surface.

5.4 Other Coating Deposition Techniques

45

5.4.3 Electrodeposition Electrochemical or electrophoretic deposition holds a number of advantages including the ability to deposit coatings and thin films on substrates that are porous or complex in shape, and the processing temperature is relatively low in comparison with plasma spraying. Furthermore, the properties of the coating can be easily controlled [67]. As a result, electrodeposition has appealed to biomedical researchers for the deposition of calcium phosphate coatings. Furthermore, these coated substrates were investigated in vitro and observations from these studies revealed electrodeposited coatings were able to support the attachment, proliferation, and osteogenic differentiation of osteoblastic-like cells, human periosteum-derived cells (hPDCs), and bone marrow-derived mesenchymal stromal cells (BSMCs) [68–81]. Evidence from early studies has suggested the unique characteristics of electrochemically deposited calcium phosphate coatings are responsible for its good biological performance and gaining an insight into that structure may lead to further fine-tuning of the deposition parameters and ultimately the fabrication of a coating with even higher quality and better clinical performance such as those required in dentistry and orthopedics [82]. In addition, findings from in vivo studies also suggested that electrodeposited coatings are also able of supporting and sustaining bone growth [70, 76–78, 80, 83]. In vivo comparative studies using animal models and human clinical trial have also been carried out in an effort to determine if there are any potential variations in the degree of osseointegration between implants and scaffolds coated with calcium phosphate deposited using electrodeposition and conventional plasma spraying technique [77, 80, 83]. Observations from one study found the magnitude of bone-implant contact ratio was higher for electrochemically deposited coatings with a plate-like nanostructure than plasma-sprayed nanotopography-free coatings after six weeks post-implantation and the authors postulated that it might be related to the early promotion of osteogenic expression and differentiation induced by the nanoplatelike surface. However, new bone formation and the bone-implant interfacial strength were higher for plasma-sprayed implants than electrochemically deposited implants after 12 weeks post-implantation [77]. Similar findings were made in a later study where calcium phosphate coatings were synthesized on 3D-printed Ti scaffolds using electrochemical deposition and plasma spraying [80]. Coatings produced by electrochemical deposition appeared to be lamellar or plate-like while plasma spraying created coatings that were relatively smooth. Micro-CT revealed the amount of new bone tissues formed within and around the plasma-sprayed scaffolds was greater than electrochemical-deposited scaffolds at four- and twelve-weeks post-implantation, suggesting plasma-sprayed scaffolds have a better ability to promote bone regeneration and integration. Additionally, observations from bone mineral density, bone volume fraction, and histological analyses also supported the notion that plasmasprayed scaffolds have better ability to promote bone regeneration at an early and late stage as well as increase the formation of mineralized matrix faster and further accelerate bone defect repair than electrochemical-deposited scaffolds.

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5 Coating Deposition Techniques

In 2016, a randomized controlled clinical trial was carried to examine whether there are any significant variations in medium-term stability and periprosthetic bone remodeling can be detected in cementless femoral stem coated with calcium phosphate deposited using either electrochemical deposition or conventional plasma spraying [83]. The authors compared both types of calcium phosphate coating in a clinical setting and did not discover any obvious differences that would affect clinical performance. They also stated that even though there was a statistically significantly higher degree of retroversion in the electrochemically deposited stems during the initial period, both groups remained completely stable during the late period. In addition, both coatings appear to provide the necessary stability to obtain secondary fixation through bone ingrowth or on-growth.

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Chapter 6

Cellular Responses

Studies based on animal models and human trials have demonstrated that the deposition of a thin HAp and other calcium phosphate coatings on surfaces of implants accelerated early bone formation as well as an increase in bone strength at the interface between the implant and bone tissue [1–10]. The significances of calcium phosphate coatings on osseointegration were first highlighted during the 1990s using in vivo animal studies [1, 2]. Clinical observations from the study by Kohri et al. showed calcium phosphate-coated implants were adherent to the surrounding bone tissue after four to six months post-implantation, while uncoated titanium implants were easily removed from the bone. Furthermore, high-magnification scanning electron micrographs showed that a portion of the interface between the calcium phosphatecoated implant and the bone displayed no void; while on the other hand, uncoated surfaces generated gaps at the bone–implant interface [1]. Later, a study demonstrated implants coated with calcium phosphate was found to perform better in mechanical pushout tests (used to determine the bone–implant interface strength) as well as the amount of bone coverage on the surface of the implant in comparison with other surface-modified implants such as blasting and high-temperature acid etching, indicating superior osseointegration after 12 weeks in situ [2]. As discussed earlier, implants coated with nanocalcium phosphate demonstrate enhancements in bone attachment and heal faster due to the nanostructured grains only if the coating is manufactured in an acceptable manner [11–17]. Observations carried out at four weeks post-implantation further reinforced the use of nanostructured calcium phosphate coatings can create significantly higher bone-to-implant contact compared to uncoated surfaces after implantation into the femur of an in vivo rabbit model [7]. Furthermore, gene expression and conventional removal torque assessment were compared in tissues around coated and uncoated implants using real-time reverse transcription in a rabbit model. Significant upregulation in alkaline phosphatase and osteoprogenitor activity was observed after two weeks postimplantation. Moreover, a notable improvement in the tissue quality was recorded around the coated implants at three weeks using nanoindentation testing. Progressive © The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2023 A. H. Choi and B. Ben-Nissan, Calcium Phosphate Nanocoatings for Bone Regeneration, Tissue Repair and Reconstruction, https://doi.org/10.1007/978-981-99-5506-0_6

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mineralization of the bone encapsulating the coated implants was observed after four weeks post-implantation based on the discovery of osteocalcin. Additionally, adenosine triphosphatase was significantly higher for the coated implants demonstrating steady resorption of the calcium phosphate coating [8]. Similar results were observed in another study where it was discovered that the interfacial area or the healing zone between nanostructured calcium phosphate coating and mature bone was rich in collagen and mesenchymal stem cells after two weeks post-implantation into a rat femur, which differentiated into osteoblast phenotype and developed non-mineralized bone osteoid signifying the beginning of bone regeneration [9]. Moreover, the deposition of nanometer-sized particles of calcium phosphate has been suggested to be beneficial in reducing the duration of the healing phase and subsequently providing earlier fixation and loading protocols after implantation based on the histomorphometric and histologic observations of new bone formation around the surfaces of dental implants treated with calcium phosphate nanoparticles. After two months post-implantation and healing in the posterior maxilla, firm contact of peri-implant bone to the surfaces of nanocalcium phosphate treated implants were observed in addition to enhanced adaptation to the threads compared to implants without the surface treatment. Moreover, 3D reconstruction of sections obtained using confocal laser scanning microscopy showed the intimacy of the contact between bone and treated surface through the entire thickness of the implants [6]. Aside from the influence of surface chemistry and surface morphology, a study also suggests that the different dissolution and reprecipitation properties of calcium phosphate micro- and nanocoatings with various degrees of crystallinity may enhance early bonding and bone formation [4]. This hypothesis was emphasized by the findings from an in vitro cellular study in which low crystalline, nanostructured calcium phosphate coatings were able to stimulate the attachment and proliferation of NCTC murine fibroblasts and primary dental pulp stem cells [10]. Furthermore, the influence of coating properties such as surface roughness and crystallinity on the responses of human osteosarcoma HOS TE85 cells were investigated in vitro and the finding revealed the cells attached and proliferated well and expressed alkaline phosphatase and osteocalcin to a greater degree on coatings with higher crystallinity when compared to the poorly crystallized coatings. In addition, cell attachment was enhanced on rough coatings but similar results on the alkaline phosphatase and osteocalcin expression levels were found between rough coatings and smooth coatings [3]. The bioactivity and cellular responses of coatings can also be improved by substituting or doping calcium phosphate with trace elements or therapeutic ions [18, 19]. It has been suggested that ionic substitutions show promise as an alternative to conventional calcium phosphate coatings by providing additional properties such as osteoinductivity or antibiotic activity. Consequently, the deposition of coatings doped with therapeutic ions such as strontium, magnesium, or silicon could offer improvements in biomechanical fixation and osseointegration in patients with skeletal disorder such as osteoporosis [19, 20].

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53

As a biomaterial, magnesium has been doped into a number of bioceramics including tricalcium phosphate and HAp and it was postulated that its utilization could change the adsorption of cell adhesion and tissue regeneration proteins [21, 22]. In addition, it has been suggested that doping calcium phosphate with a small quantity of magnesium could potentially decrease its solubility in vitro as well as provide improvements in cell attachment and growth based on osteoblast culture studies [23]. Similar observations were made in a later study where MC3T3-E1 cells cultured on magnesium-substituted calcium phosphate coatings demonstrated an increase in cell number, alkaline phosphatase activity, and osteocalcin secretion in comparison with pure calcium phosphate-coated samples at all time points [24]. In vivo implantation of magnesium-doped calcium phosphate into animal models revealed a notable increase in bone volume and osteoconductivity in comparison with undoped calcium phosphate and could potentially enhance the osseointegration process during the early stages of bone healing [25, 26]. Another chemical element that has attraction attention is strontium due to its ability to increase bone formation and at the same time inhibit and decrease osteoclastogenesis and bone resorption [27, 28]. These characteristics have led to the application of strontium in the treatment of osteoporosis [29, 30]. Based on findings from animal models, it has been postulated that the systemic administration of strontium might enhance the osseointegration of implants and peri-implant bone quality [31–35]. In addition, strontium can be introduced into the calcium phosphate structure due to its close chemical resemblance to calcium, and this resulted in the investigation of the correlation between strontium-substituted calcium phosphate and the osseointegration of implants and tissue engineering scaffolds in both soft and hard tissues [36–44]. Furthermore, observations from in vivo studies also suggested that the doping of calcium phosphate with up to 20 mol.% strontium could be beneficial in enhancing implant osseointegration in osteoporotic bones [36, 38, 39]. Silicon has been shown to be essential during the growth and skeletal development of chicks as well as in the formation of bone, articular cartilage, and connective tissue [45–48]. Findings from the study by Carlisle suggested that the major effect of silicon appears to be on the collagen content of the connective tissue matrix and a deficiency will result in abnormal skull matrix formation. In addition, the study also supported the hypothesis concerning the involvement of silicon during the initial stages of bone formation [47]. It has also been reported that silicon is situated at active calcification sites in mammalian bones and contributes to the mineralization process of bone growth [48, 49]. Recently, an in vitro response of human mesenchymal stem cells to different concentrations of silicon either adsorbed onto the surface of the coating after precipitation or incorporated into the coating during precipitation was examined [50]. The data revealed stem cells responded to the presence of silicon in the coating in a dose-dependent fashion. The increase in silicon concentration also caused an increase in the expression of markers of osteogenic differentiation by the stem cells. More importantly, the incorporation of silicon expressed higher levels of alkaline phosphatase and osteopontin in comparison with the adsorption approach. Consequently, a number of in vivo examinations were carried out to explore the correlation between implants coated with silicon-doped calcium phosphate and the

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effect of osseointegration [37, 42, 51–55]. Porter et al. postulated that the incorporation of silicate ions into calcium phosphate could enhance its bioactivity through an increase in the number of defect structures, which are specific sites within the bioceramic that are most susceptible to dissolution. Hence, the rate of osseointegration increases by increasing the number of defect structures and the solubility of the bioceramic in a biological environment [51]. Studies were also performed to investigate the influence of different doping ions such as strontium and silicon on bone ingrowth and osseointegration [37, 42]. Other chemical elements such as silver [56, 57] and zinc [58, 59] have also been studied for their possible capacity to improve implant osseointegration or the degree of bone-to-implant contact but more research and in vivo assessments are needed to ascertain their efficacy.

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36. Li Y, Li Q, Zhu S et al (2010) The effect of strontium-substituted hydroxyapatite coating on implant fixation in ovariectomized rats. Biomaterials 31:9006–9014 37. Ballo AM, Xia W, Palmquist A et al (2012) Bone tissue reactions to biomimetic ion-substituted apatite surfaces on titanium implants. J R Soc Interface 9:1615–1624 38. Zhang J, Liu L, Zhao S et al (2015) Characterization and in vivo evaluation of trace elementloaded implant surfaces in ovariectomized rats. Int J Oral Maxillofac Implants 30:1105–1112 39. Tao ZS, Bai BL, He XW et al (2016) A comparative study of strontium-substituted hydroxyapatite coating on implant’s osseointegration for osteopenic rats. Med Biol Eng Comput 54:1959–1968 40. Li J, Yang L, Guo X et al (2017) Osteogenesis effects of strontium-substituted hydroxyapatite coatings on true bone ceramic surfaces in vitro and in vivo. Biomed Mater 13:015018. https:// doi.org/10.1088/1748-605X/aa89af 41. Ehret C, Aid-Launais R, Sagardoy T et al (2017) Strontium-doped hydroxyapatite polysaccharide materials effect on ectopic bone formation. PLoS ONE 12:e0184663. https://doi.org/10. 1371/journal.pone.0184663 42. Mumith A, Cheong VS, Fromme P et al (2020) The effect of strontium and silicon substituted hydroxyapatite electrochemical coatings on bone ingrowth and osseointegration of selective laser sintered porous metal implants. PLoS ONE 15:e0227232. https://doi.org/10.1371/journal. pone.0227232 43. Ma P, Chen T, Wu X et al (2021) Effects of bioactive strontium-substituted hydroxyapatite on osseointegration of polyethylene terephthalate artificial ligaments. J Mater Chem B 9:6600– 6613 44. Su S, Chen W, Zheng M et al (2022) Facile fabrication of 3D-printed porous Ti6Al4V scaffolds with a Sr-CaP coating for bone regeneration. ACS Omega 7:8391–8402 45. Carlisle EM (1972) Silicon: an essential element for the chick. Science 178:619–621 46. Carlisle EM (1976) In vivo requirement for silicon in articular cartilage and connective tissue formation in the chick. J Nutr 106:478–484 47. Carlisle EM (1980) A silicon requirement for normal skull formation in chicks. J Nutr 110:352– 359 48. Carlisle EM (1980) Biochemical and morphological changes associated with long bone abnormalities in silicon deficiency. J Nutr 110:1046–1056 49. Carlisle EM (1970) Silicon: a possible factor in bone calcification. Science 167:279–280 50. Rodrigues AI, Reis RL, van Blitterswijk CA et al (2017) Calcium phosphates and silicon: exploring methods of incorporation. Biomater Res 21:6. https://doi.org/10.1186/s40824-0170092-8 51. Porter AE, Patel N, Skepper JN et al (2003) Comparison of in vivo dissolution processes in hydroxyapatite and silicon-substituted hydroxyapatite bioceramics. Biomaterials 24:4609– 4620 52. Zhang E, Zou C (2009) Porous titanium and silicon-substituted hydroxyapatite biomodification prepared by a biomimetic process: characterization and in vivo evaluation. Acta Biomater 5:1732–1741 53. Izquierdo-Barba I, Santos-Ruiz L, Becerra J et al (2019) Synergistic effect of Si-hydroxyapatite coating and VEGF adsorption on Ti6Al4V-ELI scaffolds for bone regeneration in an osteoporotic bone environment. Acta Biomater 83:456–466 54. Ilea A, Vrabie OG, B˘abt, an AM et al (2019) Osseointegration of titanium scaffolds manufactured by selective laser melting in rabbit femur defect model. J Mater Sci Mater Med 30:26. https:// doi.org/10.1007/s10856-019-6227-9 55. Liu L, Wang X, Zhou Y et al (2020) The synergistic promotion of osseointegration by nanostructure design and silicon substitution of hydroxyapatite coatings in a diabetic model. J Mater Chem B 8:2754–2767 56. Łapaj Ł, Wo´zniak W, Markuszewski J (2019) Osseointegration of hydroxyapatite coatings doped with silver nanoparticles: scanning electron microscopy studies on a rabbit model. Folia Morphol 78:107–113

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57. Köse N, Bayrak ÇH, Köse AA et al (2020) The use of orthopaedic implants with silver iondoped ceramic coating in the prevention of implant-related infections. Orthop Proc 102-B:74. https://doi.org/10.1302/1358-992X.2020.11.074 58. Mistry S, Burman S, Roy S et al (2021) Surface characteristics of titanium dental implants with improved microdesigns: an in vivo study of their osseointegration performance in goat mandible. J Biomater Appl 35:799–813 59. Bhattacharjee A, Bandyopadhyay A, Bose S (2022) Plasma sprayed fluoride and zinc doped hydroxyapatite coated titanium for load-bearing implants. Surf Coat Technol 440:128464. https://doi.org/10.1016/j.surfcoat.2022.128464

Chapter 7

Enhancing Implant Osseointegration Through Nanocomposite Coatings

The research and production of new nanocomposites and nanolaminates that is suitable for a variety of applications such as surface-modified dental and orthopedic implants and prostheses for enhanced soft and hard tissue attachment and scaffolding materials with increased bioactivity for tissue repair and regeneration has been ongoing. The developments of nanoceramic composite coatings based on calcium phosphate have been the focus for biomedical and dental researchers since 2000. In general, the thickness of a single-layered coating is less than 100 nm. Coatings that contain multiple layers with suitable biological, physical, chemical, and mechanical properties can be synthesized with relative ease and can be deposited with different compositions as multilayered gradient coatings or nanolaminates [1– 3]. As such, one category of nanocomposite coating is a multilayered, nanolaminated mixed nanocoatings and this can be produced by laminating different nanocoatings together to achieve the desired thickness, structure, and properties [1–3]. The properties of calcium phosphate-based nanocomposite coatings and nanolaminates are dependent on the secondary or dispersion phase. At the moment, the development of new generations of nanocomposite coatings containing synthetic and natural nanomaterials such as bioglass, collagen, and chitosan are being examined for their capacity to promote osseointegration [4]. Polymeric materials have also been utilized as the secondary phase in nanocomposites with calcium phosphate. A study has suggested that cellular responses at the interface between bone tissue and titanium implant can be induced if a maleic polyelectrolyte such as sodium maleatevinyl acetate copolymer is added to calcium phosphate. Cells grown on the coated surfaces induced a higher proliferation rate and the inclusion of a sodium maleate copolymer enhanced surface bioadhesion in comparison to just calcium phosphate [5]. Later, an in vivo biocompatibility study was attempted in which a composite coating comprised of nanocalcium phosphate and polyamide was deposited onto injection-molded polyamide substrate and implanted into the trochlea of rabbit femurs. Results from the push-out test at 12 weeks post-implantation showed that the

© The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2023 A. H. Choi and B. Ben-Nissan, Calcium Phosphate Nanocoatings for Bone Regeneration, Tissue Repair and Reconstruction, https://doi.org/10.1007/978-981-99-5506-0_7

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composite coated implants produced significantly higher shear strength value than compared to uncoated implants [6]. A greater emphasis will be placed on the incorporation of molecular and nanoscale-based biological materials and pharmaceutics such as bone morphogenetic proteins, growth factors, stem cells, osteopontin, and simvastatin into calcium phosphate as multifunctional nanocomposite coatings in an effort to reduce the timeframe needed for implant integration as well as enhancing and promoting osseointegration of dental and orthopedic implants and prostheses.

7.1 Collagen Considered as one of the most useful biomaterials for applications such as tissue scaffolding based on the fact that human bone tissue is essentially a composite consists of an organic collagen matrix reinforced by inorganic calcium phosphate crystals [4]. This has led researchers to the idea of creating a collagen-calcium phosphate composite for bone regeneration and according to the observations from numerous investigations that such a composite could encourage bony growth into the porous structure [7, 8]. In 2008, a study was carried out to create foam-like nanocomposite bone scaffolds that could be recognized and remodeled by natural bone once implanted. The scaffold consisted of a microstructure similar to trabecular bone together with an apatitic phase that has similar chemical composition, crystalline phase and grain size as the trabecular bone apatite. Based on in vivo observations, the authors stated that the nanocomposite scaffold was osteoconductive and could potentially be used to heal non-union fractures as well as critical-sized defects [7]. Later, a study was attempted to examine the effect of nanoporosity on cell-material interactions when bone marrow-derived mesenchymal stem cells were cultured onto the nanocomposites consisted of nanoHAp, reconstituted collagen, and nanoamorphous calcium phosphate [8]. The authors claimed that the nanocomposite significantly facilitated the osteogenic differentiation of mesenchymal stem cells. The biological performance and bone regeneration capacity of collagen-calcium phosphate nanocomposites were further examined in a later study using a tibial bone defect model in rats. The study also attempted to examine and compare the effects of utilizing calcium phosphate fibers instead of calcium phosphate powders based on the hypothesis that fibrous material would have a more appropriate morphology and the amalgamation of collagen with calcium phosphate fibers would present a better structure toward attracting bone cell growth resulting in bone formation [9]. Greater quantities of newly formed bone in the region of the defect on top of a reduction in the composite material were observed when calcium phosphate was presented in the fiber form. It has also been suggested that a composite coating of collagen and calcium phosphate can be beneficial over the individual components as the ductile properties of collagen counterbalance the poor fracture toughness of calcium phosphate [10– 15]. In 2010, a study has revealed the osteogenic behavior of rat bone marrow cells

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could be stimulated in vitro by calcium phosphate-collagen composite coatings using a calcium phosphate/collagen ratio comparable to that of native bone tissue even if the thickness of the composite coating is less than 100 nm. In addition, the composite coating is also capable of enhancing osteoblast differentiation in comparison to pure calcium phosphate coatings based on accelerated mineral deposition and a decrease in proliferation [10]. Recently, a study was attempted to determine the possibility of producing an orthopedic implant coating that is capable of preventing bone infection while at the same time promote osseointegration [14]. Two separate in vivo animal models were used in their study in which a rat model was used to investigate the effectiveness of the collagen-calcium phosphate layer to deliver antibiotics such as vancomycin in an effort to prevent bone infection; while a minipig model with terminated skeletal growth was used to study the osseointegration behavior. Observations from the antimicrobial investigation at six-week post-implantation suggested that impregnating the collagen-calcium phosphate layer with 10 wt.% vancomycin was sufficient to maintain the cortical bone porosity and calcium/phosphate ratio at physiological levels following the introduction of bacterial contamination using a clinical isolate of Staphylococcus epidermidis during the surgical implantation. In addition, quantified ratio of new bone recorded in the pig model six months post-implantation signaled enhancements in osseointegration and this was based on a two-fold bone ingrowth was observed on the coated implants in comparison to uncoated implants.

7.2 Chitosan Despite the fact that it is one of the most extensively investigated material in tissue engineering, the bioactivity of chitosan will need enhancements for certain tissues and subsequently making them not as ideal if used as a stand-alone material [16]. A study has suggested that chitosan could possibly be utilized as a bioactive coating with the intention of enhancing osseointegration of implants. However, the bonding strength of chitosan coating deposited onto the surfaces of titanium implants were less than those reported for plasma-sprayed calcium phosphate coatings [17]. Subsequently, it is added to calcium phosphate in the form of a composite, which also addresses the issue of chitosan’s poor mechanical properties and this has limited their use in loadbearing applications [4]. Studies have demonstrated that such a composite coating could enhance cell adhesion and differentiation in vitro and potentially be beneficial in promoting osseointegration in vivo [18–20].

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7.3 Bone Morphogenetic Proteins (BMPs) Commercially available calcium phosphate bioceramics will display a very low specific surface area and poor surface reactivity after sintering at an elevated temperature. Despite being an exceptional osteoconductive material, the osteoinductive properties of calcium phosphate seems relatively insufficient. It has been suggested that the deposition of coatings composed of nanocrystalline carbonate apatite onto the surfaces of calcium phosphate would improve its surface and biological properties and such a layer could encourage biological activity by stimulating the creation of nanopores and increasing the specific surface area [21]. Adsorption examination using recombinant human BMP-2 (rhBMP-2), an osteogenic growth factor, confirmed enhancements in the surface reactivity. The study also indicated the bioceramic is able of adsorbing more of the protein and releasing it in a prolonged manner. It was concluded that the combination of the coated bioceramic and rhBMP-2 could improve bone formation as observed in an in vivo ovine animal model. During the last five years, a number of studies have been carried out to examine the bone regeneration capacity of rhBMP-2 in oral and maxillofacial surgery such as alveolar grafting, mandibular reconstruction, and peri-implant ridge and maxillary sinus floor augmentation [2, 22–30]. BMPs are multifunctional growth factors belonging to the superfamily of transforming growth factor β (TGFβ), and their roles in cellular functions in postnatal and adult animals have been extensively investigated in recent years [31, 32]. In 1965, the activities of BMPs were first recognized by Urist [33]. However, it was not until the last 1980s after the duplication of human BMP-2 and BMP-4 as well as the purification and sequencing of bovine BMP-3 (osteogenin) that the protein responsible for bone induction was identified [34–36]. RhBMP-2, isolated originally from demineralized bone matrix, has the ability to stimulate cell differentiation into chondroblasts and osteoblasts to initiate the process of generating new bone tissue and cartilage [36–38]. In vivo studies have demonstrated the deposition of calcium phosphate coatings, regardless if synthesized biomimetically or using conventional coating technique such as plasma spraying, can potentially be utilized to deliver growth factors such as BMPs in an effort to improve peri-implant bone regeneration and their gradual release from the coating has been theorized to be the consequence of cell-mediated degradation [39–46]. Moreover, BMP can be incorporated and signaling factors be steadily released from the coating during in vivo degradation [39–42]. The deposition of a calcium phosphate coating based on precipitation in a simulated body fluid solution has also generated new opportunities for the incorporation of BMP-2 to improve the osseointegration of dental implants. Studies were conducted to gain insights into the in vivo osteoinductive capability of these biomimetic or “self-assembled” coatings containing BMP-2 deposited onto zirconia [44] as well as on titanium [39, 42].

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Fig. 7.1 Peri-implant bone formation at 6 weeks post-implantation into rabbit models revealing implant coated with HAp plus collagen (c) generated greater bone formation than uncoated implant (a), implant coated with only HAp (b), and implant coated with HAp plus collagen plus BMP-2 (d). The composite coating was synthesized by immersing the HAp-coated titanium implant into a self-assembled collagen solution followed by a BMP-2 solution and subsequently freeze-dried. Reprint with permission from [47]

On the other hand, a number of studies have raised the concern that the effectiveness of BMP-2 on improving osseointegration and bone formation is governed by the way in which the protein is delivered (Fig. 7.1) [40, 41, 47]. Findings from the study by Liu et al. showed that implant osteoconductivity was most severely compromised when the protein was adsorbed onto the coating instead of being incorporated into the coating, indicating that the incorporation approach will deliver the protein via a slow-release system instead of a burst-release profile [40]. This notion was later supported by the in vivo observations from the study by Yoo et al. in which significant increases in bone-implant contact and bone area fraction occupancy were observed at three weeks post-implantation when rhBMP-2 was applied to the coated surfaces, whereas the protein did not as effectively improve the bone response when adsorbed directly onto the surfaces but still noticeably higher than uncoated implants [41]. As discussed earlier, the combination of osteoinductive BMP-2 and osteoconductive apatite can produce a much greater osteogenic surface environment on titanium and one of the primary concerns regarding BMP-2 is obtaining a suitable and reliable system that can ensure the protein is delivered to target sites and released in a controlled and sustained manner while maintaining its bioactivity [48, 49]. In 2012, a study was attempted to develop a nanocomplex nanoparticle by coupling BMP2 with negatively charged chondroitin sulfate to be used as a delivery vehicle and it was suggested that growth factors contained in such manner would be released in controlled mode in comparison to those in a free form [48]. The nanocomplex nanoparticles were precipitated on biomimetic calcium phosphate-coated titanium surfaces. Observations showed that the immobilized BMP-2 was released in a moderate rate for four weeks without an initial burst of BMP-2. Mouse osteoblast cells seeded on the coated surface containing the nanocomplex nanoparticles displayed

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faster cell proliferation as compared to titanium alone or calcium phosphate-coated titanium. In addition, the gene expression of osteocalcin and type I collagen was significantly upregulated with the use of the nanocomplex nanoparticle. Similar result was observed in the measurement of alkaline phosphatase activity [48]. Later, BMP2 was incorporated into biomimetic calcium phosphate coating deposited on porous titanium alloy treated with micro arc oxidation and it was hypothesized that it would provide a method in which BMP-2 will be released in a prolonged manner with enhanced cell proliferation and differentiation [49]. The authors claimed that this approach could release BMP-2 in a sustained manner over 35 days and the slow release at the bone–implant interface for an extended period resulted in improvements in osseointegration when the implants were examined in vivo by implantation into the calvarial bone of animal models. The results also revealed enhanced bone regeneration was identified at 12 weeks post-implantation [49]. In addition to rhBMP-2, researchers have also examined the use of another protein, the rhBMP-7, which is also referred to a recombinant human osteogenic protein1 (rhOP-1) [50–53]. The efficacies of rhBMP-7 in encouraging osseointegration of dental implants were examined using an in vivo canine model. Both calcium phosphate-coated and uncoated implants were placed in fresh extraction sites together with the protein. At twelve weeks post-implantation, observations revealed extraction sites treated with the protein were filled completely with new bone and were well incorporated with the host bone. Small increases in osseointegration and adjacent new bone formation were also apparent surrounding implants placed with the protein. The results also demonstrated that new bone had filled the untreated defect sites but the quantity, density, and magnitude of remodeling were less than those observed in the sites treated with the protein [50]. Later, a study was attempted to examine if rhBMP7 delivered by a porous calcium phosphate coating could enhance bone formation and ingrowth at the bone–implant interface without the use of bone grafts. Even though in vivo observations showed such a combination provided better bone ingrowth, the presence of rhBMP-7 did not affect the magnitude of bone apposition and callus height [51]. Another study suggested that nanometer-thick calcium phosphate coating could potentially enhance new bone formation in the presence of rhBMP-7 [52]. It has also been hypothesized that if rhBMP-7 was released slowly from a delayed drug-delivery system such as biomimetic calcium phosphate coating over time at near-physiological concentrations should be able to actively promote and accelerate the formation of new bone in the peri-implant space in a hostile osteopenic environment. In addition, the authors further postulated that such a system would permit for a reduction in the required dosage of osteogenic growth factor as well as improving the safety of the protein [53]. The release of growth factors such as BMP-2 and the angiogenic vascular endothelial growth factor (VEGF) from bioceramics delivery systems such as calcium phosphate has been well documented to influence the magnitude of bone formation in animal models, and observations from in vivo studies have suggested that the utilization of calcium phosphate for the combined delivery of both BMP-2 and VEGF could accelerate new bone formation as well as improvements in osteogenic potential [54, 55]. On the other hand, it remains unclear on the effectiveness of utilizing calcium

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phosphate coatings to deliver growth factors with the intention of enhancing osseointegration at the bone–implant interface. The efficacy of dual delivery of rhBMP-2 and recombinant human VEGF using biomimetically octacalcium phosphate-coated implants on osseointegration was examined [56, 57]. Bone volume density values were enhanced after two weeks post-implantation into frontal skulls of domestic pigs but bone-implant contact did not improve significantly after four weeks postimplantation [56]. The authors also suggested based on the observations of a later in vivo study that the dual delivery of growth factors favored bone mineralization in addition to expression of vital bone matrix proteins [57].

7.4 Peptides Apart from utilizing BMPs to biofunctionalize implant surfaces with the intention of enhancing osseointegration, studies were also carried out on biomimetic peptides such as RGD and P-15 for their potential applications in enhancing the cellular interactions with biomaterials such as calcium phosphate. More importantly, they can be produced synthetically and purification can be conducted relatively easily [58–61]. Within these biomimetic active peptides, only a cell-binding sequence is contained and by biologically mimicking cell-binding sites, the RGD (Arg-Gly-Asp) peptide has been postulated to promote cell adhesion [62]. The combination of RGD-containing peptide and calcium phosphate deposited on implant surfaces could potentially improve the attachment and differentiation of osteoblasts. As postulated, the osteoconductivity of calcium phosphate was strengthened by the availability of peptide [63]. The protein absorption on the surfaces of calcium phosphate is altered after the peptide pretreatment once implanted into bone [60]. The improvement in osseointegration may be the result of the preferential attachment of cells such as osteoprogenitors to the surfaces of calcium phosphate. It has also been suggested that the immobilization of RGD on anodized titanium via chemical grafting could improve the osseointegration of implants [64]. On the other hand, observations from an in vivo animal study that investigated the influence of Ti-6Al-4V implants with collagen coating and covalently bound RGD peptides on peri-implant bone formation and bone-implant contact suggested there is only weak evidence that implants coated with the peptide would increase the magnitude of new bone formed in the alveolar process. In spite of this, the study revealed the presence of RGD peptides improved the amount of bone-implant contact from one month to three months post-implantation in comparison to uncoated titanium implants [65]. A similar observation was noticed in an in vivo investigation that although the deposition of a cyclic RGD coating on titanium did present a significant bone stimulating effect at the bone–implant interface four weeks postimplantation, an increase in bone density was not observed further away from the implant. The authors also stated that there were no differences in bone volume between the coated and uncoated implants at the 0–750 μm zone adjacent to the implant surface, indicating the effect of RGD was primarily at the interface [66].

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A study in 2008 also questioned whether functionalizing calcium phosphate with RGD peptides would have any benefit on the performance of coated implants [67]. Uncoated and RGD-coated calcium phosphate disks implanted into rat tibiae for five days revealed the presence of the peptide significantly inhibited the amount of new bone directly contacting the implant perimeter in addition to the total bone formation [67]. These marginal healing responses has been postulated to be the consequence of uncontrolled signaling responses at the bone–implant interface by unregulated or sub-optimal integrin binding [68]. Furthermore, observations from another study suggested structural changes in peptides adsorbed onto titanium as a response to its surface characteristics or low peptide adsorption decreases the chance of osteoblast adhesion to be promoted [69]. Due to the contradictory clinical results generated by the use of RGD peptides, researchers began to search for an alternative peptide sequence that can be applied as an implant coating. Studies have demonstrated that a biomimetic bone matrix referred to as P-15 was more potent in comparison to RGD-containing peptides when competing with collagen for cell binding [58, 59, 70]. This biomimetic matrix is a synthetic, 15-amino-acid residue peptide that is identical to the 766 GTPGPQGIAGQRGVV780 sequence of the type I collagen α1(I) chain [70]. A study was conducted in 2010 to investigate the hypothesis that faster osseointegration process can be achieved if surfaces of dental implants were coated with calcium phosphate and P-15 peptides. A significant higher percentage of bone-toimplant contact was observed on implants containing high concentration of P-15 after 14 and 30 days post-implantation into the forehead regions of 12 adult pigs as revealed by the results of histomorphometric analysis. Furthermore, an increase in peri-implant bone density at 30 days was recorded on implants containing both low and high concentrations of P-15 peptide [59]. These observations were later endorsed in an in vivo animal model study confirming the postulation that the bioactivity of implants coated with calcium phosphate can be improved by the presence of P-15 and its influence were particularly noticeable during the early stages of the healing period [58]. However, an in vivo study in 2016 suggested that there is no benefit in the early phase of osseointegration could be established for dental implants with P-15 containing surface when they compared the bone-to-implant contacts between implants with different surfaces including those that consisted of electrochemically deposited calcium phosphate and P-15. Observations revealed there were no statistical differences between surface types following 2 and 7 days of healing, even though implants with calcium phosphate and P-15 displayed superiority in bone-to-implant contact after 6 months of healing [71].

7.5 Stem Cells The method in which mesenchymal stem cells (MSCs) are obtained from healthy tissues is the main challenge related to its clinical application. The oral cavity in recent times has contributed significantly as a critical source of MSCs. In addition to

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the efficacy of cells being isolated from dental tissues such as freshly extracted teeth, the accessibility to the dental clinicians and surgeons render the clinical exploitation of oral-derived stem cells such as dental follicle progenitor cells, dental pulp stem cells, and periodontal ligament stem cells extremely appealing [72–74]. Stem cells derived from dental pulp offer an appropriate model to examine bone differentiation due to their osteogenic capability in comparison to other types of cells collected from the adult human body [75]. In vitro studies have been carried out to examine the interactions between dental pulp stem cells and its interactions with the surfaces of dental implants [75–77]. Observations from in vitro studies have led researchers to postulate the possibility of using dental pulp stem cells together with calcium phosphate in the form of a tissue scaffold for the repair of critical-size defects in animal models [78, 79] or as a coating to potentially enhance the osseointegration process with adequate levels of bone-implant contact [80, 81]. Being a specialized connective tissue that support and maintain teeth in situ as well as providing an easily accessible tissue resource that can be expanded ex vivo, findings from a study suggested that stem cells contained within periodontal ligament possess the capacity to generate periodontal ligament/cementum-like tissue in vivo [82]. A study in 2018 examined the possibility of creating periodontal structures using cultured human periodontal ligament-derived stem cells on calcium phosphatecoated titanium surface implanted into a xenogeneic athymic rat femur model and in an autologous canine mandible model. The authors claimed that a periodontal-like structure was formed around the implant [83]. Human deciduous teeth have also been suggested as another source of stem cells due to their ability to differentiate into odontoblasts, neural cells, osteoinductive cells, and adipocytes [84]. In addition, stem cells from human exfoliated deciduous teeth were discovered to be able of inducing bone formation after in vivo implantation into animal models [84–88]. In vivo animal studies have hypothesized that if titanium implants were treated with stem cells prior to implantation could offer the possibility of early osseointegration (Figs. 7.2 and 7.3) [89, 90]. Human periapical inflammatory cysts, which are a biological waste and intended to be surgically removed in order to avoid disabling pathological conditions within the oral cavity, also display MSC-like properties. Studies have suggested that MSCs isolated from human periapical cysts are able to differentiate into osteoblasts and adipocytes based on their self-renewal capacity and multilineage differentiation potency [91, 92]. A study has suggested that the combination of a bioactive material such as calcium phosphate and autologous stem cells from periapical cysts could offer a solution which encourages the regenerative healing of oral structures such as alveolar bone [92].

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Fig. 7.2 Topology of titanium implant surface analyzed using SEM. a and e Plasma-untreated Ti on which PBS was immobilized (N-PBS) implants. b and f Plasma-treated Ti on which PBS was immobilized (P-PBS) implants. c and d Plasma-untreated and plasma-treated titanium on which stem cells from human exfoliated deciduous teeth-conditioned medium was immobilized under 10,000-times magnification, respectively. g and h The same implants under 30,000-times magnification. Reprint with permission from [89]

7.6 Functionalization Using Osteopontin An interaction begins rapidly between the surface of bioceramic and extracellular matrix proteins as soon as the bioceramic is implanted. Biological signals are produced by the proteins depending on the chemical features and surface topography of the implant. The biological response as a result is taken from the cells against the surface features, and this response provides cell adhesion stability and triggers cell proliferation and differentiation [93]. Osteopontin is an extracellular matrix protein that has been shown to play a role in bone mineralization, foreign body response, and cell adhesion and differentiation [94, 95]. It has been suggested that a deficiency in osteopontin will disrupt direct osteogenesis resulting in a delay in osseointegration after immediate placement of endosseous implants [96]. Furthermore, in vivo studies have also suggested that functionalization of titanium implant surfaces with osteopontin could be beneficial in enhancing bone-implant integration [97–99]. An in vitro study was carried out to examine the adsorption of bovine osteopontin on calcium phosphate and the influence of the resulting osteopontin protein coating on the adsorption of human mesenchymal stem cells. In comparison to uncoated calcium phosphate, the presence of a protein layer on the calcium phosphate surface triggered a larger and faster cell spreading in addition to higher cell motility [100]. Later, a canine endosseous 0.75-mm gap implant model was used to examine the osseointegrative effect of a composite coating composed of poly-d,l-lactic-acid and calcium phosphate nanoparticles pre-adsorbed

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Fig. 7.3 Four weeks post-implantation of implants with pre-adhered stem cells from human exfoliated deciduous teeth into adult Beagle dogs. In the group containing stem cells from human exfoliated deciduous teeth (d–f at 12.5-, 40-, and 100-times magnification, respectively), there was an absence of intervening fibrous tissue and more bone was observed in comparison to the control group, which received implants and phosphate-buffered saline (PBS) (a–c). In addition, close contact between the bone and the implant surface was observed (f). Reprint with permission from [90]

with bovine osteopontin deposited onto a commercially available titanium implant. A significant increase in the formation of new bone in the porosities of the implant was observed but no differences were observed in the gap between pure composite coated implant and osteopontin functionalized implants [101]. Recently, observations from an animal study examining the healing patterns at the bone–implant interface after immediately placed implantation in the maxilla also suggested that the use of osteopontin on calcium phosphate could improve direct osteogenesis during osseointegration following implantation [95].

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7.7 Simvastatin Mundy et al. suggested the administration of statins in appropriate doses could enhance new bone formation. This was based on the findings that an increase in bone formation was observed when simvastatin was injected subcutaneously over the calvaria of mice. Furthermore, an increase in cancellous bone volume was recorded in rats when simvastatin was administered orally. They suggested that this effect was associated with increased expression of the BMP-2 gene in bone cells [102]. Later, an in vivo study investigated the effect of daily administration of simvastatin on the promotion of osteogenesis around titanium implants inserted into the tibias of rats, revealing the bone contact ratio and bone density around the implants were significantly greater in comparison to the control group which did not receive simvastatin [103]. Similar observations were made in other in vivo animal studies suggesting simvastatin has the potential to enhance implant osseointegration under normal and osteoporotic bone conditions [104, 105]. In addition, in vivo studies were also carried out to examine the possibility of utilizing calcium phosphate coatings deposited onto titanium implants as a way to delivery simvastatin locally [106–109]. It has been suggested that modifying titanium surfaces with nanosized calcium phosphate and simvastatin could enhance the osteogenic differentiation of MC3T3-E1 cells in vitro and improve bone formation at the bone–implant interface when inserted into the proximal tibia and femoral head of rabbits. Furthermore, the authors claimed that such a system could release simvastatin continually for up to 28 days [107]. Similar observations were also made in a recent study based on the findings from histological evaluation, micro-CT, and biomechanical pull-out tests (Fig. 7.4) [108]. As discussed earlier, it has been postulated that the systemic administration of strontium might enhance the osseointegration of implants under osteoporotic conditions. An in vivo study postulated that an implant coating composed of simvastatin, strontium, and calcium phosphate would be beneficial in osteoporotic conditions through the combined efforts of the excellent osteoconductivity of calcium phosphate, the increased degradation rate of strontium-doped calcium phosphate, and the physiological effects of strontium and simvastatin [109].

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Fig. 7.4 Histological evaluation of the simvastatin-HAp-Ti-6Al-4V in vivo. a Representative histological images of hematoxylin and eosin staining and Giemsa staining of rat femurs. The black arrows represent massive destruction of cortical bone. The green arrows indicate new bone formation around implants. The red arrows indicate bacteria remaining in the bone tissue. b Hematoxylin and eosin staining analyses of different organs, including the heart, liver, spleen, lung, and kidney. Reprint with permission from [108]

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42. Hunziker EB, Jovanovic J, Horner A et al (2016) Optimisation of BMP-2 dosage for the osseointegration of porous titanium implants in an ovine model. Eur Cell Mater 32:241–256 43. Liu Y, Schouten C, Boerman O et al (2018) The kinetics and mechanism of bone morphogenetic protein 2 release from calcium phosphate-based implant-coatings. J Biomed Mater Res A 106:2363–2371 44. Teng F, Zheng Y, Wu G et al (2019) Bone tissue responses to zirconia implants modified by biomimetic coating incorporated with BMP-2. Int J Periodontics Restorative Dent 39:371–379 45. Teng F, Wei L, Yu D et al (2020) Vertical bone augmentation with simultaneous implantation using deproteinized bovine bone block functionalized with a slow delivery of BMP-2. Clin Oral Implants Res 31:215–228 46. Pang K, Seo YK, Lee JH (2021) Effects of the combination of bone morphogenetic protein-2 and nano-hydroxyapatite on the osseointegration of dental implants. J Korean Assoc Oral Maxillofac Surg 47:454–464 47. Lee SW, Hahn BD, Kang TY et al (2014) Hydroxyapatite and collagen combination-coated dental implants display better bone formation in the peri-implant area than the same combination plus bone morphogenetic protein-2-coated implants, hydroxyapatite only coated implants, and uncoated implants. J Oral Maxillofac Surg 72:53–60 48. Bae SE, Choi J, Joung YK et al (2012) Controlled release of bone morphogenetic protein (BMP)-2 from nanocomplex incorporated on hydroxyapatite-formed titanium surface. J Control Release 160:676–684 49. Teng FY, Tai IC, Ho ML et al (2019) Controlled release of BMP-2 from titanium with electrodeposition modification enhancing critical size bone formation. Mater Sci Eng C Mater Biol Appl 105:109879 50. Cook SD, Salkeld SL, Rueger DC (1995) Evaluation of recombinant human osteogenic protein-1 (rhOP-1) placed with dental implants in fresh extraction sites. J Oral Implantol 21:281–289 51. Saran N, Zhang R, Turcotte RE (2011) Osteogenic protein-1 delivered by hydroxyapatitecoated implants improves bone ingrowth in extracortical bone bridging. Clin Orthop Relat Res 469:1470–1478 52. Bonato RS, Fernandes GVO, Calasans-Maia MD et al (2022) The influence of rhBMP-7 associated with nanometric hydroxyapatite coatings titanium implant on the osseointegration: a pre-clinical study. Polymers 14:4030. https://doi.org/10.3390/polym14194030 53. Hunziker EB, Liu Y, Muff M et al (2021) The slow release of BMP-7 at a low dose accelerates dental implant healing in an osteopenic environment. Eur Cell Mater 41:170–183 54. Chen S, Shi Y, Zhang X et al (2020) Evaluation of BMP-2 and VEGF loaded 3D printed hydroxyapatite composite scaffolds with enhanced osteogenic capacity in vitro and in vivo. Mater Sci Eng C Mater Biol Appl 112:110893. https://doi.org/10.1016/j.msec.2020.110893 55. Kim TW, Ahn WB, Kim JM et al (2020) Combined delivery of two different bioactive factors incorporated in hydroxyapatite microcarrier for bone regeneration. Tissue Eng Regen Med 17:607–624 56. Ramazanoglu M, Lutz R, Ergun C et al (2011) The effect of combined delivery of recombinant human bone morphogenetic protein-2 and recombinant human vascular endothelial growth factor 165 from biomimetic calcium-phosphate-coated implants on osseointegration. Clin Oral Implants Res 22:1433–1439 57. Ramazanoglu M, Lutz R, Rusche P et al (2013) Bone response to biomimetic implants delivering BMP-2 and VEGF: an immunohistochemical study. J Craniomaxillofac Surg 41:826–835 58. Coelho PG, Teixeira HS, Marin C et al (2014) The in vivo effect of P-15 coating on early osseointegration. J Biomed Mater Res B Appl Biomater 102:430–440 59. Lutz R, Srour S, Nonhoff J et al (2010) Biofunctionalization of titanium implants with a biomimetic active peptide (P-15) promotes early osseointegration. Clin Oral Implants Res 21:726–734 60. Itoh D, Yoneda S, Kuroda S et al (2002) Enhancement of osteogenesis on hydroxyapatite surface coated with synthetic peptide (EEEEEEEPRGDT) in vitro. J Biomed Mater Res 62:292–298

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Chapter 8

Calcium Phosphate Nanocoated Coralline Apatite

Highly functional architectural structures with open pores that are interconnected can be discovered easily within the marine environment and these structures are ideal for human implantation in its original form or converted to materials more suitable for biomedical applications as a result of their chemical compositions and high mechanical strength. In addition, off-the-shelf inorganic and organic marine skeletons contains a perfect environment for the proliferation of added MSC populations and promoting bone formation which is appropriate from a clinical perspective [1, 2]. Coralline apatites and converted coral skeletons are perfect examples [3–5]. Corals, either in their natural form or as hybridized synthetic forms, offer great opportunities in bone tissue engineering. Natural coral exoskeletons have been suggested and extensively utilized in dental and maxillofacial surgery and orthopedics as bone replacements due to their capacity to establish chemical bonds with bone and soft tissues in vivo on top of their open and interconnected porosity and excellent mechanical properties. Corals and coralline structures have been found to possess similar or superior mechanical properties in comparison to most porous calcium-based bioceramics. Moreover, their rates of resorption have been observed to be the same as the formation of new host bone tissues [5]. It has been suggested that the biocompatibility and mechanical properties of corals are regulated strongly by its organic composition. The abundance and composition of the organic matrices are responsible for the successful biological integration of natural unconverted coral with human host bone [6]. Molecules essential for the regulation and guiding bone morphogenesis and in particular the actions related to the mineral metabolism and deposition were found in the earliest marine organisms. This is based on the notion that they represent the first molecular components known for calcification, morphogenesis, and wound healing. The design and availability of marine materials have played a pivotal role in the creation of one of the simplest solutions to important problems in regenerative medicine and in providing opportunities and frameworks of mineralized proteins, © The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2023 A. H. Choi and B. Ben-Nissan, Calcium Phosphate Nanocoatings for Bone Regeneration, Tissue Repair and Reconstruction, https://doi.org/10.1007/978-981-99-5506-0_8

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osteopromotive analogues, nanofibers, macro- and microspheres. This is demonstrated by the biological efficiency of marine structures such as corals, shells, and sponge skeletons to accommodate self-sustaining musculoskeletal tissues and to promote bone formation through the utilization of nacre seashells and sponging extracts [3, 4]. Naturally occurring biomatrices such as sponge skeletons and marine shells with wide-ranging structural similarities and chemical homologies to human extracellular matrices and whole tissues have thus far been identified as candidates in our quest to find scaffolding materials for bone tissue engineering. Due to its chemical, crystallographic, and structural similarities to native human bone, natural and converted corals are the primary source of natural marine skeletons that has been utilized in the regeneration of human bone [3–5, 7]. The application of the hydrothermal processing has permitted natural skeletons to be employed directly as a scaffold for growing cells into tissues and eventually in the generation of new bone tissue [7]. It has been widely acknowledged that the beginning of the coral life cycle starts with the absorption of carbonic acid and calcium ions found in seawater by the polyps and uses them to synthesize calcium carbonate in the form of aragonite crystals, which represents 97–99% of the coral skeleton. Depending on the environmental conditions, the remaining composition contains various elements but mostly consists of trace elements of fluorine, strontium, phosphorous in the form of phosphate, and strontium [5]. These elements found in the coral exoskeleton structure are known to play a crucial role in the bone mineralization process as well as in the activation of key enzymes related to bone cell remodeling. The exoskeleton of marine madreporic corals is used as a starting material for the synthesis of natural coral graft substitutes. In 1929, the use of coral for dental applications were first attempted and since then, corals were recognized and examined in animals during the early 1970s and later in humans in 1979 as a potential candidate as bone graft substitutes. A commonly used coral, the structure of Porites is comparable to that of human cancellous bone. In addition, its initial mechanical properties are also similar. Biomineralization and structural examinations of coral can also be used to improve the development of new advanced functional scaffolds owing to the unique nanoscale organization of organic mineral and tissue. For example, the characterization of the ultrastructure of deep-sea Bamboo coral (Anthozoa: Gorgonacea: Isididae) revealed the internodes demonstrate bone-like mechanical and biochemical properties. Moreover, the organic matrix of the Bamboo coral, which is comprised of an acidic fibrillar protein framework, displays potential as a model for future applications in tissue engineering. The growth of both osteoblast and osteoclast are supported by this organic matrix. It has been hypothesized that implants of blood vessels can be generated using the gorgonin or collagen matrix of this coral based on its exceptional bioelastomeric properties. Quinones can be used to cross-link and harden collagenous gorgonin proteins and the resultant product closely resembles to human keratin [1, 2]. Coral grafts, in spite of not being osteoinductive or osteogenic, are ideal to serve as delivery systems for growth factors as well as allowing the attachment, growth, differentiation, and spread of cells. If they are applied in a proper manner, natural

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coral exoskeletons have been observed to be an excellent substitute as bone grafts. The combination of coral skeleton with in vitro expanded human bone marrow stromal cells demonstrated enhancements in osteogenesis at levels greater than those observed with pure scaffold or fresh marrow incorporated into the scaffold [8]. Results from clinical studies of orthopedic and maxillofacial surgical cases that utilized in vivo large animal segmental defect revealed there was a complete re-corticalization and formation of a medullary canal with mature lamellar cortical bone and on lay graft for contour augmentation of the face, resulting in clinical union in a majority of cases [9, 10]. Most importantly, it is imperative to select the most appropriate coral so that its resorption rate is the same as the rate of bone formation at the implantation site. It was discovered based on the observations from a number of studies that the activities of osteoblasts during the mineralization process can be stimulated through the presence of strontium, while at the same time inhibiting osteoclasts. Likewise, fluorine was also found to assist in bone formation through similar stimulatory effect on the proliferation of osteoblast [6]. Similarly, the availability of magnesium is also beneficial during bone remodeling as it has been demonstrated to increase the mechanical properties of newly formed bone [11]. Evidently, most of the elements discovered in human bone can be detected in different corals, but they differ in their distribution and quantities. Numerous synthetic and natural bone graft materials that are currently being used are manufactured from coralline HAp. However, commercial coralline HAp has retained calcium carbonate (CaCO3 ) or coral as a consequence of the conversion process. This in turn results in high dissolution rates since the structure possesses nanopores within the inter-pore trabeculae. These features, under certain conditions, decrease strength and durability and are not suited for applications where high structural strength is required. This issue can be resolved through the use of sol–gel-derived nanocoatings, which permits the strength of corals to be enhanced and subsequently enables them to be used more frequently at various skeletal locations [1–5, 12–15]. The approach involves a two-stage application process and in the first stage, a complete conversion of coral to 100% HAp is achieved. One of these conversion processes is commonly referred to as the hydrothermal exchange conversion technique developed in 1974 [16]. Basically, the process involves the exchange of carbonate component of the coral to phosphate to create calcium phosphates and its derivatives at temperatures between 200 and 260 °C for a period of 24–48 h. Different forms of calcium phosphates can be produced by adjusting the molar ratio of calcium to phosphate. In comparison to other calcium phosphate compositions, HAp or tricalcium phosphates are more suitable as bone grafts under certain circumstances. Furthermore, tricalcium phosphate has been extensively investigated and applied as bone grafts mainly due to its relatively faster dissolution rate compared to HAp [5]. The hydrothermal conversion to tricalcium phosphate from calcium carbonate exoskeletons would require a calcium to phosphate molar ration of 1.5. Depending on the size of the material being converted, the time needed for the conversion to take place is a crucial factor. Carbonated tricalcium phosphate is generated

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if the conversion time is less than 24 h. Conversely, complete transformation will occur if the conversion process exceeds 48 h. During the second stage, a sol–gel derived calcium phosphate nanocoating is deposited directly to cover the meso- and nanopores within the intra-pore material, while maintaining the large pores suitable for bone growth [1–5, 12–15]. As a consequence of this unique double treatment, mechanical properties such as Young’s modulus, compression and biaxial strengths, and fracture toughness were each enhanced. This was demonstrated by the findings of a mechanical study in which nanocrystalline ceramic coatings were examined using a standard four-point bend test. More importantly, this two-stage approach is expected to result in improvements in bioactivity due to the nanograin size, and hence the large surface area which increases the reactivity of the nanocoating [3–5].

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